Encyclopedia of Biomedical Engineering (vol. 1-3) 0128048298, 9780128048290

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ENCYCLOPEDIA OF BIOMEDICAL ENGINEERING

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ENCYCLOPEDIA OF BIOMEDICAL ENGINEERING EDITOR IN CHIEF

Roger Narayan University of North Carolina at Chapel Hill, Chapel Hill, NC, United States

VOLUME 1

Section Editors Min Wang The University of Hong Kong, Pokfulam, Hong Kong

Cato Laurencin University of Connecticut Health Center, Farmington, CT, United States

Xiaojun Yu Stevens Institute of Technology, Hoboken, NJ, United States

Amsterdam • Boston • Heidelberg • London • New York • Oxford Paris • San Diego • San Francisco • Singapore • Sydney • Tokyo

Elsevier Radarweg 29, PO Box 211, 1000 AE Amsterdam, Netherlands The Boulevard, Langford Lane, Kidlington, Oxford OX5 1GB, United Kingdom 50 Hampshire Street, 5th Floor, Cambridge, MA 02139, United States Copyright Ó 2019 Elsevier Inc. All rights reserved. No part of this publication may be reproduced or transmitted in any form or by any means, electronic or mechanical, including photocopying, recording, or any information storage and retrieval system, without permission in writing from the publisher. Details on how to seek permission, further information about the Publisher’s permissions policies and our arrangements with organizations such as the Copyright Clearance Center and the Copyright Licensing Agency, can be found at our website: www.elsevier.com/permissions. This book and the individual contributions contained in it are protected under copyright by the Publisher (other than as may be noted herein). Notice Knowledge and best practice in this field are constantly changing. As new research and experience broaden our understanding, changes in research methods, professional practices, or medical treatment may become necessary. Practitioners and researchers may always rely on their own experience and knowledge in evaluating and using any information, methods, compounds, or experiments described herein. In using such information or methods they should be mindful of their own safety and the safety of others, including parties for whom they have a professional responsibility. To the fullest extent of the law, neither the Publisher nor the authors, contributors, or editors, assume any liability for any injury and/or damage to persons or property as a matter of products liability, negligence or otherwise, or from any use or operation of any methods, products, instructions, or ideas contained in the material herein. Library of Congress Cataloging-in-Publication Data A catalog record for this book is available from the Library of Congress British Library Cataloguing-in-Publication Data A catalogue record for this book is available from the British Library ISBN 978-0-12-804829-0

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EDITORIAL BOARD

EDITOR IN CHIEF Roger Narayan University of North Carolina at Chapel Hill, Chapel Hill, NC, United States

SECTION EDITORS

Levi Hargrove Rehabilitation Institute of Chicago, Chicago, IL, United States

Christian Hellmich TU Wien, Vienna University of Technology, Vienna, Austria

Sri Krishnan Ryerson University, Toronto, ON, Canada

Cato Laurencin University of Connecticut Health Center, Farmington, CT, United States

Diego Mantovani Laval University, Quebec City, QC, Canada

William Z Rymer Rehabilitation Institute of Chicago, Chicago, IL, United States

Pankaj Vadgama Queen Mary University of London, London, United Kingdom

Min Wang The University of Hong Kong, Pokfulam, Hong Kong

Alexander Wong University of Waterloo, Waterloo, ON, Canada

Xiaojun Yu Stevens Institute of Technology, Hoboken, NJ, United States

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EDITOR IN CHIEF Roger Narayan Dr. Roger Narayan is a professor in the Joint Department of Biomedical Engineering at the University of North Carolina and North Carolina State University. He is an author of over 200 publications as well as several book chapters on processing, characterization, and modeling of biological and biomedical materials. Dr. Narayan has edited several books, including Biomedical Materials, Printed Biomaterials, Computer Aided Biomanufacturing, Diamond-Based Materials for Biomedical Applications, Medical Biosensors for Point of Care (POC) Applications, Monitoring and Evaluation of Biomaterials and their Performance In Vivo, Nanobiomaterials: Nanostructured Materials for Biomedical Applications, and the ASM Handbook on Materials for Medical Devices. He has previously served as chair of the Functional Materials Division of The Minerals, Metals & Materials Society and is currently chair-elect of the Bioceramics Division of American Ceramics Society. Dr. Narayan has received several honors for his research activities, including the North Carolina State University Alcoa Foundation Engineering Research Achievement Award, the North Carolina State University Sigma Xi Faculty Research Award, the University of North Carolina Jefferson-Pilot Fellowship in Academic Medicine, the National Science Faculty Early Career Development Award, the Office of Naval Research Young Investigator Award, the American Ceramic Society Richard M. Fulrath Award, the Royal Academy of Engineering Distinguished Visiting Fellowship, and TMS Brimacombe Medal. He has served as Fulbright Scholar at the University of Otago, the National Polytechnic Institute (Mexico City), and the University of Sao Paulo. He has been elected as Fellow of ASM International, the American Association for the Advancement of Science, the American Ceramic Society, and the American Institute for Medical and Biological Engineering.

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SECTION EDITORS Levi Hargrove Dr. Hargrove is currently the Director of the Center for Bionic Medicine and of the Neural Engineering for Prosthetic and Orthotics Laboratory at the Shirley Ryan AbilityLab. He is also an Associate Professor in the Departments of Physical Medicine and Rehabilitation and the McCormick School of Engineering at Northwestern University. A major goal of his research is to develop clinically realizable myoelectric control systems that can be made available to persons with limb loss in the near future. His research addresses all levels of amputation and has been published in the Journal of the American Medical Association and the New England Journal of Medicine, and multiple patents. Key projects include the development of advanced and adaptive control systems for prosthetic legs, improving control of robotic hand prostheses, and intramuscular EMG signal processing. In 2012, Dr. Hargrove cofounded Coapt, a company to transition advanced rehabilitation technologies from the research laboratory to patients’ homes.

Christian Hellmich Dr. Christian Hellmich, Full Professor at the Department of Civil Engineering of the Vienna University of Technology (TU Wien), is the director of the Institute for Mechanics of Materials and Structures. At TU Wien, he received his engineering, Ph.D., and habilitation degrees (in 1995, 1999, and 2004, respectively). From 2000 to 2002, he was a Max Kade Postdoctoral Fellow in the Department of Civil and Environmental Engineering at the Massachusetts Institute of Technology. His work is strongly focused on well-validated material and (micro)structural models, in terms of theoretical foundations and applications to concrete, soil, rock, wood, bone, and biomedical implants, up the structural level (tunnels, pipelines, bridges, biological organs such as the skeleton)dwith complementary experimental activities if necessary. He has led several projects for the tunnel, railway, and pipeline industries, as well as international research activities sponsored by the European Commission, including the coordination of the mixed industry-academia consortium “BIO-CT-EXPLOIT” at the crossroads of numerical simulation and computer tomography, or the cross-domain COST action NAMABIO integrating engineers, physicists, (stem) cell biologists, and medical doctors across the European continent and beyond. He has published more than 130 papers in international refereed scientific journals in the fields of engineering mechanics, materials science, and theoretical biology, more than 20 book chapters, and more than 120 papers in refereed conference proceedings. Dr. Hellmich has served as the Chairman of both the Properties of Materials Committee of the Engineering Mechanics Division of the American Society of Civil Engineers (ASCE), and the Poromechanics and Biomechanics Committees of the Engineering Mechanics Institute (EMI), as associate editor of the Journal of Engineering Mechanics (ASCE), and as Coeditor in Chief of the Journal of Nanomechanics and Micromechanics (ASCE). As community service, he has (co)chaired and/or supported more than 50 international conferences (including chairmanship of the 2013 Biot Conference on Poromechanics and the 2015 CONCREEP conference; both EMI-ASCE supported), and he has reviewed for 128 different scientific journals and 15 science foundations. He was awarded the Kardinal Innitzer Science Award of the Archbishopry of Vienna in 2004 (for his habilitation thesis), the Science Award of the State of Lower Austria in 2005 (for his achievements in the micromechanics of hierarchical composites), and he was the recipient of the 2008 Zienkiewicz Award for Young Scientists in Computational Engineering Sciences, sponsored by the European Community on Computational Methods in Applied Sciences (ECCOMAS). For further activities in the multiscale poromicromechanics of bone materials, he received one of the highly prestigious ERC Grants of the

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European Research Council in 2010; and he was elected member of the Young Academy of the Austrian Academy of Sciences in 2011. In 2012, he was rewarded the prestigious Walter L. Huber Research Prize of the ASCE, for his contributions to the microporomechanics of hierarchical geomaterials and biomaterials; he was elected Fellow of EMI in 2014 and was corecipient of the 2017 Kajal Mallick Memorial Award of the Institution of Civil Engineers (United Kingdom).

Sri Krishnan Sridhar (Sri) Krishnan received B.E. degree in Electronics and Communication Engineering from the College of Engineering, Guindy, Anna University, Chennai, India, in 1993, and M.Sc. and Ph.D. degrees (with student fellowship from Alberta Heritage Foundation for Medical Research) in Electrical and Computer Engineering from The University of Calgary, Calgary, Alberta, Canada, in 1996 and 1999, respectively. Sri Krishnan joined Ryerson University in July 1999 and is currently a Professor in the Department of Electrical and Computer Engineering. Since July 2011, he is an Associate Dean (Research, Development and External Partnerships) for the Faculty of Engineering and Architectural Science. He is also the Codirector of the Institute for Biomedical Engineering, Science and Technology (iBEST) and an affiliate scientist at the Keenan Research Centre in St. Michael’s Hospital, Toronto. Since January 2002 Sri Krishnan held various administrative leadership positions in the Department of Electrical and Computer Engineering and the Faculty of Engineering and Architectural Science. In 2010–2011, Sri Krishnan held Visiting Appointments in University of Rennes 1 (France), Grenoble Institute of Technology (France) and Indian Institute of Technology (Madras). Sri Krishnan is a registered professional engineer in the Province of Ontario and is a senior member of IEEE (EMBS and SP societies). He was the Founding Chair (2005–2015) of IEEE Signal Processing Society, Toronto Section and Region 7 (Canada), and a Founding Member of the IEEE Engineering in Medicine and Biology Society, Toronto Section. He currently serves as a Technical Committee Member (Biomedical Signal Processing) of IEEE EMBS. Sri Krishnan held the Canada Research Chair position (2007–2017) in Biomedical Signal Analysis. Sri Krishnan has successfully supervised/trained 10 postdoc fellows, 10 Ph.D., 30 Masters (thesis), 9 Masters (project), 42 RAs, and 20 Visiting RAs. Sri Krishnan’s research interests include adaptive signal representations and analysis and their applications in biomedicine, multimedia (audio), and biometrics. He has published 295 papers in refereed journals and conferences, filed 10 invention disclosures, and has one US patent. He has presented keynote/plenary/invited talks in more than 35 international conferences and workshops. Sri Krishnan also serves as a reviewer, committee member, and chair for many international conferences, journals, and granting bodies. Sri Krishnan’s academic interests include (interdisciplinary) curriculum design, experiential learning, and innovation. Sri Krishnan serves in the advisory boards of research institutes, innovation centers, incubator zones, and business organizations. Sri Krishnan is a recipient/awarded Outstanding Canadian Biomedical Engineer Award 2016; Certificate of Appreciation from PEO York Chapter 2016; Fellow of Canadian Academy of Engineering in 2014; 2014 Exemplary Service Award from IEEE Toronto Section; 2014 Certificate of Merit from IEEE Signal Processing Society; 2013 Achievement in Innovation Award from Innovate Calgary; 2011 Sarwan Sahota Distinguished Scholar Award; 2011 Certificate of Appreciation from IEEE Signal Processing Society; 2010 Shastri Visiting Professorship; 2010 French Embassy Visiting Researcher; 2008 Ontario Research Innovation Award from Biodiscovery Toronto; 2007 Canadian Engineers’ Young Engineer Achievement Award from Engineers’ Canada; 2006 New Pioneers Award in Science and Technology; 2006 South Asian Community Achiever Award; 2006 IEEE Toronto Section Best Chapter Chair Award; 2005 IEEE AESS Best Chapter Chair Award; 2005 IEEE Certificate of Appreciation from Six Societies; Six Best Research Paper Awards coauthored with his graduate students in International Conferences; and 2005 FEAS Research Excellence Award.

Cato Laurencin Cato T. Laurencin, M.D., Ph.D. is the University Professor at UCONN. He is the eighth designated in UCONN’s history. He is Professor of Chemical Engineering, Professor of Materials Science and Engineering, and Professor of Biomedical Engineering, and the Van Dusen Distinguished Endowed Professor of Orthopaedic Surgery. He directs the Institute for Regenerative Engineering and the Raymond and Beverly Sackler Center at the University of Connecticut. Dr. Laurencin earned his B.S.E. degree in Chemical Engineering from Princeton University. He earned his Ph.D. in Biochemical Engineering/Biotechnology from the Massachusetts Institute of Technology where he was named a Hugh Hampton Young Fellow. At the same time, he earned his M.D., Magna Cum Laude from the Harvard Medical School where he received the Robinson Award for Surgery. Dr. Laurencin is an expert in biomaterials, nanotechnology, stem cell science, and, the new field he has pioneered, Regenerative Engineering. He is a fellow of American Institute of Chemical

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Engineers and was named one of the 100 Engineers of the Modern Era by the AICHE. He received the Percy Julian Medal from National Organization of Black Chemists and Chemical Engineers, and the Pierre Galletti Award from the American Institute of Medical and Biological Engineering. He has received the NIH Director’s Pioneer Award and the National Science Foundation Emerging Frontiers in Research and Innovation Award for his research in Regenerative Engineering. Dr. Laurencin is an elected member of the National Academy of Engineering, the National Academy of Medicine, the Indian National Academy of Engineering, the Indian National Academy of Sciences, and the African Academy of Sciences. He is an academician and foreign member of the Chinese Academy of Engineering. Dr. Laurencin has two awards named in his honor. The W. Montague Cobb Institute and the National Medical Association established the Cato T. Laurencin Lifetime Research Achievement Award, while the Society for Biomaterials established The Cato T. Laurencin, M.D., Ph.D. Travel Fellowship Award. Dr. Laurencin received the Presidential Faculty Fellow Award from President Bill Clinton and the Presidential Award for Excellence in Science, Mathematics, and Engineering Mentoring from President Barack Obama. He is the recipient of the National Medal of Technology and Innovation, America’s highest award for technological achievement from President Barack Obama in ceremonies at the White House.

Diego Mantovani, Ph.D., FBSE. Prof. Diego Mantovani is the director of Laboratory for Biomaterials and Bioengineering at Laval University, in Canada, and senior scientist of the Regenerative Medicine Division of the Quebec University Hospital Research Centre. He received his doctoral degree jointly from University of Technology of Compiègne, France, and Laval University in 1999 and his joint Diploma in Engineering from Politecnico di Milano and the University of Technology of Compiegne, France, in 1993. After an industrial postdoc (1999), he becomes professor at Laval University School of Science and Engineering in 2000. Since the beginning he established is Laboratory at the University Hospital Research Center in Quebec City. Within his team, researches focus on surface modifications by plasma, thin polymer functional films, cell–materials interactions, degradable metals, scaffolds, and bioreactors for the replacement and regeneration of cardiovascular tissue. He has authored more than 260 original articles, holds 5 patents, and presented more than 185 keynotes, invited and seminar lectures worldwide. His H-index is 43 (June 2018), and his works were cited more than 7000 times. He was President of the Canadian Society for Biomaterials (2008–2009), and Executive Cochair of the World Biomaterials Congress in 2016 in Montreal, Canada. In 2012, he was elected Fellow of the World Biomaterials Science and Engineering Society. Since 2012, he is the holder of the Canada Research Chair 1 in Biomaterials and Bioengineering for the Innovation in Surgery. He was member of ad hoc panels at FDA, ISO, and Health Canada and member of a number of funding, regulatory and scientific committees worldwide. He is Adjunct Professor at Politecnico di Milano and Universita del Piemonte Orientale in Italy, as well as at the Vellore Institute of Technology, in India. He was invited professor in several universities worldwide, including Campinas, Brasil (2012–2015), Bologna (2015), Bordeaux (2014), Siao Tong West, China (2012), Cergy-Pontoise (2012), ParisTech (2011), Buenos Aires (2010), Namur, Belgium (2008), Tor Vergata, Italy (2007), Ankara, Turkey (2006), and others. He is member of the editorial board of five scientific journals in the field and of the advisory board of three medical devices consortia worldwide.

William Z Rymer Professor William Z Rymer is Professor of Physical Medicine and Rehabilitation and Physiology at the Rehabilitation Institute of Chicago, Chicago, IL, United States. His focus of work includes pathophysiology, stroke, spinal cord injury, spinal circuits, biomedical engineering, and neural signal processing.

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Pankaj Vadgama Pankaj Vadgama qualified in Degree in Medicine and then in Chemistry at the University of Newcastle upon Tyne, United Kingdom, with a First Class Honors BSc. He is a chemical pathologist, becoming a Fellow of the Royal College of Pathology. He completed his Ph.D. on medical biosensors as an MRC Fellow at Newcastle, and while there, he was made Director of the Biosensors Group and later appointed as Professor of Clinical Biochemistry at the University of Manchester, subsequently becoming Research Dean for the Faculty of Medicine. He was appointed Director of the Interdisciplinary Research Centre in Biomedical Materials at Queen Mary, University of London and was, until recently, Head of the Department of Clinical Biochemistry, Barts Health NHS Trust. His main interests are variously biosensors, applied bioelectrochemistry, point-of-care testing, and membrane technology. He has published over 200 papers. He is also Fellow of the Royal Society of Medicine, Institute of Physics, Royal Society of Chemistry, the Institute of Materials Minerals and Mining, and the Royal Society of Biology. He was given the Foundation Award of the Association of Clinical Biochemistry and Laboratory Medicine, has been a Sandoz Lecturer of the British Geriatric Society, and delivered the Latner lecture at the University of Newcastle. He has served on various UK Research Council grants award committees and is at present member of the Institute of Materials Minerals and Mining Smart Materials and Nano Committees and the Biomedical Materials Application Division. He sits on various BSI committees and was Chair of the ISO subpanel on nanomedicine nomenclature. He sits on various editorial boards and is Editor in Chief of Bioelectrochemistry. He is Deputy Chair of the Council for the Frontiers of Science based in Uganda directed at research training in East Africa.

Min Wang Min Wang is a Full Professor at The University of Hong Kong (HKU), and as Programme Director (2013–2018), he has led HKU’s Medical Engineering Programme (which is retitled to “Biomedical Engineering Programme” in 2018). He has worked in universities in the United Kingdom (1991–1997), Singapore (1997–2002), and Hong Kong (2002–Present) and has been a Guest Professor or Adjunct Professor of several universities in mainland China (Shanghai Jiao Tong University, Zhejiang University, Tianjin University, Southwest Jiao tong University, etc.). He was awarded BSc (1985) and Ph.D. (1991), both in Materials Science and Engineering, by Shanghai Jiao Tong University and University of London, respectively. He is a chartered engineer (CEng, 1995; UK) and chartered scientist (CSci, 2005; UK). He is an elected fellow of professional societies in the United Kingdom, Hong Kong, United States, and internationally (FIMMM, 2001; FIMechE, 2007; FHKIE, 2010; FBSE, 2011; FAIMBE, 2012; WAC Academician, 2013). Since 1991, he has been conducting research in biomaterials and tissue engineering and developing new biomaterials using the composite/hybridization approach. He was a founding member of UK’s Interdisciplinary Research Centre (IRC) in Biomedical Materials at the University of London. His biomaterials research has covered metals, polymers, ceramics, and composites and includes surface modification of materials or scaffolds. In recent years, he has focused on nanobiomaterials, electrospinning, and 3D printing. He and his research staff/students have won many awards at international conferences. He has authored a large number of research papers as well as many book chapters. His research has been widely cited by other researchers around the world. He has given many conference presentations, including more than 150 invited talks at international conferences. He has also given more than 110 seminars in universities, research institutes, and hospitals in Europe, North America, Asia, and Australia. He has been Chairman/Organizer of many conferences and has served in committees of more than 70 international conferences. He is the Founding Series Editor of Springer Series in Biomaterials Science and Engineering books and has been Editor, Associate Editor, or member of the Editorial Board of 20 international, printed journals, including International Materials Reviews, Composites Science and Technology, Surface and Coatings Technology, Journal of Materials Science: Materials in Medicine, and Journal of the Royal Society Interface. He has acted as a referee for more than 110 international journals in the fields of materials science and engineering, biomaterials and tissue engineering, physics, chemistry, medicine, dentistry, medical devices, biofabrication, nanoscience, nanotechnology, and 3D printing. He has been active in professional society activities and has served in various roles in these societies. He was Chairman of the Biomedical Division of Hong Kong Institution of Engineers (HKIE). He serves/has served in the Nomination Committee of World Academy of Ceramics (WAC) and the ICF-BSE Steering Committee of the International College of Fellows of the International Union of Societies for Biomaterials Science and Engineering (IUS-BSE). He has been an elected Council Member of Chinese Society for Biomaterials, Hong Kong Institution of Engineers, Asian Biomaterials Federation, World Association for Chinese Biomedical Engineers (WACBE), and Administrative Council of International Federation for Medical and Biological Engineering (IFMBE). (http://web.hku.hk/memwang/).

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Alexander Wong Alexander Wong, P.Eng., is currently the Canada Research Chair in Artificial Intelligence and Medical Imaging, Codirector of the Vision and Image Processing Research Group, and an Associate Professor in the Department of Systems Design Engineering at the University of Waterloo. He had previously received the B.A.Sc. degree in Computer Engineering from the University of Waterloo, Waterloo, ON, Canada, in 2005, the M.A.Sc. degree in Electrical and Computer Engineering from the University of Waterloo, Waterloo, ON, Canada, in 2007, and Ph.D. degree in Systems Design Engineering from the University of Waterloo, ON, Canada, in 2010. He was also an NSERC postdoctoral research fellow at Sunnybrook Health Sciences Centre. He has published over 400 refereed journal and conference papers, as well as patents, in various fields such as computational imaging, artificial intelligence, computer vision, and medical imaging, and has received numerous awards such as 13 paper awards at international conference and an Early Researcher Award from the Ministry of Economic Development and Innovation.

Xiaojun Yu Dr. Yu is Associate Professor, Biomedical Engineering at Stevens Institute of Technology, Hoboken, NJ, United States. Dr. Yu’s primary research interests focus on tissue engineering, polymeric biomaterials and drug delivery. His current research activities include nano- and microscale functionalization of biomimic three-dimensional scaffolds for neural and musculoskeletal tissue repair and regeneration, investigation of cell and material interactions in bioreactors, development of controlled release systems for the delivery of growth factors and drugs, and manipulation of microenvironment for stem cell proliferation and differentiation.

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CONTRIBUTORS TO VOLUME 1 Tyler Ackley UConn Health, Farmington, CT, United States

M Bohner RMS Foundation, Bettlach, Switzerland

Rafiq Ahmad Chonbuk National University, Jeonju-si, Jeollabuk-do, Republic of Korea

M E Bronner California Institute of Technology, Pasadena, CA, USA

Song Ih Ahn George W. Woodruff School of Mechanical Engineering, Wallace H. Coulter Department of Biomedical Engineering, Institute for Electronics and Nanotechnology, Parker H. Petit Institute for Bioengineering and Bioscience, Georgia Institute of Technology, Atlanta, GA, United States Khiam Aik Khor Nanyang Technological University, Singapore James M Anderson Case Western Reserve University, Cleveland, OH, United States Lavinia Cosmina Ardelean “Victor Babes” University of Medicine and Pharmacy Timisoara, Timisoara, Romania A Atala Wake Forest Baptist Medical Center, WinstoneSalem, NC, USA S T Avecilla New York Presbyterian Hospital, Weill Cornell Medical College, New York, NY, USA Hani A Awad University of Rochester, Rochester, NY, United States Besim Ben-Nissan University of Technology Sydney, Sydney, NSW, Australia

Ashley C Brown North Carolina State University and University of North Carolina at Chapel Hill, Raleigh, NC, United States; and North Carolina State University, Raleigh, NC, United States Arnaud Bruyas Stanford University, Stanford, CA, United States S A Busch Athersys, Inc., Cleveland, OH, USA Ruth Cameron University of Cambridge, Cambridge, United Kingdom Sophie Cazalbou UMR 5085 UPS-INPT-CNRS, Toulouse, France Paul Z Chen University of Waterloo, Waterloo, ON, Canada Shiyu Cheng Beijing Engineering Research Center for BioNanotechnology and CAS Key Laboratory for Biomedical Effects of Nanomaterials and Nanosafety, CAS Center for Excellence in Nanoscience, National Center for NanoScience and Technology, Beijing, P. R. China Daniel Chester North Carolina State University and University of North Carolina at Chapel Hill, Raleigh, NC, United States; and North Carolina State University, Raleigh, NC, United States

Serena Best University of Cambridge, Cambridge, United Kingdom

Andy H Choi University of Technology Sydney, Sydney, NSW, Australia

Aldo R Boccaccini University of Erlangen-Nuremberg, Erlangen, Germany

Ezharul Hoque Chowdhury Monash University, Clayton, VIC, Australia

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Contributors to Volume 1

Tzahi Cohen-Karni Carnegie Mellon University, Pittsburgh, PA, United States M Csete Huntington Medical Research Institutes, Pasadena, CA, USA M M Cushing New York Presbyterian Hospital, Weill Cornell Medical College, New York, NY, USA Michael A Daniele North Carolina State University and University of North Carolina at Chapel Hill, Raleigh, NC, United States; and North Carolina State University, Raleigh, NC, United States

Margaret A T Freeberg University of Rochester, Rochester, NY, United States Katie Gailiunas University of Connecticut, Storrs, CT, United States Emmanuel Gibon Stanford University, Stanford, CA, United States Stuart B Goodman Stanford University, Stanford, CA, United States Paul Frank Gratzer School of Biomedical Engineering, Dalhousie University, Halifax, NS, Canada Frank X Gu University of Waterloo, Waterloo, ON, Canada

Natalia Davidenko University of Cambridge, Cambridge, United Kingdom

Vincenzo Guarino National Research Council of Italy, Naples, Italy

Caroline N Dealy UConn Health, Farmington, CT, United States

Lin Guo The University of Hong Kong, Pokfulam, Hong Kong

Jinqi Deng Beijing Engineering Research Center for BioNanotechnology and CAS Key Laboratory for Biomedical Effects of Nanomaterials and Nanosafety, CAS Center for Excellence in Nanoscience, National Center for NanoScience and Technology, Beijing, P. R. China; and Sino-Danish College, University of Chinese Academy of Sciences, Beijing, P. R. China

Babak Hassan Beygi The Hong Kong Polytechnic University of Hong Kong, Hung Hom, Hong Kong Y Hayashi Department of Dental Regenerative Medicine, Center of Advanced Medicine for Dental and Oral Diseases, National Center for Geriatrics and Gerontology, Research Institute, Obu, Japan

Kyle G Doherty University of Liverpool, Liverpool, United Kingdom

J S Hayes NUI Galway, Galway, Ireland

Dionysios Douroumis University of Greenwich, Greenwich, United Kingdom

Michelle Hobert University of Connecticut, Storrs, CT, United States

Bin Duan University of Nebraska Medical Center, Omaha, NE, United States

T Hochgreb-Hägele California Institute of Technology, Pasadena, CA, USA

Felipe Eltit The University of British Columbia, Vancouver, BC, Canada Jorge Luis Escobar Ivirico University of Connecticut, Storrs, CT, United States; and University of Connecticut Health Center, Farmington, CT, United States Bing Fang University of Delaware, Newark, DE, United States Melanie Fisher UConn Health, Farmington, CT, United States Kate Fox RMIT University, Melbourne, VIC, Australia

Rong Huang Queensland University of Technology, Brisbane, QLD, Australia K Ishida Tokyo University of Science, Noda, Chiba, Japan Manisha Jassal Stevens Institute of Technology, Hoboken, NJ, United States Wenkai Jia Michigan Technological University, Houghton, MI, United States Sirui Jiang Case Western Reserve University, Cleveland, OH, United States

Contributors to Volume 1

Xingyu Jiang Beijing Engineering Research Center for BioNanotechnology and CAS Key Laboratory for Biomedical Effects of Nanomaterials and Nanosafety, CAS Center for Excellence in Nanoscience, National Center for NanoScience and Technology, Beijing, P. R. China; and Sino-Danish College, University of Chinese Academy of Sciences, Beijing, P. R. China Radoslaw Junka Stevens Institute of Technology, Hoboken, NJ, United States Jacob G Kallenbach University of Rochester, Rochester, NY, United States Victoria R Kearns University of Liverpool, Liverpool, United Kingdom Stephnie M Kennedy University of Liverpool, Liverpool, United Kingdom Yusuf Khan University of Connecticut Health Center, Farmington, CT, United States Gilson Khang Chonbuk National University, Jeonju-si, Jeollabuk-do, Republic of Korea Kristi L Kiick University of Delaware, Newark, DE, United States Sungwoo Kim Stanford University, Stanford, CA, United States YongTae Kim George W. Woodruff School of Mechanical Engineering, Wallace H. Coulter Department of Biomedical Engineering, Institute for Electronics and Nanotechnology, Parker H. Petit Institute for Bioengineering and Bioscience, Georgia Institute of Technology, Atlanta, GA, United States Nicholas J Kohrs University of Kentucky, Lexington, KY, United States D S Koslov Wake Forest Baptist Medical Center, WinstoneSalem, NC, USA Sangamesh G Kumbar University of Connecticut Health, Farmington, CT, United States; and University of Connecticut, Storrs, CT, United States Vincenzo La Carrubba University of Palermo, Palermo, Italy Rebecca Lace University of Liverpool, Liverpool, United Kingdom

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Dimitrios A Lamprou University of Kent, Canterbury, United Kingdom Cato T Laurencin University of Connecticut Health Center, Farmington, CT, United States; and University of Connecticut, Storrs, CT, United States Yunki Lee George W. Woodruff School of Mechanical Engineering, Wallace H. Coulter Department of Biomedical Engineering, Institute for Electronics and Nanotechnology, Parker H. Petit Institute for Bioengineering and Bioscience, Georgia Institute of Technology, Atlanta, GA, United States Andreas Lendlein Institute of Biomaterial Science and Berlin Brandenburg Center for Regenerative Therapies, Teltow, Germany Hannah J Levis University of Liverpool, Liverpool, United Kingdom Jiao Jiao Li University of Sydney, Sydney, NSW, Australia; Kolling Institute, Northern Sydney Local Health District, St Leonards, NSW, Australia; and Sydney Medical School Northern, University of Sydney, St Leonards, NSW, Australia Zhong Li Nanyang Technological University, Singapore Liliana Liverani University of Erlangen-Nuremberg, Erlangen, Germany Thilanga Liyanage University of Kentucky, Lexington, KY, United States Tianzhi Luo University of Delaware, Newark, DE, United States Christopher Mancuso University of Connecticut, Storrs, CT, United States Ohan S Manoukian University of Connecticut Health, Farmington, CT, United States; and University of Connecticut, Storrs, CT, United States Ethan A Marrow North Carolina State University and University of North Carolina at Chapel Hill, Raleigh, NC, United States; and North Carolina State University, Raleigh, NC, United States Z Master Albany Medical College, Albany, NY, USA; and University of Alberta, Edmonton, AB, Canada

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Contributors to Volume 1

Stefani Mazzitelli University of Ferrara, Ferrara, Italy

David A Puleo University of Kentucky, Lexington, KY, United States

R Morita Tokyo University of Science, Noda, Chiba, Japan

Zichen Qian Michigan Technological University, Houghton, MI, United States

Mahboob Morshed Independent University, Dhaka, Bangladesh Amir Najarzadeh University of Kentucky, Lexington, KY, United States M Nakashima Department of Dental Regenerative Medicine, Center of Advanced Medicine for Dental and Oral Diseases, National Center for Geriatrics and Gerontology, Research Institute, Obu, Japan Claudio Nastruzzi University of Ferrara, Ferrara, Italy Inn Chuan Ng National University of Singapore, Singapore D H Nguyen Shiley Eye Center and Institute for Genomic Medicine, University of California at San Diego, La Jolla, CA, USA Mitsuo Niinomi Tohoku University, Sendai, Japan; Osaka University, Osaka, Japan; Meijo University, Nagoya, Japan; and Nagoya University, Nagoya, Japan

Ru Qing Yu The University of Hong Kong, Hong Kong Daniel Radke Michigan Technological University, Houghton, MI, United States P Rajan The Scripps Research Institute, La Jolla, CA, USA Sahil Kumar Rastogi Carnegie Mellon University, Pittsburgh, PA, United States Muhammad Y Razzaq Institute of Biomaterial Science and Berlin Brandenburg Center for Regenerative Therapies, Teltow, Germany Lucien Reclaru VVSA, branch of Richemont International SA Varinor Innovation, Delémont, Switzerland Markus Reinthaler Institute of Biomaterial Science and Berlin Brandenburg Center for Regenerative Therapies, Teltow, Germany

Bridget Oei UConn Health, Farmington, CT, United States

Nicholas P Rhodes University of Liverpool, Liverpool, United Kingdom

Anurag Ojha University of Connecticut, Storrs, CT, United States

Kelsey Richard UConn Health, Farmington, CT, United States

M Oshima Tokyo University of Science, Noda, Chiba, Japan

R G Richards AO Research Institute, Davos, Switzerland

H Ouyang Shiley Eye Center and Institute for Genomic Medicine, University of California at San Diego, La Jolla, CA, USA

Aaqil Rifai RMIT University, Melbourne, VIC, Australia

Chi-Chun Pan Stanford University, Stanford, CA, United States Pornteera Pawijit National University of Singapore, Singapore Aura Penalosa University of Connecticut, Storrs, CT, United States D Pergament Case Western Reserve University School of Law, Cleveland, OH, USA; and Children’s Law Group, LLC, Chicago, IL, USA Elena Pirogova RMIT University, Melbourne, VIC, Australia

Steven A Ross University of Greenwich, Greenwich, United Kingdom M Saito Tokyo University of Science, Noda, Chiba, Japan Naseem Sardashti University of Connecticut Health, Farmington, CT, United States; and University of Connecticut, Storrs, CT, United States Mark Schröder Institute of Biomaterial Science and Berlin Brandenburg Center for Regenerative Therapies, Teltow, Germany Nikolaos Scoutaris University of Greenwich, Greenwich, United Kingdom

Contributors to Volume 1

Swaminathan Sethuraman SASTRA Deemed University, Thanjavur, India

Mian Wang Northeastern University, Boston, MA, United States

Jeong Eun Song Chonbuk National University, Jeonju-si, Jeollabuk-do, Republic of Korea

Min Wang The University of Hong Kong, Pokfulam, Hong Kong

Alexander Martin Stahl Stanford University, Stanford, CA, United States

Qiong Wang The University of British Columbia, Vancouver, BC, Canada

Teagen Stedman University of Connecticut Health, Farmington, CT, United States

Rizhi Wang The University of British Columbia, Vancouver, BC, Canada

Anuradha Subramanian SASTRA Deemed University, Thanjavur, India

Thomas J Webster Northeastern University, Boston, MA, United States; and Wenzhou Medical University, Wenzhou, China

Millicent O Sullivan University of Delaware, Newark, DE, United States Haoran Sun The University of Hong Kong, Pokfulam, Hong Kong Mitch Tahtinen Michigan Technological University, Houghton, MI, United States Jordon Tan Temasek Polytechnic, Singapore A E Ting Athersys, Inc., Cleveland, OH, USA Nirmalya Tripathy University of Washington, Seattle, WA, United States T Tsuji Tokyo University of Science, Noda, Chiba, Japan; and Organ Technologies Inc., Tokyo, Japan N Turovets University of California, Irvine, CA, USA Morgan A Urello University of Delaware, Newark, DE, United States Antonio Valdevit Stevens Institute of Technology, Hoboken, NJ, United States; and SEA Limited, Columbus, OH, United States Nandakumar Venkatesan University of Kentucky, Lexington, KY, United States Chong Wang Dongguan University of Technology, Dongguan, China Guifang Wang Michigan Technological University, Houghton, MI, United States Jing Yi Wang The University of Hong Kong, Hong Kong

xix

Rachel L Williams University of Liverpool, Liverpool, United Kingdom Christian Wischke Institute of Biomaterial Science and Berlin Brandenburg Center for Regenerative Therapies, Teltow, Germany Man Sang Wong The Hong Kong Polytechnic University of Hong Kong, Hung Hom, Hong Kong Yin Xiao Queensland University of Technology, Brisbane, QLD, Australia Yunzhi Yang Stanford University, Stanford, CA, United States Hanry Yu National University of Singapore, Singapore; Agency for Science, Technology and Research (A*STAR), Singapore; BioSyM, Singapore-MIT Alliance for Research and Technology, Singapore; and Nanfang Hospital, Southern Medical University, Guangzhou, China Xiaojun Yu Stevens Institute of Technology, Hoboken, NJ, United States K Zhang Shiley Eye Center and Institute for Genomic Medicine, University of California at San Diego, La Jolla, CA, USA Feng Zhao Michigan Technological University, Houghton, MI, United States Qilong Zhao Shenzhen Institutes of Advanced Technology, Chinese Academy of Sciences, Shenzhen, China

xx

Contributors to Volume 1

Li Wu Zheng Prince Philip Dental Hospital, The University of Hong Kong, Hong Kong

Yu Zheng The Hong Kong Polytechnic University of Hong Kong, Hung Hom, Hong Kong

Wenfu Zheng Beijing Engineering Research Center for BioNanotechnology and CAS Key Laboratory for Biomedical Effects of Nanomaterials and Nanosafety, CAS Center for Excellence in Nanoscience, National Center for NanoScience and Technology, Beijing, P. R. China

Yinghong Zhou Queensland University of Technology, Brisbane, QLD, Australia Hala Zreiqat University of Sydney, Sydney, NSW, Australia

CONTENTS OF VOLUME 1 Editorial Board

v

Editor in Chief

vii

Section Editors

ix

Contents of All Volumes Preface

xxv xxxv

Biomaterials: Science and Engineering Alternative Processing Techniques for CoCr Dental Alloys Lucien Reclaru and Lavinia Cosmina Ardelean

1

Bioceramics Besim Ben-Nissan, Sophie Cazalbou, and Andy H Choi

16

Biomedical Composites Min Wang and Qilong Zhao

34

Bulk Properties of Biomaterials and Testing Techniques Min Wang and Chong Wang

53

Corrosion of Orthopedic Implants Qiong Wang, Felipe Eltit, and Rizhi Wang

65

Decellularized Extracellular Matrix Paul Frank Gratzer

86

Diamond, Carbon Nanotubes and Graphene for Biomedical Applications Aaqil Rifai, Elena Pirogova, and Kate Fox

97

Gold Nanoparticles for Colorimetric Detection of Pathogens Paul Z Chen and Frank X Gu

108

Manufacture of Biomaterials Min Wang, Lin Guo, and Haoran Sun

116

Materials and Their Biomedical Applications Min Wang and Bin Duan

135

Nano-Biomaterials and their Applications Mian Wang and Thomas J Webster

153

xxi

xxii

Contents of Volume 1

Natural Biopolymers for Biomedical Applications Natalia Davidenko, Ruth Cameron, and Serena Best

162

Polymeric Coatings and Their Fabrication for Medical Devices Dimitrios A Lamprou, Nikolaos Scoutaris, Steven A Ross, and Dionysios Douroumis

177

Porous Biomaterials and Scaffolds for Tissue Engineering Liliana Liverani, Vincenzo Guarino, Vincenzo La Carrubba, and Aldo R Boccaccini

188

Preparation and Properties of Coatings and Thin Films on Metal Implants Zhong Li and Khiam Aik Khor

203

Titanium Alloys Mitsuo Niinomi

213

Biomaterials: In Vitro and in Vivo Studies of Biomaterials Anatomy and Physiology for Biomaterials Research and Development Inn Chuan Ng, Pornteera Pawijit, Jordon Tan, and Hanry Yu

225

Animal Models in Biomaterial Development James M Anderson and Sirui Jiang

237

Blood–Biomaterial Interactions Nicholas P Rhodes

242

Interaction Between Mesenchymal Stem Cells and Immune Cells in Tissue Engineering Rong Huang, Yinghong Zhou, and Yin Xiao

249

Osseointegration of Permanent and Temporary Orthopedic Implants J S Hayes and R G Richards

257

Tissue Response to Biomaterials Jiao Jiao Li and Hala Zreiqat

270

Biomaterials: Biomaterial Applications and Advanced Medical Technologies Biomaterials in Dentistry Li Wu Zheng, Jing Yi Wang, and Ru Qing Yu

278

Biomaterials in Ophthalmology 289 Rachel L Williams, Hannah J Levis, Rebecca Lace, Kyle G Doherty, Stephnie M Kennedy, and Victoria R Kearns Biomaterials in Orthopaedics Emmanuel Gibon and Stuart B Goodman

301

Cell Encapsulation and Delivery Stefani Mazzitelli and Claudio Nastruzzi

308

Drug Delivery Systems and Controlled Release Nicholas J Kohrs, Thilanga Liyanage, Nandakumar Venkatesan, Amir Najarzadeh, and David A Puleo

316

Electrospinning and Electrospray for Biomedical Applications Min Wang and Qilong Zhao

330

Gene Delivery and Clinical Applications Mahboob Morshed and Ezharul Hoque Chowdhury

345

Materials for Exoskeletal Orthotic and Prosthetic Systems Man Sang Wong, Babak Hassan Beygi, and Yu Zheng

352

Contents of Volume 1

xxiii

Microfluidics for Biomedical Applications Shiyu Cheng, Jinqi Deng, Wenfu Zheng, and Xingyu Jiang

368

Organs-on-Chips Yunki Lee, Song Ih Ahn, and YongTae Kim

384

Shape-Memory Polymer Medical Devices Muhammad Y Razzaq, Markus Reinthaler, Mark Schröder, Christian Wischke, and Andreas Lendlein

394

Regenerative Engineering Adult Bone Marrow-Derived Stem Cells: Immunomodulation in the Context of Disease and Injury A E Ting and S A Busch

406

Assessment of Cellular Responses of Tissue Constructs in vitro in Regenerative Engineering Margaret A T Freeberg, Jacob G Kallenbach, and Hani A Awad

414

Assessment of Tissue Constructs In Vivo in Regenerative Engineering Anuradha Subramanian and Swaminathan Sethuraman

427

Bioengineered Kidney and Bladder D S Koslov and A Atala

432

Bioengineering Scaffolds for Regenerative Engineering Zichen Qian, Daniel Radke, Wenkai Jia, Mitch Tahtinen, Guifang Wang, and Feng Zhao

444

Biomaterials for Tissue Engineering and Regenerative Medicine Ohan S Manoukian, Naseem Sardashti, Teagen Stedman, Katie Gailiunas, Anurag Ojha, Aura Penalosa, Christopher Mancuso, Michelle Hobert, and Sangamesh G Kumbar

462

Biomimetic Approaches for Regenerative Engineering Nirmalya Tripathy, Rafiq Ahmad, Jeong Eun Song, and Gilson Khang

483

Bioreactors: System Design and Application for Regenerative Engineering Antonio Valdevit

496

Bone Substitute Materials M Bohner

513

Case Studies for Soft Tissue Regenerative Engineering Jorge Luis Escobar Ivirico and Cato T Laurencin

530

Characterizing the Properties of Tissue Constructs for Regenerative Engineering Yusuf Khan

537

Clinical and Laboratory Aspects of Hematopoietic Stem Cell Transplantation S T Avecilla and M M Cushing

546

Dental Stem Cells M Nakashima and Y Hayashi

554

Drug and Gene Delivery for Regenerative Engineering Morgan A Urello, Tianzhi Luo, Bing Fang, Kristi L Kiick, and Millicent O Sullivan

565

Ethics of Issues and Stem Cell Research: the Unresolved Issues Z Master

584

Eye Diseases and Stem Cells H Ouyang, D H Nguyen, and K Zhang

598

xxiv

Contents of Volume 1

Human Parthenogenetic Pluripotent Stem Cells N Turovets and M Csete

608

Human Pluripotent Stem Cells P Rajan

618

Introduction to Regenerative Engineering Manisha Jassal, Radoslaw Junka, Cato T Laurencin, and Xiaojun Yu

624

Nanoelectronics for Neuroscience Sahil Kumar Rastogi and Tzahi Cohen-Karni

631

Neural Crest Stem Cells T Hochgreb-Hägele and M E Bronner

650

Osteoarthritis at the Cellular Level: Mechanisms, Clinical Perspectives, and Insights From Development 660 Melanie Fisher, Tyler Ackley, Kelsey Richard, Bridget Oei, and Caroline N Dealy Reproductive Technologies, Assisted D Pergament

677

Tooth Regenerative Therapy: Tooth Tissue Repair and Whole Tooth Replacement M Oshima, K Ishida, R Morita, M Saito, and T Tsuji

686

Vascularized Tissue Regenerative Engineering Using 3D Bioprinting Technology Sungwoo Kim, Arnaud Bruyas, Chi-Chun Pan, Alexander Martin Stahl, and Yunzhi Yang

696

Wound Healing and the Host Response in Regenerative Engineering Daniel Chester, Ethan A Marrow, Michael A Daniele, and Ashley C Brown

707

CONTENTS OF ALL VOLUMES VOLUME 1 Biomaterials: Science and Engineering Alternative Processing Techniques for CoCr Dental Alloys Lucien Reclaru and Lavinia Cosmina Ardelean

1

Bioceramics Besim Ben-Nissan, Sophie Cazalbou, and Andy H Choi

16

Biomedical Composites Min Wang and Qilong Zhao

34

Bulk Properties of Biomaterials and Testing Techniques Min Wang and Chong Wang

53

Corrosion of Orthopedic Implants Qiong Wang, Felipe Eltit, and Rizhi Wang

65

Decellularized Extracellular Matrix Paul Frank Gratzer

86

Diamond, Carbon Nanotubes and Graphene for Biomedical Applications Aaqil Rifai, Elena Pirogova, and Kate Fox

97

Gold Nanoparticles for Colorimetric Detection of Pathogens Paul Z Chen and Frank X Gu

108

Manufacture of Biomaterials Min Wang, Lin Guo, and Haoran Sun

116

Materials and Their Biomedical Applications Min Wang and Bin Duan

135

Nano-Biomaterials and their Applications Mian Wang and Thomas J Webster

153

Natural Biopolymers for Biomedical Applications Natalia Davidenko, Ruth Cameron, and Serena Best

162

Polymeric Coatings and Their Fabrication for Medical Devices Dimitrios A Lamprou, Nikolaos Scoutaris, Steven A Ross, and Dionysios Douroumis

177

xxv

xxvi

Contents of All Volumes

Porous Biomaterials and Scaffolds for Tissue Engineering Liliana Liverani, Vincenzo Guarino, Vincenzo La Carrubba, and Aldo R Boccaccini

188

Preparation and Properties of Coatings and Thin Films on Metal Implants Zhong Li and Khiam Aik Khor

203

Titanium Alloys Mitsuo Niinomi

213

Biomaterials: In Vitro and in Vivo Studies of Biomaterials Anatomy and Physiology for Biomaterials Research and Development Inn Chuan Ng, Pornteera Pawijit, Jordon Tan, and Hanry Yu

225

Animal Models in Biomaterial Development James M Anderson and Sirui Jiang

237

Blood–Biomaterial Interactions Nicholas P Rhodes

242

Interaction Between Mesenchymal Stem Cells and Immune Cells in Tissue Engineering Rong Huang, Yinghong Zhou, and Yin Xiao

249

Osseointegration of Permanent and Temporary Orthopedic Implants J S Hayes and R G Richards

257

Tissue Response to Biomaterials Jiao Jiao Li and Hala Zreiqat

270

Biomaterials: Biomaterial Applications and Advanced Medical Technologies Biomaterials in Dentistry Li Wu Zheng, Jing Yi Wang, and Ru Qing Yu

278

Biomaterials in Ophthalmology 289 Rachel L Williams, Hannah J Levis, Rebecca Lace, Kyle G Doherty, Stephnie M Kennedy, and Victoria R Kearns Biomaterials in Orthopaedics Emmanuel Gibon and Stuart B Goodman

301

Cell Encapsulation and Delivery Stefani Mazzitelli and Claudio Nastruzzi

308

Drug Delivery Systems and Controlled Release Nicholas J Kohrs, Thilanga Liyanage, Nandakumar Venkatesan, Amir Najarzadeh, and David A Puleo

316

Electrospinning and Electrospray for Biomedical Applications Min Wang and Qilong Zhao

330

Gene Delivery and Clinical Applications Mahboob Morshed and Ezharul Hoque Chowdhury

345

Materials for Exoskeletal Orthotic and Prosthetic Systems Man Sang Wong, Babak Hassan Beygi, and Yu Zheng

352

Microfluidics for Biomedical Applications Shiyu Cheng, Jinqi Deng, Wenfu Zheng, and Xingyu Jiang

368

Organs-on-Chips Yunki Lee, Song Ih Ahn, and YongTae Kim

384

Contents of All Volumes

Shape-Memory Polymer Medical Devices Muhammad Y Razzaq, Markus Reinthaler, Mark Schröder, Christian Wischke, and Andreas Lendlein

xxvii

394

Regenerative Engineering Adult Bone Marrow-Derived Stem Cells: Immunomodulation in the Context of Disease and Injury A E Ting and S A Busch

406

Assessment of Cellular Responses of Tissue Constructs in vitro in Regenerative Engineering Margaret A T Freeberg, Jacob G Kallenbach, and Hani A Awad

414

Assessment of Tissue Constructs In Vivo in Regenerative Engineering Anuradha Subramanian and Swaminathan Sethuraman

427

Bioengineered Kidney and Bladder D S Koslov and A Atala

432

Bioengineering Scaffolds for Regenerative Engineering Zichen Qian, Daniel Radke, Wenkai Jia, Mitch Tahtinen, Guifang Wang, and Feng Zhao

444

Biomaterials for Tissue Engineering and Regenerative Medicine Ohan S Manoukian, Naseem Sardashti, Teagen Stedman, Katie Gailiunas, Anurag Ojha, Aura Penalosa, Christopher Mancuso, Michelle Hobert, and Sangamesh G Kumbar

462

Biomimetic Approaches for Regenerative Engineering Nirmalya Tripathy, Rafiq Ahmad, Jeong Eun Song, and Gilson Khang

483

Bioreactors: System Design and Application for Regenerative Engineering Antonio Valdevit

496

Bone Substitute Materials M Bohner

513

Case Studies for Soft Tissue Regenerative Engineering Jorge Luis Escobar Ivirico and Cato T Laurencin

530

Characterizing the Properties of Tissue Constructs for Regenerative Engineering Yusuf Khan

537

Clinical and Laboratory Aspects of Hematopoietic Stem Cell Transplantation S T Avecilla and M M Cushing

546

Dental Stem Cells M Nakashima and Y Hayashi

554

Drug and Gene Delivery for Regenerative Engineering Morgan A Urello, Tianzhi Luo, Bing Fang, Kristi L Kiick, and Millicent O Sullivan

565

Ethics of Issues and Stem Cell Research: the Unresolved Issues Z Master

584

Eye Diseases and Stem Cells H Ouyang, D H Nguyen, and K Zhang

598

Human Parthenogenetic Pluripotent Stem Cells N Turovets and M Csete

608

Human Pluripotent Stem Cells P Rajan

618

xxviii

Contents of All Volumes

Introduction to Regenerative Engineering Manisha Jassal, Radoslaw Junka, Cato T Laurencin, and Xiaojun Yu

624

Nanoelectronics for Neuroscience Sahil Kumar Rastogi and Tzahi Cohen-Karni

631

Neural Crest Stem Cells T Hochgreb-Hägele and M E Bronner

650

Osteoarthritis at the Cellular Level: Mechanisms, Clinical Perspectives, and Insights From Development 660 Melanie Fisher, Tyler Ackley, Kelsey Richard, Bridget Oei, and Caroline N Dealy Reproductive Technologies, Assisted D Pergament

677

Tooth Regenerative Therapy: Tooth Tissue Repair and Whole Tooth Replacement M Oshima, K Ishida, R Morita, M Saito, and T Tsuji

686

Vascularized Tissue Regenerative Engineering Using 3D Bioprinting Technology Sungwoo Kim, Arnaud Bruyas, Chi-Chun Pan, Alexander Martin Stahl, and Yunzhi Yang

696

Wound Healing and the Host Response in Regenerative Engineering Daniel Chester, Ethan A Marrow, Michael A Daniele, and Ashley C Brown

707

VOLUME 2 Biomechanics Biomechanics of Cells as Potential Biomarkers for Diseases: A New Tool in Mechanobiology Dinesh R Katti, Kalpana S Katti, Shahjahan Molla, and Sumanta Kar

1

Bone Micro- and Nanomechanics Caitlyn J Collins, Orestis G Andriotis, Vedran Nedelkovski, Martin Frank, Orestis L Katsamenis, and Philipp J Thurner

22

Cell Adhesion: Basic Principles and Computational Modeling Diego A Vargas and Hans Van Oosterwyck

45

Centrifugation and Hypergravity in the Bone Carmelo Mastrandrea and Laurence Vico

59

Computational Modeling of Respiratory Biomechanics Christian J Roth, Lena Yoshihara, and Wolfgang A Wall

70

Constitutive Modeling of Soft Tissues Michele Marino

81

Continuum Mechanics and Its Practical Applications at the Level of Scaling Laws Ko Okumura

111

CT-Based Bone and Muscle Assessment in Normal and Pathological Conditions Paolo Gargiulo, Magnus K Gislason, Kyle J Edmunds, Jonathan Pitocchi, Ugo Carraro, Luca Esposito, Massimiliano Fraldi, Paolo Bifulco, Mario Cesarelli, and Halldór Jónsson

119

Knowledge Extraction From Medical Imaging for Advanced Patient-Specific Musculoskeletal Models Marie-Christine Ho Ba Tho and Tien Tuan Dao

135

Contents of All Volumes

xxix

Mathematical Quantification of the Impact of Microstructure on the Various Effective Properties of Bones Miao-Jung Y Ou, Annalisa De Paolis, and Luis Cardoso

143

Multiphase Porous Media Models for Mechanics in Medicine: Applications to Transport Oncophysics and Diabetic Foot Pietro Mascheroni, Raffaella Santagiuliana, and Bernhard Schrefler

155

Multiscale Bone Mechanobiology Stefan Scheiner, Maria-Ioana Pastrama, Peter Pivonka, and Christian Hellmich

167

Multiscale Mechanical Behavior of Large Arteries Claire Morin, Witold Krasny, and Stéphane Avril

180

Nanoindentation-Based Characterization of Hard and Soft Tissues Pasquale Vena and Dario Gastaldi

203

Nanomechanical Raman Spectroscopy in Biological Materials Yang Zhang, Ming Gan, and Vikas Tomar

215

On the Use of Population-Based Statistical Models in Biomechanics Justin Fernandez, Shasha Yeung, Alex Swee, Marco Schneider, Thor Besier, and Ju Zhang

229

Poroelasticity of Living Tissues Andrea Malandrino and Emad Moeendarbary

238

Structural and Material Changes of Human Cortical Bone With Age: Lessons from the Melbourne Femur Research Collection Romane Blanchard, C David L Thomas, Rita Hardiman, John G Clement, David C Cooper, and Peter Pivonka Vascular Tissue Biomechanics: Constitutive Modeling of the Arterial Wall Thomas Christian Gasser

246

265

Medical Devices 3D Printing in the Biomedical Field Alexander K Nguyen, Roger J Narayan, and Ashkan Shafiee Biocompatibility Evaluation of Orthopedic Biomaterials and Medical Devices: A Review of Safety and Efficacy Models Michel Assad and Nicolette Jackson

275

281

Biological Grafts: Surgical Use and Vascular Tissue Engineering Options for Peripheral Vascular Implants 310 Francesca Boccafoschi, Martina Ramella, Luca Fusaro, Marta C Catoira, and Francesco Casella Current Advancements and Challenges in Stent-Mediated Gene Therapy Shounak Ghosh, Katari Venkatesh, and Dwaipayan Sen

322

Dentistry: Restorative and Regenerative Approaches Geetha Manivasagam, Aakash Reddy, Dwaipayan Sen, Sunita Nayak, Mathew T Mathew, and Asokami Rajamanikam

332

Ephemeral Biogels: Potential Applications as Active Dressings and Drug Delivery Devices Larreta-Garde Véronique, Picard Julien, and Giraudier Sébastien

348

Immunological Responses in Orthopedics and Transplantation Caroline D Hoemann and Martin Guimond

359

xxx

Contents of All Volumes

Iron-Based Degradable Implants Sergio Loffredo, Carlo Paternoster, and Diego Mantovani

374

Medical Devices: Coronary Stents Vanessa Montaño-Machado, Malgorzata Sikora-Jasinska, Carolina Catanio Bortolan, Pascale Chevallier, and Diego Mantovani

386

Medical Devices in Otorhinolaryngology Paolo Aluffi Valletti, Massimiliano Garzaro, and Valeria Dell’Era

399

Medical Devices in Neurology Abbas Z Kouzani and Roy V Sillitoe

409

Obstetrics and Gynecology: Hysteroscopy Antonio Santos-Paulo

414

Orthopedic Implants Weihong Jin and Paul K Chu

425

Pharmacology: Drug Delivery Frédéric Chaubet, Violeta Rodriguez-Ruiz, Michel Boissière, and Diego Velasquez

440

Prosthetic Aortic Valves Anne-Sophie Zenses, Philippe Pibarot, Marie-Annick Clavel, Ezequiel Guzzetti, Nancy Cote, and Erwan Salaun

454

Urology and Nephrology: Regenerative Medicine Applications Ingrid Saba, Stéphane Chabaud, Sophie Ramsay, Hazem Orabi, and Stéphane Bolduc

467

Zinc-Based Degradable Implants Ehsan Mostaed, Malgorzata Sikora-Jasinska, and Maurizio Vedani

478

Medical Imaging Biomechanics Imaging and Analysis Reza Sharif Razavian, Sara Greenberg, and John McPhee

488

Breast Imaging: Mammography, Digital Tomosynthesis, Dynamic Contrast Enhancement Mehran Ebrahimi

501

Diffusion Magnetic Resonance Imaging Jennifer Shane Williamson Campbell and Gilbert Bruce Pike

505

Digital Holographic Microscopy Farnoud Kazemzadeh and Alexander Wong

519

Digital Pathology Matthew G Hanna and Liron Pantanowitz

524

Functional Magnetic Resonance Imaging Jean Chen and Julien Cohen-Adad

533

Hemodynamic Imaging Robert Amelard and Alexander Wong

545

Imaging Informatics David A Koff and Thomas E Doyle

551

Macroscopic Pigmented Skin Lesion Prescreening Eliezer Bernart, Eliezer Soares Flores, and Jacob Scharcanski

561

Contents of All Volumes

xxxi

Magnetic Resonance Imaging Rachel W Chan, Justin Y C Lau, Wilfred W Lam, and Angus Z Lau

574

Perceptual Quality Assessment of Medical Images Hantao Liu and Zhou Wang

588

Radiomics Farzad Khalvati, Yucheng Zhang, Alexander Wong, and Masoom A Haider

597

Ultrasound Elastography Hyock Ju Kwon and Bonghun Shin

604

Rehabilitation Engineering and Integrative Technologies Functional Electric Stimulation Therapy Dejan B Popovic

614

ProsthesesdAssistive TechnologydSports Bryce T J Dyer

621

ProsthesesdAssistive TechnologydUpper Jonathon W Sensinger, Wendy Hill, and Michelle Sybring

632

Robotics: Exoskeletons Daniel P Ferris, Bryan R Schlink, and Aaron J Young

645

RoboticsdSoft Robotics Gursel Alici

652

VOLUME 3 Mathematical Techniques in Biomedical Engineering Cardiac Modeling Edward Vigmond and Gernot Plank

1

Mathematical Approaches for Medical Image Registration Barbara Zitova

21

Mathematical Modeling of Gene Networks Lakshmi Sugavaneswaran

33

Mathematical Modeling Tools and Software for BME Applications Fred J Vermolen and Amit Gefen

56

Mathematical Techniques for Biomedical Image Segmentation Roberto Rodríguez and Juan H Sossa

64

Mathematical Techniques for Circulatory Systems Jason Carson, Raoul Van Loon, and Perumal Nithiarasu

79

Mathematical Techniques for Noninvasive Muscle Signal Analysis and Interpretation Roberto Merletti, Ales Holobar, and Dario Farina

95

Optimization Techniques in BME Jeevan Kumar Pant

112

xxxii

Contents of All Volumes

Single-Cell-Based In Silico Models: A Tool for Dissecting Tumor Heterogeneity Aleksandra Karolak, Saharsh Agrawal, Samantha Lee, and Katarzyna A Rejniak

130

Spectrotemporal Modeling of Biomedical Signals: Theoretical Foundation and Applications Raymundo Cassani and Tiago H Falk

144

Statistical Modeling in Biomedical Engineering Yunfeng Wu and Pinnan Chen

164

Time–Frequency Distributions in Biomedical Signal Processing Yashodhan Athavale and Sridhar Krishnan

177

Wavelets in Biomedical Signal Processing and Analysis Babak Azmoudeh and Dean Cvetkovic

193

Bioinstrumentation and Bioinformatics A Systematic Workflow for Design and Computational Analysis of Protein Microarrays Jonatan Fernández-García, Rodrigo García-Valiente, Javier Carabias-Sánchez, Alicia Landeira-Viñuela, Rafael Góngora, María Gonzalez-Gonzalez, and Manuel Fuentes

213

Ambulatory EEG Monitoring Bernard Grundlehner and Vojkan Mihajlovic

223

Automated EEG Analysis for Neonatal Intensive Care Nathan Stevenson and Anton Tokariev

240

Big Data Calls for Machine Learning Andreas Holzinger

258

Bioimpedance Spectroscopy Processing and Applications Herschel Caytak, Alistair Boyle, Andy Adler, and Miodrag Bolic

265

Bioinformatics in Design of Antiviral Vaccines Ashesh Nandy and Subhash C Basak

280

Bioinformatics in Disease Classification Miguel Ángel Medina

291

Biopotential Monitoring Julián Oreggioni, Angel A Caputi, and Fernando Silveira

296

Blood Gas Analysis and Instrumentation Rebecca Symons, Robindro Chatterji, Kirsty Whenan, Rita Horvath, and Paul S Thomas

305

Computational Approaches in microRNA Biology Ulf Schmitz, Shailendra K Gupta, Julio Vera, and Olaf Wolkenhauer

317

Detection and Classification of Breast Lesions Using Ultrasound-Based Imaging Modalities Md Kamrul Hasan and Sharmin R Ara

331

DNA Microarrays: Fundamentals, Data Integration and Applications Eduardo Valente and Miguel Rocha

349

ECG Monitoring: Present Status and Future Trend Saurabh Pal

363

Genetic Algorithms for Breast Cancer Diagnostics Florin Gorunescu and Smaranda Belciug

380

Contents of All Volumes

xxxiii

Machine Learning in Biomedical Informatics Carlos Fernandez-Lozano, Adrián Carballal, Cristian R Munteanu, Marcos Gestal, Víctor Maojo, and Alejandro Pazos

389

Matrix Assisted Laser Desorption/Ionization as a New Cancer Diagnostic Tool Bozena Hosnedlova, Marta Kepinska, Branislav Ruttkay-Nedecky, Carlos Fernandez, Tomas Parak, Halina Milnerowicz, Jiri Sochor, Geir Bjørklund, and Rene Kizek

400

Medical Utility of NIR Monitoring Zuzana Kovacsova, Gemma Bale, and Ilias Tachtsidis

415

Metabolomics in Biomaterial Research Ana M Gil, Maria H Fernandes, and Iola F Duarte

432

Nucleic-Acid Sequencing G Dorado, S Gálvez, H Budak, T Unver, and P Hernández

443

Optical Techniques for Blood and Tissue Oxygenation Panayiotis Kyriacou, Karthik Budidha, and Tomas Y Abay

461

Polymerase Chain Reaction (PCR) G Dorado, G Besnard, T Unver, and P Hernández

473

Single-Photon Emission Computed Tomography: Principles and Applications Yong Du and Habib Zaidi

493

Translational Bioinformatics: Informatics, Medicine, and -Omics Sergio Paraiso-Medina, David Perez-Rey, Raul Alonso-Calvo, Cristian R Munteanu, Alejandro Pazos, Casimir A Kulikowski, and Victor Maojo

507

Ultrasonic Imaging in Biomedical Applications Roman Gr Maev and Fedar M Severin

515

Index

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PREFACE The use by man of available technology to treat damaged or diseased tissue is older than the written historical record. For example, the Mayan people created artificial teeth out of nacre, which were shown to be fused to the bone (Bobbio, 1972; Westbroek & Marin, 1998). Giovanni Borelli’s studies of the cardiovascular system (e.g., the elasticity of arteries), which were published in De Motu Animalium (On Animal Motion), can be considered as one of the foundations of the field of biomechanics (Parker, 2009). The hypothesis of an intrinsic ’animal electricity’ by Luigi Galvani in the 18th century led to the development of the field of electrophysiology (Piccolino, 1997). In the 19th century, the development of the antiseptic approach to surgical procedures by Joseph Lister made implantation of medical devices without certain postoperative infection possible; for example, Lister described the use of silver wire for treatment of a fractured patella (Worboys, 2013). The discovery of Xrays by Roentgen at the end of the 19th century was rapidly translated for medical imaging (Rowland, 1896; Schuster, 1896). In our lifetime, the work by W. T. Green on generating new cartilage by seeding of chondrocytes as well as by John Burke and Ioannis Yannas on generating skin substitutes is recognized as the birth of the field of regenerative engineering (Vacanti, 2006). At the beginning of the 21st century, the American Institute for Medical and Biological Engineering identified several research areas for the field of biomedical engineering. These include: (a) functional genomics and proteomics, (b) nanotechnology, (c) targeted drug delivery, (d) tissue engineering, and (e) the development of new types of medical instrumentation (Hendee, Chien, Maynard, & Dean, 2002). As some of these research areas have matured, others have become more prominent. Over the past few years, the use of 3D printing and bioprinting technologies to create medical devices has become more prominent. One benefit of utilizing 3D printing and bioprinting for patient care is that medical imaging data (e.g., magnetic resonance imaging and computed tomography data) may be employed to fashion prostheses or artificial tissues with submicroscale features that conform with the requirements of the patient (Narayan, Doraiswamy, Chrisey, & Chichkov, 2010; Skoog & Narayan, 2013). Another technology that will likely transform the field of biomedical engineering over the coming decades involves the use of clustered regularly interspaced short palindromic repeats (CRISPR)/Cas9 for engineering of the human genome. The interface between biomedical engineering and the new field of genome engineering has already spawned research into new technologies for delivery of genome editing tools into the body; the synergy between these fields will only grow over time (Wright, Nuñez, & Doudna, 2016). The goal of the Encyclopedia of Biomedical Engineering is to consider the principles and technologies that underlie the field of biomedical engineering. The encyclopedia is divided into three volumes. The first volume contains a section on biomaterials, which was edited by Min Wang at the University of Hong Kong, and a section on regenerative engineering, which was edited by Cato Laurencin at the University of Connecticut and Xiaojun Yu at the Stevens Institute of Technology. The second volume contains a section on rehabilitation engineering and integrative technologies, which was edited by William Rymer and Levi Hargrove at Northwestern University, a section on biomechanics, which was edited by Christian Hellmich at the Vienna University of Technology, a section on medical devices, which was edited by Diego Mantovani at the University of Laval, and a section on medical imaging, which was edited by Alexander Wong at the University of Waterloo. The third volume contains a section on mathematical techniques in biomedical engineering, which was edited by Sri Krishnan at Ryerson University, and a section on bioinstrumentation and bioinformatics, which was edited by Pankaj Vadgama at Queen Mary University of London. I would like express my sincere appreciation to the section editors and authors for all of their efforts on the encyclopedia. I would also like thank Beckie Brand, Susan Dennis, Becky Gelson, Ginny Mills, Blerina Osmanaj,

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Laura Escalante Santos, and Will Smaldon at Elsevier for their outstanding efforts to bring the encyclopedia to publication. I hope that this work serves the biomedical engineering community by providing a resource that considers topics at the interface of the biological sciences and engineering. Roger J Narayan, M.D., Ph.D. UNC/NCSU Joint Department of Biomedical Engineering. References Bobbio, A. (1972). The first endosseous alloplastic implant in the history of man. Bull. Hist. Dent, 20, 1–6. Hendee, W. R., Chien, S., Maynard, C. D., & Dean, D. J. (2002). The National Institute of biomedical imaging and Bioengineering: history, status, and potential impact. Annals of Biomedical Engineering, 30, 2–10. Narayan, R. J., Doraiswamy, A., Chrisey, D. B., & Chichkov, B. N. (2010). Medical prototyping using two photon polymerization. Materials Today, 13, 44–50. Parker, K. H. (February 2009). A brief history of arterial wave mechanics. Medical & Biological Engineering & Computing, 47(2), 111–118. Piccolino, M. (October 1997). Luigi Galvani and animal electricity: two centuries after the foundation of electrophysiology. Trends in Neurosciences, 20(10), 443–448. Rowland, S. (March 7, 1896). Report on the Application of the new Photography to medicine and surgery. Br Med J, 1(1836), 620–622. Schuster, A. (January 18, 1896). On the new Kind of Radiation. Br Med J, 1(1829), 172–173. Skoog, S. A., & Narayan, R. J. (2013). Stereolithography in medical device fabrication. Advanced Materials & Processes, 171, 32–36. Vacanti, C. A. (July 2006). The history of tissue engineering. J Cell Mol Med, 10(3), 569–576. Westbroek, P., & Marin, F. (1998). A marriage of bone and nacre. Nature, 392, 861–862. Worboys, M. (September 20, 2013). Joseph Lister and the performance of antiseptic surgery. Notes and Records of the Royal Society of London, 67(3), 199–209. Wright, A. V., Nuñez, J. K., & Doudna, J. A. (January 14, 2016). Biology and Applications of CRISPR systems: Harnessing Nature’s Toolbox for genome engineering. Cell, 164(1–2), 29–44.

PERMISSION ACKNOWLEDGMENTS The following material is reproduced with kind permission of Taylor & Francis Figure 1 On the Use of Population-Based Statistical Models in Biomechanics Figure 2 On the Use of Population-Based Statistical Models in Biomechanics Figure 3 On the Use of Population-Based Statistical Models in Biomechanics www.taylorandfrancisgroup.com The following material is reproduced with kind permission of American Association for the Advancement of Science Figure 8a Nanotechnology for Regenerative Engineering www.aaas.org The following material is reproduced with kind permission of Oxford University Press Figure 1 Translational Bioinformatics for Personalised Medicine www.oup.com The following material is reproduced with kind permission of Nature Publishing Group Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure

5 Continuum Mechanics and Its Practical Applications at the Level of Scaling Laws 6 Continuum Mechanics and Its Practical Applications at the Level of Scaling Laws 7 Continuum Mechanics and Its Practical Applications at the Level of Scaling Laws 8 Continuum Mechanics and Its Practical Applications at the Level of Scaling Laws 9 Continuum Mechanics and Its Practical Applications at the Level of Scaling Laws 2 Multi-scale Bone Mechanobiology 3 Multi-scale Bone Mechanobiology 5 Functional Magnetic Resonance Imaging 6 Functional Magnetic Resonance Imaging 4b Cell Mechanics and Cell Adhesion - Basic Principles and Computational Modeling 9 Diffusion Tensor Imaging 8 Corrosion of Biomaterials 8 Biomaterials for Tissue Engineering and Regenerative Medicine 9 Biomaterials for Tissue Engineering and Regenerative Medicine 12 Biomaterials for Tissue Engineering and Regenerative Medicine 3c Nanotechnology for Regenerative Engineering 4a Nanotechnology for Regenerative Engineering 4b Nanotechnology for Regenerative Engineering 5b Nanotechnology for Regenerative Engineering 5c Nanotechnology for Regenerative Engineering 6 Nanotechnology for Regenerative Engineering 7b Nanotechnology for Regenerative Engineering

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Figure 7c Nanotechnology for Regenerative Engineering Figure 8b Nanotechnology for Regenerative Engineering Figure 5 Drug and Gene Delivery for Regenerative Engineering Figure 3 Holographic Microscopy Figure 4 Radiomics http://www.nature.com

BIOMATERIALS: SCIENCE AND ENGINEERING Alternative Processing Techniques for CoCr Dental Alloys

Lucien Reclaru, VVSA, branch of Richemont International SA Varinor Innovation, Delémont, Switzerland Lavinia Cosmina Ardelean, “Victor Babes” University of Medicine and Pharmacy Timisoara, Timisoara, Romania © 2019 Elsevier Inc. All rights reserved.

Introduction Manufacturing Technologies for CoCr Alloys Evaluation of CoCr Alloys Manufactured by Different Technologies Chemical Composition of the Alloys Metallographical Evaluation Corrosion Evaluation Toxicological Aspects Conclusions References Relevant Websites

1 2 5 5 5 8 13 15 15 15

Glossary Castability The ease of forming a quality casting. Electrochemical corrosion testing A process for studying various forms of metallic corrosion, which provides information about the extent of corrosion activity. Polarization curve Plot of current density versus electrode potential for a specific electrode-electrolyte combination. Potentiodynamic polarization Probably the most commonly used polarization testing method for measuring corrosion resistance; a technique where the potential of the electrode is varied at a selected rate by application of a current through the electrolyte. Sintering The process of compacting and forming a solid mass of material using heat or pressure without melting it.

Introduction Since the early 1900s, a wide range of metals and their alloys are used in surgically implanted medical devices, prostheses and dental materials, in order to provide improved physical and chemical properties, such as strength, durability and corrosion resistance (Ardelean et al., 2015). The classes of metals used in medical devices and dental materials include stainless steels, cobalt–chromium alloys, and titanium (as alloys and unalloyed) (Wassell et al., 2002). In addition, dental casting alloys are based on precious metals (gold, platinum, palladium, and silver), nickel, and copper and may in some cases contain smaller amounts of many other elements, added to improve castability, handling, ceramic bond, or other physical properties. Alloys used in dental applications classify as shown in Table 1. Despite the wide range of dental alloys, each one of them has its shortcomings. Due to a number of factors, the use of CoCr based alloys increased for the last years. Noble alloys became too expensive for the average population. Cobalt alloys are considered an economic alternative to nonprecious nickel-based alloys, known as potential allergens. After a period of time when manufacturers tried to produce better dental alloys, for example, CoCr based alloys doped with precious metals, which were quite a disappointment (Ardelean et al., 2015, 2010, 2016a), today the trend changed to improving the manufacturing process, as an alternative, and not developing new compositions of the dental alloys. CoCr based alloys are known to have excellent corrosion resistance. Because of their outstanding mechanical properties (e.g., high stiffness) these alloys are mainly used for the fabrication of removable partial dentures, but also for metal-ceramic fixed dentures, where fine frameworks constructions are needed (Ardelean et al., 2015, 2016a). Unfortunately, they are difficult to treat and to process for the dental technician and for the dentist in traditional casting technique because of high hardness. Handling has considerably improved by using the new technologies which appeared in the last years, 3D printing being considered as the “next industrial revolution.” In our “high-speed” and “low-cost” world these effortless and time economically technologies, suitable for manufacturing the cheap CoCr alloys seem to be the best alternative available.

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Table 1

Classification of dental alloys

Precious alloys

Nonprecious alloys

Conventional crown and bridge alloys High gold Low gold Ag-based Alloys for the porcelain-fused-to metal technique High gold Low gold (NIOM Type B) Pd-based: Pd Ag, Pd Sn, Pd Cu, Pd Cr, Pd Co

Nickel-based Type A-Base Ni–Cr, major secondary element: Be Type B-Base Ni–Cr, major secondary element: Fe Type C-Base Ni–Cr, major secondary element: Mo Type D-Base Ni–Cr, major secondary element: Mn Type E-Base Ni–Cr–Fe CoCr-based, modified CoCr þ Pt, Ru, Nb, Au, In, Fe-CoCr Cu-based (bronze) Titanium: Ti grade 1-4 , Ti-6Al-4V, Ti-6Al-7Nb, Ti- Mo, Ti-40Zr, Ti-Pd-Co, Ti-50Ni, Ti-42.5Ni-7.5Pd, Ti-5Al-13Ta, Ti-5Ag, Ti-20Ag

Manufacturing Technologies for CoCr Alloys Manufacturing technologies currently available for CoCr alloys are: a. Traditional casting process (origin: bulk metal) This traditional laborious and time consuming method implies castability problems and porosities which often appear when specific flowing parameters of the casting machine are not respected. b. Milling process (origin: bulk metal) CAD/CAM technology implies obtaining a virtual image of the dental prosthesis based on an impression or a 3D image. The CAD software virtually designs the prosthetic element and the milling unit physically manufactures it (Fig. 1). Milling CoCr blanks (Fig. 2) may be difficult because its hardness and high demands are placed on the manufacturing unit (coolant delivery, rigidity of the machine etc.) (Ardelean et al., 2016b). A great progress was made by using milling units with four or five axes. In case of these milling units, the kinematics of the machine is optimized, with large angulations of the fourth and fifth axes, more than 30 degrees. Thus it allows the milling and the dry or wet grinding of very good quality dental prosthesis. On the other hand, the elements obtained by milling do not show any structure defects such as porosities, cracks etc. (Fig. 3). c. CAD/CAM sintering (origin: powder) This is a new technology developed by Amann Girrbach, which uses nonprecious metal blanks with a wax-like texture and allows them to be effortlessly dry milled in system’s benchtop machines. The milling is done in the green body state (unsintered metal powder held together by a binder). These dry millable CrCr blanks are similar to partially sintered zirconia blanks and can be easily processed in the preliminary state. After the required frameworks have been milled from the blank they are debinded and densely sintered in a downstream process. The sintering process of the milled CoCr structures is easy carried out in a special furnace and apparently gives a result with good structure quality: homogeneous, distortion-free frameworks without contraction cavities (www.amanngirrbach.com). d. 3D Printing Technologies (origin: powder) In principle these technologies are very much alike, the differences come mainly in the process itself. 3D printers seamlessly integrate with computer-assisted design (CAD) software and other digital files like magnetic resonance imaging.

Fig. 1

Milling of prosthetic elements according to CAD files.

Biomaterials: Science and Engineering j Alternative Processing Techniques for CoCr Dental Alloys

Fig. 2

CoCr blank for milling.

Fig. 3

CoCr prosthetic elements obtained by milling.

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When talking about CoCr alloys manufacturing by 3D Printing, some alternatives are available:

•

Selective laser sintering (SLS), developed by Dr. Carl Deckard in Austin USA.

In the mid-1980’s Deckard uses a moving laser beam to trace and selectively sinter powdered materials (including metals) into successive cross-sections of a three-dimensional part. Selective laser sintering (SLS) technique uses a high power laser (CO2 laser) to fuse small particles of metal powders into a mass representing a desired 3D object. Based on a virtual image, the various powders (CoCr, NiCr or ceramic) are slowly built, layer by layer, as the 3D CAD software measures thousands of cross-sections of each prosthetic element to determine exactly how each layer is to be constructed. As in all rapid prototyping processes, the parts are built upon a platform that adjusts in height equal to the thickness of the layer being built. After each cross-section is scanned, the powder bed is lowered by one layer thickness. Additional powder is deposited on top of each solidified layer and sintered. The powder is maintained at an elevated temperature so that it fuses easily upon exposure to the laser (www.popular3dprinters.com). The laser selectively fuses metal powders by scanning cross-sections generated from a 3D digital description of the part, a CAD file or scan data on the surface of a powder bed (Figs. 4 and 5) (Ardelean et al., 2016a; Reclaru et al., 2012a). The laser sintering technique makes possible the manufacture of extremely accurate prosthetic elements with mechanical properties that correspond to any clinical requirement (Fig. 6).

•

Direct metal laser sintering (DMLS), developed jointly by Rapid Product Innovations (RPI) and EOS GmbH in Munich, Germany.

Starting in 1994, is the first commercial rapid prototyping method to produce metal parts in a single process. Metal powder, free of binder or fluxing agent, is completely melted by the scanning of a high power laser beam to build the part with properties of the original material. The absence of the polymer binder avoids the burn-off and infiltration steps, and results in a 95% dense steel part compared to roughly 70% density in case of SLS. In addition DMLS has higher detail resolution than SLS, because of thinner layers, enabled by a smaller powder diameter (www.popular3dprinters.com).

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Biomaterials: Science and Engineering j Alternative Processing Techniques for CoCr Dental Alloys

Fig. 4

3D digital description of the computer-based program for a fabrication tray.

Fig. 5

The laser selectively fuses the CoCr metal powder.

Fig. 6

Laser sintered prosthetic elements.

•

Selective laser melting (SLM), started at the Fraunhofer Institute ILT in Aachen, Germany.

It is an additive manufacturing method that uses high powered laser to melt metallic powders together to shape the product from a 3D CAD data. Renishaw uses a high powered ytterbium fiber laser to fuse metal powders (www.renishaw.com/en/metal-3dprinting-for-healthcare). The recoater sweeps a layer of fine material powder and makes it ready for the laser to fuse them according to the 2D cross section of each layer under a tightly controlled inert atmosphere. When the part is made completely, it goes for the required heat treatment and postprocessing (www.popular3dprinters.com).

•

Electron beam melting (EBM), developed by ARCAM AB in Sweden-General Electric-2016.

This is another type of additive manufacturing for metal parts, often classified as a rapid manufacturing method. The technology manufactures parts by melting metal powder layer by layer with an electron beam in a high vacuum. Unlike some metal sintering techniques, the parts are fully dense, void-free, and extremely strong (www.popular3dprinters.com).

Biomaterials: Science and Engineering j Alternative Processing Techniques for CoCr Dental Alloys

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Evaluation of CoCr Alloys Manufactured by Different Technologies Chemical Composition of the Alloys The chemical composition of the CoCr alloys is shown in Table 2 (compositions as given by the manufacturers). When comparing chemical composition, CoCr alloys for casting, milling, CAD/CAM sinter process and SLS do not show great differences. In addition, the Ceramill Sintron alloy blanks contain an organic binder. Ceramill Sintron sinter alloy has comparable and, in the case of some parameters, even superior strength properties than cast CoCr. Similar evaluations are available in case of lasermelted alloys compared with cast CoCr alloys (Geis-Gerstorfer et al., 2013).

Metallographical Evaluation The CoCr samples were embedded in a cold-curing resin on a methyl methacrylate basis (Technovit, Kulzer), then polished with SiC paper (grit 320/500/1200/1400) and finally with diamond spray (6/3/1 mm) (Struers). Electrolytic etching has been done in a bath of 100 mL H2O dest., 10 mL HCl conc. and 5 g chromium (VI)-oxide during 5 s under 0.4 V and 0.3 A. The microstructures of the alloys were observed using a metallographical microscope (Polyvar Met, Reichert-Jung). Two scanning electron microscopes (JEOL JSM 6300) equipped with an EDX system (Oxford, INCA) for local phase analysis and SEM Sigma Zeiss with an Oxford X-MAX EDX Instrument) for microanalysis were used. The analyzed samples were covered with a gold flash. a. Traditional casting process (origin: bulk metal) The use of CoCr alloys is traditionally carried out by casting, which is quite an unwieldy process. The cast structure is a classical dendritic one (Fig. 7). b. Milling process (origin: bulk metal) Figs. 8–10 show the micrographical structures of the CoCr alloy without and with chemical attack. These are specific CoCr alloy gross flow structures, obtained either by continuous casting of molten metal, or by hot lamination and cut into a disc form. There are no abnormalities to be noticed. The chemical composition of the phases is given in Table 3. c. CAD/CAM sintering process (origin: powder) The micrographical structure without chemical attack shows micrometric porosities, homogeneously distributed in sections. No significant defects may be seen (Fig. 11).

Fig. 7

Table 2

Composition of the alloys (wt.%)

Element

Cast

Milled

Ceramill Sintron

SLS

Co Cr Mo Mn Si W Fe

63.7 28.9 5.3 0.8

60 24 4.5

0.1 0.4

8.5

66 28 5 HVTi-6Al-4V > HVTi.

Nanoindentation In nanoindentation, extremely smaller loads and a hard indenter tip (usually a diamond tip) with a geometry of three-sided pyramid (the so-called “Berkovich tip”) are used (Fig. 4B). The indented areas are very small, being only be a few square micrometers or even nanometers. During the course of the instrumented nanoindentation process, unlike conventional hardness or microhardness testing which only record the load applied, the depth of penetration of the indenter in the material being tested is continuously recorded. The area of the indent is determined using the known geometry of the indenter tip. During nanoindentation, the applied load and depth of penetration are measured to provide a load-displacement curve, which can be used to extract mechanical properties of the material being tested. Wang et al. used nanoindentation to study the mechanical properties of bone-like apatite formed in vitro on various biomaterials (Wang et al., 2002). They have demonstrated that nanoindentation is a powerful tool in determining mechanical properties of a very thin layer of bioceramic on biomaterials surface. Across the width of the bone-like apatite layer, the elastic modulus of bone-like apatite does not vary significantly. Fig. 6 displays a nanoindentation on bone-like apatite and a typical nanoindentation load-displacement curve for bone-like apatite. Using nanoindentation, Sun and Tong

Biomaterials: Science and Engineering j Bulk Properties of Biomaterials and Testing Techniques

Fig. 4 Testing at the micro- or nanoscale: (A) microhardness testing using a Vickers indenter, (B) indenter tip for nanoindentation, and (C) single molecule testing using optical tweezers.

Fig. 5

An indentation on sintered dense hydroxyapatite made by a Vickers indenter.

Fig. 6 Nanoindentation of bone-like apatite: (A) a nanoindentation made by a Berkovich indenter, and (B) a typical load-displacement curve for bone-like apatite.

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investigated fracture properties of three different natural materials: nacre, bovine hoof wall and beetle cuticle (Sun and Tong, 2007). Micro/nanoscale cracks are generated by nanoindentation using a Berkovich tip. The nanoindentation results in pile-up around the indent. The fracture toughness of bovine hoof wall is the highest, the second is nacre and the elytra cuticle of dung beetle is the lowest.

Tests for Single Molecules Mechanical characterization of single molecules or biological cells is very challenging, and cellular and molecular biomechanics have made significant advances over the past two decades due to the invention and innovation in testing techniques and testing equipment. Optical tweezers are now widely used to study various matters such as intramolecular and intermolecular interactions in the folding/mis-folding of proteins and RNAs and the dynamics underlying the function of molecular motors and complex macromolecular assemblies, in which mechanical forces can be applied to individual molecules by a force probe. Optical tweezers are a powerful tool in biomechanical studies owing to the very high resolution it can achieve and also its high sensitivity to rare and transient events. In a typical test setup using optical tweezers, the molecule of interest is attached to polystyrene beads at the two ends of the molecule, with one bead being trapped by a laser beam which holds the bead in position at this end. The coverslips that hold the bead at the other end of the molecule move with the motorized X–Y stage and hence the molecule is stretched during testing (i.e., tensile testing) (Fig. 4C). Two main measurement techniques can be used: first, non-equilibrium measurements, such as force-ramp or force-jump; and second, equilibrium measurements, where the molecule is held under a constant force and/or the traps are kept at a constant position and then the extension of the molecule is measured. Sun et al. stretched type II collagen molecules with optical tweezers and directly investigated the mechanical properties in a physiologically relevant condition (Sun et al., 2004). The molecule is stretched by moving the two-bead complex away from the optical trapping center at a constant rate through the piezo-stage. The displacement of the trapped bead is monitored by the use of a build-in interferometer. The highly nonlinear force-extension curves for type II collagen molecules are thus obtained. It is found that the force increases sharply when the molecules are stretched around their contour length. The molecular parameters of collagen are obtained by fitting force-extension curves using the worm-like chain elasticity model. The observations in this study indicate that collagen molecules are flexible molecules, rather than rigid rod, in the neutral pH solution.

Bonding Strength and Testing Methods Tensile Bonding Tests Surface properties are highly important for all biomaterials. Coatings made of biocompatible metals, polymers, ceramics or composites are often used for modifying the surface of current biomaterials or biomaterials under development. One of the important properties of coatings is the bonding strengthdtensile bonding strength and/or shear bonding strengthdwith the substrate (metal, polymer, ceramic, or composite). Because of the importance of bonding strength in nearly all branches of surface engineering, standard testing methods for bonding strength, albeit they are not perfect and sometimes yield questionable results due to inherent problems, have been long established. ASTM Standard D907 defines adhesion as “the state in which two surfaces are held together by interfacial forces which may consist of valence forces or interlocking forces or both.” ASTM C633 standard specifies test details for determining tensile bonding strength and shear bonding strength. It basically requires the detection of the load or stress to fail a standard joint. These tests need a strong bonding agent to attach the coating to a loading fixture. The test results suffer from epoxy penetration into the coating (when epoxy is used as a bonding agent in the test fixture). Another problem for both tensile and shear bonding strength tests is that the fracture mode can be ambiguous because mixed-mode failure often occurs in the tests, which make it indistinguishable for adhesive strength and cohesive strength of the coating. Apart from biomedical coatings, the bonding strength of biomaterials to body tissues is also important and similar test strategy can be adopted for the bonding strength determination/comparison. A key issue for the successful application of new resin-based composites as filling materials in dentistry is the bonding and performance of the composites. The composites need to infiltrate and impregnate the entire demineralized and opened collagen network of dentin to overcome humidity and to form a high crosslinking polymer inside the collagen network and hence the composite-dentin hybrid layer can have a high resistance to degradation. El-Malky and Abdelaziz studied the effect of pre-curing waiting time of different bonding resins on the tensile bonding strength to dentin (El-Malky and Abdelaziz, 2015). The methacrylate-based two-steps-total-etch bonding system and two-steps-self-etch bonding system show significant higher strength values than that of the one-step-self-etch bonding system. A 30-s pre-curing waiting time for bonding resins appears to provide high tensile bonding strength to dentin.

Shear Bonding Tests Biomedical coatings are formed on biomaterials (metals, polymers, ceramics or composites) for either enhancing biological properties, for example, biocompatibility and bioactivity, or improving physical, mechanical and/or other properties, for example, corrosion-resistance and reduced wear, of the biomaterial (the substrate). As discussed above, the shear bonding strength for biomedical coatings can also be measured using standard test methods. But there are problems in interpreting the test results

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when mixed-mode failure occurs in the tests. Again, in many situations, the shear bonding strength of a biomaterial to the host tissue is highly important. For example, dental adhesives must maintain their stability when shear masticatory forces are frequently applied. Dental adhesives are composed of resins or low-filler-content composites, which function to bond dental materials to teeth. The strength and durability of dental adhesives are the key factors for ensuring a successful and durable dental restoration. Biocompatible dental adhesives with enhanced performances are required and different types of adhesives are investigated. Sun et al. incorporated surface modified titania nanoparticles into dental adhesives (Sun et al., 2017). It is found that the addition of nanoparticles significantly enhance shear bonding strength between teeth and the titania-filled adhesives.

Peel Tests One important property of bioadhesives such as those used in surgical tapes is their adhesive strength. ASTM D1876 provides a standard for testing peel resistance for adhesives. The goal of a peel test is to determine the adhesive strength of a material or the strength of the adhesive bond between two materials. The adhesive strength may be referred to as the “stickiness” of an adhesive as it is a measure of the resistance of the adhesive to separation from the substrate after the adhesive has been applied. Peel tests for measuring the adhesive strength include T-peel, 90 degree peel, and 180 degree peel. The T-peel test is a type of tensile test performed upon two flexible substrates that have been bonded together and placed into peel test grips such that one substrate sticks up and the other sticks down while the bonded area sticks out horizontally so that the entire setup forms a “T” shape. The 90 degree test requires a 90 degree peel test fixture to determine the adhesive strength between a flexible (tape) and rigid substrate (plate), where the plate lies horizontally with the gripped end of the tape sticking up perpendicularly while the rest is bonded to the plate so that it forms an “L” shape. The 180 degree test is similar to the 90 degree peel test except that the bonded area between the tape and plate is placed vertically between the peel test grips while the free end of the of plate is gripped by the bottom and the free end of the tape is gripped by the top so that it forms a tight “U” shape. Lin et al. investigated drug-loaded wound dressing with thermoresponsive, adhesive, absorptive and easy peeling properties (Lin et al., 2001). The tack property of the EudragitE films increases with an increasing content of poly(N-isopropylacrylamide) (PNIPAAm) microgel beads, and the peel strength of Eudragit E films initially decreases with the addition of PNIPAAm microgel beads but increases to a maximum value when PNIPAAm microgel beads are added from 4% to 7.6%.

Summary Short-term mechanical tests for tension, compression, shear, bending, torsion, etc. provide fundamental mechanical properties that can be used to compare different biomaterials for targeted applications or during the development of new biomaterials, to study (bio)materials processing-structure-property relationships and screen out materials. Long-term mechanical tests are highly important for load-bearing implants/devices and hence creep (for biomedical polymers and polymeric composites, as metallic and ceramic biomaterials do not creep at the human body temperature), fatigue, and wear tests need to be conducted under simulated human body conditions. Bonding strength determination is needed for biomedical coatings and bioadhesives as well as for assessing the biomaterial-host tissue bond. Mechanical tests at the micro- and nanoscale are performed not only for scientific research but also for new biomaterials development. For all types of mechanical tests, if a testing standard has been developed and is available, the test must be conducted according to the standard. This will ensure that test results can be compared and used by different groups of people (researchers, manufacturer, medical device end users, etc.) across national boundaries.

Acknowledgements Min Wang thanks Nanyang Technological University, Hong Kong Polytechnic University, The University of Hong Kong, UK’s Engineering and Physical Sciences Research Council, Singapore’s Ministry of Education, Hong Kong Research Grants Council and the National Natural Science Foundation of China for research funding. Chong Wang thanks the Office of Education of Guangdong Province for funding his Young Innovative Talent Project and Dongguan University of Technology for support with the High-level Talents (Innovation Team) Research Project.

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Panigrahi, A., Sulkowski, B., Waitz, T., Ozaltin, K., Chrominski, W., Pukenas, A., Horky, J., Lewandowska, M., Skrotzki, W., & Zehetbauer, M. (2016). Journal of the Mechanical Behavior of Biomedical Materials, 62, 93–105. Puertolas, J. A., Vadillo, J. L., Sanchez-Salcedo, V. S., Nieto, A., & Vallet-Regi, M. (2011). Acta Biomaterialia, 7, 841–847. Sun, J., & Tong, J. (2007). Journal of Bionic Engineering, 4, 11–17. Sun, Y. L., Luo, Z. P., Fertala, A., & An, K. N. (2004). Journal of Biomechanics, 37, 1665–1669. Sun, Y., Huang, Y. J., Fan, H. B., Liu, F. Y., & Chen, J. J. (2014). Journal of Non-Crystalline Solids, 406, 144–150. Sun, J., Petersen, E. J., Watson, S. S., Sims, C. M., Kassman, A., Frukhtbeyn, S., Skrtic, D., Ok, M. T., Jacobs, D. S., Reipa, V., Ye, Q., & Nelson, B. C. (2017). Acta Biomaterialia, 53, 585–597. Takahashi, Y., Tateiwa, T., Shishido, T., Masaoka, T., & Yamamoto, K. (2016). Journal of the Mechanical Behavior of Biomedical Materials, 62, 399–406. Wang, M., Chen, L. J., & Ni, J. (2002). Transactions of the Society For Biomaterials 28th Annual Meeting, Tampa, Florida, USA (p. p. 259). Wei, C. K., & Ding, S. J. (2016). Journal of the Mechanical Behavior of Biomedical Materials, 62, 366–383. Yang, X., & Hutchinson, C. R. (2016). Acta Biomaterialia, 42, 429–439. Zhao, X. F., Niinomi, M., Nakai, M., & Hieda, J. (2012). Acta Biomaterialia, 8, 1990–1997.

Further Reading Callister, W. D., Jr., & Rethwisch, D. G. (2014). Materials science and engineeringdAn introduction (9th edn.). John Wiley & Sons. Hertzberg, R. W., Vinci, R. P., & Hertzberg, J. L. (2013). Deformation and fracture mechanics of engineering materials (5th edn.). John Wiley & Sons. Meyers, M. A., & Chawla, K. K. (2008). Mechanical behavior of materials (2nd edn.). Cambridge University Press.

Corrosion of Orthopedic Implants Qiong Wang, Felipe Eltit, and Rizhi Wang, The University of British Columbia, Vancouver, BC, Canada © 2019 Elsevier Inc. All rights reserved.

Introduction Metals in Hip Implants Fundamentals of Corrosion Mechanisms General Concepts Types of Corrosions in Hip Implants In Vivo Corrosion Environment Corrosion-Induced Implant Fracture Clinical Fracture of Implants Corrosion in Fatigue Crack Initiation Corrosion-Induced Tissue Failure Ion and Particle Release Mechanisms Ion release mechanisms Particulates release mechanisms Biological Reactions to Corrosion Products Adverse local tissue reaction to metal implants Histopathology of ALTRs Roles of metal ions in tissues and fluids Biological effect of metal particles New Biomaterial Designs Future Perspectives References

65 66 67 67 68 70 70 70 71 72 72 72 73 74 74 75 75 77 79 79 80

Introduction Metals and their alloys have been successfully used as biomaterials in a wide range of implantable medical devices, from coronary stents and artificial hearts to dental implants and orthopedic implants (Ratner et al., 2004; Williams, 2014). Their essential roles as core structural materials in implants are likely to remain the same for decades to come. This is because metals have the best combination of stiffness, strength, toughness, and formability among all engineering materials including polymers, ceramics, and their composites. Despite their generally robust performance as biomaterials, metals suffer from low corrosion resistance (as compared with ceramics). Under the combination of the corrosive physiological environment and the challenging biomechanical loading, metallic implants may degrade. The consequences of implant corrosion can be serious and are twofold. First, corrosion may eventually lead to the destruction and mechanical failure of the implants. Second, metal ions and particles released from the implants may cause adverse tissue reactions (ALTRs) that will ultimately require surgical revision to the implants. The purpose of this article is thus to review the corrosion of metallic implants, from mechanisms to clinical issues. We further focus our discussion on total hip implants in orthopedics. Orthopedic implants comprised nearly half of all medical device implants in use (Moore et al., 1991). Among them, over 1 million total hip replacements (THR) are performed annually worldwide to relieve pain and restore mobility of patients with osteoarthritis (Bozic et al., 2009; Singh, 2011; Learmonth et al., 2007). Replacing the mobile hips also represents the most challenging biomechanical application of metallic biomaterials. It is our hope that the readers will find the basic mechanisms applicable to other medical implants. A THR system consists of a metallic femoral stem (mainly a Ti alloy), attached to a hemispherical femoral head articulating against an acetabular cup with a liner (Fig. 1). One important design characteristic in hip replacements is the combination of materials that make up the bearing surface of the implanted hip joint, namely the material used for the articulating femoral head and the acetabular liner. The femoral head is usually made of materials with high wear resistance such as CoCr alloys, and ceramics. The liner is mainly made of polyethylene in addition to metals and ceramics. Usually, the material combination used at articulating surfaces was used to abbreviate the types of implants. For example metal-on-polyethylene (MoP) implant means a type of implant consisting of a metal femoral head articulating against a polyethylene liner. Modularity has been commonly used in THR allowing the surgeons to select components independently based on patient’s individual anatomy, offering greater variety of solutions during implantation, easy access in revision procedures, and reduced inventory (Marlowe et al., 2013). Primary THR surgery has > 90% success rate within a 10-year follow-up, which is one of the greatest advances in the healthcare of the 20th century (Learmonth et al., 2007; Older, 2002; Malchau et al., 2002; Williams et al., 2002). Historically, THRs fail due to aseptic loosening, infection, and dislocation (Malchau et al., 2002; Ulrich et al., 2008; The NJR Editorial Board, 2016). However, with the increasing use of modular components, complications related to corrosion emerged, leading to failure of modular systems

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Fig. 1

The components of a MoP hip replacement system.

and consequently increased revision rate. One complication is the ALTRs developed to some latest designs that require revision surgeries. As a result, regulatory agencies including Health Canada and FDA have issued alerts related to MoM hip implants (Healthy Canadians, 2012; U.S. Food and Drug Administration, 2013). ALTRs have also been observed recently in patients with MoP implants, the most commonly used design today (Whitehouse et al., 2015; Cooper et al., 2012; Picardo et al., 2011; Perino et al., 2014). Although the exact mechanisms of ALTR are unclear at present, the role of corrosion products released from the metallic materials has been widely accepted (Meyer et al., 2012; Cooper et al., 2013; Langton et al., 2010). In addition to ALTRs, corrosion may also induce mechanical failures such as fracture and loosening of components. Studies have linked these mechanical failures to corrosion at modular interfaces (Ellman and Levine, 2013; Nima et al., 2013). These clinical failures raised concerns over implant corrosion, and call for in-depth investigation of the underlying mechanisms. In the following sections, we first introduce different types of metallic biomaterials, followed by common corrosion attacks and their corresponding mechanisms. The role of corrosion in two clinical failuresdthe mechanical failure of the implants and the ALTRsdis elaborated. Development of metallic materials used in hip implants to improve their clinical performances against corrosion is also discussed.

Metals in Hip Implants Metals are the most used materials in hip replacement due to their mechanical properties. In Canada, > 90% of femoral heads used in hip replacements are made of metals (Canadian Institute for Health Information, 2015). In the United Kingdom, the metallic femoral heads are used in 65% of hip implants (The NJR Editorial Board, 2016). An ideal metallic material used in hip replacement is expected to have great biocompatibility, excellent resistance against corrosion and wear, acceptable strength to endure the cyclic loading, and a low Young’s modulus to minimize stress shielding to bone (Park and Lakes, 2007). Table 1 summarizes metallic materials commonly used in hip implants and the required mechanical properties in the regulation. Stainless steels (SS) are iron-based alloys containing a minimum of 12% Cr to prevent corrosion in the atmosphere. SS is easy to work and has high strength and low price. Type 316L used in hip replacement is an austenitic SS and contains around 17% Cr, 12% Ni, and 2.5% Mo with carbon below 0.03%. CoCr alloys are commonly used as the bearing surface due to their high corrosion and wear resistance. Forged CoCr is also widely used in the femoral stem for its excellent fatigue resistance. The CoCr alloys are composed of 27%–30%Cr, 5.0%–7.0% Mo, and small amount of other elements (Mn, Si, Ni, Fe, and C) with balanced Co (ASTM, 2014a,b). Chromium is expected to form a passive film spontaneously on the surface of the metal, offering the required corrosion resistance (Contu et al., 2003; Yan et al., 2006). Molybdenum is added to increase pitting resistance and to produce finer grains with precipitates of hard carbide phases together with chromium (Liao et al., 2012; Sugimoto and Sawada, 1977). Based on carbon contents, CoCr alloys are categorized into high carbon alloys with 0.05–0.35 wt% carbon, and low carbon alloys with carbon concentration < 0.05 wt%. Higher carbon content usually is associated with more carbide phase precipitates in the alloy, and hence superior wear resistance (Affatato et al., 2011). However, chemical and microstructural inhomogeneity caused by carbide precipitates compromises the alloy’s

Biomaterials: Science and Engineering j Corrosion of Orthopedic Implants Table 1

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Mechanical properties of metals used in hip implants

Materials

ASTM#

Condition

E(GPa)

sys min (MPa)

suts min (MPa)

d(%) min

Stainless steel

F138

Co–28Cr–6Mo Co–20Cr–15W–10Ni

F75 F799 F90

cp-Ti Ti–6Al–4V Ti–6Al–7Nb Ta

F67 F136 F1295 F560

Annealed 30% worked Cast forged Annealed Cold worked Grade 4 Wrought Wrought Cold worked Annealed

200 200 210 210 210 210 110 110 105 186 186

190 690 450 827 310 760 483 795 800 345 140

490 860 655 1172 860 1250 550 860 900 480 210

40 12 8 12 30 15 15 10 10 1 8

resistance to localized corrosion (Montero-Ocampo and Rodriguez, 1995). In the application of hip replacement, thermal treatment such as fast cooling is used to produce FCC (face-centered-cubic) structure in CoCr (Brooks, 1982; Davis, 1990). And the FCC phase remains metastable due to very low kinetic transformation from FCC to HCP (hexagonal-close-packed) structure at room temperature (López and Saldivar-Garcia, 2007). Ti alloys are commonly used for its strong corrosion resistance and high biocompatibility. Commercially pure titanium (cp-Ti) has two allotropic forms (Welsch et al., 1993). From room temperature up to 882 C, titanium exists as a phase of hexagonal-close-packed (HCP) crystal structure. Above 882 C, it transfers to b, which has a body-centered-cubic crystal structure. The mechanical properties depend on the relative amounts and distribution of a and b phase. To stabilize these two phases, elements such as aluminum, tin, or zirconium are commonly added to stabilize a phase, while molybdenum, vanadium, or niobium are often added to stabilize b phase. The most commonly used a þ b dual-phase Ti alloy in biomedical device is Ti–6Al–4V (Rack and Qazi, 2006). Tantalum is also used as for enhanced osseointegration, strong corrosion resistance, and lower elastic modulus (Davis, 2003; Bobyn et al., 1999). Due to the mismatch of elastic modulus between bone (10–30 GPa) and metals used in femoral implant, bone is not sufficiently loaded and becomes stress shielded, which eventually leads to bone resorption and compromise of clinical performance. To solve this problem, low elastic modulus of porous titanium and tantalum is developed to allow an effective load transfer and bone preservation (Ryan et al., 2006; Balla et al., 2010; Levine et al., 2006).

Fundamentals of Corrosion Mechanisms General Concepts Corrosion is an electrochemical reaction involving electron transfers. An electrochemical reaction consists of two half-reactions. Anodic reaction is a half-reaction of oxidization releasing electrons, for example, a metal turns into metal ions in Reaction (1). The other half-reaction named cathodic reaction is a reduction process gaining electrons, for example, a reduction of protons forms hydrogen gas in Reaction (2). The two half-reactions coupled together form an overall electrochemical reaction occurring in corrosion: M # Mþ þ e

(1)

2Hþ þ 2e # H2(g)

(2)

The corrosion reaction can be examined by two basic theories. The first is the thermodynamic theory characterizing the driving forces for corrosion reactions, revealing whether the reactions may happen. The other is the kinetic process related to the reaction rates, revealing how fast the corrosion reaction happens. For thermodynamics, the driving force for the reaction is the change in Gibbs free energy (D G). When D G is negative, its corresponding reaction will proceed spontaneously. D G is also associated with electrode potential in this expression D G ¼  nFE, where n is the number of electrons transferred, F is the Faraday constant, and E is the electrode potential (reduction expression) which electrons need to overcome to transfer in between. For an overall electrochemical reaction, similar D G expression can be written, where E is the cell potential given by Ecell ¼ Ecathode  Eanode. Based on thermodynamics theory, the more negative of its electrode potential, the more active it is as an anode, and the more likely it is to be oxidized. The other factor governing the corrosion is the kinetic process that determines how fast it happens. The well-known kinetic barrier is the passive film (usually an oxide layer) formed on a metal surface. Cr and Ti, two commonly used elements in implant alloys, are very active according to the thermodynamics (Table 2) (Haynes, 2004). However, the passive films spontaneously

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Biomaterials: Science and Engineering j Corrosion of Orthopedic Implants Standard electrode potentials (Haynes, 2004)

Table 2 Half-reaction

E (V)

2Hþ þ 2e # H2(g) 2Ni2 þ þ 2e # Ni(s) 2Fe2 þ þ 2e # Fe (s) Cr3 þ þ 3e # Cr(s) 2Ti2 þ þ 2e # Ti (s)

0 0.25 0.44 0.74 1.63

formed on these two metals are very dense and intact, preventing the active dissolution of metal substrates. This passivation of Cr and Ti enables alloys such as 316L SS, CoCr, and Ti alloys to possess strong corrosion resistance.

Types of Corrosions in Hip Implants Corrosion may be general or localized. General corrosion involves the uniform dissolution of the metal surface, while localized attack occurs on specific sites of a passive metal surface where there are high local dissolution rates. For alloys used in hip implants, the most relevant types are localized corrosion such as galvanic corrosion, pitting, crevice corrosion, and mechanically assisted corrosion. Galvanic corrosion: Galvanic corrosion tends to occur when dissimilar materials are connected electrically and exposed to an electrolyte. The active material with lower electrode potential is preferentially corroded to another due to the driving force of the electrode potential difference between these dissimilar metals. In most cases, however, galvanic corrosion is not a major concern on multi-alloy hip prostheses. Electrochemical studies showed low corrosion rate of 0.02 mA/cm2 when cobalt and titanium alloys were coupled and no instances of extensive corrosion were found on cobalt–titanium interfaces of the retrieved implants (Lucas et al., 1981). No detectable galvanic corrosion has been observed on (SS)-Ti alloy mixed interfaces in vitro (Serhan et al., 2004). Without pitting or crevice corrosion of the SS, the coupling of SS-CoCr alloy remains in the stable passive state, and no significant change of corrosion behavior was detected on the materials (Reclaru et al., 2002). Since passive film formed on the metal surface remains intact in static environment serving as an efficient barrier to corrosion, the risk of galvanic corrosion is negligible for alloys used in a biomedical device (Virtanen et al., 2008; Mears, 1975). Crevice corrosion and pitting: Passive oxide films are remarkable in their ability to protect a wide variety of metals and alloys from corrosion. However, passive films often break down in certain areas where localized corrosion attack happens. The most common forms of localized attack are crevice corrosion and pitting. Crevice corrosion occurs within confined space or under shielded metal surfaces. Pitting is the localized breakdown of passive films usually caused by halogen ions such as chloride, leading to the subsequent formation of pits at these sites. Crevice corrosion and pitting differ in their mechanisms of initiation but share similar mechanism of propagation (Fig. 2) (McCafferty, 2010). Crevice corrosion occurs in geometrical clearances such as the modular interfaces of hip implants. At the very beginning of crevice corrosion, oxygen reduction occurs both on the metal surface exposed to the bulk electrolyte and on the portion of the metal surface within the crevice. However, the external metal exposed to the bulk electrolyte has abundant supply of oxygen from the atmosphere so that a steady-state concentration of O2 is maintained. On the contrary, within the narrow clearance of the crevice, the oxygen becomes depleted due to the long diffusion path formed by the crevice. This difference in O2 concentration between the bulk solution and crevice becomes the driving force which makes the metal exposed to lower concentration of oxygen to have a more negative potential for oxygen reduction. This electrode potential difference between internal and external metal will initiate corrosion of metals in the crevice. The electrochemical reactions of the crevice reactions are as follows: Cathode (on metal outside of crevice): O2 þ 2H2 O þ 4 e #4 OH Anode (on metal in the crevice): Cr#Cr 3þ þ 3 e Cr 3þ þ 2H2 O#Cr ðOHÞ3 þ 3Hþ As the anodic reaction shows, the electrolyte composition within the crevice will become acidic due to the hydrolysis of metal cations. This aggressive electrolyte in the crevice will break down the passive layer and continue to dissolve the metallic element. Pitting is another localized corrosion attack. Pitting is caused by the presence of aggressive anions such as halogen ions, usually chloride in the electrolyte. The chloride ion is a strong electron donor with relatively small volume and high diffusivity. And it tends to interact with electron acceptors such as metal cations, resulting in the breakdown of the passive film. Crevice corrosion and pitting share similar mechanisms in propagation. The electrolyte composition within the crevice and pits will become acidic due to the hydrolysis of metal cations (anodic reaction above). And it will also contain concentrated amounts of

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Fig. 2 Schematic illustration of crevice corrosion (A), (B), (C), and pitting (D). (D) is the possible pitting happening when Cl is present in the crevice. Pictures are modified based on McCafferty, E. (2010). Crevice corrosion and pitting. In Introduction to corrosion science. New York, NY: Springer New York; pp. 1–575.

cations discharged from the alloy. In chloride solutions, the crevice and pits will also become concentrated in chloride ions to maintain the charge neutral. This internal electrolyte is sufficiently aggressive to break down the passive film on the metal, which will further develop corrosion. SS 316L shows higher susceptibility to pitting and crevice corrosion than CoCr and Ti alloys (Gurappa, 2002; Walczak et al., 1998). The resistance to pitting corrosion depends on Cr and Mo content. Moreover, relatively small variations in steel composition or surface condition such as roughness may significantly influence the pitting behavior (Virtanen et al., 2008). Studies on retrieved implants show that > 90% of the failure of 316L SS is due to the pitting and crevice corrosion attack (Sivakumar et al., 1995). Since CoCr alloy has relatively higher content of Cr and Mo, pitting corrosion is not common. High-carbon CoCr alloy is more susceptible to pitting corrosion due to the segregation of hard phases at the grain boundary (Panigrahi et al., 2014). Compared with CoCr alloy, titanium alloys possess much higher resistance against corrosion, especially crevice corrosion (Gurappa, 2002; Gurrappa, 2003). Indeed crevice attack of titanium alloys does not generally occur at temperature < 70 C regardless of solution pH or chloride concentration (Welsch et al., 1993). Mechanically assisted corrosion: In modular hip implants, corrosion usually occurs when the passive films are damaged by mechanical wear. The commonly observed mechanically assisted corrosions at modular implants are fretting corrosion and tribocorrosion. Fretting corrosion is a type of corrosion assisted by micromotion-induced wear (Gilbert et al., 1993). This wear disrupts the passive film, exposing the alloy that is susceptible of corrosion until a new passive layer is formed (Hallab and Jacobs, 2003; Urban et al., 2006; Swaminathan and Gilbert, 2012). This cyclic process of wear-exposure-repassivation happens at the modular junctions of the implants, such as the head–neck junction and the neck-body junction, leading to the active corrosion and generation of debris. The amplitude of displacement in fretting corrosion is governed by the applied bending moment and the rigidity of the taper in implants (Gilbert and Mali, 2012). As Fig. 3 shows, the two points (solid triangle) within the taper regions are considered rigidly connected to the opposite surface. The elastic strain associated with cyclic bending will give rise to a displacement (D) within the taper interface in the range of 20–40 mm for typical tapers used today. At the same time, the confined taper interface is a crevice geometry where crevice corrosion is favored, resulting in accelerated corrosion (Marlowe et al., 2013). Studies have confirmed that modular designs are susceptible to fretting corrosion in vivo. In a study, 16 retrieved modular titanium alloy stems from cementless hip replacement were analyzed, and evident corrosion sites were characterized with fretting scars, pitting, and etching (Urban et al., 2006). Most of the retrieved modular tapers showed severe fretting corrosion both at head–neck taper and neck-stem interfaces (Rodrigues et al., 2009; Wright et al., 2010; Lakstein et al., 2011; Brown et al., 1995). Fretting played a major role in the initiation of the corrosion at the modular interfaces, and the long head–neck extension was more prone to fretting corrosion due to the larger fretting amplitude (Brown et al., 1995). It has also been proposed that combining different alloys in a modular junction (e.g., Ti alloy combined with CoCr alloy) increases the risk of fretting corrosion at the interface between them (Gilbert et al., 1993; Collier et al., 1992; Cook et al., 1994).

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Fig. 3 Schematic illustration of fretting at the modular junction: (A) under cyclic bending, two points (solid triangle) are at rigid contact and the elastic strain gives rise to a displacement (D) within the taper, crevice geometry is formed at the junction. (B) Zoom-in on the contact region with the modular taper, metal-oxide surfaces in asperity contact cause contact stresses, local surface deformation and oxide debris (Gilbert & Mali, 2012).

Tribocorrosion is a general term used to describe wear-assisted corrosion. Technically, fretting corrosion is a subcategory of tribocorrosion (Mathew et al., 2009). Here, tribocorrosion is used to describe the specific wear-assisted corrosion at articulating surfaces. Compared to fretting corrosion, tribocorrosion is characterized with large amplitude of wear without assistance of evident crevice corrosion. The direct consequence of tribocorrosion is the release of wear debris from the bearing materials (Wimmer et al., 2013; Büscher et al., 2005; Büscher and Fischer, 2005; Illgen et al., 2008; Catelas et al., 2011; Doorn et al., 1996).

In Vivo Corrosion Environment The normal body fluid is a relatively mild environment of buffered solution at around pH of 7.4 with temperature of 37 C (Hanawa, 2004). However, there are some specific in vivo corrosive elements for metallic materials such as chloride ions, localized pH drop, proteins, and cell-induced oxidizers. The concentration of chloride ion in synovial fluid is similar to that in serum (Perman, 1980). Drop in pH values was observed in revised implants due to crevice corrosion (Willert et al., 1996) and periprosthetic tissues because of osteoclasts activities during bone remodeling (Konttinen et al., 2001). The decreased pH may accelerate the corrosion of alloys used in the hip implants. The presence of proteins in the physiological environment will affect the corrosion of metallic materials via changing the mechanics and kinetics of the corrosion reactions on the surface (Valero Vidal et al., 2010; Igual Munoz et al., 2015; Okazaki and Gotoh, 2005; Goldberg and Gilbert, 1997; Hallab et al., 2003). Additionally, active oxygen species released by immunological cells such as macrophages when responding to a foreign body at acute inflammation are potential oxidizers assisting corrosion process (Hanawa, 2004; Gilbert et al., 2015; Liu and Gilbert, 2017). Therefore, it is extremely challenging to simulate the exact in vivo environment in a laboratory setup.

Corrosion-Induced Implant Fracture Clinical Fracture of Implants The fatigue strengths of metallic biomaterials have been well documented and are listed in Fig. 4 (Niinomi, 2007). The majority of these alloys (especially forged CoCr and Ti alloys) have fatigue strength > 500 MPa (in air) (Niinomi, 2007; Black and Hastings, 1998), which is much higher than the strength of human bones (100–200 MPa) (Teoh, 2000). Historically, fracture of femoral stem was not common in THR. However, in the recent years, increasing cases of fracture associated with modular junctions of hip replacements have been reported. Prior to the introduction of modularity in the stems of hip implants, the prevalence of fracture in the femoral component was estimated to be 0.23% (Wright et al., 2010; Lakstein et al., 2011; Charnley, 1975). However, the introduction of modularity in the femoral stems has raised the risk of mechanical failure up to 1.4% (Krishnan et al., 2013; Grupp et al., 2010), and some models have been recalled from the market due to their elevated failure rate (U.S. Food and Drug Administration, 2012). Fatigue fracture of titanium alloys: Most of the fractures of titanium components occurred at the stem-taper junction in the form of fatigue (Wright et al., 2010; Grupp et al., 2010; Dunbar, 2010). In a study of 5000 cases of patients with modular titanium femoral

Biomaterials: Science and Engineering j Corrosion of Orthopedic Implants

Fig. 4

71

Fatigue strength at 107 cycles of biomedical alloys and bone (Niinomi, 2007).

neck, fatigue fracture was observed in up to 1.4% of them within the first 2 years of implantation (Grupp et al., 2010). Investigation of fracture has revealed different contributing factors such as overweight of the patient (Ellman and Levine, 2013), inadequate bone support (Lakstein et al., 2011), and stress concentration due to design (Huot Carlson et al., 2012). The most widely reported mechanism in the initiation of fatigue fracture of the Ti alloy components in modular hip systems is fretting (Ellman and Levine, 2013; Huot Carlson et al., 2012; Atwood et al., 2010; Paliwal et al., 2010). Nevertheless, corrosion-induced fracture of Ti alloy component has been reported (Gilbert et al., 2012). Fatigue fracture of CoCr alloys: CoCr alloys have higher fatigue resistance compared to Ti alloys so that the risk of CoCr fatigue occurrence is supposed to be lower. Unfortunately, fatigue fracture has also been observed in CoCr alloys. One mechanism for fatigue fracture on CoCr is the crack nucleation at microstructural defects such as voids, inclusions, and laser etching accidently introduced during manufacturing (Galante, 1980; Della Valle et al., 2005; Woolson et al., 1997; Lee and Kim, 2001). Another possible reason for fatigue fracture of CoCr stem is the overload due to undersized component in some designs (Norman et al., 2014; Ishaque et al., 2011). The most common factor triggering fatigue fracture of CoCr alloys in modular hip replacement system is corrosion, such as fretting corrosion and intergranular corrosions (Norman et al., 2014; Wang et al., 2017; Mencière et al., 2014; Gilbert et al., 1994).

Corrosion in Fatigue Crack Initiation Corrosion is a key factor in the fatigue of alloys used in hip replacement. Corrosion attacks in a physiological environment create surface defects and accelerate crack nucleation and propagation; hence, the fatigue strength of alloys can be significantly reduced leading to increased risk of fracture (Niinomi, 2007; Antunes and de Oliveira, 2012). An important question regarding corrosion fatigue is how surface cracks are initiated in the presence of corrosion. Pitting corrosion: Pitting is the major corrosion factor responsible for crack initiation as surface defects (pits) are local stress concentration sites under tensile stress. Although pitting corrosion in CoCr is seldom observed in laboratory polarization studies in simulated body fluid, since it typically fails by transpassive dissolution (Bettini et al., 2011), pitting in retrieved CoCr alloys has been reported (Gilbert et al., 1993, 1994). In one study, pitting induced by intergranular corrosion led to the removal of particles from the surface (Gilbert et al., 1993). Recent laboratory tests suggested that pitting corrosion could happen at carbide boundaries and high-energy grain boundaries in electrochemically corroded CoCr samples (Panigrahi et al., 2014; Bettini et al., 2012; Wimmer et al., 2001), especially in the presence of initial high contact pressure (Hanawa et al., 2001) and proteins (Mathew et al., 2010). The stress might tear the carbides off the surface (Hanawa et al., 2001) and the presence of proteins lowered the pitting potential, leading to localized chemical attacks (Mathew et al., 2010). Pitting on Ti alloys in a physiological environment is extremely rare as Ti shows

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very strong resistance against pitting (Gilbert et al., 2015). Corrosion features similar to pitting were observed in a retrieved Ti alloy stem and were associated with possible hydrogen embrittlement in the alloy (Rodrigues et al., 2009). Fretting corrosion: Another key factor in crack initiation is fretting corrosion. Fretting fatigue has been widely reported in the fracture of Ti alloy components in modular hip systems (Ellman and Levine, 2013; Huot Carlson et al., 2012; Atwood et al., 2010; Paliwal et al., 2010). However, fretting corrosion-induced fracture in modular CoCr stems is less commonly documented in the literature (Mencière et al., 2014). The residual stress and plastic deformation caused by fretting may create possible sites for crack initiation. The released debris due to fretting at the surface will accelerate the crack initiation stage (Hoeppner and Chandrasekaran, 1994). Fretting could also trigger the release of carbides from the surface, which may further enhance intergranular pitting around the carbides. These corrosion pits served as preferential sites for microcrack formation (Wang et al., 2017). Fretting in modular hip implants is usually accompanied by crevice corrosion. Electrochemical reactions in the crevice of the fretting interface might change the electrolyte composition and drop the pH, resulting in an aggressive environment enhancing other types of corrosion such as pitting to further facilitate crack initiation. After a crack is formed, the trapped fretting corrosion products would prevent the crack from closing under fluctuating stresses, which may facilitate further corrosion attack at the crack tip (Wang et al., 2017; Fatigue of Structures and Materials, 2009). Stress-assisted corrosion cracking: Under normal physiological loading conditions, the stem of a THR is subject to fluctuating tensile stress on the lateral side and compressive stress on medial side (Hampton et al., 1980; Cicero et al., 2007). Cyclic tensile stress in the dynamic loading of the hip implants assists the growth of surface pits into cluster of microcracks and the coalescence of those cracks (Wang et al., 2017). Such a stress-assisted crack initiation process is similar to stress corrosion cracking (SCC), except that the tensile stress involved is dynamic instead of static (Fatigue of Structures and Materials, 2009). Historically, SCC was often observed on SS since they are relatively more susceptible to pitting attack than CoCr and Ti alloys (Bundy et al., 1983; Sivakumar and Rajeswari, 1992; Bombara and Cavallini, 1977). Typical SCC responsible for fatigue crack nucleation on Ti alloys was rarely reported, although Gilbert et al. (2012) proposed a mechanism similar to SCC on a Ti–6Al–4V taper, in which the cracks propagated due to the crack tip stress arising from the oxide formation rather than externally tensile stress. Wang et al. reported the branching and the coalescence of cracks on the external surfaces of the CoCr adaptor and their association with the tensile stress suggesting a similar crack initiation mechanism to SCC (Wang et al., 2017). The risk of corrosion cracking under tensile stress (either static or dynamic) is that it could facilitate subcritical crack growth for later fatigue fracture at a stress level that is significantly lower than the yield strength of the material. Especially, when CoCr components are used together with Ti alloy under dynamic tensile stress, stress-assisted corrosion cracking may happen in the presence of crevice pitting and fretting, which may accelerate the crack nucleation process leading to early fracture. Corrosion-induced fracture is usually a result of the combination of multiple factors. Fretting corrosion at the modular interface accelerates crevice corrosion and pitting. The local aggressive electrolyte resulting from crevice corrosion and pitting will further facilitate fretting corrosion. The surface defects due to multiple corrosion attacks are preferred locations for crack initiation, especially on the lateral side of the implants subjected to dynamic tensile stress. Some of these cracks coalesced and propagated to a critical size under dynamic tensile loading, leading to the later corrosion fatigue fracture.

Corrosion-Induced Tissue Failure The described mechanisms of corrosion in the hip replacements systems, especially fretting corrosion and tribocorrosion that are present at the metal interfaces of the different implant designs, release large amounts of metal species from the implant, in the form of particulate materials and ions, to the synovial fluid, periprosthetic tissue, and blood of patients. The evolution of these corrosion products and the adverse reactions of peri-implant tissues are reviewed in this section.

Ion and Particle Release Mechanisms Ion release mechanisms Cobalt release: Cobalt is preferentially released when passive film is formed on the CoCr surface. CoCr alloys passivate spontaneously in air, resulting in the formation of a thin oxide film (around 2 nm) containing mainly Cr 3 þ, with a minor content of Co and Mo (Hanawa et al., 2001; Li et al., 1999; Milosev and Strehblow, 2003). The composition of the oxide film depends on the applied potential. In the simulated physiological solution (SPS), the passive layer consists predominantly of Cr2O3 and Cr(OH)3 at potential < 0.3 V, but at higher potentials both Co and Mo oxides can be found in the passive layer (Milosev and Strehblow, 2003). After immersion in Hanks’ solution, cobalt is dissolved, leaving Cr 3þ and Mo 4 þ, 5þ, 6þ in the surface oxide (Hanawa et al., 2001). Complexation of Co to proteins such as albumin will form a precipitate so that measuring Co ion concentration in physiological solution may underestimate the total cobalt release (Hedberg and Odnevall Wallinder, 2014). Chromium release: Chromium oxides are the main components of the passive film formed on the surface of CoCr alloys used in hip replacement systems. In static immersion tests, the amount of Cr ions released from the implants is extremely low compared with Co and Mo (Metikos-Hukovic et al., 2006). However, after long-term implantation of CoCr alloys, elevated levels of Co and Cr ions were observed in blood and urine of patients with metal hip replacements (Hanawa, 2004; Metikos-Hukovic et al., 2006; Lhotka et al., 2003). The release of chromium is believed to be associated with disruption of passive film during the mechanical wear.

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Molybdenum release: Pourbaix diagram shows that Mo species are more stable in passivity at low pH. A pH higher than 7.0 is favorable for Mo dissolution thermodynamically (Martin et al., 2013; Takeno, 2005). Moreover, molybdenum oxides formed in air on the metal surfaces disappear quickly after immersion in 0.15 M NaCl (Li et al., 1999). The interaction with proteins seems to have an important role in the dynamics of molybdenum release. Latest studies revealed that Mo played an important role in metal–protein interaction (Martin et al., 2013; Simoes et al., 2016). Electrochemical quartz crystal microbalance was used to study the mass change in different physiological solutions with proteins. Protein deposition was only observed on Mo surface rather than Co and Cr (Martin et al., 2013). Similarly, accelerated release of molybdenum ion was observed when bovine serum albumin was present in the solution (Simoes et al., 2016). It should be noted that active dissolution of metal element in the static passivation of CoCr is different from the in vivo corrosion process. First, in vivo active dissolution is often associated with mechanically assisted corrosion at interfaces. Mechanical abrasion could damage the Cr-rich passivation layer, reduce the pH, and accelerate the corrosion process (Gilbert et al., 1993; Cummings and Nordby, 1966). Patients with loosen implants have substantially higher release of cobalt and chromium from possible fretting corrosion at interfaces (Sunderman et al., 1989). Second, the presence of proteins in the physiological environment can affect the active corrosion of CoCr, via changes in the mechanics and kinetics of the corrosion reactions on the surfaces of the metal implants (Valero Vidal et al., 2010; Igual Munoz et al., 2015; Okazaki and Gotoh, 2005; Goldberg and Gilbert, 1997; Hallab et al., 2003). However, it has not yet been clarified whether biomolecules could accelerate or inhibit corrosion (Igual Munoz et al., 2015; Okazaki and Gotoh, 2005; Goldberg and Gilbert, 1997; Hallab et al., 2003). Serum proteins of  140 KDa preferentially bind to the CoCr alloy surface (Hallab et al., 2003) and this metal–protein interaction theoretically would enhance the dissolution rate (Steinemann, 1996). Additionally, when fretting occurs, the deposition of proteins on the CoCr alloy will form a barrier layer retarding Cr passivation and accelerating the active dissolution of metals (Goldberg and Gilbert, 1997). On the other hand, CoCr shows a thicker protective passive film when proteins are present, suggesting a higher resistance against metal release (Valero Vidal and Igual Muñoz, 2008). However, studies have shown that the difference in the amount of Co released from the CoCr casting alloy in immersion test was relatively small regardless of proteins (Okazaki and Gotoh, 2005). Moreover, electrochemical measurements of CoCr in human synovial fluid have revealed that the corrosion rates varied greatly depending on patients rather than the presence of proteins (Igual Munoz et al., 2015). Another possible factor affecting the active corrosion “in vivo” is the inflammation induced by the immunological reactions to the presence of metal implants. The OCP (open circuit potential) measured correlated well with the estimated inflammation of the synovial membrane (Igual Munoz et al., 2015). Cells such as macrophages generate reactive oxygen species as part of their mechanisms of response to a foreign agent, increasing the oxidative stress in the inflammatory environment (Hanawa, 2004). Recent studies suggested that the oxygen-derived species produced by macrophages had a role in the corrosion process of metal alloys (Gilbert et al., 2015; Liu and Gilbert, 2017).

Particulates release mechanisms The generation of particles is the result of the combined action of wear and corrosion. Mechanical abrasion disrupts material surface leading to the release of debris. Corrosion changes the nature of the debris. The migration and aggregation of debris accelerate the abrasion in return. Although many laboratory simulation tests have been conducted on the corrosion behavior of modular implants (Swaminathan and Gilbert, 2012; Blau et al., 2005; Roy et al., 2010; Sivakumar et al., 2011), the in vivo corrosion process has not been fully understood at present. In the following sections, the nature of corrosion debris from hip implants is reviewed to help us better understand the debris release mechanisms (Table 3). Fretting corrosion products at modular interfaces: Although corrosion has been widely observed at the interface of the modular components of the hip replacements (Gilbert et al., 1993; Collier et al., 1992; Cook et al., 1994), only a few material analyses have been reported on the solid corrosion products at the head–neck interface in total hip implants. An early case study described a large amount of black deposits at the head–neck interface, which were rich in Cr, Ca, and P and depleted in Co (Mathiesen et al., 1991). Further studies have demonstrated that regardless of the design, chemical differences have been observed on the corrosion Table 3

Summary of reported corrosion debris from CoCr hip implants

Locations

Corrosion debris and representative reports

Head–neck interface

Inside: Crystalline metal oxides containing Cr, Mo with depleted Co (Urban et al., 1994, 1997) At the opening: Amorphous chromium(III) phosphate hydrate (Urban et al., 1994, 1997) MoM implant: Nano-sized metallic particles (Catelas et al., 2004; Firkins et al., 2001) MoP implant: Mainly polyethylene debris (Illgen et al., 2008; Catelas et al., 2011; Doorn et al., 1996; Bitounis et al., 2016) Nano-sized crystalline chromium oxide (Bryant et al., 2013, 2014) MoM implants: Oxide particles containing Cr 3 þ (Catelas et al., 2004; Catelas et al., 2006; Goode et al., 2012); nano-sized metallic particles (Catelas et al., 2004; Bitounis et al., 2016; Catelas et al., 2006; Goode et al., 2012; Topolovec et al., 2013; Doorn et al., 1998); chromium(III) phosphate (Huber et al., 2009; Hart et al., 2010, 2012) MoP implants: Oxide particles containing Cr with low Co (Topolovec et al., 2013; Shahgaldi et al., 1995; Urban et al., 2000); sub-micron metallic particles (Topolovec et al., 2013; Shahgaldi et al., 1995; Urban et al., 2000)

Bearing surfaces Implant–cement interface Peri-implant tissues

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products depending on the location in the head–neck junction. Inside the internal part of the head–neck interface, the corrosion products were composed of highly crystalline metal oxides containing chromium and molybdenum with depleted cobalt (Urban et al., 1994, 1997), while at the opening of the modular junction, the corrosion products were composed of amorphous chromium(III) phosphate hydrate (Urban et al., 1994, 1997). The formation mechanisms of these products and their chemical differences are still unclear. Tribocorrosion products at articulating surfaces: Tribocorrosion is usually identified as wear tracks on the bearing surfaces (Kwon et al., 2010a). In MoM implants, nano-sized metallic particles and metal oxides have been observed after testing in hip simulator (Catelas et al., 2004; Firkins et al., 2001). However, in MoP implants, there are few reports on the particles released from metallic heads since the predominant wear debris are PE particles (Illgen et al., 2008; Catelas et al., 2011; Doorn et al., 1996; Bitounis et al., 2016). Higher surface roughness was measured on the SS femoral head in MoP implants compared to CoCr femoral head in MoM, suggesting a higher volume loss of SS due to its inferior wear resistance (Topolovec et al., 2014). Fretting corrosion at implant–cement interface: High acidity as a result of crevice corrosion is present at the interface of bone cement and Ti stem. The corrosion products have been characterized as mixed titanium oxide and hydroxide (Willert et al., 1996). On a retrieved study of 16 modular titanium alloy stems from cementless hip replacements, evident corrosion was characterized with fretting scars, pitting, and etching. Mixed titanium oxides were found in the corrosion deposits (Urban et al., 2006). A layer of nano-sized crystalline chromium oxide was found at the CoCr stem–cement interface (Bryant et al., 2013, 2014). Corrosion products in peri-implant tissues: In retrieved MoM implants, particles found were mainly oxide particles containing Cr (Catelas et al., 2004, 2006; Goode et al., 2012). Specifically, STEM-EELS revealed that majority of debris in tissues from MoM implant were nanoparticles of Cr 3 þ with trace amount of oxidized Co (Goode et al., 2012). In addition, metallic nanoparticles were also observed (Catelas et al., 2004, 2006; Bitounis et al., 2016; Goode et al., 2012; Topolovec et al., 2013; Doorn et al., 1998). The presence of metallic materials agrees with in vitro studies using hip simulator (Catelas et al., 2004; Firkins et al., 2001). The relative proportion of these two types of particles may be different (Topolovec et al., 2013; Doorn et al., 1998). A third type of corrosion particle rich in chromium and phosphate was also observed in femoral interface tissues from retrieved MoM implants of a specific model (Huber et al., 2009). Synchrotron radiation (XRF and XAS) analyses showed that particles in the tissues surrounding MoM implants were predominantly CrPO4 (Hart et al., 2010). Higher Co/Cr ratio was also found in tissues in this specific type of MoM hip implants compared to other models, implying cobalt ions to be responsible for ALTRs in this type of MoM implant (Hart et al., 2012). In retrieved MoP implants, chromium- and phosphate-rich particles were the primary debris detected in the tissues and these particles were associated with severe fretting corrosion between CoCr taper and Ti stem (Cooper et al., 2013; Hsu et al., 2012). Corrosion in MoP implants may release various types of particles into the tissue. CoCr metallic particles as well as chromiumrich particles with reduced cobalt were observed in tissue surrounding MoP implants suffering from aseptic loosening (Topolovec et al., 2013; Shahgaldi et al., 1995; Urban et al., 2000). The formation of these various particles in these loosened MoP implants is a result of abnormal wear at unintended metal interfaces, such as interface of loose metal components against bone, cement, or bearing surface due to wear through the polyethylene liner. Those processes will result in an extensive distribution of submicron metallic wear particles in periprosthetic tissues and abdominal lymph nodes (Urban et al., 2000). Possible debris releasing mechanisms: The nano-sized CoCr metal particles generated by tribocorrosion are probably a direct product of wear in the CoCr alloy. The presence of a layer of nanocrystals has been observed on the articulating surface of MoM implants (Büscher and Fischer, 2005; Sun et al., 2009; Wimmer et al., 2010), presumably as a result of recrystallization/straininduced phase transformation due to wear. This layer seems to be the releasing source of nano-sized metallic particles. On the contrary, the chromium oxide particles produced by fretting corrosion are not the direct products of fretting debris from passive films, as the nature of fretting corrosion particles is different from the passive film formed on the alloy surface (size, crystal structure). At stem–cement interface where fretting corrosion happens, a layer of nanocrystalline chromium oxide was found (Bryant et al., 2013, 2014), which is different from metallic layer formed on bearing surfaces. Latest few reports studied crosssectional tribofilms at different locations inside head–neck interface of retrieved MoP implants, revealing multiple layers with different crystal structures (Zeng et al., 2015a,b). Laboratory fretting corrosion test (ball on plate in tribometer) showed that the layer of nanocrystalline chromium oxide formed on CoCr surface was much thinner compared with the retrieved analysis (Zeng et al., 2015b). At this stage, the role of this chromium oxide nano-layer in the formation of fretting corrosion products is unclear. In theory, the formation of fretting corrosion particles is assisted by the aggressive crevice corrosion environment. The low pH and depleted oxygen will probably change the corrosion behavior at fretting interface and result in a different form of particles.

Biological Reactions to Corrosion Products As described previously, the corrosion process on implant surfaces releases products in the form of ions or particulate materials. These products have been associated with adverse reactions that affect the prognosis and general health of patients with metal implants. The interaction of these products with the tissues and cells, as well as the effects of these products on patient’s health, is discussed in this section.

Adverse local tissue reaction to metal implants Adverse local tissue reactions (ALTRs) are aseptic lesions that can arise in the periprosthetic tissue of patients with metallic hip prostheses. Despite their benign character, they can be locally aggressive causing the destruction of muscle and ligaments, and pressure

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effects on veins and nerves, affecting the prognosis of further clinical solutions (Daniel et al., 2012; Almousa et al., 2013). ALTRs are normally related to pain, swelling, and discomfort, but they can also be found in asymptomatic patients (Williams, 2011). The prevalence of ALTRs is unknown, but it is estimated that between 20% and 40% of patients with MoM implants and up to 5% of patients with MoP implants could develop ALTRs (Daniel et al., 2012; Williams et al., 2011; Feeley and Voreacos, 2013; Meier, 2013). ALTRs are the main reason of failure of metal-on-metal implants, and a significant number of MoP systems, estimating a failure rate of 9.7/ 1000 patient-years due to ALTRs for MoM systems, and < 1/1000 for MoP systems, with a peak of occurrence between 5 and 7 years after implantation (The NJR Editorial Board, 2016). Due to the high prevalence of these lesions in MoM articulation, it was speculated that the wear from the bearing surfaces was the etiologic factor that triggers ALTRs, but the presence of similar lesions in MoP systems suggests that corrosion on different implant surfaces, especially in the modular junctions, could be involved in the development of ALTRs (Cooper et al., 2013).

Histopathology of ALTRs Histologically, ALTRs have a layered structure consisting of three basic layers: an ulcerated synovial surface that may or may not be covered by fibrin deposition with sub-superficial necrosis, a second layer characterized by the infiltration of macrophages in a dense connective matrix, and externally an area with intense mononuclear infiltration that can vary from diffuse cellular infiltration to large perivascular aggregates of lymphocytes (Campbell et al., 2010; Mahendra et al., 2009; Natu et al., 2012). These characteristics vary with the severity of the lesions: lesions with limited necrosis usually present abundant macrophages and minimal lymphocytes. Moderate necrosis is associated with a mixed infiltration of macrophages and T lymphocytes, which are a mix of C4 þ and CD8 þ lymphocytes. In cases of extensive necrosis, the number of macrophages is reduced and lymphocytes are abundant forming highly organized tertiary lymphoid organs in the periphery of the tissues with highly endothelial veins, and T and B cells that can be organized in germinate centers (Mahendra et al., 2009; Mittal et al., 2013) (Fig. 5). The role of metals and their corrosion products in the etiology of ALTRs is supported by the facts that ALTRs have been observed in different implant designs (MoM THRs and resurfacing, and MoP), and associated with elevated concentrations of metals in serum and synovial fluid, and evident corrosion debris on implant surfaces and in the tissues (Kwon et al., 2014; Matharu et al., 2016). Based on previous histological description, it is evident that the pathogenesis of ALTRs involves the development of chronic inflammation (of which lymphocytes and macrophages are the main characteristics), with elevated concentration of proinflammatory cytokines (Kolatat et al., 2015; Eltit et al., 2017), which lead to a non-resolved inflammatory episode. A type IV hypersensitivity was proposed as the mechanism of activation of the immune system in ALTRs (Huber et al., 2009; Willert, 2005). However, this hypothesis was not supported in other reports (Granchi et al., 2012; Kwon et al., 2010b). The extensive necrosis of the capsule in ALTRs could be a result of the direct effect of Co and Cr ions in the tissues. “In vitro” studies have demonstrated that Co can induce apoptosis in different cell populations at a concentration of 6 mg/L, which is similar to the concentration of Co in synovial fluid in some patients with ALTRs, although the majority of them present much lower values (Catelas et al., 2006; Baskey et al., 2017; Granchi et al., 1998). Another possible mechanism is the degeneration and destruction of blood vessels of the affected tissue, generated by the prolonged inflammation (Fig. 6). This phenomenon would generate a hypoxic environment triggering the massive cell death observed in the ALTR tissues (Perino et al., 2014).

Roles of metal ions in tissues and fluids The free implant surfaces are exposed to joint fluid, a transudate composed of water and hydrophilic molecules such as hyaluronic acid, that provides the adequate lubrication to the sliding surfaces of the joint. The synovial fluid is contained in the joint capsule, a dense connective tissue envelope, which synthesizes the synovial fluid. In the inflammatory environment, the properties of this synovial fluid are expected to change, and therefore its lubricant capacity could be altered. The observed concentration of Co and Cr in synovial fluid of patients without implants on their joint is around 1 mg/L and 4 mg/L, respectively. The concentration of these metals can increase to a range between 50 and 100 mg/L in patients with implants in good conditions, and could reach between 500 and 10,000 mg/L in patients with failed implants (Eltit et al., 2017; Lass et al., 2014). The values of cobalt and chromium in blood have been widely studied in order to monitor the performance of hip implants. Blood has the advantage of being easy to obtain with low risk and minimum equipment. Its analysis is based on the rationale that the metals released from the metal implants will drain to the blood vessels in the surrounding tissues and will be found in circulation. The UK Medicines and Healthcare Regulatory Agency has issued a blood cobalt and chromium guidance value of 7 mg/L to identify MOM hip implant patients who may need further surveillance on excessive implant wear. Similarly, in a statement of the American Association of Hip and Knee Surgeons, the American Academy of Orthopedic Surgeons, and the Hip Society, these institutions recommend the systematic evaluation of patients with dual modular neck systems, to include the analysis of serum ion levels together with clinical and radiological evaluation to optimize the management of these patients (Kwon et al., 2014). Although the serum content of cobalt and chromium has been proposed as a monitoring tool for the performance of metal implants, it has been demonstrated that their analysis has only a modest sensitivity and specificity (60%) in identifying patients with ALTRs in MoM (Matharu et al., 2016; Malek et al., 2012), and no standards have been established for other types of bearings. Interestingly, the ratio of Co/Cr changes depending on the bearing surfaces. In patients with metal-on-metal hip resurfacings, in which there is a bearing surface subject to tribocorrosion but not a modular taper subject to fretting corrosion, the concentration of chromium in blood is slightly higher than cobalt (1.2–2.5-fold), in patients with metal-on-metal THRs, in which both events are present (metal-bearing surfaces and modular joint), the concentration of cobalt is higher than chromium (1.5–2.5 fold), while in the MoP THRs, in which there is no metal-to-metal-bearing surfaces, but there exist modular interfaces, the concentration of cobalt

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Fig. 5 Histology of ALTRs: (A) ALTR with low degree of necrosis, synovial epithelium is absent in the joint surface (arrows), but sub-superficial necrosis is localized within 100 mm. Abundant macrophages infiltrate (asterisks) and diffuse distribution of lymphocytes is seen (arrowheads). (B) ALTR with extensive necrosis, the synovial surface is covered by fibrin (arrows), and the sub-superficial necrosis is evident. Scarce macrophages are observed infiltrating the tissue (asterisks) and lymphocytes organize in densely packed structures (arrowheads).

is 2.5–5 times higher than chromium, revealing that the presence of cobalt in blood is mainly associated with the corrosion at the modular taper junctions of the implants, while the higher chromium concentration in blood is associated to surface wear in the articulations (Lavigne et al., 2011; Garbuz et al., 2010; Vendittoli et al., 2007). The mechanisms of cell damage by hexavalent chromium (Cr 6 þ) have been largely studied due to its association with cancer. Cr 6þ penetrates cells through sulfate channels and is intracellularly reduced to Cr 3 þ that combines with DNA and proteins, generating mutations and reducing the cell functionality. In peri-implant tissues, Cr 6þ is not observed, and Cr 3þ (the observed form) has no carcinogenic effect due to its difficulty to penetrate the cell membrane. However, in vitro studies have demonstrated that it can induce cell death in high concentrations due to the induction of hypoxia in the cells exposed to Co (Vengellur and LaPres, 2004; Nyga et al., 2015; Samelko et al., 2013). There is no higher risk of cancer in patients with Co/Cr alloy implants (Smith et al., 2012), and Cr 3 þ is not considered a health hazard (Eastmond et al., 2008). Cobalt ions have the capability of penetrating cells by using different ionic channels in the cell membrane. It is an important cofactor of vitamin B12 (cobalamin), which has an important role in the metabolism of the nervous systems and the formation of red blood cells. However, the excess of cobalt in cells has deleterious effects through affecting cell metabolism and DNA replication and repair machinery, due to the binding of cobalt ions to specific proteins and affecting their functionality and consequently the cell metabolism as it has been demonstrated through affinity chromatography (Kwon et al., 2009; Scharf et al., 2014).

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Fig. 6 Vascular changes in ALTRs: (A) initial changes in the blood vessels: intense lymphocyte infiltration of the vessel walls (arrows); however, the lumen has normal appearance (asterisk). (B) The endothelium presents a cubic shape (arrows) (should be squamous), and the vessel’s lumen is observed compressed with an amorphous content (asterisk).

Biological effect of metal particles As was described previously, the particles observed in the tissues correspond mainly to wear particles, histologically observed as dark spots of nano-micron size that tend to be accumulated by the process of phagocytosis carried out by the macrophages, and corrosion particles that are histologically observed as large (tens to hundreds of microns) translucent structures under the light microscope (Ricciardi et al., 2016) (Fig. 7). The negative effects of particles in the tissues could be due to the presence of particles itself, or to the leaching of ions from the particle to the surrounding tissues. It has been demonstrated that under exposure of similar amount of particulate material, cells are more susceptible to damage under nano-sized particles than micron-size particles, presumably due to the increased surface area for metal leaching from the smaller particles (Papageorgiou et al., 2007). The particles as a foreign structure can trigger a foreign body response, in which cells capable of performing phagocytosis (mostly macrophages) are recruited in the zone aiming to destroy the foreign agent. In the tissues, the macrophages are usually seen loaded with small metal particles contained in vesicles in their cytoplasm (Figs. 7 and 8), and “in vitro” studies have shown that CoCr particles activate macrophages increasing the secretion of cytokines (Posada et al., 2015). Due to their incapability of degrading these particles, they synthesize a higher amount of enzymes and oxidative molecules (peroxides and superoxides), generating an elevated oxidative stress that led to cell damage, and accelerating the ion leaching from the metal particles, especially of cobalt, thus increasing the local concentration of this ion at the time that a Cr-rich core remains insoluble in the cytoplasm of the macrophages (Goode et al., 2012; Xia et al., 2011). Large particles of metal oxides, on the other hand, are much bigger than a single macrophage. To be able to phagocyte them, multiple macrophages fuse, generating giant multinucleated cells, where the multi-nuclei correspond to various individual nuclei of the fused macrophages (Fig. 7). Although above arguments can explain the final effects of the development of ALTRs, the relationship between metal particles and ALTRs has not been proved. In vitro studies have tried to link the development of immune response to metal particles of CoCr alloy, demonstrating by cell culture that they can induce phagocytosis in monocyte-derived cells, and different levels of cell damage

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Fig. 7 Phagocytosis of corrosion products in ATLRs. (A) Phagocytosis of small metallic particles by macrophages. The metallic particles can be seen in the cytoplasm of macrophages (arrows). (B) Large corrosion products (asterisks) induce the formation of giant multinucleated cells (arrows). The giant multinucleated cells are the resultant of the fusion of macrophages, and perform the phagocytosis of large particles as seen in the image.

Fig. 8

Metal particles in intracellular vesicles. TEM of macrophages with metal particles in endosomic compartments (Scharf et al., 2014).

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in different cell populations (Samelko et al., 2013; Papageorgiou et al., 2007); however, histopathology analysis has failed in demonstrating association between the amount of particles and the development and/or severity of ALTRs (Ebramzadeh et al., 2014).

New Biomaterial Designs Although Ti alloys and CoCr alloys discussed in this article have been successfully used in hip implants, there are still concerns over their clinic performance against corrosion, specially wear-assisted corrosion. Hence, efforts are being made to improve their mechanical properties and corrosion resistance. Ti alloys: The most commonly used Ti alloy in orthopedic implants has been Ti–6Al–4V (Rack and Qazi, 2006). Ti–6Al–4V is classified as a a þ b alloy for its microstructure, in which Al acts as an a stabilizer while V as a b phase stabilizer. With increasing concerns on the biological response of potentially toxic vanadium (Maurer et al., 1994), niobium and iron are used to substitute vanadium, leading to the development and introduction of Ti–6Al–7Nb and Ti–5Al–2.5Fe alloys (Semlitsch et al., 1985, 1992; Niinomi, 1998). And the replacement doesn’t compromise the alloy’s electrochemical and corrosion behavior in SPS (Choubey et al., 2004). Both show excellent biocompatibility and similar mechanical properties to Ti–6Al–4V. Another concern of the Ti a þ b alloy is its high modulus (110 GPa) compared to bone (20–30 GPa). This creates stress shielding effect due to high stiffness of Ti alloy, leading to bone resorption. Hence, metastable b titanium alloys such as Ti–15Mo (Zardiackas et al., 1996), Ti–12Mo– 6Zr–2Fe (TMZFTM), Ti–15Mo–3Nb–0.3O (21SRx) (Wang, 1996), Ti–13Nb–13Zr (Davidson et al., 1994) and Ti–Nb–Zr–Ta alloys (TNZT) (Nag et al., 2005) have been developed to lower the elastic modulus. At the same time, these b titanium alloys exclude Al, which eliminates the possible risks of Al-associated toxicity. The most challenging application of Ti alloys is their use at interfaces involving wear. Ti alloys have poor wear resistance and the friction at the interface leads to a faster degradation of the alloy, which limit their application in articulating surfaces. Attempts have been made on the surface modification of Ti alloys in order to improve their wear resistance. Different hard coatings such as TiN, ZrN, and amorphous carbon have been applied and tested on Ti alloy surfaces (Hendry and Pilliar, 2001). However, retrieved analyses revealed the presence of wear debris and delaminated asperities of TiN-coated Ti–6Al–4V articulating surface, pointing to a weak coating-substrate interface (Lass et al., 2014). To improve the interfacial strength, new processing technologies such as PIRAC (Powder Immersion Reaction Assisted Coating) technology were used to reduce the residual stresses at the interface (Shenhar et al., 2000). Studies have shown that TiN coating prepared with PIRAC was still intact and the wear rate of the polyethylene cups was significantly lower after up to 4 million cycles wear test (Gutmanas and Gotman, 2004). Another way to improve the interface strength is to make an interface with gradual change in composition. Up to 86% CoCr alloy could be obtained on the surface of Ti–6Al–4V alloy and this microstructure increased surface hardness and reduced the wear rated in wear test (Vamsi Krishna et al., 2008; Nychka et al., 2011). CoCr alloys: CoCr alloys are mainly used at the articulating surfaces because of their high wear resistance. However, there are concerns over ALTRs associated with metal release from the corrosion of CoCr alloys. To improve biocompatibility, toxic elements such as Ni used in CoCr alloy were reduced (Sonofuchi et al., 2016). Another attempt is to apply a hard and biocompatible coating on the surface of CoCr alloy to prevent toxic elements from releasing and to retain high wear resistance at the same time. Different engineered CrN, TiN, DLC (diamond like carbon), and tantalum carbide coatings and their combinations on CoCr alloys have been studied (Ludema et al., 2011; Türkan et al., 2006; Pham et al., 2011; Balagna et al., 2012; Liu et al., 2013). Coating delamination remains to be a key issue to be solved (Ludema et al., 2011; Liu et al., 2013). An alternative to CoCr alloy used in bearing surface is ceramics such as alumina and zirconia. Although ceramics have superior wear resistance and no risk of toxic ion release, their brittle nature makes them vulnerable to fracture (Semlitsch and Willert, 1997; Lerouge et al., 1997; Fritsch and Gleitz, 1996). A promising design is a composite material combining advantages of the superior wear resistance in ceramics with high toughness in metals. One example is Oxinium (Smith & Nephew, Memphis, Tennessee), a zirconium–niobium alloy with around 5 mm oxidized layer formed on the surface by thermal diffusion (Sonntag et al., 2012). The metal substrate is tough, and the hard oxide surface has similar wear resistance to ceramics. Oxinium has lower wear rate than CoCr alloy with similar fracture toughness (Good et al., 2003; Bourne et al., 2005). Another example is the zirconia reinforced with biocompatible metals such as Nb and Ta (Bartolomé et al., 2016; Smirnov et al., 2017). This ceramic–metal composite exhibits excellent wear resistance, toughness, and biocompatibility (Bartolomé et al., 2007, 2016; Rodriguez-Suarez et al., 2012; GutiérrezGonzález et al., 2012).

Future Perspectives Corrosion-induced mechanical and biological failures affect the clinical performance of current orthopedic implants, but also present opportunities for future biomaterial development. The most commonly used metallic alloys used in implants today were developed more than half century ago initially for other engineering applications. Alloys were not specifically designed to take biocompatibility into consideration. However, such a situation is about to change. Today’s biomedical implants are sitting at the interface between two rapidly growing fields: biology and materials science (Ratner et al., 2004; Williams, 2014). Progress in biology and medicine will provide a clear picture on the structure and functions of various tissues down to cellular and molecular levels. In-depth understanding of tissue-biomaterial interaction will guide the design of new alloys and surfaces. This is further

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strengthened by promising development in materials science and engineering, as represented by the Materials Genome Initiative launched by the US Government in 2011 (National Science and Technology Council, 2014). Past breakthroughs in material modeling, theory, and data mining make it possible to significantly accelerate the discovery and development of advanced materials including new biomaterials.

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Decellularized Extracellular Matrix Paul Frank Gratzer, School of Biomedical Engineering, Dalhousie University, Halifax, NS, Canada © 2019 Elsevier Inc. All rights reserved.

Introduction Decellularization Techniques Physical Chemical Inhibition of Enzymes Example Decellularization Processes Assessing Decellularization Sterilization Issues Example Decellularization Applications The Future of Decellularization Further Reading

86 87 87 87 89 90 91 93 93 96 96

Glossary Allograft Donor tissue or organ transplanted from one human to another. Critical micelle concentration The concentration at which detergents form a spherical cluster. Decellularization Techniques or processes whereby tissues or organs are treated to remove cellular materials while preserving the non-living, structural scaffold. Monoclonal antibody An antibody produced by a single clone of cells or cell line and consisting of identical antibody molecules. Xenograft Donor tissue or organ transplanted from an animal to a human.

Abbreviations CMC Critical micelle concentration DAPI 40,6-Diamidino-2-phenylindole DSC Differential scanning calorimetry ECM Extracellular matrix EO Ethylene oxide FITC Fluorescein GAG Glycosaminoglycan GI Gastrointestinal H&E Hematoxylin and eosin HIT Hydrothermal isometric temperature testing HLA Human leukocyte antigens IPSC Induced pluripotent stem cells MHC Major histocompatibility complexes PMSF Phenylmethanesulfonylfluoride SDS Sodium dodecyl sulphate SIS Small intestinal submucosa TnBP Tri-butyl phosphate UBM Urinary bladder matrix

Introduction The term “decellularization” is used to describe techniques or processes whereby tissues or organs obtained from human (allograft) or animal (xenograft) donors are treated to remove cellular materials while preserving the non-living, structural scaffolddreferred

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to as the extracellular matrix (ECM). By doing so, the main source of rejection of the donor materialdthe cellsdare removed. Further, by maintaining the ECM of the tissue or organ, the remaining non-living donor material can be used as a scaffold to regenerate living, functional replacements that can be used to treat patients with diseased, dysfunctional, or damaged parts of their body without being rejected. The idea of decellularization has been around since 1975, however, it has seen increased attention and research activity within the last 20 years as the areas of Tissue Engineering and Regenerative Medicine look to develop living replacements for parts of the human body.

Decellularization Techniques The main objectives of decellularization techniques are to remove as much donor cellular materials (genetic, cell membrane, cytoskeletal proteins) as possible while at the same time keeping the native structure and composition of the ECM as intact as possible. These objectives, in theory, seem relatively simple and straight forward, however the implementation of techniques to achieve the goals of decellularization are far more difficult. The goal of decellularization is akin to removing the white and yolk (i.e., cellular materials) from an egg without damaging or disrupting the egg shell (i.e., the ECM). Decellularization techniques, as described below and summarized in Table 1, are combined in various configurations to yield decellularization processes that are applied to donated tissues and organs. Almost every decellularization process that has been developed by researchers to date involves at least two basic techniques: (i) a lysis step to burst and break up cells and (ii) washing step(s) to remove disrupted and exposed cellular materials from within the ECM.

Physical At the start of most decellularization procedures, cells are lysed physically using osmotic gradients, mechanical compression/ massage, or freeze-thaw cycles. Hypertonic (high salt concentration) and hypotonic (low salt concentration) treatments put hydrostatic pressure on cell membranes in an aqueous environment causing them to burst and release their cell contents. The contents of the cell are then more accessible to later treatments with detergents or isotonic (balanced salt) washout procedures. Mechanical compression or massage is used to encourage membrane degradation and the gradual exposure of more cell membranes to extraction solutions. Freeze-thaw cycles are used to kill cells and then fracture their cell membranes so that subsequent washout procedures can access internal cell contents and fragmented membranes. Vacuum techniques, use of ultrasound (sonication), and increased pressure, for example use of supercritical liquid CO2, can also be used to lyse cells or aid in the removal of cellular components. Care must be taken with cell lysis methods, however, in order to avoid damaging the ECM. This is especially true for mechanical and freeze/thaw techniques.

Chemical Critical to the success of many decellularization strategies are the use of detergents or surfactants. Detergents are a special group of phospho-lipid compounds that are soluble in aqueous environments. Detergents are amphiphiles in that they have both a polar and a non-polar domain. They are useful for decellularization when they are added in a sufficient concentration to form micelles. A micelle is a cluster of detergent monomers, often spherical, that is oriented so that the non-polar domains of the detergent molecules are interacting internally, and the polar domains are interacting with water molecules externally. The concentration at which detergents form micelles is called the critical micelle concentration (CMC). The CMC varies with decellularization conditions, including ionic strength, pH, temperature and the presence of protein and lipids (including other detergent molecules). Micelles occupy space in the aqueous environment, and as such, must disrupt hydrogen bonds between water molecules. The interaction of polar head groups of the detergent monomers in the micelle is usually more than sufficient to satisfy the energy requirements for keeping water molecules apart, thus maintaining the micelle as an intact stable structure in the fluid. The “protected” non-polar tails of the detergent molecules in the center of the micelle essentially provide a second phase where nonpolar lipid and non-polar protein domains can “dissolve.” During decellularization, the extraction of cell components, including lipids and proteins from tissue, is accomplished when these micelles enter the tissue, dissolve non-polar components, and are washed out with solution changes. Detergents can be classified by one of three designations: ionic, nonionic, and zwitterionic. Ionic detergents are either anionic or cationic, although cationic detergents are not used in decellularization processes due to their strong protein denaturing tendencies. Anionic detergents contain a negatively charged (polar) head group. Anionic detergents, such as sodium dodecyl sulphate (SDS), are highly efficient at solubilizing proteins and are generally denaturing to some extent. A subgroup of ionic detergents are the bile acid salts, found in the intestine where they solubilize fats. Bile acid salts, such as sodium deoxycholate, have a rigid steroidal backbone and lack a defined polar head group. Instead, they have their polar residues arranged all on one side of the rigid steroid backbone and form kidney-shaped micelles, unlike the spherical micelles formed by detergents with linear non-polar backbones. Bile acids are milder than other anionic detergents such as SDS. Nonionic detergents, such as Triton X-100, have neutral polar head groups and are nondenaturing to proteins. Nonionic detergents break lipid–lipid and lipid–protein interactions. Finally, zwitterionic detergents, such as CHAPS, have properties of both ionic and nonionic detergents. Zwitterionic detergents are generally milder than ionic detergents and more denaturing to proteins than nonionic detergents.

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Table 1

Decellularization techniques and reagents

Physical methods Osmotic gradient Freeze/thaw Mechanical delamination Agitation/compression/vacuum Chemical methods Anionic detergents Bile acid salt Sodium deoxycholate Synthetic SDS (sodium dodecyl sulfate)

Mode of action

Potential drawbacks

Bursts or contracts cells, disrupts cell membranes Disrupts cell membranes Separates tissue layers along natural planes of dissection Increases exposure of cell membranes to extraction solutions

Lyses cells, releasing proteases that can attack ECM Ice crystals can destroy ECM (flash freezing) Excessive forces can damage ECM

Solubilizes phospholipids Disrupts protein–lipid interactions Generally non-denaturing, mild

Precipitates at pH 6.9 to form bile acid Potential growth factor removal

Disrupts protein–protein interactions Denatures and solubilizes proteins Binds virus particles (incl. HIV)

May denature collagen and elastin in ECM Removes GAGs, growth factors May contribute to swelling Reagent is cytotoxic May need to use concurrently with zwitterionic detergents for highest effectiveness, potential growth factor removal May leave cell remnants, potential growth factor removal

Triton X-200 (alkylaryl polyether sulfonate)

Solubilizes cellular components

N-lauroyl-sarcosinate

Ruptures cells Solubilizes cell membrane proteins

Nonionic detergents Triton X-100 (PEG tert-octylphenyl ether)

Non-denaturing protein solubilization

Tween 20 or 80 (PEG-sorbitan monolaurate/ oleate)

Solubilizes peripheral membrane proteins

Igepal CA630 (formerly sold as Nonidet P40) (Octylphenyl-PEG)

Non-denaturing protein solubilization

N-octyl-b-D glucopyranoside Zwitterionic detergents Sulfobetain-10 and -16 CHAPS Alcohols Glycerol

Solubilization of membrane proteins Help decrease micelle size with anionic detergents Retain zwitterionic nature over wide pH range Disrupts protein–protein interactions Non-denaturing, forms small micelles

Ethanol

Destroys bacteria Solubilizes cell components Disrupts cell membranes, dissolves lipids Destroys bacteria and viruses Destroys bacteria

1-Butanol

Extracts lipids

Isopropanol

Acids Peracetic acid Bases NaOH NH4OH Chelators

ECM can be damaged by sonication or excessive forces

Low removal efficiency in some tissues May disrupt ECM organization, potential growth factor removal Low efficiency in isolation Tween 20 less effective than Tween 80 Potential growth factor removal Less hydrophilic than Triton X-100 May sediment during storage Potential growth factor removal Requires cold storage Potential growth factor removal Hygroscopic Potential growth factor removal High critical micelle concentration (6–10 mM, depending on temperature) Potential growth factor removal

Volatile, ECM mechanical property changes Volatile, causes matrix shrinkage, change in mechanical properties Severely contracts ECM, change in mechanical properties

Solubilizes cytoplasm components Disrupts DNA

Can denature ECM components and growth factors (especially above 0.1% v/v)

Disrupt DNA/RNA, cell membranes Destroy viruses, inactivates prions Weakens lipid–protein bonds

Loosens, frays, and fragments elastin fibers High vapor pressure

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Decellularization techniques and reagentsdcont'd

EDTA EGTA Enzymatic methods Membrane/attachment enzymes Trypsin

DISPASE DISPASE II Phospholipase Antigen-targeted enzymes Thermolysin a-galactosidase Exonucleases DNAse I RNAse A Endonucleases Benzonase

Mode of action

Potential drawbacks

Bind to metallic cofactors, inhibiting enzymes Bind ions cells need to attach to substrates

Can inhibit enzymes used in decellularization Processes

Disrupts desmosomes, focal adhesions Cleaves peptide bonds on Arg and Lys Attacks type IV collagen, helps lift cells

Does not remove all active MMPs Removes fibronectin, laminin, elastin, GAGs, growth factors Can damage ECM after prolonged exposure Can completely disrupt basement membranes

Digests phospholipids in cell membranes

Requires special conditions (pH, temperature)

Attacks antigen in the hemidesmosome of the basal layer of keratinocytes Digests a-1,3-galactose antigen

Requires special conditions (pH, temperature)

Facilitates hydrolysis of terminal DNA strands Facilitates hydrolysis of terminal RNA strands Facilitates degradation of internal bonds in DNA (single and double stranded) and RNA

Sensitive to mechanical denaturation Presence of EDTA inhibits activity May be degraded by trypsin or thermolysin Presence of EDTA inhibits activity

Solvents, such as tributyl phosphate (TnBP), can be useful in decellularization as they are capable of disrupting protein and lipid interactions by destabilizing hydrophobic interactions. An added benefit to the use of TnBP in a decellularization protocol is its proven anti-viral action. While not widely used in current decellularization protocols, other solvents have also shown good decellularization performance with very little damage to the ECM. Alkaline bases have been used to denature DNA. Commonly used alkaline bases include ammonium hydroxide, sodium sulfide, sodium hydroxide, and calcium hydroxide. Such compounds have been used in the decellularization of dense tissues such as dermis, but alkalines can degrade structural component of the ECM including collagen. Alkaline treatment results in a dramatic reduction in GAG content and altered mechanical properties of tissues. Similarly, acids are used to dissociate DNA from the ECM by solubilizing cytoplasmic components and disrupting nucleic acids. Acids can also denature ECM proteins including GAGs, collagen, and growth factors. It is important to optimize the dose and exposure time when acids are used for decellularization. Peracetic acid, applied at 0.1% (v/v) in a single wash for 2 h, and combined with appropriate mechanical methods and rinsing, can thoroughly decellularize thin tissues such as small intestinal submucosa (SIS) and urinary bladder matrix (UBM). Acids commonly used for decellularization include deoxycholic acid and acetic acid. However, acetic acid has been shown to cause damage to and removal of collagens from the ECM. Alcohols can be effective in decellularization processes if cell membranes are permeabilized. If the polar hydroxyl groups of alcohols can diffuse into the cell, the alcohols replace intracellular water and lyse the cell by dehydration. Ethanol or methanol may also be used as a final wash to remove residual nucleic acids from tissue. The non-polar carbon chain of alcohols can also help to dissolve non-polar substances such as lipids. Care must be taken, however, at the use of alcohols has been shown to significantly alter the mechanical properties of the ECM.

Inhibition of Enzymes During decellularization, the lysis of cells releases the contents of lysosomes into the extracellular space. Many of the enzymes released, especially proteases, are highly effective at degrading the ECM. The goal of decellularization is to preserve the intact ECM while removing cellular components, so the action of degradative enzymes is counterproductive. Most decellularization procedures make use of protease inhibitors at least during the initial cell lysis phase of the procedure while some continue the use of protease inhibition throughout. Protease inhibition is often achieved through a combination of methods. Many procedures inhibit lysosomal proteases with elevated pH ( pH 8). Chelating agents, such as EDTA or 1,10-phenanthroline, that bind metallic enzyme cofactors such as magnesium, iron, or zinc are also used. Chemical inhibition of proteases can be achieved by the addition of any one of a number of common protease inhibitors. Chemical inhibitors include phenylmethanesulfonylfluoride (PMSF), aprotinin, or leupeptin. PMSF (irreversible binding) and aprotinin (reversible binding) inhibit serine proteases such as trypsin and chymotrypsin, whereas leupeptin (low stability in aqueous environments) inhibits both serine and cysteine proteases. The inhibition of bacteria is also essential, as most tissues processed with decellularization are not procured under aseptic conditions. Common

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antibiotics used for decellularization treatments include gram positive bacteria cell wall synthesis inhibitors (penicillin or vancomycin), bacteria and mycoplasma protein synthesis inhibitors (streptomycin or gentamicin) and agents that induce ribosomal miscoding in gram positive/negative bacteria and mycoplasma (kanamycin). Sodium azide (NaN3) is also often added to decellularization solutions to inhibit microbial growth.

Example Decellularization Processes At first glance, there appears to be a bewildering array of different decellularization processes presented in the scientific literature. Fortunately, there are only a few truly unique decellularization processes with others being variations of these basic methods. Most decellularization processes can be classified as a (i) physically-based process, (ii) chemically-based process, or (iii) biochemicallybased process. There are many protocols that borrow from these three basic processes and include combinations of various decelluarlization techniques outlined above. An example of a biochemically-based decellularization protocol involves the use of enzyme(s) such as Trypsin. Trypsin is commonly used in laboratories throughout the world to release cells from substrates during routine cell culture. It is also used in enzymatic assays studying the “nativity” or how intact collagen is within tissues, as it will preferentially degrade only denatured collagen. It has been used in decellularization protocols as it digests desmosome complexes holding cells together and disrupts cell adhesions that link cells to the ECM. A representative protocol using trypsin for decellularization of tissues is shown in Table 2A. In general, studies using trypsin for cell extraction have shown poor extraction efficiency and significant ECM damage when compared with detergent processes. Mechanical delamination can be used to accomplish decellularization. Tissues from the gastrointestinal (GI) tract and the urinary system have been treated by this process. Tissues from the GI tract and the bladder have the same basic organization from deep to superficial of mucosa, submucosa, muscularis, and serosa. In most strategies using GI tissues, the tubular tissue is opened longitudinally and the mucosa delaminated. The remaining serosa and muscularis are then also delaminated, leaving Table 2

Example decellularizaation protocols Treatment

(A) Representative trypsin-based decellularization protocol 1 0.05%–0.5% trypsin digest (optional repeat) 2 DNAse/RNAse digest (optional) 3 PBS washout and storage

Conditions

Duration

pH 7.4, 0.02%–0.2% EDTA HBSS, pH 7.4, 37 C 4 C, antibiotics

24–96 h 2h Discretionary

Treatment

Conditions

(B) Representative mechanical delamination decellularization protocol 1 Bladder: osmotic gradient removes urothelial cells 2 Delamination: Bladder: delaminate serosa, muscularis, and submucosa (three superficial layers) GI tract: split longitudinally, delaminate mucosa, turn over GI tract: delaminate muscularis and serosa (two superficial layers) 3 Cell lysis/disinfection of Bladder: urinary bladder matrix, UBM (mucosa and lamina propria remain) GI tract: small intestinal submucosa, SIS (submucosa and lamina propria remain) 4 Phosphate buffered saline rinse 5 Distilled water rinse Treatment

1 N NaCl

0.1% peracetic acid 4% ethanol Sterile water

pH 7.4, 2  15 min 2  15 min

Conditions

(C) Representative Triton/SDS detergent decellularization protocol 1 Hypotonic tris buffer pH 8, 4 C-room temp., 5 mM EDTA, PMSF, antibiotics, agitation 2 Hypertonic tris buffer þ nonionic detergent: pH 8, 4 C-room temp., 1.5 M KCl, PMSF, antibiotics, 1% (v/v) Triton X-100 agitation 3 DNAse/RNAse digest Hank’s Balance Salt Solution (HBSS), pH 7.4, 37 C 4 Anionic or nonionic detergent: pH 8, antibiotics, agitation 1% (v/v) SDS or Triton X-100 in tris buffer 5 Washout: ethanol or HBSS Antibiotics, agitation 6 PBS washout (optional) and storage 4 C, antibiotics

Duration 24–36 h 24–48 h 1–5 h 24–48 h 48 h Discretionary

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the submucosal layer and the basement membrane of the mucosa. In the case of the bladder, the submucosa, muscularis, and serosa are often microdissected away from the mucosa with the help of mechanical delamination. The few cells that remain in the submucosa/lamina propria (small intestinal submucosa, SIS) or the mucosa/lamina propria (urinary bladder matrix, UBM) are lysed and removed in a washout procedure and peracetic acid treatments. A representative protocol useful for processing intestine or bladder tissue is presented in Table 2B. Such methods, however, lead invariably to the damage of the ultrastructure and the integrity of the ECM scaffold. Due to the thin sheet-like nature of SIS materials, specialized load-bearing tissues such as heart valves and ligaments may not be the ideal application for this class of material. It has recently been reported, however, that the mechanical strength of SIS can be improved by layering of the material. SIS and UBM are ideally suited to patching applications such an in the urinary bladder and general tissue augmentation. The most widely used decellularization protocols are based on synthetic detergents that solubilize proteins. The anionic detergent SDS and the non-ionic detergent Triton X-100 are the most widely used synthetic detergents for decellularization. A representative protocol is given in Table 2C. This decellularization concept has been widely used and is one of the most studied to date in terms of its effects on the properties of the ECM. The effectiveness of cellular extraction with Triton-SDS procedures is generally excellent, although this treatment is not as mild as that obtained with sodium deoxycholate. There are also numerous other examples of the combination or replacement of Triton X-100 and/or SDS with other detergents such as CHAPS, Tween 20/80, sodium deoxycholate, trypsin and the organic solvent tri-n-butyl phosphate (TnBP). Many of these protocols are reported to achieve high levels of cellular extraction with varying levels of associated changes in the properties and composition of the ECM. No one decelllarization process can be used as a universal treatment for all tissues or organs. Generally, each tissue or organ is unique in terms of the conditions and techniques used to successfully decellularize them. While successful decellularization treatment processes have been found for specific applications, these have largely been determined by empirical knowledge and experience.

Assessing Decellularization The most important step in assessing the effectiveness of a new decellularization process or the application of a decellularization process to a new tissue or organ source requires the use of multiple characterization assays and techniques. There have not been any regulations or standards developed or accepted for describing what constitutes minimal or even optimal requirements to ensure that any tissue or organ processed with decellularization will be safe and effective as a replacement for a part of the human body. That being said, over the last 20–25 years, research has shown that there are areas of measurement to assess the effects of decellularization on tissues and organs that directly affect performance when implanted into a recipient. These measurement criteria for decellularized tissues and organs should include: 1. the lack of visible nuclear material in a histological section stained with hematoxylin and eosin (H&E) or 40,6-diamidino-2phenylindole (DAPI) 2. < 50 ng ds DNA per mg of dry weight of ECM 3. < 200 bp DNA fragment length 4. the lack of intracellular components (e.g., cytoskeletal materials, cell organelles) 5. the lack of cell membrane elements, specifically known immunogenic ones such as major histocompatibility complexes (MHC) or human leukocyte antigens (HLA) 6. the maintenance of ECM elements in their native state and quantities (collagen, glycosaminoglycans (GAGs), elastin, etc.) 7. the resulting ECM scaffold does not contain nor release cytotoxic elements The removal of cellular nuclear elements in decellularized materials is important as DNA has been shown to elicit immune responses and promote calcification in certain applications. At minimum, routine histologic staining can be used as a rough detection of DNA/RNA. Histologic stains, such as H&E or trichrome, provide a relatively insensitive measure for identifying DNA within the ECM (Fig. 1). A second more accurate and sensitive quantifiable method of DNA detection can be achieved using commercially available dsDNA intercalators, such as PicoGreen, propidium iodide, or bisbenzimide, as well as by gel electrophoresis. As shown in Fig. 2, DNA content of tissues treated with different decellularization processes can vary dramatically. Not only is the removal of cellular genetic materials important, but the removal of intracellular and cytoplasmic debris, such as cytoskeletal proteins a- and b-actins and vimentin, are important as they may be involved in adverse host reactions such as increased inflammation. Further, cytoskeletal protein removal has been shown as a good indicator of a high level of decellularization having been achieved. Most essential to the reduction of immunogenicity in allo- and xenograft materials via decellularization, however, is the complete removal of cell membrane elements (i.e., MHC or HLA). Both MHC and HLA cytoplasmic proteins have been shown to be directly responsible for adverse immune and inflammatory reactions in transplanted donor materials. Methods best suited to detect immunogenic cytoplasmic proteins in decellularized materials involve the use of antibodies that recognize them and chromogenic or immunofluorescent markers to detect the antibodies such as peroxidase or fluorescein (FITC) based staining (Fig. 3). Equally as important as removing immunogenic donor cellular materials, is the maintenance of the composition and structure of the ECM of decellularized tissues and organs. Untreated tissues and organs are used as the comparative control to determine if decellularization has affected the ECM. As stated previously, certain decellularization techniques can have particular detrimental effects on ECM elements such as collagen, GAGs, and elastin that could remove or damaged them during processing. Maintaining

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Fig. 1 Pig bone-anterior cruciate ligament (ACL)-bone graft before (left) and after (right) decellularization processing. Note loss of red/pink coloration (above right) as cells and blood products are removed during decellularization resulting in a whitish color (above left). The removal of cells is also indicated by the loss of blue/purple stained cell nuclei in histological sections of untreated tissue (below left) when compared to decellularize tissue (below right) using the stain hematoxylin and eosin.

Fig. 2 DNA content for commercially available and laboratory produced ECM scaffold materials as determined with the PicoGreen assay. (Color version of figure is available online.) From Journal of Surgical Research 152, 135–139 (2009), “Quantification of DNA in Biologic Scaffold Materials”, Thomas W. Gilbert, Ph.D., John M. Freund, B.S., and Stephen F. Badylak, D.V.M., Ph.D., M.D.

the native composition and structure of the original donated tissue or organ is key to the decellularized ECM interacting with and guiding cells, either in vitro or in vivo, to repopulate the ECM matrix and reform living functional tissues. The role of the ECM includes providing a repository for growth factors via binding to matrix molecules, cell instructive cues via cell integrins, and mechanical and physical properties of tissues The assessment of the composition and degree of intact structure of tissues and organs should include biochemical, immunohistochemical, structural, and mechanical property measurements where appropriate. Biochemical assays may include hydroxy-proline content for collagen quantitation, nin-hydrin elastin detection, or di-methylene blue assay for GAGs. Monoclonal antibodies may also be used to assess the quantity and location of specified ECM elements in the processed ECM matrix. For the assessment of possible changes to the native structure of ECM components, methods that detect macro- and micro-molecular changes can be used. For example, conducting mechanical tests, transmission electron microscopy (TEM), scanning electron microscopy (SEM), or histological microscopy using stains such as Masson’s trichrome or Van Geisen elastin stain can provide information on the macromolecular structure of ECM components.

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Fresh

DermGEN™

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GraftJacket™

HLA-DR

HLA-A,B,C

Fig. 3 Immunohistochemical staining (brown peroxidase) for the presence of two main cellular immunogenic proteins responsible for allograft skin rejection and increased inflammation: HLA-DR and HLA-A,B,C. Note significant staining for both fresh (unprocessed) human skin and Graftjacket™ (Acelity) Decellularized Human Dermal Matrix. In contrast, DermGEN™ (DeCell Technologies Inc.) decellularized human dermal matrix shows complete removal of all immunogenic cell components. 400 magnification.

For micromolecular structural information, techniques such as thermal denaturation assays for collagen stability (e.g., hydrothermal isometric testing (HIT) or differential scanning calorimetry (DSC)) can be used. Further, more sensitive measures of collagen structural changes can be detected using enzyme-based assays (e.g., trypsin or chymotrypsin) where the enzyme only interacts with and solubilizes collagen with a non-native structure. A final criterion for ensuring a positive outcome when using decellularized materials is the verification of a non-toxic material. Cytotoxic effects are mainly attributed to residual chemicals or reagents used in the decellularization process. For example, sodium dodecyl sulphate (SDS) has been shown to be cytotoxic and it binds very tightly with ECM components. Therefore, its removal from the decellularized ECM is both critical and difficult to achieve without disruption of protein structures in the ECM. Assays conducted to asses possible cytotoxicity of decellularized tissues and organs can be done with samples of intact material or with homogenized extract samples. Multiple standard cell lines, such as mouse derived immortalized 3T3 cells, can be cultured in the presence of decellularized materials and commercial cell viability assays (e.g., MTT (3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide)) can be used to detect possible cytotoxicity.

Sterilization Issues As with any material implanted into the human body, sterility is of the utmost concern to ensure patient safety. Unfortunately, most standard methods used to sterilize medical instruments and devices, such as surgical instruments, cannot be used with biological materials. Use of exposure to high energy sources, such as gamma irradiation, has been shown to have detrimental effects on the structure and chemical composition of ECM components. For example, collagen is shown to lose thermal stability indicative of it denaturing whereby it goes from a very organized helical form to a random globular (gelatin) structure (Fig. 4). This loss of collagen’s native structure in irradiated tissue is also indicated by its increased susceptibility to degradation by the enzyme trypsin (Fig. 4). Other nonenergetic standard sterilization techniques, such as treatment with the chemical ethylene oxide (EO) in gaseous form, have been shown to bind and react chemically to proteins in the ECM thereby altering their structure and chemical composition. These changes have been linked to immune responses, increased inflammation, and accelerated degradation of EO treated tissues. Further, both irradiation and EO sterilization have been shown to significantly alter the mechanical properties of tissue-based materials. One method of sterilization recently developed for decellularized materials that has been shown to be both effective and nondetrimental to its properties is the chemical reagent peracetic acid. Peracetic acid has been and is still used to sterilize donated human bone. The concentrations and conditions that are used to treat bone however, destroy soft tissues and organs. Therefore, a combination of unique treatment conditions and lower peracetic acid concentrations have been developed to successfully sterilize decellularized materials without altering their structure or properties.

Example Decellularization Applications Initially, decellularization processes as far back as the 1970s were aimed at soft-tissues such as skin, heart valves, arteries and veins, ligaments, and tendons. Organ decellularization, however, has only developed in the last decade. Even so, to date, almost

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Fig. 4 Rabbit trachea before (A) and after (B) decellularization. Trachea begins with red coloration indicating presence of cells and blood components (A) which are removed during decellularization resulting in a whitish appearance (B). Histological sections (Masson’s trichrome staining) of untreated (C) and decellularized trachea (D). Note removal of blue/black nuclei of cells after decellularization and the maintenance of the pink/purple stained extracellular matrix.

every soft-tissue and organ in the mammalian body has been exposed to a decellularization process. For example, ligament (Fig. 1), skin (Fig. 3), trachea (Fig. 4), ear cartilage (Fig. 5), Heart (Fig. 6) and Liver (Fig. 6). Clinical application and commercialization of decellularized materials has so far remained in the area of soft-tissue with applications mainly in the areas of wound care and reconstructive surgeries. Commercial wound care products for the treatment of chronic wounds such as diabetic ulcers, venous stasis ulcers, and pressure ulcers include GraftjacketÔ (Acelity), DermamatrixÔ (Synthes), DermacellÔ (Lifenet), DermapureÔ (Tissue Regenix), and DermGENÔ (DeCell Technologies Inc.) (Fig. 7). Other commercial products are being used for breast reconstruction (e.g., AllodermÔ (Aceltity)), Orthopedic surgery (e.g., OrthopureÔ (Tissue Regenix), ArthroflexÔ (Arthrex)), and dental surgery (AllodermÔ (Acelity), OracellÔ (Lifenet)). To date, the use of decellularized organs has not passed animal model studies.

Fig. 5 Human ear cartilage histological sections (hemotoxylin and eosin staining) before (A) and after (B) decellularization. Note removal of blue nuclei of cells after decellularization and the maintenance of the pink stained extracellular matrix.

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Fig. 6 (Above) Decellularization of an adult porcine heart over 48 h. A 6-month-old porcine heart undergoing perfusion decellularization over a period of 48 h. The native structure and vasculature are preserved after decellularization. After 48 h, the heart is completely decellularized. (Below) Medical decellularization of an adult porcine liver over 24 h. A 6-month-old porcine liver undergoing perfusion decellularization over a period of 24 h. The native structure and vasculature remain preserved after decellularization. This figure is taken an Elsevier publication: Eliminating the organ transplant waiting list: The future with perfusion decellularized organs. Dominique Seetapun, PhD, Jeffrey J. Ross, PhD. Surgery Volume 161, Issue 6, Pages 1474–1478 (June 2017).

Fig. 7 Application of decellularized human dermal graft (DermGEN™) on chronic diabetic foot ulcers. Patient 1 (upper) presented with a large nonhealing ulcer on the heel. After DermGEN™ was applied, the ulcer was completely healed in 8 weeks. Patient 2 presented with a small, deep ulcer on the top of the big toe. After DermGEN™ was applied, the ulcer was close after 2 weeks and remained closed at 7 weeks with continued epithelial renewal.

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The Future of Decellularization The future for the application of decellularization technology is promising. Already, successful applications of decellularization technology to solve clinical problems exist. As a better understanding develops as to what is required for the removal of immunogenic donor cellular materials and the preservation of the remaining ECM, more effective decellularization processes will be developed. Further, as new cell sourcesdsuch as adult stem cells and induced pluripotent stem cells (IPSC)dare used for the repopulation of decellularized tissues and organs, more successful applications of decellularization technology to treat human disease and dysfunction.

Further Reading Badylak, S. F., & Gilbert, T. W. (2008). Immune response to biologic scaffold materials. Seminars in Immunology, 20, 109–116. Dyck, C. R., & Gratzer, P. F. (2007). Decellularized tissues in tissue engineering. In D. R. Bloomington (Ed.), Chapter 11, New research on biomaterials (pp. 281–320). Hauppauge, NY: Nova Science Publishers Inc. Chapter 11. Fu, R.-H., Wang, Y.-C., Liu, S.-P., et al. (2014). Decellularization and recellularization technologies in tissue engineering. Cell Transplantation, 23, 621–630. Gilpin, A., & Yang, Y. (2017). Decellularization strategies for regenerative medicine: From processing techniques to applications. BioMed Research International, 2017, 1–13. Gratzer, P.F. (2017) U.S. Pat. 9 566 369 B2 (Methods for tissue decellularization) issued February 14, 2017. Kawecki, M., Labu, W., Klama-Baryla, A., et al. (2018). A review of decellurization methods caused by an urgent need for quality control of cell-free extracellular matrix scaffolds and their role in regenerative medicine. Journal of Biomedial Materials Research Part B Applied Biomaterials, 106B, 909–923. Keane, T. J., Swinehart, I. T., & Badylak, S. F. (2015). Methods of tissue decellularization used for preparation of biologic scaffolds and in vivo relevance. Methods, 84, 25–34. Morris, A. H., Chang, J., & Kyriakides, T. R. (2016). Inadequate processing of decellularized dermal matrix reduces cell viability in vitro and increases apoptosis and acute inflammation in vivo. BioResearch Open Access, 5(1), 177–187. Seetapun, D., & Ross, J. J. (2017). Eliminating the organ transplant waiting list: The future with perfusion decellularized organs. Surgery, 161, 1474–1478.

Diamond, Carbon Nanotubes and Graphene for Biomedical Applications Aaqil Rifai, Elena Pirogova, and Kate Fox, RMIT University, Melbourne, VIC, Australia © 2019 Elsevier Inc. All rights reserved.

Diamond Common Methods of Diamond Production Diamond in Medicine CVD diamond Nanodiamonds Diamond-Like Carbon Carbon Nanotubes Graphene Specific Applications Antibacterial Applications Hard Tissue Implants Drug Delivery Bio-Imaging and Bio-Sensing Conclusion Further Reading

98 98 99 99 100 101 101 103 104 104 105 106 106 106 107

Abbreviations CVD Chemical vapor deposition CNT Carbon nanotube DLC Diamond like carbon MSC Mesenchymal stem cells MWCNT Multiwalled carbon nanotube PCD Polycrystalline diamond PDMS Polydimethylsiloxane PVD Physical vapor deposition SWCNT Single walled carbon nanotube UNCD Ultrananocrystalline diamond

Emerging trends show that carbon-based materials are rapidly taking place in medical applications. Carbon materials in current biomedical use are diamond, diamond-like carbon, graphene, and carbon nanotubes. These allotropes of carbon have distinctive chemical and physical characteristics owing to their spatial arrangement of the elements. Materials like diamond, carbon nanotubes and graphene are used as a material for optics, medical electronics, tissue engineering, medical implants, medical devices, sensors and other biomedical applications. In this reference module, we focus on broad applications of diamond, carbon nanotubes and graphene. Diamond is used in advanced biomedical applications due to its hardness, wear resistance and biocompatibility properties. Diamond like carbon (DLC) is used for coating hip joints and other articulating surfaces, with an improvement to the overall longevity of the implants. Silicon carbide (SiC) has been utilized in heart stents as well pyrolitic carbon used within artificial hearts. Carbon nanotubes are more commonly used for drug delivery and sensing capabilities. Graphene is used in biomedicine with focus on drug delivery, cancer therapy and biological imaging as well as in optics as an electrically insulating material. In addition, carbon is popular due to its capacity to be formed across both the nano and micro scales. Carbon nanomaterials are revolutionizing nanomedicine due to their controllable chemistries and properties. Zero-dimensional (0-D) carbon allotrope materials have all the dimensions in the nanoscale such as particulate diamonds, fullerenes and carbon blacks. Alternatively, onedimensional materials (1-D) and two-dimensional materials (2-D) have two and one nanoscale dimension, respectively, and these are usually restricted to the thickness or diameter of the carbon material. An example of each may be carbon nanotubes (CNT), diamond nanorods or carbon nanofibers for 1-D and graphene or diamond nanoplatelets for 2-D. A large number of the biomedical carbons are however three-dimensional materials, for example, nanocrystalline diamond (NCD) films, nanodiamond, nanostructured diamond-like carbon (DLC) films and fullerene. Amongst these materials mentioned, diamond, graphene, and CNT have gained the most attention in medicine. A number of deposition techniques have been explored to produce various allotropes of carbon depending on the bonding nature (sp3 or sp2) and whether it is crystalline or amorphous. The carbon bonds can be

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Comparison of the advantages and disadvantages of some typical carbon materials used in biomedical applications

Material

Advantages

Disadvantages

Amorphous carbon (Graphite-like carbon GLC, diamond-like carbon DLC)


Moderate hardness
Wear resistance
Lower friction than diamond
High load bearing capacity
Good Adhesion
Low internal stress
Hard
Low friction
Corrosion resistant
Chemical inertness
High electrical resistance/resistivity
Optical transparency
Biocompatible
Biocompatible
High surface area
High sensitivity
Soluble
Can be functionalized
Flexible
Optical transparency
Conductive
Hard


Hardness is lower for GLC than for DLC

Diamond

CNTs

Graphene


Brittle
Difficult to upscale


Safety concerns
Mostly impure
Lack of selectivity
High sensitivity to humidity
Cost
Difficult to upscale

classified in terms of amorphous carbon (DLC), nanocrystalline diamond (NCD), and ultra nanocrystalline diamond (UNCD). Referring to Table 1, it is clear that each carbon material offers different properties and benefits towards biomedical implantation. Biomedical materials or biomaterials can be used to assist, treat, repair, or replace any function in the tissue, organ or body. Carbon-based biomaterials are known to be biocompatible and have become increasingly common within the past decade. Polymers comprising carbon are also of interest in biomedical applications. For instance, Poly-L-lactide (PLLA) is a polymer commonly used for biodegradable coronary stents and bone plates. Contact lenses made from poly(methyl methacrylate, PMMA) and polymers such as polycaprolactone (PCL) are used for resorbable applications. As nanomedicine has gained traction, these polymers are readily being used as housing for nanoparticulate and drug delivery. However, for the purpose of this reference work we have not focused on polymers rather the forms of carbon graphene, diamond and carbon nanotubes. This article emphasizes on the importance of these materials and presents the major research undertaken in this area.

Diamond Diamond is an exciting material for biomedical applications with broad and diverse medical applications from orthopedics to medical bionics. Diamond is a material composed of carbon in its sp3 form. Diamond has a reported biocompatibility and bioactivity making it a very suitable material for biointerfacing applications in areas such as orthopedics, dental and cardiovascular engineering. The first instance where diamond provided the clear advantage to more common metallic biomaterials was in wear resistant coatings for hip implants. Diamond is used in medicine in a number of different materials, predominantly nanocrystalline diamond (NCD), polycrystalline (PCD), particulate nanodiamonds (PNDs), and nanodiamonds. PNDs are commonly fabricated using detonation synthesis, however other forms of nanodiamonds can be synthesized to a high purity in to obtain specifically tuned materials. The unique characteristics of PND are hardness, high thermal conductivity, chemical stability, fluorescence and optical properties. More importantly, PNDs are highly biocompatible and non-toxic compared to other materials. Nanodiamonds are also highly versatile types of diamond as they can be functionalized and optimized according to the type of bioapplication required.

Common Methods of Diamond Production Diamond is predominantly applied as a biomedical coating. These coatings are applied using the chemical vapor deposition (CVD) technique in which materials to be coated are exposed to a plasma containing hydrogen, and carbon to deposit a secondary diamond coating. In order to do so, nanodiamond powder is required to seed or provide a thin film in order to nucleate to form the thicker CVD diamond coating. Where a freestanding diamond sheet is required sacrificial templates can be used (usually silicon) and removed using acid dissolution to leave the remaining diamond. Nanodiamonds are generated by either the bottom up

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Types of diamond Crystal size

Appearance

Biomedical use

Single crystals

mm–cm

Clear–colored

• No current biomedical application but possibilities in

Microcrystalline diamond (MCD) films

0.5–50 mm

Black–clear

Nano crystalline diamond (NCD)

10–500 nm

Black

Ultra-nano crystalline diamond (UNCD)

5–10 nm

Black

Nanodiamond (ND)

5 nmþ

Black

• Implantable electrodes • Implant encapsulation • Heat dissipation • Wear resistant coating • Wear resistance • Anti-corrosion barrier films • Barrier film (biomedical) • Implantable electrodes • Wear resistance • Barrier film (biomedical) • Antibacterial coatings • Sensing • Drug delivery

sensing

Modified from Garrett, D. J., Tong, W., Simpson, D. A. and Meffin, H. (2016). Diamond for neural interfacing: A review. Carbon 102, 437–454. doi: https://doi.org/10.1016/ j.carbon.2016.02.059 with permission.

approach of TNT detonation in a pressurized container or by the top down approach of milling a larger diamond. CVD presents a form of diamond growth in micro and nano-crystalline form, in which the crystalline form can define the diamond surface morphology. Outside of the requirement for homogeneous diamond coatings, nanodiamonds themselves can provide an excellent biomaterial. Nanodiamonds are however restricted to cellular uptake and tracking applications as far. As an example, fluorescent nanodiamonds have been used successfully in biomarking and cellular tracking applications. Diamond coatings using CVD are limited to line of sight as well as chamber size. Interestingly, however, the CVD technique of diamond growth has provided superior applications for medical bionics applications. By changing the elemental content within the gas plasma to include elements such as nitrogen and boron, conductive channels can be produced in the atomic diamond lattice to change what is normally an insulating material into a conductive diamond material. However, as with any secondary coating, the thermal mismatch between diamond and its underlying substrate material provides an opportunity for coating delamination, particularly when coated using a high temperature, high pressure environment such as the CVD chamber. A table summarizing the main types of diamond and its common medical applications are shown in Table 2.

Diamond in Medicine CVD diamond Examples of CVD diamond coatings in current medical implant applications are seen in tissue contacting areas mandibular plates, heart valves, neural stimulators and joint replacements where biological interaction is necessary. The use of diamond is however somewhat restricted by its inherent properties. In particular, its extreme hardness provides limitations in its interfacing with hard tissue in load bearing applications such as that of hip implant stems or fracture fixation where stress-sheilding would be a problem. As a result, its hardness restricts it to contact surfaces where mechanical wear may occur or where the implant requires its material to provide an avenue to guard against corrosion. Furthermore, despite its hardness, diamond is a brittle material. As a result, it has found a role in tissue engineering (see e.g., heart valves) where biocompatibility and biointerfacing is paramount due to diamond’s reported bioactivity, which provides an improved interaction between diamond-containing implants and the neighboring soft tissue. Both nanostructured diamond coatings such as ultra-smooth nanodiamond and PCD have been used to reduce the wear debris in orthopedic and dental implants providing high hardness, low surface roughness and excellent fracture toughness and adhesion. Interestingly, these characteristics are especially favored towards titanium alloys compared to other common biomedical metals such as cobalt chromium. The reason behind this is unknown, however, it is possible that the diamond has difficulty forming a carbide layer with some metals. Typically, in total hip replacement surgeries, cobalt chromium alloys have been used, due to their superior qualities mechanically, resistance to corrosion and biodegradation. Cobalt chromium alloys however are susceptible to ion release under prolonged wear. Diamond has been shown to limit the release of toxic ions and therefore increase the longevity of the implant. Although diamond offers many opportunities for biomedical use, the progress in diamond and diamond-coated implantables remains at a steady pace, with future avenues for diamond-breakthroughs being in bionics and in hermetic capsules where micro/nano devices are currently at high need. The advantages of diamond in medical bionics were highlighted by its selection for the electrode material in Australia’s contribution to the race towards the bionic eye. Diamond was selected as the electrode material due to the fact that it could be made to be both conducting and insulating within the same substrate (Fig. 1). The diamond electrode array, which interfaced directly with the retinal ganglion cells, was the key implanted material enable when placed to epiretinally interface with

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Fig. 1 Diamond based bionic eye device using ultrananocrystalline diamond for the electrodes and polycrystalline diamond for the implantable capsule: (A) shows the diamond electrode; (B) the packaging system; and (C) the implanted epi-retinal component. Reproduced with permission from Ahnood, A., Meffin, H., Garrett, D. et al. (2017). Diamond devices for high acuity prosthetic vision. Advanced Biosystems 1(1), 1600003.

both the neural cells but also with head mounted camera offering the best performance and safety for patients. The stimulating array is entirely fabricated using diamond to maximize the longevity and to increase biocompatibility due to the increased electrochemical surface area offered by the diamond itself when compared to traditional materials such as platinum. In order to fabricate the device, insulating polycrystalline diamond was used for the housing and nitrogen-incorporate ultrananocrystalline diamond was grown via CVD into laser drilled feedthrough channels. By using the CVD technique and with both conducting and insulating components made from diamond, hermetic sealing of the electrodes was achieved. As diamond can be difficult to fabricate, the limitations of diamond-based device in retinal prosthesis are apparent, however, this technique presents a suitable method to produce a large number of conductive diamond feedthroughs in monolithic polycrystalline films (256 electrodes). Biosensing is vastly expanding in the medical field, especially, with increased pressure for biomedical solutions for detection and biomarking. Previously, CVD diamond has been commercialized outside of the biomedical field for applications in heat sinks, X-ray windows, particle detectors, solar-blind UV detectors and electrodes. However, translation of these devices towards medical use has been limited. Not unlike the retinal electrodes, the potential of diamond has not been explored fully due to the difficulty in producing wafer-size devices. Fabrication of wafer-size crystals may potentially enhance the properties of diamond for its further biomedical applications. Researchers showed that nanocrystalline diamond can exhibit improved properties in micromechanical machines (MEMS and NEMS devices), surface acoustic wave devices, biological cell cultures and for DNA detection.

Nanodiamonds As with diamond films, nanodiamonds exhibit favorable characteristics in the biomedical applications due to its enhanced cellular adhesion. Nanodiamonds (ND) are emerging as a novel class of nanomaterials due to their inexpensive cost, fluorescent capability, non-cytotoxicity and biocompatibility. In fact, of all the nanocarbons, nanodiamond has the reported highest biocompatibility compared to the alternative carbon nanomaterials such as carbon black, single walled nanotubes and multiwalled nanotubes. NDs can be largely scaled as well as tunable with functionalized surfaces. Common biomedical avenues of NDs include drug delivery, cell interaction, cancer therapy, protein and gene delivery, biomarking and antibacterial applications. These NDs are highly versatile and are able to deliver antigens, water insoluble drugs, antibodies, nucleic acids and imaging agents into target cells, where the therapeutic or imaging/diagnostic molecules would be released, and efficiently utilized. Nanodiamonds have found a specific role in biomedical applications for their capacity to provide bio-sensing within individual cells. Intersecting between quantum physics and biology, the capability to track the vacancies within the diamond lattice using a laser provides revolutionary capacities for nanodiamonds in vivo. This will provide an excellent bio-sensing platform as individual nanodiamonds can specifically interact with specific cells providing information as to cellular behavior and as a drug-delivery vehicle. These biomedical systems can be scaled up for industrial production and easily altered to create designated functionality and terminations. Studies comparing nanodiamond powder to nanodiamond films suggest that both provide a solid substrate for cellular interactions. In a study by Lechleitner, published in the esteemed journal biomaterials (see Further Reading) it was shown that for a borosilicate glass sample coated with either nanoparticulate diamond powder or with nanocrystalline diamond film, both diamond coating techniques provided an improved bioscaffold for epithelial cells, with neither producing any changes to epithelial phenotype (Fig. 2). As discussed however, previously in this article, the termination of the nanodiamond powder and films dictated the epithelial response. Both hydrogen terminated and oxygen terminated powder were assessed finding that hydrogen termination provides a restrictive and inhibiting cell attachment surface compared to the more biosuitable oxygen terminated surface. As a result, the functional polar groups of nanodiamonds provide a method of controlling cell adhesion. This is important in biomedical environment where both hydrophilic and hydrophobic surfaces have a role.

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Fig. 2 Fluorescence images of the cell–surface interface of epithelial (HK-2) cells grown on glass, nanoparticulate diamond powder (DP) and oxygen terminated nanocrystalline diamond (NCD-O). Images on the left side: Alexafluor 594 labeled activated Focal adhesion kinase stain. Images on the right side: Rhodamin/Phalloidin labeled actin filaments (taken 24 h after seeding). Bars indicate 10 mm. Adapted with permission from Lechleitner, T., Klauser, F., Seppi, T., et al. (2008). The surface properties of nanocrystalline diamond and nanoparticulate diamond powder and their suitability as cell growth support surfaces. Biomaterials 29(32), 4275–4284.

Diamond-Like Carbon Amorphous carbon with a level varying in sp3/sp2 content is called diamond-like carbon (DLC). As the sp3/sp2 ratio increases in the favor of sp3, the amorphous carbon material shows diamond-like properties. Various methods exist to form DLC, but like diamond, CVD and physical vapor deposition (PVD) are favored. DLC shows promise for biomedical applications with reported excellent wear resistant properties. However the strength of DLC remains a concern. DLC exists in numerous forms of carbon, such as amorphous carbon with hydrogen (a-C:H), amorphous carbon without hydrogen (a-C) and amorphous carbon pure sp3hybridized (ta-C). The hardness between the carbons forms also differs. The biggest downfall of DLC coating has been the load bearing, residual stress and the level of adhesion with substrate materials with coating delamination prevalent. DLC has been shown to have excellent bio- and haemocompatibility. It is thought that this may be due to the carbon bonds of DLC react with oxygen removing the harmful super-oxide radicals often linked to tissue damage, strokes and cancer.

Carbon Nanotubes Carbon-based nanotechnological advances made use of such a platform for a variety of biomedical applications. Carbon nanotubes (CNTs) are composed of carbon atoms and arranged in a benzene ring forming sheets to display a seamless cylinder. CNT have superior structural, mechanical and electronic properties by diameter, length and chirality or twist. Fundamentally, CNTs are a series of cylindrical carbon tubes with a high aspect ratio (Fig. 3). CNTs can be composed of a single tube, commonly called

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Fig. 3 Scanning electron micrograph of multi-walled carbon nanotubes showing that the tubular structure is a nanometer in diameter and a micrometer in length. Adapted with permission from Harrison, B. S. and Atala, A. (2007). Carbon nanotube applications for tissue engineering. Biomaterials 28(2), 344–353. https://doi.org/10.1016/j.biomaterials.2006.07.044.

the single-walled carbon nanotube (SWNT) or the multiple cylinders, commonly known as multiwalled carbon nanotubes (MWNTs). These two types of CNT differ in term of the sheet orientation. Some of the common applications of CNTs include DNA and protein biosensors, ion channel blockers, bioreceptors and biocatalysts. As a result of the current trend in miniaturization and nanomedicine, CNTs are becoming more prominent in both neuroscience and tissue engineering research. CNT are aligned in the top tier of carbon-based materials for both research and industrial applications. CNT are ordered, hollow nanostructures composing of carbon atoms bonded around each other through sp2 bonds. There is a high demand for the use of CNTs in biomedical applications for drug delivery, as antioxidants, for biosensors and for mechanical reinforcements. CNTs are typically produced using the CVD technique. CVD has benefits in CNT production, however limitations exist in the technique when repeatability is sought. CNTs appear promising as a scaffold for neuronal and ligamentous tissue growth for the CNS and orthopedic loci. Further, CNTs have also been used as new structures capable of detecting antibodies (e.g., in autoimmune disease) or in combination with DNA or peptide nucleic acid used for ultrasensitive complementary DNA strand detection. CNTs appear to be well suited towards biomaterials and may become more interesting as a tool for tissue engineering. The CNTs have the capacity to be combined in cellular imaging, biological and chemical sensing, bioactive agent delivery, and matrix engineering. While there is still potential and development of CNT for biomedical applications, concerns arise from their cytotoxicity. The lacking aspect of CNT is the nonbiodegradability, however excretion from the body is not a problem as it was shown in vivo. In addition, the capability of creating nanosized sensors may provide key information regarding tissue microenvironments, thus yielding new insight into the cell–matrix interface. CNT have been shown as promising material for drug delivery, electrical stimulus and cellular feedback. In summary, CNT serve as a carbon-based material for novel applications on the rise including tissue engineering. CNTs for tissue engineering represent a challenging but potentially satisfying opportunity to develop the next generation of engineered biomaterials. Improved tissue engineered mechanisms are new and versatile in tunability. The possibilities of nanosensors monitoring tissues would open up pathways for tracking the performance, tissue and cellular responses. For example, the ability to monitor diseases, and changes in critical biochemical processes such as apoptosis and angiogenesis leading to progressive disorders with high spatial resolution would be beneficial for tissue engineering and sensing. An approach to monitoring engineered tissues might be to take advantage of implantable sensors capable of relaying information outside the body. Real-time data is ideal for relating to the physiological relevant variables such pH, PO2 (partial pressure of oxygen) and glucose levels. Sensing particles in the nanoscale can be a significant advantage due to a number of life threatening diseases. The sensor also reduces a need for revision surgeries and the cost of this procedure. Several features make CNTs ideal mechanisms for nanosensors including their electrical properties, large surface area, and the capacity to immobilize DNA or other proteins. It should be noted, however, that there remains doubt over the safety of CNTs within the body. CNTs are generally insoluble in all types of solvents and it remains unknown how this translates to its solubility within the body itself. Introducing foreign biomaterials into the body can be challenging with risks of cytotoxicity involved. At any particular time, understanding the organism’s response to the foreign substance is crucial. Currently, there are a number of studies that debate whether fullerene nanomaterials such as buckyballs and CNTs are cytotoxic. Up to date research is contradictory and patchy, with some research studies reporting that CNTs are toxic while others stating CNTs demonstrate excellent properties for cellular growth. Several in vitro studies, however, have indicated that CNTs may be cytotoxic. For instance, Harrison and Atala reported in Biomaterials their 2007 study which showed that when SWNTs and MWNTs were incubated with the macrophages of alveolar, a significant increase in cytotoxicity was shown after 6 h of growth. The impairment of phagocytosis of alveolar macrophage was shown to be significant with low doses of SWNT and with high doses of MWNT, necrosis and degeneration were observed.

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In contrast, however, CNTs have been shown to also improve both neural signal transfer and support dendrite elongation. Further subdermal implantation showed minimal inflammation or noted irritation. As a result, it remains clouded as to the real toxicity of CNTs. Nanotoxicity is of great interest at the moment. Consequently, many new reports are forming on the in vivo toxicity. It is currently considered that the CNTs themselves are non-toxic but that toxicity is generated by accumulation and agglomeration of CNTs (as well as other non’carbon nanoparticles used today in nanomedicine). The toxicity due to CNT agglomeration, in airways and vessels is not unlike the buildup of cholesterol within the body. Whilst there are concerns of the toxic effects of CNT, newly developed fabrication methods are likely able of mitigating some of the risk. By functionalizing CNTs with glycopolymers, the CNTs have been shown to be indistinguishable. This development gives rise to the use of CNTs for advanced tissue engineering applications.

Graphene Unlike diamond, graphene is a 2D atomic crystal that only comprises of a single layer of carbon atoms and typically has a honeycomb structure. Graphene is the individual two-dimensional crystal structures made from a layer of atomic layers of graphite (Fig. 4). Since the discovery of graphene sheet crystals, the exploration into its properties has been exponentially increasing. There have been many recent advances in biomedical graphene as the material has the attributes of supreme mechanical stiffness, elasticity, strength, thermal and electrical conductivity, all features sought after in medicine. Further, graphene has been used for advanced applications in optics, electronics and thermal therapy. Another notable recent development is graphene oxide. Graphene oxide is derived from sheets with oxygen functional groups. The most common method of graphene oxide production is known by Hummer’s Method that involves the oxidation of graphite in potassium permanganate in sulfuric acid. Graphene is composed of entirely sp2 hybridized carbon. The van der Waals forces keep the graphene sheets tightly packed together in graphite. Some of the most popular biomedical applications of graphene oxide are schematically represented in Fig. 5. Biologically, graphene is recognized for the high surface area and delocalized electrons serving as a site for drug delivery for cancer. Due to the oxidation of graphene, there are several groups in graphene that provide suitability. Like other carbon biomaterials, currently graphene is synthesized using the CVD technique. Graphene however exerts some non-desirable features when implanted in the biomedical environment. The van der Waal forces in graphene results in aggression under aqueous conditions. Therefore, caution must be taken when using graphene as opposed to other carbon-based materials for biomedical applications. Alternatively, the oxidative functional groups are hydrophilic and the water molecules can communicate with each other. The interaction of graphene oxide and water is advantageous when using as a bio-lubricant.

Fig. 4 The different structures of graphene from the 2D sheet to 3D graphite (right) or 1D CNTs (center). Figure reproduced with permission from Geim, A K. and Novoselov, K S. (2007). The rise of graphene. Nature Materials 6(3), 183–191.

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Fig. 5 Most common biomedical applications of graphene related materials including sensing applications, drug delivery, and photothermal therapy. Reproduced with permission from Bitounis, D., Ali-Boucetta, H., Hong, B. H., Min, D. H. and Kostarelos, K. (2013). Prospects and challenges of graphene in biomedical applications. Advanced Materials 25(16), 2258–2268.

In the field of nanotechnology, new materials are continuously being sought for advanced biosensing purposes where accurate, sensitive and selective detection of biomarkers are becoming a necessity. Recently, graphene-based sensors have shown great potential. Graphene-based sensors, used to detect thrombin and the protein, caspase-3, are effective tools for diagnosis and monitoring of chronic and acute conditions. Also, they are effective in screening for genetic disorders such as Alzheimer’s disease. Here using graphene, abnormalities can be traced for early identification and determination of the level of severity of a genetic disorder. Graphenebased nanomaterial solutions have two alternate technologies. One that involves the use of a probe molecule on the graphene sheet and other uses a label-free approach with measurements for electrical properties and interactions with an analyte. In addition to biosensing applications, graphene-based materials were also shown a great promise for different applications in tissue engineering. Graphene oxide, as an example, has been reported to show good biocompatibility (cellular growth and proliferation) with osteoblast cells as well as increased bone bioactivity and hydroxyapatite formation in vitro. Graphene materials have also provided some advantage to both stem cell growth and differentiation. Comparative studies have shown that the proliferation and mineralization of mesenchymal stem cells (MSC) on graphene and graphene oxide substrates was higher when compared to the standard PDMS substrates. Hence, it is likely that graphene provides an improvement in osteogenic properties than more traditional substrates. For a more detailed analysis of graphene’s interactions with other cell types (e.g., human mesenchymal stem cells, human neural stem cells, mouse hippocampal neurons) refer to the studies listed in the references and further readings at the end of this article. Interestingly, graphene appears to provide an advanced material in neural sensing with graphene reportedly promoting adherence and the inducement of neural stem cells towards neurons instead of ganglion cells as well as increasing neuronal outgrowth in the early stages of development. Fig. 6 shows the enhancement of neural stem cells on fluorinated graphene substrates showing the preferential attachment of mesenchymal stem cells on graphene. As a result, graphene-based substrates can be designed to be an advanced biomedical material for the next generation of neural chips and implantable substrates for neurodegenerative diseases. Overall, the mechanical, electrical, cellular properties of graphene can be significant for developing biomedical systems, scaffolds, and other applications.

Specific Applications Antibacterial Applications Antibacterial materials or even materials with some antibacterial advantage are critical for biomedical applications. Antibacterial coatings can be classified into two groups, passive where only upon contact will the coating effect bacterial adhesion or active where antibacterial agents are delivered locally. Reducing the risk of bacterial adhesion is highly recommended in the modern day

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Fig. 6 Graphene nanomaterials in tissue engineering. (A) Enhanced neuronal differentiation of neural stem cells (NSCs) on fluorinated graphene in 2D; (i) schematic showing mesenchymal stem cell (MSC) patterning by printing PDMS barriers on graphene films directly; (ii and iii) MSCs preferentially attached on the fluorinated graphene with aligned F-actin (red) and expression of neural specific markers-Tuj1 and MAP2 (green) (scale bar 50 mm). (B) (i) SEM of 3D graphene foam; (ii) fluorescence images of NSCs proliferated on 3D-GFs for 5 days stained for nectin (green), nuclei (blue) and (iii) fluorescence images of differentiated NSCs under differentiation conditions, Tuj-1 for neuron (green), GFAP for astrocyte (red), O4 for oligodendrocyte (green) and DAPI for nuclei (blue). Reproduced with permission from Goenka, S., Sant, V. and Sant, S. (2014). Graphene-based nanomaterials for drug delivery and tissue engineering. Journal of Controlled Release 173, 75–88.

orthopedic implantation. As carbon materials have become more prevalent in biomedical implantation, carbon nanomaterials have shown toxicity in some environments to both eukaryotic (cells with a nucleus) and prokaryotic (cells without a nucleus like bacteria) cells. It is unknown whether this is due to the hydrophilicity of these materials or due to their nanosize which both enable the nanoparticles to cross the cell membrane. As a result, these moderately toxic carbon nanomaterials such as fullerene CNTs and graphene oxide have gained momentum for antibacterial applications. The antibacterial effect appears to be related to surface treatment, reactivity and mechanical interaction which is more likely to affect bacterial adhesion than the carbon material itself. In any case, antibacterial effects have been reported for most carbon biomaterials, particularly graphene and diamond. Graphene oxide “nanowalls” showed some toxicity to bacteria whilst graphene, graphite, graphene oxide, reduced graphene have a reported effect on Escherichia coli bacterial growth. Diamond, on the other hand, appears to provide resistance to bacterial colonization when compared to more orthodox biomaterials (e.g., stainless steel and titanium).

Hard Tissue Implants Bone comprises highly porous hard tissues with variable mechanical properties. Where hard tissue replacement is required, tissues such as bone and teeth can be replicated using carbon-biomaterials. For hard tissue implants, the biggest concern is trauma, infection, tumor, and general wear. Carbon biomaterials, such as diamond are strong applicants for an implant due to its hardness, strength, biocompatibility and chemical stability to provide the longevity for implantation, particularly as a coating material. As a coating material, diamond shows osteogenic differentiation, biocompatibility and osseointegration. Graphene has also showed similar osteogenic capabilities with reported accelerated bone growth upon these substrates. When combined with polymers like polycaprolactone, functionalized MWCNTs suggest similar osteogenic properties with reported increases in osteoblast proliferation. These carbon materials may provide both the implant and the fixture to support bone formation. Similarly, nanodiamond can also improve the mechanical properties of the scaffold. Functional termination changes the properties of the material. Well dispersed oxygen terminated nanodiamond can increase the hardness of the underlying substrate for implantation. This modification can significantly increase the cellular attachment as well as the bone formation. The osteogenic, cytotoxic and mechanical properties are key criteria’s that must be evaluated against to promote a scaffold. Good osteogenic factors will improve the bonding strength of the implant and the bone tissue. With a sound fixation, the bone tissue repair as well as the wear will be long lasting. Without a good interface between implant and bone, the implant can be a subject to aseptic loosening and subsequent implant failure. Further, both diamond and CNT coatings onto more traditional metal implants can provide a protective anticorrosive layer. The anticorrosive nature of the diamond and CNTs serves as a critical element to reduce metallic ions that may be lost during the

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Fig. 7 Demonstration of the use of the NV center in the diamond lattice for sensing applications. Reproduced with permission from Suter, D. and Jelezko, F. (2017). Single-spin magnetic resonance in the nitrogen-vacancy center of diamond. Progress in Nuclear Magnetic Resonance Spectroscopy 98–99, 50–62.

wear of implants. Some carbon-based nanomaterials play a crucial role for treatment of those who suffer from arthritis. The water soluble nature of carbon molecules can act to prevent the loss of fluids in joints. Ultimately, carbon-based materials have shown benefits that supersede other materials in hard tissue implants.

Drug Delivery For drug-delivery, nanocarbon biomaterials are widely used due to their nanoscale size and unique physiochemical properties allowing the material to be taken up by individual cells. As a result, many reports exist of carbon-based materials being used for the treatment of cancer. The most significant carbon materials used for drug delivery appear to be nanodiamond, CNT and graphene. Given the strong advantages carbon appears to offer in hard tissue applications, it is logical that nanocarbons have a role in tracking bone morphogenetic protein-2 (BMP-2). Using graphene oxide coatings on titanium, appears to influence BMP-2 drug delivery. As a result, in vivo bone formation can be enhanced through graphene use. Similarly, particulate nanodiamonds are also capable of delivering BMP-2 which can induce the differentiation and enhancement of bone formation. The protein release may be regulated by the pH or delayed responses, which give surgeons more control during an operation. The steady release of the protein makes it desirable for areas that require longer bone healing. For the repair of tissue, carbon nanotubes (SWCNT) can be used. In the case of Alzheimer’s disease, nanocarbon materials have been used to deliver particular drugs such as acetylcholine in mice brains. Likewise, functionalized MWCNT can deliver growth factors through the use of non-covalent grafting and by dispersing the growth factors provide a reduction in toxicity.

Bio-Imaging and Bio-Sensing Carbon-based materials are typically presented in form of bio-imaging agents or biosensors for diagnostic purposes. CNTs have amazing electrochemical properties, making them a popular form of biosensors. Unlike conventional glassy carbon, CNTs have the ability to undergo fast electron transfer. Hence, they are able to negate most disadvantages faced by the other carbon or metal electrodes for amperometric or voltammetric analyte detection. CNT have also used optical biosensors to design new infrared fluorescence and Raman scattering of the sensors. The other material with good electrochemical and conductive properties is graphene. Graphene uses two types of systems for bio-sensing. The first uses a probe molecule which communicates with the analyte, whilst the second measures the change in the electrical properties and how it interacts with the analyte. The major advantage, when using graphene-based electrochemical sensors, is that the measurements are only affected by the electrodes. Further, carbon nanomaterials in bio-sensing platforms can be utilized for the detection of cancer biomarkers. Due to the biocompatibility of carbon based materials, it is easier to disperse the agents and markers in the body. They have lower chances of toxicity with other cells. The most exciting breakthrough however is the recent intersection of quantum physics and medicine. Diamond has long being of interest in quantum computing due to the vacancies in the diamond lattice which exhibit a detectable color center. At the present, nanodiamonds have been used to track individual cells that have taken up the single nanodiamond through phagocytosis. As a result of tracking the color center within that diamond, individual cells can be probed for activity and movement (Fig. 7). In summary, as the research evolves diamond- and graphene-based bio-sensors will be more prevalent.

Conclusion Carbon-based materials are becoming extensively used in biomedical applications. As reported throughout this article, various diamond, carbon nanotubes and graphene materials can be fabricated and manipulated to generate unique properties and provide

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a unique biomedical material. The unique properties of carbon material have the potential to improve the current status-quo in biomaterial design, tissue engineering, drug delivery, cytotoxicity, biosensing, bioelectronics and many more biomedical applications. The use of carbons in biosensing, bioelectronics and implants have risen. Advances such as the utilization of diamond for stimulating retinal electrodes demonstrate the optimization of carbon for biomedical performance. Such research complements other advances that have followed breakthroughs in the biomedical field. Carbon nanotubes provide a novel material for applications in drug delivery, whilst functionalized CNTs produce highly soluble biomaterials. Although there are some safety concerns for the nanocarbons, there remains benefits for exploring carbon in nanomedicine pathways. Similarly, graphene has thrived in biomedical industries with advantages over other non’carbon materials in implantable biosensors. Overall, numerous biomedical applications can be envisaged as the growth of carbon-based materials continues to thrive.

Further Reading Ahnood, A., Meffin, H., Garrett, D., et al. (2017). Diamond devices for high acuity prosthetic vision. Advanced Biosystems, 1(1), 1600003. Bitounis, D., Ali-Boucetta, H., Hong, B. H., Min, D. H., & Kostarelos, K. (2013). Prospects and challenges of graphene in biomedical applications. Advanced Materials, 25(16), 2258–2268. Cha, C., Shin, S. R., Annabi, N., Dokmeci, M. R., & Khademhosseini, A. (2013). Carbon-based nanomaterials: Multifunctional materials for biomedical engineering. ACS Nano, 7(4), 2891–2897. Chen, H., Müller, M. B., Gilmore, K. J., Wallace, G. G., & Li, D. (2008). Mechanically strong, electrically conductive, and biocompatible graphene paper. Advanced Materials, 20(18), 3557–3561. Correa-Duarte, M. A., Wagner, N., Rojas-Chapana, J., et al. (2004). Fabrication and biocompatibility of carbon nanotube-based 3D networks as scaffolds for cell seeding and growth. Nano Letters, 4(11), 2233–2236. Garrett, D. J., Ganesan, K., Stacey, A., et al. (2011). Ultra-nanocrystalline diamond electrodes: Optimization towards neural stimulation applications. Journal of Neural Engineering, 9(1), 016002. Garrett, D. J., Tong, W., Simpson, D. A., & Meffin, H. (2016). Diamond for neural interfacing: A review. Carbon, 102, 437–454. https://doi.org/10.1016/j.carbon.2016.02.059. Geim, A. K., & Novoselov, K. S. (2007). The rise of graphene. Nature Materials, 6(3), 183–191. Goenka, S., Sant, V., & Sant, S. (2014). Graphene-based nanomaterials for drug delivery and tissue engineering. Journal of Controlled Release, 173, 75–88. Hadjinicolaou, A. E., Leung, R. T., Garrett, D. J., et al. (2012). Electrical stimulation of retinal ganglion cells with diamond and the development of an all diamond retinal prosthesis. Biomaterials, 33(24), 5812–5820. Harrison, B. S., & Atala, A. (2007). Carbon nanotube applications for tissue engineering. Biomaterials, 28(2), 344–353. https://doi.org/10.1016/j.biomaterials.2006.07.044. Hauert, R. (2004). An overview on the tribological behavior of diamond-like carbon in technical and medical applications. Tribology International, 37(11), 991–1003. Jaatinen, J. J., Korhonen, R. K., Pelttari, A., et al. (2011). Early bone growth on the surface of titanium implants in rat femur is enhanced by an amorphous diamond coating. Acta Orthopaedica, 82(4), 499–503. Kim, S., Ku, S. H., Lim, S. Y., Kim, J. H., & Park, C. B. (2011). Graphene–biomineral hybrid materials. Advanced Materials, 23(17), 2009–2014. Lechleitner, T., Klauser, F., Seppi, T., et al. (2008). The surface properties of nanocrystalline diamond and nanoparticulate diamond powder and their suitability as cell growth support surfaces. Biomaterials, 29(32), 4275–4284. Lee, W. C., Lim, C. H. Y., Shi, H., et al. (2011). Origin of enhanced stem cell growth and differentiation on graphene and graphene oxide. ACS Nano, 5(9), 7334–7341. Li, N., Zhang, X., Song, Q., et al. (2011). The promotion of neurite sprouting and outgrowth of mouse hippocampal cells in culture by graphene substrates. Biomaterials, 32(35), 9374–9382. Lin, X., Clasky, A., Lai, K., & Yang, L. (2016). Carbon-based nano biomaterials: Design, fabrication and application. In , 49. In: Biomedical nanomaterials: From design to implementation. London: Institution of Engineering and Technology. McGuinness, L. P., Yan, Y., Stacey, A., Simpson, D., Hall, L., et al. (2011). Quantum measurement and orientation tracking of fluorescent nanodiamonds inside living cells. Nature Nanotechnology, 6(6), 358–363. Park, S. Y., Park, J., Sim, S. H., et al. (2011). Enhanced differentiation of human neural stem cells into neurons on graphene. Advanced Materials, 23(36), H263–H267. Passeri, D., Rinaldi, F., Ingallina, C., et al. (2015). Biomedical applications of nanodiamonds: An overview. Journal of Nanoscience and Nanotechnology, 15(2), 972–988. Siew, P. S., Loh, K. P., Poh, W. C., & Zhang, H. (2005). Biosensing properties of nanocrystalline diamond film grown on polycrystalline diamond electrodes. Diamond and Related Materials, 14(3), 426–431. https://doi.org/10.1016/j.diamond.2004.11.016. Suter, D., & Jelezko, F. (2017). Single-spin magnetic resonance in the nitrogen-vacancy center of diamond. Progress in Nuclear Magnetic Resonance Spectroscopy, 98-99, 50–62. Tang, L., Tsai, C., Gerberich, W., Kruckeberg, L., & Kania, D. (1995). Biocompatibility of chemical-vapor-deposited diamond. Biomaterials, 16(6), 483–488.

Gold Nanoparticles for Colorimetric Detection of Pathogens Paul Z Chen and Frank X Gu, University of Waterloo, Waterloo, ON, Canada © 2019 Elsevier Inc. All rights reserved.

Introduction Classical Methods for Pathogen Detection Principles of Biosensing Gold Nanoparticles as Colorimetric Biosensing Agents Nucleic Acid-Based Pathogen Detection Complementary Base Pairing Aptamers Antibody-Based Pathogen Detection Nonfunctionalized Pathogen Detection Outlook Acknowledgements Further Reading

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Glossary Absorbance Energy loss of photons passing through a material due to dissipation through inelastic processes. Extinction The combined effects of absorbance and scattering. For nanoparticles, especially those greater than 40 nm, a spectrophotometer generally measures extinction rather than absorbance due to nonnegligible scattering processes. Multiplex Synonymous with “many”. Used as an adjective to describe the simultaneous detection of multiple pathogens, usually through the use of a corresponding number of specific biorecognition elements. Scattering An incident photon energy causes a photon to be emitted at the same frequency or a shifted frequency. Sensitivity A commonly used homonym: In analytical contexts, it is the smallest concentration change that can be detected, which is sometimes, but not always, the same as the limit of detection. In diagnostic contexts, it is the true positive rate. Surface plasmon resonance The resonant, or maximal, oscillation of conduction electrons at a negative permittivity/positive permittivity interface under illumination. Dielectric/metal interfaces satisfy the permittivity requirement, which leads to a significant enhancement in the absorption and scattering properties of metals. UV/Vis spectrophotometer A key instrument used to characterize colorimetric nanoparticles. It is used for the measurement, at a single wavelength or spectral range, of the attenuation of a beam of light when passing through or reflecting off of a sample.

Abbreviations Au NP Gold nanoparticle Au NR Gold nanorod bp Base pair CTAB Cetyltrimethylammonium bromide DNA Deoxyribonucleic acid LSPR Localized surface plasmon resonance mL Milliliter, 10
3 L nm Nanometer, 10
9 m PCR Polymerase chain reaction PNA Peptide nucleic acid POC Point-of-care RNA Ribonucleic acid SPR Surface plasmon resonance ssDNA Single-stranded deoxyribonucleic acid ssRNA Single-stranded ribonucleic acid

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Introduction Over the last 20 years, nanomaterial-based pathogen detection has witnessed an explosion in interest. Publications in the area increase yearly significantly each year; intense research has fuelled academic and clinical significance explorations. Seminal studies from the late 1990’s and early 2000’s have laid the foundation for a number of novel methods for pathogen detection, of which some are undergoing clinical validation or are even available commercially. Such progression is due to two primary reasons: First, the study of nanomaterials has itself exploded to uncover desirable properties that are unavailable to bulk-scale materials. As one example, at the bulk scale gold is, simply, the color gold. As nanoparticles, however, their color can be tuned throughout the visible spectrum (red, orange, green, purple, blue, etc.) and beyond. Second, infectious diseases continue to be one of the world’s most pressing issues. Once thought to be a solved problem, mortality and morbidity mount. Infectious diseases still account for 20% of global deaths and 30% of disability-adjusted life years, a measure of overall disease burden. Moreover, the crisis of drug resistance threatens the efficacy of treatment. Effective diagnosisdusually through the detection of pathogens, the causative agents of infectious diseasedis classically insensitive, slow, unreliable and difficult but remains as the key step in improving outcomes and reducing drug resistance.

Classical Methods for Pathogen Detection Detecting pathogens has been a formidable challenge that has thus far limited disease outcomes. Classically, pathogen detection is preformed using one of three methods: culturing and inspection, immunological assays or polymerase chain reaction (PCR)-based assays (Fig. 1). Culturing and inspection isolates and grows pathogens from a clinical specimen. For detection, cells are then grown in selective media and/or undergo an indefinite series of biochemical tests (Gram staining, catalase test, hemolysis, etc.). Thus, turnaround times are slow (generally days, sometimes 2 weeks for certain pathogens), and the method is prone to contamination and inaccurate results. Moreover, some pathogens cannot be cultured in the lab. The second classical method, immunological assay, is primarily preformed in the clinic for biomolecule detection but is occasionally used for pathogens. The enzyme-linked immunosorbent assay (ELISA) is a well-known example. ELISA uses antibodies bound on a substrate, usually a microwell plate, to capture and isolate the target analyte. In sandwich ELIZA, the most popular format, secondary antibodiesdusually a polyclonal antibodydthat target a surface epitope of the analyte are then added. These secondary antibodies are conjugated to enzyme that catalyzes the conversion of chromogenic substrate for a colorimetric response. The intensity of colorimetric response is proportional to the concentration of target analyte. Immunological assays are, thus, highly specific but are costly and still require many operational steps by trained personnel and have a long sampleto-answer time. PCR-based assays amplify and then detect a target nucleic acid sequence through the design of primer and probe nucleic acid sequences. If the target sequence is specific to a species or even strain of pathogen, then detection can be conferred. A number of variations of PCR have been developed. Reverse transcription-PCR (RT-PCR) can be used to detect RNA viruses. It also takes advantage of the rapid degrading nature of RNA to distinguish between live and dead cells. Quantitative realtime PCR (qPCR) uses fluorescent molecular probes to monitor amplification in real-time and thus can confer detection during amplification. PCR-based methods can decrease run time but require extensive sample preparation by highly trained personnel and are susceptible to inaccuracies due to contamination. Interestingly, culturing and inspection remains the “gold standard” for pathogen detection. Some hospitals have begun to adopt qPCR but a clear need exists for improved methods.

Culturing & Inspection

Immunological assays

PCR-based assays Fig. 1

Schematic representations of the three classical methods for pathogen detection.

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Principles of Biosensing A biosensor is comprised of a biorecognition element coupled to a transducer. The former enables binding of the nanomaterial to target pathogen; the latter converts a binding event to some measurable readout (e.g., color). Generally, nanomaterials are functionalized with a biorecognition element and employed as transducer or a component of the transduction process. The performance of a biosensor depends on six key characteristics: sensitivity, dynamic range, accuracy, sample-to-answer time, robustness and amenability. Sensitivity is the smallest concentration change that can be detected. Sometimes, but not always, it is equivalent to the limit of detection, the lowest concentration that can be detected. Dynamic range is the working, linear scoped from lowest concentration to highest concentrationdwhich the biosensor operates. Accuracy is a key factor for validity and considers false results. Sample-to-answer time is the integral metric for rapidity, as it considers the entire method (sample preparation, amplification, detection time, etc.), which is of greatest diagnostic importance. A robust biosensor is not easily perturbed by potential changes in conditions and can dictate overall performance. Amenability depends on a host of factors such as ease of use, requirements for trained personnel, number of operational steps and amenability for use at the point-of-care (POC). The introduction of nanomaterials has led to novel, improved biosensors. These biosensors harness unique properties of nanomaterials for improved pathogen detection. Reports have demonstrated them to be more sensitive, accurate, rapid, robust and utilizable.

Gold Nanoparticles as Colorimetric Biosensing Agents For biosensing, gold nanoparticles (Au NPs) are the most prevalent nanomaterial used. In general, Au NPs are the most stable type of metal nanoparticle. They have also garnered much attention due to facilely for synthesis and modification. Importantly, Au NPs can be prepared in a high quality, high yield and size-controllable manner with notable colloidal stability. In addition, the formation of spontaneous, covalent thiol’gold bonds is often used for modification. For example, antibodies or nucleic acids with exposed thiol groups can chemisorb onto and functionalize Au NPs in a facile manner. Interesting optical properties of Au NPs further increase their attractiveness for pathogen detection. Many of these optical properties rely on surface plasmon resonance (SPR). In SPR, coherent electron oscillations (called plasmons) are excited by the electric field of incident light. To transfer momentum from a photon to a plasmon, a negative permittivity/positive permittivity interface must be present. Metal-dielectric interfaces satisfy this requirement (i.e., the phenomenon occurs at a surface), and these excitations are strongest at a resonant frequency, hence SPR. For nanoparticles, plasmons are confined or localized to the significantly smaller nanoparticle surface, hence LSPR, which leads to an enhanced, local electromagnetic field. Nanoparticles composed of a noble metal, such as gold, show strong LSPR bands in the visible range thereby enabling colorimetric utility. In addition to composition, shape, size, local environment and interparticle distance can strongly influence LSPR (Fig. 2). Changes in the latter two factors, through binding or aggregation upon recognition of analyte, are exploited in the design of colorimetric biosensors. Colorimetric biosensing enables a number of modalities. For example, color change in response to a pathogen can be observed by eye, increasing amenability. Moreover, it can be assessed quantitatively via camera or spectrophotometer for more robust measurements. Furthermore, small changes in local environment and interparticle distance are observed with real-time changes in visual color to potentially increase rapidity. For colorimetric pathogen detection, and general biosensing, Au NPs are (B)

Increasing size

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Fig. 2 (A) Transmission electron micrographs of gold nanoparticles with varying sizes and shapes. Scale bars are 50 nm each. (B) Corresponding image of visual colors of those gold nanoparticles. Reproduced from Verma, M. S., Chen, P. Z., Jones, L. and Gu, F. X. (2014). Branching and size of CTAB-coated gold nanostars control the colorimetric detection of bacteria. RSC Advances 4, 10660, with permission from the Royal Society of Chemistry.

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functionalized with nucleic acids or antibodies or are nonfunctionalized. The following sections will elucidate these three methods and while highlighting seminal, impactful and promising reports of Au NPs.

Nucleic Acid-Based Pathogen Detection Complementary Base Pairing Biosensors can utilize the highly conserved, specific binding between nucleic acid base pairs (bps) for pathogen detection. Through thiol’gold chemistry, Au NPs can be functionalized with probes of DNA, RNA or peptide nucleic acid (PNA), noncharged oligomers with a peptide bond-linked backbone that present various purine or pyrimidine bases. A target pathogen can be detected through binding of a species- or strain-specific complementary sequence. Small changes in refractive index local to the surface of the Au NPs can lead to a colorimetric response. Such an approach, however, is insensitive. That is, innumerably large concentrations of target nucleic acid must be present for a sufficient colorimetric response. In addition, nonspecific binding or simple changes in the composition of solution can lead to false positive readings. For greater sensitivity, specificity and robustness, an aggregation-based approach utilizing the dependence of LSPR on interparticle distance can be employed. In a seminal report, Mirkin et al. showed three key results: Au NPs can be stably functionalized with ssDNA probes; recognition of target sequence can lead to a visually discernible colorimetric response; and this response is highly specific, distinguishing target from single-bp mismatches. Spontaneous aggregation is achieved through the use of two ssDNA probes that are complementary to each half of the target sequence (Fig. 3). Detection is conferred only upon binding to both halves and formation of dsDNA between Au NPs. This enables the facile colocalization of Au NPs. A significantly stronger and more sensitive colorimetric response can be observed. Since small, spherical Au NPs were used, the characteristic red-to-blue color change occurs. Many reported biosensors have employed this foundational strategy, two probe-functionalized Au NPs for target recognition. Some modifications have used a portable spectrophotometer for quantitative POC analysis. Other use camera phones and apps for image-based colorimetric analysis. For clinically relevant pathogen detection, however, additional factors must be considered. First, the target sequence must be accessible. DNA or RNA of the pathogen (bacteria, virus, fungi, etc.) must be extracted. Second, amplification must be preformed. Although, a relatively sensitive method, target sequences are usually too dilute at clinically relevant concentrations. Targeting sequences with many copies per pathogen (e.g., 16S rRNA) can help but even these nucleic acids are too dilute. In addition, reducing sample volumes (e.g., in very small microfluidic devices) can improve sensitivity, but this comes with a trade-off in the observable colorimetric response and a smaller dynamic range. Amplification before detection is usually preformed via PCR or a related technique, such as rolling circle amplification (RCA). Amplification after binding can be preformed by enriching the signal of Au NPs, for example, by catalytically growing the AuNPs. Third is a constraint on all colorimetric approaches: changes in color must be observable. Many clinical samples (e.g., blood) have their own color and varying degrees of opacity. Thus, for these, pathogens must be first be isolated.

(A)

Target DNA

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Fig. 3 (A) Schematic representation of nucleic acid-functionalized gold nanoparticles used for the colorimetric detection of complementary sequence. (B) Visual image of the colorimetric response after aggregation in response to (i) target nucleic acid, (ii) no target, (iii) complementary to one probe, (iv) a 6-bp deletion, (v) a 1-bp mismatch, and (vi) a 2-bp mismatch. Reproduced from Elghanian, R., Storhoff, J. J., Mucic, R. C., Letsinger, R. L. and Mirkin, C. A. (1997). Selective colorimetric detection of polynucleotides based on the distance-dependent optical properties of gold nanoparticles. Science 277, 1078–1081, with permission from the American Association for the Advancement of Science.

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Interestingly, the use of PNA probes generally increases sensitivity, accuracy and robustness. Their neutral backbone enables PNA-DNA binding without the electrostatic repulsion that accompanies DNA–DNA binding. In addition, PNAs are more resistant to nucleases and proteases. But, some PNAs may be slightly more hydrophobic than their DNA or RNA counterpart, although this does not typically affect Au NP stability. Approaches alternative to the two half-complementary functionalized probes exist for pathogen detection. In one approach, after amplification, free ssDNA probe can be added to a sample before the addition of Au NPs and salt. ssDNAs, generally adsorb more significantly onto the surface of Au NPs than do dsDNAs. Au NPs coated with ssDNA have improved stability over bare Au NPs. Thus, in the presence of salt, which is applied at a concentration surpassing colloidal stability, bare Au NPs aggregate while ssDNA-coated ones do not. In other words, a solution changing from red to blue indicates the detection of target pathogen. A similar approach uses short thiol-modified probes. Rather than absorb onto the surface of Au NPs, the thiol allows for covalently binding. Target sequence-binding leads to a much longer dsDNA. Thus, in the presence of salt, dsDNA-coated Au NPs remain red whereas aggregation and blue color indicate a lack of target. A third approach uses DNAzymes, nucleic acids that can catalyze the cleavage of other nucleic acids. DNAzyme-functionalized Au NPs can cleave DNA or RNA. The cleaved nucleic acids can then induce aggregation in the presence of salt and heat.

Aptamers Aptamers are synthetic oligomers of ssDNA, ssRNA or peptides that bind to some target with high specificity and affinity. Analogous to antibodies, aptamer binding leads to structurally conform to their target, with the bound state as thermodynamically favorable. They are often used for the detection of small molecules or metal ions, which are difficult to detect using antibodies (too large compared to small molecules or metal ions) or nonfunctionalized approaches (no specificity for targets at that size scale). And, indeed, aptamers can be used to detect pathogen byproducts enabling indirectly detection of pathogens, such as for the Shiga toxin-producing Escherichia coli. However, their specificity and high stabilitydrelative to antibodiesdalso allow them to be used for direct pathogen detection. In one approach, aptamers first adsorb onto the surface of Au NPs. In the presence of target pathogen, aptamers dissociate from the surface of gold nanoparticles. Thus, their specificity and higher affinity for target enables a red-to-blue color change that signals the presence of pathogen. In another approach, Au NPs can be modified with two ssDNA probes. Each probe is complementary to part of the DNA aptamer to yield an initial blue color. Importantly, some aptamer bases remain unpaired. In the presence of target pathogen, the aptamer binds and changes conformation thereby leading to dissociation and therefore de-aggregation of the Au NPs. A color change to red signals pathogen detection.

Antibody-Based Pathogen Detection Antibodies can be stably functionalized, without loss of structure and while presenting paratope, onto Au NPs. Antibodies can be modified to present thiol groups or Au NPs can be modified to present carboxylate or amine groups. The biotin–avidin or biotin– streptavidin method can also be used to functionalize modified Au NP. Thus, while anchored by a Au NP, binding antigen leads to a change in the local refractive index and associated colorimetric response. Increased packing of Au NPs on the surface of pathogens can lead to interparticle coupling for a corresponding increase in sensitivity; dynamic range may increase as well. Similar factors must be considered for antibody-based pathogen detection. A colorimetric response must be observable in sample of interest. Additionally, like DNA and RNA, antibodies are prone to denaturation from heat. Thus, antibody-based approaches must considered elevated temperatures. Due to the amplification step present in DNA-based approaches, when compared antibody-based ones are generally less sensitive. Colorimetrically, antibody-functionalized Au NPs have been used for the detection of foodborne infections. Food matrices are usually amenable for nanoparticles to retain stability, and after slight preparation, samples are void enough of conflicting color for the direct detection of pathogens. Thus, the steps required for nucleic acidbased detection (extraction, isolation, amplification) are not necessary for such applications if the Au NPs are sensitive enough. Antibody-functionalized Au NPs are more often used for direct detection without a large number of preliminary steps. As previously mentioned, the structure of Au NPs is highly tunable. For example, rather than a spherical shape, rods can be synthesized. These are called gold nanorods (Au NRs). Since LSPR strongly depends on it, different structures have different optical properties. Thus, two distinctly different structures can be used for multiplex detection (Fig. 4). That is, Au NPs, or more specifically Au NRs in this case, that contrast largely can detect one or more pathogens from the same sample. As one example, the aspect ratio of Au NRs (length vs. width) can be increased to change extinction spectra. With nonoverlapping signals and appropriately functionalized each Au NR with antibody, multiplex pathogen detection can be achieved based on aggregation around the surface of the pathogen(s) (where the epitope[s] should be presented). Visual detection for this method is possible; but the presence of multiple Au NRs in the same solution usually necessitates the use of an instrument for accurate utility. An interesting format using antibody-functionalized Au NPs is the lateral flow immunochromatographic assay, a POC assay that is colloquially known from the home pregnancy test. Indeed, the colored lines (usually red) that appear on some home pregnancy tests are due to the collection of Au NPs. In these products, urine is applied to the sample zone. Once soaked, fluid travels through capillary action. As it travels up the pad a collection of antibody-functionalized Au NPs is rehydrated and bind to target protein. If present, target protein is then captured on the test zone in a sandwich format by pad-bound antibodies. Thus, the development of

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Pathogen 1

Fig. 4 (Top) Schematic representation of two antibody-functionalized gold nanorods used for the multiplex detection of whole-cell pathogens. Gold nanorods vary in aspect ratio to enable. Spectrophotometric responses (A) before and (B) after recognition of a target pathogen. Transmission electron micrographs showing the aggregation of gold nanorods with (C, D) differing surface coverage around target bacteria. Reproduced from Wang, C. and Irudayaraj, J. (2008). Gold nanorod probes for the detection of multiple pathogens. Small 4, 2204–2208, with permission from Wiley-VCH.

red color on the test zone signals the presence of target protein and thus pregnancy. Remaining Au NPs are captured on a control zone to ensure robustness. This approach has also been demonstrated for the detection of pathogens. Antibodies are replaced with ones suitable for target pathogen. The sample of interest is applied similarly to the sample zone, and rehydrated Au NPs develop color to detect pathogen.

Nonfunctionalized Pathogen Detection The last main method uses nonfunctionalized Au NPs. Without nucleic acids or antibodies presented, the surface of Au NPs are usually stabilize with a capping agent, such as citrate, surfactant, such as CTAB, or thiol-ligand, such as an alkanethiol. These agents are usually present during, to help direct, synthesis and self-assemble onto the Au NP surface. If implemented properly, these agents can also aid in the detect pathogens. Nonfunctionalized Au NPs can be used for the general detection of pathogens. For example, using a positively charged thiolligand can impart a positive surface charge onto Au NPs. b-Galactosidase, an anionic enzyme near neutral pH, can then selfassemble onto the surface. While on the surface of the Au NP, the enzyme is inhibited. Thus, even in the presence of chromogenic substrate, no color change occurs. In the presence of bacteria, which present many negative surface charges (teichoic acids, lipopolysaccharides, phospholipids, etc.), Au NPs bind, and b-galactosidase is displaced. Thus, activity is restored; substrate is quickly converted and color develops. Moreover, similar to ELISA, multiple turnovers, a characteristic feature for an enzyme, provide amplification for color development; and an impressive limit of detection can be reached (100 cells mL
1). Interestingly, these Au NPs can be printed onto paper strips for a cost-effective POC test. The previous report highlights major advantages and disadvantages of nonfunctionalized Au NPs. Interestingly modifications can enable novel, effective biosensing strategies. However, nonfunctionalized Au NPs are generally limited from utility in complex environments due to nonspecific interactions. For example, positively charged Au NPs can change color due to the presence of any sufficient negative charge (e.g., protein, salt, erythrocytes), leading to significant false positives. Relevance in some clinical samples is currently limited in contrast to nucleic acid- or antibody-based approaches. Simple environments, which are expected to be sterile and void of sufficient negative charge, such as drinking water, are best suited for these nonfunctionalized Au NPs. In addition, species- or strain-level identification, rather than general detection of pathogens, can be challenging. One interesting report employs a “chemical nose” strategy. Here, Au NPs are coated with CTAB to present a positive charge. They interact with the negative surface charge of pathogens. Since the amount and density of negative surface charge is unique for each

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Fig. 5 (A) Schematic representation of nonfunctionalized gold nanoparticles used as a “chemical nose” for the detection of whole-cell pathogens. (B) Colorimetric responses of “chemical nose” biosensor to various bacterial pathogens. Reproduced from Verma, M. S., Chen, P. Z., Jones, L. and Gu, F. X. (2014). “Chemical nose” for the visual identification of emerging ocular pathogens using gold nanostars. Biosensors & Bioelectronics 61, 386–390, with permission from Elsevier.

bacterial species, the Au NPs produce a unique degree of color change when testing ocular bacterial pathogens. Thus, colorimetric identification at the species-level can be achieved in an accurate manner using a single set of Au NPs (Fig. 5).

Outlook Au NPs enable a variety of biosensing approaches for the colorimetric detection of pathogens. Importantly, when compared with classical methods, these Au NP-based ones are more sensitive, accurate, rapid and robust. The use of colorimetry also maximizes amenability; pathogen detection can be conferred visually. And, POC formats can be developed naturally. With varying advantages and disadvantages, Au NPs can be employed in three ways: they can be functionalized with nucleic acids or antibodies or can avoid functionalization. As these products pass clinical trials and are introduced to real-world application, they will provide exceptional diagnostic capabilities and therefore improve disease outcomes while curbing the development of drug resistance.

Acknowledgements This work was financially supported by the Natural Science and Engineering Research Council of Canada (NSERC). P.Z.C. is supported by the NSERC Vanier Canada Graduate Scholarship and WIN Nanofellowship.

Further Reading Ahmed, A., Rushworth, J. V., Hirst, N. A., & Millner, P. A. (2014). Biosensors for whole-cell bacterial detection. Clinical Microbiology Reviews, 27, 631–646. Bell, S. E., & Sirimuthu, N. M. (2006). Surface-enhanced Raman spectroscopy (SERS) for sub-micromolar detection of DNA/RNA mononucleotides. Journal of the American Chemical Society, 128, 15580–15581. Elghanian, R., Storhoff, J. J., Mucic, R. C., Letsinger, R. L., & Mirkin, C. A. (1997). Selective colorimetric detection of polynucleotides based on the distance-dependent optical properties of gold nanoparticles. Science, 277, 1078–1081. Giljohann, D. A., & Mirkin, C. A. (2009). Drivers of biodiagnostic development. Nature, 462, 461–464. Hamula, C. L., Hughes, K., Fisher, B. T., et al. (2016). T2Candida provides rapid and accurate species identification in pediatric cases of candidemia. American Journal of Clinical Pathology, 145, 858–861. Huang, S. H. (2006). Gold nanoparticle-based immunochromatographic test for identification of Staphylococcus aureus from clinical specimens. Clinica Chimica Acta, 373, 139–143. Kelley, S. O., Mirkin, C. A., Walt, D. R., et al. (2014). Advancing the speed, sensitivity and accuracy of biomolecular detection using multi-length-scale engineering. Nature Nanotechnology, 9, 969–980. Kelly, K. L., Coronado, E., Zhao, L. L., & Schatz, G. C. (2003). The optical properties of metal nanoparticles: The influence of size, shape, and dielectric environment. Journal of Physical Chemistry B, 107, 668–677. Li, X., Kong, H., Mout, R., et al. (2014). Rapid identification of bacterial biofilms and biofilm wound models using a multichannel nanosensor. ACS Nano, 8, 12014–12019. Liu, J., & Lu, Y. (2006). Preparation of aptamer-linked gold nanoparticle purple aggregates for colorimetric sensing of analytes. Nature Protocols, 1, 246–252. Miranda, O. R., Li, X., Garcia-Gonzalez, L., et al. (2011). Colorimetric bacteria sensing using a supramolecular enzyme-nanoparticle biosensor. Journal of the American Chemical Society, 133, 9650–9653. Murray, C. J., Vos, T., Lozano, R., et al. (2012). Disability-adjusted life years (DALYs) for 291 diseases and injuries in 21 regions, 1990-2010: A systematic analysis for the global burden of disease study 2010. Lancet, 380, 2197–2223. Neu, H. C. (1992). The crisis in antibiotic-resistance. Science, 257, 1064–1073.

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Verma, M. S., Chen, P. Z., Jones, L., & Gu, F. X. (2014a). Branching and size of CTAB-coated gold nanostars control the colorimetric detection of bacteria. RSC Advances, 4, 10660. Verma, M. S., Chen, P. Z., Jones, L., & Gu, F. X. (2014b). “Chemical nose” for the visual identification of emerging ocular pathogens using gold nanostars. Biosensors & Bioelectronics, 61, 386–390. Verma, M. S., Chen, P. Z., Jones, L., & Gu, F. X. (2015). Controlling “chemical nose” biosensor characteristics by modulating gold nanoparticle shape and concentration. Sensing and Bio-Sensing Research, 5, 13–18. Verma, M. S., Wei, S. C., Rogowski, J. L., et al. (2016). Interactions between bacterial surface and nanoparticles govern the performance of “chemical nose” biosensors. Biosensors & Bioelectronics, 83, 115–125. Verma, M. S., Tsuji, J. M., Hall, B., et al. (2016). Towards point-of-care detection of polymicrobial infections: Rapid colorimetric response using a portable spectrophotometer. Sensing and Bio-Sensing Research, 10, 10–19. Wang, C., & Irudayaraj, J. (2008). Gold nanorod probes for the detection of multiple pathogens. Small, 4, 2204–2208.

Manufacture of Biomaterials Min Wang, Lin Guo, and Haoran Sun, The University of Hong Kong, Pokfulam, Hong Kong © 2019 Elsevier Inc. All rights reserved.

Introduction Special Requirements for Making Biomaterials Good Manufacturing Practice Manufacture of Metals for Biomedical Applications Chemical Metallurgy Metal Fabrication Techniques: Forming Operations Metal Fabrication Techniques: Casting Other Metal Fabrication Techniques Physical Metallurgy Machining of Metallic Biomaterials Manufacture of Porous Metals Manufacture of Biomedical Polymers Polymer Classification and Architecture Polymer Synthesis Step-growth polymerization Addition polymerization Copolymerization Amorphous Polymers and Semi-Crystalline Polymers Polymer Additives Polymer Blends and Interpenetrating Networks Polymer Forming Techniques Compression molding and transfer molding Injection molding Extrusion Blow molding Casting Fabrication of Hydrogels Manufacture of Porous Polymers Manufacture of Bioceramics Fabrication of Dense Bioceramics Manufacture of Porous Bioceramics Fabrication of Pyrolytic Carbon Fabrication of Bioactive Glasses Fabrication of Bioactive Glass-Ceramics Manufacture of Biomedical Composites Composite and Composite Classification Fabrication of Particulate Biomedical Composites Polymer matrix composites Metal matrix composites Ceramic matrix composites Fabrication of Fibrous Biomedical Composites Short-fiber composites Long-fiber composites Fabrication of Laminated Structures Fabrication of Porous Composite Scaffolds for Tissue Engineering Manufacture of Nano-Biomaterials Fabrication of Nano-Structured Biomaterials Fabrication of Nano-Sized Biomaterials Surface Modification Techniques for Biomaterials Surface Modification Techniques for Biomedical Polymers Surface Modification Techniques for Metallic Biomaterials Biomimetic Deposition Summary Acknowledgments Further Reading

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Introduction Special Requirements for Making Biomaterials Materials (metals, polymers, ceramics, and composites) are made and used in nearly all engineering fields. Many technologies for making these materials and their products are also suitable for the manufacture of biomaterials (metallic biomaterials, biopolymers, bioceramics and biomedical composites). Biomaterials can be considered as a sub-set of engineering materials and the paramount criterion for turning engineering materials into biomaterials is the biocompatibility of materials. Many materials meet the biocompatibility requirement and also possess required properties (mechanical, electrical, biological, etc.) and hence they are used as biomaterials. However, any or slight deviation in composition or microstructure for the material may instantly disqualify the material from the biomedical application as such changes can result in the loss of the required property or in significantly lowered performance in the biological environment. Therefore, the manufacturing technologies should be carefully selected and used for biomaterials, avoiding change in composition and/or microstructure for the finished products. Another point that must be borne in mind in biomaterials manufacture is the purity of biomaterials. Any synthesis or fabrication process should not introduce impurity into the biomaterial. Accidental inclusion of impurity/impurities in biomaterials during their fabrication can alter their biological performance to be below an acceptable level. Therefore, commercial production of biomaterials and manufacture of implants and medical devices most likely take place in the clean room, avoiding any contamination. The final point is about sterilization. Implant and medical devices are used for patients in the sterilized state, which is a major difference for biomaterials and engineering materials. The product sterilization rendered by currently used sterilization methods can cause deterioration of mechanical properties and shortened shelf-life of implants, especially polymeric devices. Manufacture of biomaterials and their sterilization may need to be considered together.

Good Manufacturing Practice Different from the production of engineering materials and their products, the manufacture of biomaterials and their products must follow good manufacturing practices (GMPs). GMPs are practices issued by government agencies that control the authorization and licensing of the manufacture and sale of medical devices. They provide minimum requirements that a manufacturer must meet to assure that their products are consistently high in quality, from batch to batch, for their intended use. They provide guidance for manufacturing, testing, and quality assurance in order to ensure that the product is safe for human use. One important aspect in following GMP is that the manufacture is well documented. In the United States, GMPs are issued and enforced by the U.S. Food and Drug Administration. GMPs, along with good laboratory practices and good clinical practices, are overseen by regulatory bodies in many countries.

Manufacture of Metals for Biomedical Applications Chemical Metallurgy Metallic biomaterials include noble metals (Au, Pt, etc.), 316L stainless steel, Co–Cr alloys, Ti and Ti alloys, and other biocompatible metals. They are essential for implants and devices in many medical fields (orthopedics, dentistry, cardiovascular, etc.). Metals exist in combination with other elements in the Earth as minerals (compounds) and need to be extracted from ores, which are dug out in mines in different geographical regions of the world. Metallurgical principles in extractive metallurgy are followed to remove a metal from an ore and refine the extracted raw metal into a purer form. Metal alloys are made by adding other elements into pure metals, resulting in materials with improved or desired properties. For example, to make 316L stainless steel with excellent corrosion resistance for orthopedic implants, iron is alloyed with specific amounts of Cr, Ni, Mo and other elements while keeping the carbon content below a limit. These processes (extraction, alloying, etc.) involve chemical reactions and this branch of science is termed chemical metallurgy. (The other branch is physical metallurgy, which deals mainly with mechanical, thermal, electrical, magnetic and other properties of metals in the solid state.) Both physical and chemical processes including mining, extraction, melting, re-melting, alloying and controlled solidification are used to make metal-containing ore to raw metal products in the bulk form such as ingots, from which stock metal shapes (powders, wires, plates, rods, tubes, etc.) are subsequently produced. The stock materials are purchased and used by medical device manufacturers to fabricate various types of products for biomedical applications. Metallurgy is distinguished from metalworking although metalworking relies on metallurgy. Metalworking processes make stock materials into useable shapes and sizes, and they include casting, forging, rolling, sintering, 3D printing, etc.

Metal Fabrication Techniques: Forming Operations Metals and alloys are made into products of different shapes (plates, rods, tubes, etc.) with desired properties by different metal fabrication techniques. These techniques include metalworking operations (e.g., forging), powder metallurgy, welding, etc. Practically, the reason for choosing a particular fabrication technique depends on several major factors such as properties of the metal, the size and shape of the finished products, and the cost. Major metalworking operations, that is, forging, rolling, extrusion and

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(C) (A)

Billet

Force

Die

Die Ram

Pre-heated metal blank

Extruded product Container Force (D)

(B)

Roll Die Tensile force

Roll Fig. 1

Schematic diagrams showing four metal forming operations: (A) forging, (B) rolling, (C) extrusion, and (D) drawing.

drawing, are generally used to form metal pieces whose shapes can be formed by large plastic deformation. The four operations are depicted in Fig. 1. Forging utilizes large plastic deformation of metals to form objects and is normally conducted at an elevated temperature. A high force is applied to the hot metal which deforms and conforms to the internal geometry of the die. Rolling can be performed at either room temperature (“cold rolling”) or an elevated temperature (“hot rolling”). Hot rolling can have the advantage of obtaining finegrain structures for the products. Metal sheets and thin films are commonly produced by rolling. In extrusion, a metal bar or rod is pushed through a die orifice by a high force acting on the ram. Metal rods and tubes with complicated cross-sectional geometries can be produced by extrusion. Drawing uses tensile force to pull a material through a die orifice. Thin metal wires can be made through drawing.

Metal Fabrication Techniques: Casting Casting is the manufacturing process of pouring fully molten metal into a mold cavity, with the liquid metal undergoing solidification into the desired shape. There are a number of casting techniques, including sand casting (e.g., for magnesium alloys), investment casting (e.g., for stainless steel alloys), die casting (e.g., for zinc alloys), and continuous casting (e.g., for aluminium alloys). Casting is extensively used for fabricating metal products with complex shapes which are difficult or uneconomical to make by other methods. Sand casting, which is the simplest and most common type of casting, allows for fabricating metal products with low tooling and equipment costs. The sand, either moist or dry, is mixed and bonded together with chemical binders and/or polymerized oils to form a sand mold with internal cavity in the shape and dimensions of the finished metal product. The metal, such as Ti alloys, is melted first and then poured into the cavity of the sand mold. Subsequently, the sand mold is separated and the solidified casting is removed. As the sand mold needs to be destroyed in order to obtain the metal product, this fabrication technique has a low production rate. Sand casting is extensively used for the fabrication of metals with complex structures (e.g., engine blocks of automobiles) or some small metal parts such as gears and pulleys. Another important technique is investment casting (also known as “lost-wax casting”). In this process, a wax pattern with full design specifications of the product is firstly made using a metal mold. The wax patterns are then assembled and coated with a refractory ceramic material (e.g., Al2O3). De-waxing is conducted to remove wax from the assembly, and the molten metal is poured into the cavity of the assembly and undergoes solidification. The ceramic shell is then knocked down and individual castings are thus produced. The whole process of investment casting is complicated as compared to simple sand casting and hence it is relatively expensive. This technique is suitable for the fabrication of products with intricate shapes for almost all alloys, for example, dental fixtures. Die casting is a fabrication process that requires pressure to force molten metal into the cavity of a mold. In this process, the steel mold is not destroyed. Die casting can be classified into different groups and die casting with high pressure is most widely used. The die casting technique is simple but the equipment, dies and other related components for die casting are expensive. It is suitable for producing very large quantities of castings with small or medium sizes for non-ferrous metals (zinc alloys, tin alloys, etc.). In continuous casting, a molten metal is poured into a mold and the semi-finished product continuously solidifies while new molten metal is constantly poured into the mold. Thus a continuous length of metals is produced as the casting keeps going

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downward. Continuous casting is time-saving as it can be fully automated. It is extensively used for the production of steel, copper or aluminium in industries.

Other Metal Fabrication Techniques Beside the metal fabrication techniques described above, several other methods such as power metallurgy and welding are also well developed. Powder metallurgy produces metal products from metal powders in basically three steps: powder blending, die compaction at room temperature, and sintering. It can be used to fabricate metal products that cannot be made through melting or formed by other methods. For example, porous Ti alloys that can be used for orthopedic implants can be made via powder metallurgy. Welding is a fabrication technique that joins metal parts by causing fusion through heat alone or together with pressure. In conventional welding, an electric arc between the metal and an electrode is created and the metals are melted in the welding area. After solidification, metal parts are joined. Shielded metal arc welding and gas metal arc welding are commonly welding techniques. Welding can be used to make joints between dissimilar metals (e.g., titanium-stainless steel dissimilar material pair) in implantable medical devices such as stents and hip replacement implants.

Physical Metallurgy Physical metallurgy is based on the fundamentals and applications of the theory of phase transformations in metals and employs heat treatment to alter microstructures of metals for achieving desired mechanical, electrical, magnetic, thermal and other properties. The words “heat treatment” are general and indicate that a cycle of heating and cooling is performed on a metal to change its property (which is caused by the microstructural change(s) in the metal). Commonly employed heat treatment techniques are annealing, quenching, tempering, normalizing, case hardening, precipitation strengthening, etc. In annealing, a metal is heated up at a pre-determined heating rate to a pre-set high temperature, dwells at this temperature for a required duration, and is finally cooled down (or quenched) at a desired cooling rate. Metals after heavy “cold work”, for example, forging at room temperature and cold rolling, have high strength but low ductility. They require heat treatment for stress relief and recovery of ductility and hence partial annealing or full annealing is performed to soften the metal. During annealing, microstructural changes (producing “recovery,” and “recrystallization”) occur in coldworked metal, resulting in mechanical property changes. The yield strength of metals depends on the average grain size (the “Hall-Petch equation”). A heat treatment above a certain temperature can increase the grain sizes of a metal (“grain growth law”) and hence change the mechanical properties of the metal. A heat treatment can also change phase(s) in metals. For example, two phases, a (ferrite, BCC crystal structure) and g (austenite, FCC crystal structure), can exist in steels. a is a soft phase and g phase has high strength. 316L stainless steel with g phase is used for making orthopedic implants. For steels, subcritical annealing can be performed at 540–650 C where there is no change of crystal structure. Intermediate annealing can take place at 650–760 C and some transformation to austenite occurs. Full annealing for complete austenitization may be conducted at 810–930 C.

Machining of Metallic Biomaterials Machining is a process in which a piece of metal (sometimes, “hard” polymer; very rarely, bioceramic) is made into desired shape and dimensions through controlled material removal by using a machine. Machining including turning, milling, drilling, etc., and cutting is the most common method. The machines must have cutting tools made of very hard materials for performing machining operations. Metallic biomaterials are often machined into final products. However, care must be taken in machining certain biomaterials, for example, Ti-based alloys, due to characteristics of these materials. Different machining conditions based on the types of cooling process and metal cutting parameters are used for metallic biomaterials, and the machining techniques are generally dry machining, wet machining, minimum quantity lubricant (MQL) machining, cryogenic machining, high speed machining (HSM), and air cooling machining. Among all these, wet machining is the most widely used material removal process for metals with the advantages of high efficiency and low production time and cost. The surface properties of metals, cutting tools and cutting fluids play important roles in the machining process of metallic biomaterials.

Manufacture of Porous Metals Porous metals were firstly attempted in 1943 when pores were intentionally introduced into aluminium by adding mercury into the molten aluminium. The concept of using porous metals in biomedical applications was investigated much later, and the earliest work was conducted in 1972 when employing porous metals for osseointegration was studied. There have been continuous efforts for using porous Ti alloys in orthopedics. In recent years, developing porous metallic scaffolds for regenerating human hard tissues is making good progress, which involves materials such as Ti alloys and Mg alloys. At present, a number of techniques are used to fabricate porous metals, which include powder metallurgy, space holder method, plasma spraying, and additive manufacturing technologies.

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Manufacture of Biomedical Polymers A polymer is a long chain macromolecule composed of repeating subunits termed “monomers.” The monomers are connected by covalent bonds in polymers. The term “polymer” is derived from ancient Greek words poly and meres, which mean “many” and “parts,” respectively. Polymers play an important role in daily life owing to their broad range of properties, availability and low cost. Polymers have been widely used for many types of biomedical devices (e.g., dental, orthopedic, and cardiovascular implants). Knowledge in polymer synthesis and product manufacture is essential in the biomedical field.

Polymer Classification and Architecture Polymers can be classified in a number of ways. They can be classified according to the origin of materials: natural or synthetic. Natural polymers such as RNA, DNA, proteins and polysaccharides have existed since life began, and they play critical roles in all forms of life. In the history of mankind, since the earliest times, people have been exploiting natural polymers for clothing, writing materials, weapons, etc. Natural polymers can possess either relatively simple structures, such as natural rubber, cellulose and starch, or very complex structures, including enzymes, proteins and nucleic acids. Synthetic polymers appeared much later as compared with natural polymers but have been playing crucial roles in the modern world. In 1909, the first synthetic polymer phenol-formaldehyde (“Bakelite”) was invented, which was soon followed by many other synthetic polymers. Polymers can also be classified into two groups, thermoplastic polymers and thermosetting polymers, based on their thermal response properties. Thermoplastic polymers can be thermally softened or plasticized repeatedly. Polymers such as polyolefins, nylons and linear polyesters belong to this group. For thermosetting polymers, in the product manufacture process, chemical changes occur upon heating of this type of polymers and convert them into an infusible mass. The curing or setting process leads to growth and cross-linking of chain molecules, producing giant molecules. After product manufacture, thermosetting polymers cannot be re-melted. Polymers including resins, urea, diene rubbers and phenolic are thermosetting polymers. On the basis of polymer chain structure, polymers can be classified as linear, branched, cross-linked, and network. Fig. 2 depicts the structures of these polymers. In linear polymers, the monomers are joined together in a linear manner, while in branched polymers, some monomers are joined as branches on the polymer backbone. If the monomer units are joined in multiple chains and form interconnections between chains, cross-linked polymers are made. When a cross-linked polymer includes plentiful interconnections between chains in 3D, a network polymer is formed. The interlinking ability of a monomer, or the number of available sites for bonding with other monomers during polymerization, is termed the functionality of a monomer. A molecule can therefore be monofunctional (one site), bifunctional (two sites), or polyfunctional (more than two sites). Monomers with bifunctional structural units can only have two linkages with other monomers. Thus the array of linkages between bifunctional monomers results in a linear polymer. Polyfunctional monomers can be linked together to form nonlinear structures. If the side growth during polymerization is terminated before the polymer chain can link up with another polymer chain, a branched polymer is formed. If the growing polymer chains of polyfunctional monomers can be linked with each other, cross-linked polymers or network polymers are produced. For biomedical applications, natural polymers such as collagen, gelatin, chitosan, hyaluronic acid and alginate are used. Synthetic polymers such as poly(methyl methacrylate) (PMMA), polytetrafluoroethylene (PTFE) and ultra-high molecular weight polyethylene (UHMWPE) are widely used for various implants and medical devices. Biodegradable polymers such as poly(lactic acid) (PLA), poly(lactic-co-glycolic acid) (PLGA) and poly(ε-caprolactone) (PCL) are extensively used for tissue engineering applications.

(A)

(B)

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(D)

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Structures of polymer chains.

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Polymer Synthesis Two main types of polymerization are used in polymer synthesis: step-growth polymerization and addition polymerization. Ringopening polymerization is also used for some polymers. If polymerization involves multiple types of monomers, copolymerization takes place, forming copolymers.

Step-growth polymerization Step-growth polymerization, or condensation polymerization, is commonly employed to produce polyesters and nylons. The polymerization reaction occurs between the functional groups of molecules. Small molecules such as water are eliminated by the chemical reaction in step-growth polymerization. Below is an example for making nylons: R  NH2 þ R 0 COOH /R 0 CONHR þ H2 O

(1)

In step-growth polymerization, one or more types of monomers can be involved, and each monomer should have at least two sites for bonding. For polymerization with more than one type of monomers, for example, involving A and B monomers, A–B stepgrowth polymerization or A–A/B–B step-growth polymerization can occur.

Addition polymerization Addition polymerization, or chain reaction polymerization, requires the monomers to have at least one double bond. In addition polymerization, no molecule is eliminated and no by-product is generated. The molecular weight of the formed polymer is exactly the same as the sum of all monomers included in the polymerization. A chain reaction links monomers together by rearranging the bonds with each monomer. For example, the reaction to synthesize ethylene is nfCH2 ¼ CH2 g/–CH2 –ðCH2 –CH2 Þ–CH2 –

(2)

The initiation of addition polymerization (in above case, the breaking of a double bond) requires catalysts, pressure, heat or UV light. The growing polymer chain is a free radical in the polymerization. Terminal radicals are required for ending the reaction. There are three steps in this type of polymerization: initiation, propagation, and termination. During initiation, the monomer acquires an active site to become a free radical. The addition of initiators or other approaches such as absorption of heat, light or irradiation can trigger the initiation process. In propagation, the initiated monomers add other monomers in rapid succession. This step continues until the active site, which is continuously relocated at the end of the growing chain during propagation, is deactivated by chain termination or chain transfer. The termination step involves the reaction of a polymer chain radical with another free radical. These three steps constitute the normal step-growth polymerization process. However, in many cases, a fourth step, namely, chain transfer, is also included. In chain transfer, the growing activity of a polymer chain is transferred to previously inert species.

Copolymerization Copolymerization is the polymerization using two or more types of monomers and produces copolymers. The composition of copolymers can vary widely, resulting in vastly different properties. Possible arrangements of monomers in different copolymers are shown in Fig. 3. In copolymerization, the number of reactions increases geometrically with the number of monomer types, that is, four reactions for two types of monomers. The composition and structure of copolymers are determined by the relative rates of different chain propagation reactions.

(A) Alternating (B) Random (C) Block

(D) Graft

(E) Random tripolymer

Monomer A

Fig. 3

Arrangements of monomers in copolymers.

Monomer B

Monomer C

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Amorphous Polymers and Semi-Crystalline Polymers The structure of polymers in the solid state can be either amorphous or semi-crystalline. (Totally crystalline polymers, e.g., polymer crystals, are only made for scientific research.) When polymers are forming the solid structure (either cooling from the molten state or concentrated from a solution), polymer molecules tend to aggregate together. For some polymers, some individual polymer chains are folded and packed in an ordered arrangement during this process while other polymer chains are not, resulting in semi-crystalline structures. Polymers with semi-crystalline structures are sometimes referred to as crystalline polymers. Since polymer chains are normally long, there are always amorphous regions in polymers. If individual chains of polymers cannot be folded and packed in an orderly fashion during the solid forming process, the chains are arranged randomly and even entangled, forming amorphous polymers. Typical examples of amorphous polymers include polycarbonate and poly(methyl methacrylate). Compared with amorphous polymers, semi-crystalline polymers tend to possess enhanced mechanical properties. Polyethylene and polyacrylonitrile are typical examples of semi-crystalline polymers.

Polymer Additives Additives are widely used to modify polymer behaviors. Only a few polymers are used in the chemically pure form in real applications. In most cases, additives are incorporated in polymers to adjust their physical or chemical properties. A major category of additives is stabilizer, which can enhance the stability of polymers against the environmentally degradative effects. Some polymers undergo rapid deterioration in mechanical properties under normal environmental conditions. The deteriorative effects are often caused by oxidation, thermal exposure or UV radiation. To prevent such deteriorative processes, antioxidants and thermal/light stabilizers are frequently incorporated in commercial polymer products. A major concern for most polymers is that they are flammable in the chemically pure form. To reduce the flammability of polymers, another type of additives, flame retardants, is used in polymer products. Flame retardants either initiate chemical reactions to cool the combustion region or interfere the combustion process to cease burning. Fillers and plasticizers are additives that are widely used to modify physical properties of polymers. Various fillers can be incorporated in polymeric materials to improve thermal stability, abrasion resistance, tensile or compressive strength, toughness, etc. Plasticizers are used to enhance the ductility, toughness, and flexibility of polymers while reducing their hardness and stiffness. Some polymers are difficult to process due to their high melt viscosity or adhesion with machine surfaces. Processing additives such as lubricants are therefore added to improve the manufacture of the products of these polymers. There are also many other types of additives, such as colorants (imparting specific colors to polymers), blowing agents (forming rigid or flexible foams), antistatic agents (reducing or eliminating static electricity), and biocide (control or destroy harmful organism).

Polymer Blends and Interpenetrating Networks Polymer blends are materials that are obtained by mixing two or more polymers. Such physical mixtures are made when the polymers are in the molten state or as polymer solutions. Polymer blends are produced to achieve desired properties, which can be tailored to meet specific application requirements. Another purpose of using polymer blends is to reduce the cost of expensive polymeric materials by introducing low-cost polymers in the blends. Polymer blends can be classified into two groups: compatible (miscible) polymer blends and incompatible (immiscible) polymer blends. Compatible polymer blend is a homogeneous (single-phase) system with a single value for a property. For example, the glass transition temperature (Tg) of a compatible polymer blend is generally the weighted average of the values of individual polymer components. Some compatible polymer blends possess superior properties than those of the individual components alone. This is a result of the high level of thermodynamic compatibility caused by strong intermolecular attraction. Incompatible polymer blends are phase-separated systems consisting of heterogeneous mixtures of polymer components. Such system has multiple values for a property. For example, multiple glass transition temperatures related with each individual components of the blend exist for these polymer blends. Factors that decide whether a particular polymer blend is a single phase or phase-separated system include thermodynamic of the system, kinetics of the mixing process, processing temperature, solvent, and incorporation of additives. The primary factor to determine miscibility of two polymers is thermodynamics which is governed by Gibbs free energy. Another approach for developing new polymeric materials with desirable properties from multiple polymer components is to create interpenetrating networks (IPNs), which are the combinations of two or more polymers in the network form. At least one polymer in an IPN is crosslinked and/or is synthesized in the presence of the other. Simply mixing two or multiple polymer components does not create IPNs. A schematic diagram for IPN structures is shown in Fig. 4. IPNs have advantages of both network polymers and polymer blends. Polymers include polystyrene, poly(ethyl acrylate), polyurethanes, and poly(methyl methacrylate) has been used to form IPNs. IPNs can be classified as several types: sequential IPN, simultaneous IPN (SIN), latex IPN, gradient IPN, thermoplastic IPN, and semi-IPN. Sequential IPN is the case that polymer networks are formed one by one. One polymer network is fabricated first, and then the monomers and their crosslinker and activator of the other polymer network are swollen into the first polymer network for in situ polymerization, forming the final IPN. In contrast, SIN is produced by simultaneous polymerization of different monomers

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Fig. 4

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A schematic diagram showing the structure of polymer interpenetrating networks.

and/or prepolymers that are mixed together before reactions. Latex IPN has a structure similar with latexes, with each particle constituting a micro-IPN. Gradient IPN is the case that the crosslinking density or overall composition varies geometrically on the macroscopic level. A general process to prepare gradient IPN is to partially swell the initially formed polymer network by the other type of monomers, followed by rapid polymerization before the larger-scale diffusion occurs. Thermoplastic IPNs are the polymer networks that are produced mainly by physical crosslinks. The physical crosslinks make thermoplastic IPNs flowable at high temperatures. Semi-IPN is the case that one or more polymers are crosslinked in the IPN, while the rest of polymers in the IPN is not crosslinked. IPNs have been investigated for many industrial applications including food packaging and fuel cells. They have high potential for applications in tissue engineering and controlled drug delivery.

Polymer Forming Techniques Various techniques have been employed in polymer forming and new methods are continuously developed. The selection of a processing method for a polymer depends on the following factors. The first factor is whether the material is thermoplastic or thermosetting. For thermoplastic polymers, the use of a processing method also depends on the temperature at which the polymer softens, the atmospheric stability of the polymer, and the geometry and size of the finished product. For thermosetting polymers, the first step is the preparation of a linear polymer, which is normally in liquid state due to its low molecular weight. The linear polymer is then cured in a mold, forming polymer product.

Compression molding and transfer molding Compression molding and transfer molding are widely used for processing thermosetting polymers. A schematic diagram of compression molding is shown as Fig. 5A. Compression molding can be used to fabricate both thermoplastics and thermosets and requires heating up of raw materials (in the form of powders or pellets) in the mold. For thermoplastics, the temperature should be higher than Tg during the forming process. A pressure is applied in compression molding and should be maintained during cooling. For compression molding, an appropriate amount of the polymer and necessary additives are thoroughly mixed and loaded between upper and lower parts of the mold. Both parts are heated up and only one is movable. After the mold is closed, heat and pressure are applied. The polymer in the mold is melted and flows into all space of the mold cavity. After a pre-set time, the mold is cooled and opened up to eject the product with desired shape. A major advantage of compression molding is the low capital cost due to the simplicity of the process. Transfer molding is a variation of compression molding for processing thermosetting polymers. In this forming process, the solid ingredients are initially melted in a heating chamber. The melted viscous polymer is then injected into the mold cavity and a pressure is applied. Transfer molding can process thermosets with complex parts.

Injection molding Injection molding of polymers is similar to die casting for metals. It converts polymers from the power or pellet form to various polymer products, such as radio cabinets, computer frames and many other commodities. It is usually used for processing thermoplastic polymers. A schematic diagram of this technique is shown as Fig. 5B. Generally, polymer pellets or powders are fed into a heating chamber and heated up until they melt. The melted viscous liquid is then injected into an enclosed mold under pressure and cooled for solidification. After the polymer becomes solid, the mold is opened and the product is ejected. Injection molding is a continuous process and can be automated. The production rate is high and hence it is used for the mass production of polymer products.

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(A) Compression moulding

Mould: heated & cooled

Mould plunger Guide pins

Moulding compounds

Mould cavity

Mould Close

Mould Open (B) Injection moulding

Hopper Heating zone Nozzle

Rotating screw

Mould Fig. 5

Schematic diagrams showing compression molding and injection molding processes for polymers.

Thermosetting polymers can also be processed through injection molding, employing what is called reaction injection molding. Curing reaction of thermosets takes place while the material is under pressure in the heated mold. Longer cycle time as compared with injection molding for thermoplastic polymers is required in reaction injection molding.

Extrusion Extrusion is for converting granules or pellets of thermoplastic materials into solids by simply injection molding through an openended die. Polymeric materials with continuous lengths and constant cross-sectional geometries can be produced through extrusion. The schematic illustration of polymer extrusion is similar to that of extrusion for metal alloys (Fig. 1C). A polymer in the form of granules or pellets is fed into an extruder and melted. The molten polymer is then forced through a die with an orifice of designed shape, which enters a sizer and cooler to develop the correct size and shape. The solid polymer is pulled by motordriven rolls and transferred to a cutter to make the final product (pellets, wires, rods, etc.).

Blow molding Blow molding is frequently used to fabricate plastic containers with thin walls. The process is similar to blowing glass bottles. Blow molding involves two essential stages: forming polymer parison, and blowing the parison into the desired hollow shape. A thermoplastic is melted first and forms the parison, which in most cases is hollow tubes, and a compressed air is blown into the soft molten parison to expand it until it reaches the contours of the mold. The product is then cooled and removed from the mold.

Casting Casting is a technique which pours a liquid material into a mold and then solidifies the material to form a rigid product that reproduces the detailed shape and dimensions of the mold cavity. In casting, the mold is not heated. The liquid for casting can be molten polymers or polymer solutions. Solidification of the material in the mold is through a physical (cooling) or chemical (polymerization) process. This technique can be used for both thermoplastic and thermosetting polymers. Solvent-casting is often used in the laboratory for developing new materials and products. In this process, polymer is dissolved by a solvent to form a polymer solution. The polymer solution is then poured into a mold. The solvent evaporates from the solution in the mold in a controlled environment and a solid polymer product is obtained.

Fabrication of Hydrogels Hydrogels, also known as gels, are crosslinked three-dimensional macromolecular networks obtained from hydrophilic polymers which can absorb and retain a considerable amount of water. Because of its favorable biocompatibility and other advantages, hydrogels have been intensively investigated in the biomedical field and find applications such as soft contact lenses and drug delivery vehicles. Several fabrication technologies have been developed to produce hydrogels, including physical crosslinking, chemical crosslinking, radiation crosslinking, and grafting polymerization. Only two fundamental methodsdphysical crosslinking and chemical crosslinkingdare briefly covered here.

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In physical crosslinking, no cross-linking agents are included in the reaction. This feature gives advantages to the physical crosslinking approach since using cross-linking agents can affect the integrity of substances to be entrapped. In certain applications, the cross-linking agents even need to be removed before application. The formation of physically crosslinked hydrogels is largely dependent on hydrocolloid type, concentration, pH, salt type, temperature, and other thermodynamic parameters. With careful selection of these parameters, a broad range of hydrogel textures and properties can be obtained. Many methods have been developed for physical crosslinking, such as heating/cooling a polymer solution, H-bonding, complex coacervation, ionic interaction, freezingthawing, etc. In chemical crosslinking, hydrogels are formed through covalent bonding. Hydrogels fabricated through chemical crosslinking cannot be dissolved in solvents unless the covalent bonds formed in the crosslinking process are cleaved. Chemical crosslinking includes three major approaches: (1) crosslinking of water-soluble polymers with irradiation, (2) crosslinking of water-soluble polymers with cross-linker, (3) copolymerization of a monomer with cross-linker. For these three approaches, many crosslinking methods have been explored, such as chemical grafting, radiation grafting, aqueous state radiation, solid state radiation, radiation in paste, etc.

Manufacture of Porous Polymers With the emergence of tissue engineering, making porous polymer structures becomes highly important. Porous polymers are also very useful in other biomedical applications, such as filtration and drug controlled release. Many methods are available for fabricating porous polymers, including gas foaming, solvent casting and porogen leaching, phase separation, templating via selfassembled structures, electrospinning, and 3D printing. With the increasing demand for porous polymers with specifically required characteristics/properties, new manufacture techniques are continuously investigated and developed.

Manufacture of Bioceramics Fabrication of Dense Bioceramics Bioceramics are biocompatible ceramic or glassy materials that are designed to repair or reconstruct the damaged parts of human bodies. Bioceramics can be fabricated into a variety of forms (e.g., powder, coating, and bulk) to provide different functions in human tissue repair or replacement. They include ceramics, glasses, glass-ceramics, and composites. Bioceramics can be mainly divided into three groups, as illustrated by Fig. 6. Quite a few members in the calcium phosphate (Ca–P) family are attractive ceramic biomaterials, such as hydroxyapatite (HA, Ca10(PO4)6(OH)2) and b-tricalcium phosphate (b-TCP, Ca3(PO4)2). Both HA and TCP are weak bioceramics and cannot be used directly for load-bearing applications. For many applications, bulk, non-porous bioceramic products are required. Bioceramic products start with the synthesis of bioceramic powders. For example, for HA, several methods can be used for making its powders, with wet synthesis being the most common technique utilized. For producing bars, rods or other types of products, bioceramic powders or powder mixtures of desired composition together with (or without) chemical additives are pressed into desired shape and size in dies using a hydraulic press to form powder compacts, known as the greenbody, which are not physically strong. The greenbody is then sintered at a high

Bioceramics

Bioinert ceramics

Fig. 6

Bioactive ceramics, glasses and glassceramics

Bioresorbable ceramics

Carbon

Hydroxyapatite

Tricalcium phosphate

Alumina

Bioglass®

Calcium sulphate

Zirconia ceramics

A-W glassceramic

Others

Others

Others

Classification of bioceramics according to their bioactivity and biodegradability.

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(A)

MO

2

SiO2

Si

O

2

SiO2

O

2

Si O

Initial stage

MO

SiO2

SiO2

MO

SiO2

MO

MO

MO

SiO2

SiO2

Si

SiO 2

SiO2

MO

MO

MO

Intermediate stage

Final stage

(B)

Al2O3

Initial stage Fig. 7

Al 2O 3

Al2O3 Al2O3

Intermediate stage

Al2O3

Al 2O 3 Al 2O 3

Al 2O 3 Al2O3

2O 3

2O 3

Al

Al2O3

Al 2O 3

2O 3

Al 2O 3

Al2O3

Al 2O 3

Al

Al 2O 3

Al2O3

Al2O3

Al

Al2O3

Final stage

Common ceramic processing technologies for making dense bioceramics: (A) liquid-phase sintering, and (B) solid-state sintering.

temperature to bond particles to form dense bioceramic products. The purpose of sintering is to use thermal energy to form chemical bonds between bioceramic particles, resulting in strong, bulk bioceramic products. There are two important conventional methods for producing bulk bioceramics, liquid-phase sintering and solid-state sintering, as illustrated in Fig. 7. Liquid-phase sintering uses a small amount of an additive (e.g., a glass, which must be biocompatible) which melts below the bioceramic melting or decomposition point. The molten additive can flow and enter the space between bioceramic particles, making dense, almost nonporous bioceramic products. Liquid-phase sintering has the advantage of sintering ceramics at relatively low temperatures, thus avoiding bioceramic decomposition, and of producing highly dense (non-porous) bioceramics and hence has been used to make bioceramics such as glass-toughened HA, which is a composite. Solid-state sintering can be conducted for both singlecomponent system and multi-component system bioceramics without using a liquid phase in sintering. It is used for making alumina or zirconia ceramic products. Solid-state sintering normally requires relatively high temperatures for sintering and may not be able to totally eliminate pores in the sintered products. Both liquid-phase sintering and solid-state sintering can be performed without the application of pressure. Hot pressing is a manufacturing method that combines the forming and sintering steps to produce bulk bioceramics with relatively simple shapes. The ceramic powder is put into a die and subjected to sintering with the simultaneous application of uniaxial pressure. The sintering temperature can be lowered due to the application of pressure, which helps to avoid decomposition of bioceramics at elevated temperatures. Hot isostatic pressing (HIPing) uses uniform pressure from all directions, instead of uniaxial pressure, to sinter bioceramics into both simple and complex shapes. The bioceramic powder in a container is put into a high pressure vessel and is subjected to both high temperature and isostatic gas pressure for sintering. The most widely used pressurizing gas is argon, an inert gas that does not cause chemical reaction. When the chamber is heated, the pressure inside the vessel increases. HIPing can use still lower temperature for sintering the bioceramic and the product can be almost non-porous. However, the HIPing equipment is expensive.

Manufacture of Porous Bioceramics Porous bioceramics can be used for non-load bearing implants in orthopedic and craniofacial applications. The porous structure can allow the tissue to infiltrate, which enhances the attachment of implant and the tissue. To introduce pores into the bioceramic bodies, a variety of methods are developed, such as incorporating pore-creating additives (e.g., amino-acid derivatives), gel-casting, freeze casting, direct foaming, 3D printing, etc. Porous calcium phosphate bioceramics are extensively studied as bone scaffolds for bone tissue engineering due to their excellent biocompatibility and bioactivity and suitable degradation behavior. When calcium phosphate bioceramics have open porous structures with three-dimensional interconnected pores, cells and biomolecules such as growth factors can penetrate into them, resulting in better bone regeneration. However, one of the shortcomings that needs to be considered is the severe reduction in

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mechanical properties when pores are introduced into bulk bioceramics, especially at the high porosity levels. Thus exploring ways to achieve suitable mechanical properties of porous bioceramics is highly important. Compared with conventional techniques for producing porous bioceramics with a random porous structure, 3D printing of porous bioceramics, which has distinctive advantages in several aspects, now gains increasing attention.

Fabrication of Pyrolytic Carbon Pyrolytic carbon is a synthetic biomaterial and the term was first introduced in 1960s. Generally, pyrolytic carbon is produced through thermal deposition of hydrocarbon compounds at its decomposition temperature without the presence of oxygen. Pyrolytic carbon can have different microstructures (isotropic or lamellar) depending on the fabrication condition (temperature, concentration of carbon precursor, gas flow rate, etc.). In the majority of biomedical applications, pyrolytic carbon is used as a versatile coating due to its good blood and tissue biocompatibility. It can be deposited onto synthetic blood vessels to form a thin layer that has excellent blood compatibility without affecting the flexibility of the grafts. It is also used in mechanical heart valves. The most important forms of pyrolytic carbon for biomedical applications are low-temperature isotropic (LTI) pyrolytic carbon and ultra low-temperature isotropic (ULTI) pyrolytic carbon. Dense and high strength LTI pyrolytic carbon is generally deposited onto implants via chemical vapor deposition (CVD) from a hydrocarbon gas at a high temperature in a fluidized bed. The ULTI pyrolytic carbon is usually deposited from a carbon precursor via a hybrid vacuum process in the presence of a catalyst.

Fabrication of Bioactive Glasses The excellent biocompatibility and bioactivity of bioactive glasses make them excellent materials for implants for repairing or reconstructing damaged bones. Bioactive glasses are also promising materials for porous scaffolds for bone tissue regeneration owing to their high osteogenic potential and in recent years, the controlled biodegradation of new bioactive glasses. Bioactive glasses were first discovered by Hench and his associates in Florida, United States, in 1969. Comparing with other bioactive materials, one of the most distinctive characteristics of bioactive glasses is the formation of an HA-like layer between the glass implant and host tissue, resulting in a firm and stable bonding which has evolved from a series of timedependent surface reactions after implantation. Currently, the most common manufacturing techniques of bioactive glasses are conventional melt quenching and the sol–gel method. The flowcharts in Fig. 8 describe the processes of these two methods. There are various bioactive glasses developed for biomedical applications. BioglassÒ was invented by Hench and is a family of bioactive glasses that contain SiO2, Na2O, CaO, and P2O5 in specific proportions. Different from traditional soda-lime-silica glasses, BioglassÒ has low SiO2 content (less than 60 mol%), high Na2O and CaO contents and a high CaO/P2O5 ratio. Hence BioglassÒ is a weak bioceramic and cannot be used for load-bearing applications. BioglassÒ can be fabricated using the conventional glass-making technique via melt quenching. Certain amounts of weighted raw materials including SiO2, Na2CO3, CaCO3, and Ca2P2O7 are thoroughly mixed first, followed by melting at 1300–1450 C and annealing at 450–550 C. Impurities such as Al2O3 must be avoided. Bulk BioglassÒ can be formed by casting or molding using graphite or steel molds.

(A) Melt quenching

method

Raw materials

(B) Sol-gel

method

Raw materials

Mixing

Preparation of sol

Melting

Gelation

Casting or injection moulding

Aging Drying

Annealing Stabilization Bioactive glass

Sintering Bioactive glass nanoparticles

Fig. 8

Production processes for bioactive glasses: (A) melt quenching technique, and (B) sol–gel method.

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Fabrication of Bioactive Glass-Ceramics Bioactive glass-ceramics are crystal-glass nanocomposites which have one or more crystalline phases with controlled size and morphology in a glass matrix and are made through controlled crystallization of glasses. Consequently, they possess the advantages of glasses as well as specific properties of sintered crystalline ceramics. They can be stronger and tougher than the parent glass or glass matrix owing to the reinforcement and toughening by the small crystals. Various types of glass-bioceramics have been developed after the invention of BioglassÒ, and CeraboneÒ A-W glass-ceramic which contains apatite and b-wollastonite (CaO$SiO) crystals was invented Kokubo and his co-workers in Japan in the early 1980s. A-W glass-ceramic is considered the most outstanding bioactive glass-ceramics for hard tissue repair. A-W glass-ceramic (with the commercial brand name of CeraboneÒ) is manufactured from a parent glass in a pseudoternary system 3CaO $ P2O5  CaO $ SiO2  MgO $ CaO $ 2SiO2 and has the contents of 38 wt% apatite, 34 wt% wollastonite and 28 wt% residual glass. The parent glass is prepared using the conventional melt-quenching method, and nano-sized oxyfluoroapatite crystals and b-wollastonite crystals are precipitated during controlled crystallization of the parent glass at 870 C and 900 C and are homogenously distributed in the glass matrix to form the homogeneous bioactive glass-ceramic. The material is slowly cooled from high temperature to avoid cracking. A-W glass-ceramic possesses excellent mechanical properties owing to the reinforcing effect of wollastonite as well as good bioactivity rendered by apatite. These characteristics make A-W glass-ceramic implants very successful for bone tissue repair.

Manufacture of Biomedical Composites Composite and Composite Classification Composites are heterogeneous materials consisting of two or more homogeneous phases which are separated by interface(s). The distinctive constituent phases are bonded together by chemical or physical bonds at the interface. Unlike traditional materials, the physical, chemical, biological and other properties of composites can be tailored according to the requirements in specific applications. This is a distinctive and highly important advantage for this group of materials and is greatly enhanced by carefully modulating the composition, interfacial bonding and physical arrangement of constituent phases in composites. This advantage should be utilized to the fullest extent when developing new materials for targeted applications. A composite is designed to combine the best properties of each constituent in the composite or to create new materials with properties that current materials cannot provide. In most cases, a composite consist of two distinct phases: the matrix phase and the reinforcing phase (or, “second phase” or “dispersed phase”). The matrix phase is a continuous phase that surrounds the other phase and provides the overall form. The reinforcing phase is a dispersed phase and is generally stronger and stiffer than the matrix phase. Many human body tissues are natural composites such as bone. Artificial composites have been investigated such as composites for dental filling and reinforced bone cement. In biomedical applications, each component in a composite must be biocompatible. If the composite degrades in the body, the degradation products must also be biocompatible. There are different ways to classify composite materials. One common way is to categorize composites according to the geometry of the reinforcing phase. Hence there are particulate composites, fibrous composites and laminated structures (Fig. 9). In each category, composites are further divided into different sub-groups. For example, fibrous composites can belong to either of the two groups: continuous fibers and short/chopped fibers, on the basis of the ratio of fiber length over its diameter (the so-called “aspect

Composite

Particle-reinforced

Fig. 9

Fiber-reinforced

Structural

Large particles

Continuous long fibers

Laminates

Dispersion strengthened

Short fibers / Chopped fibers

Sandwich panels

Others

Others

Others

Classification of composite materials according to the geometry of the reinforcing phase.

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ratio”). Continuous fibers generally have a ratio greater than 1  105, while aspect ratios of short/chopped fibers are between 5 and 200. Composites can also be categorized by the matrix material used. Therefore, there are metal matrix composites, ceramic matrix composites, and organic matrix composites. Among these three groups, polymer matrix composites are more commonly developed for biomedical applications. The trend now is to develop biodegradable biomedical composites for human tissue repair or regeneration.

Fabrication of Particulate Biomedical Composites Polymer matrix composites Particulate polymer-matrix composites are the largest group of biomedical composites. Using standard equipment in the plastics industry, they can be fabricated using one of the three methods: extrusion, injection molding, and compression molding. These three techniques have already been introduced in the previous section on the manufacture of biomedical polymers. Extrusion and injection molding are limited to the manufacture of composites with particulates or short/chopped fibers only. Injection molding is suitable for the mass production of particulate or short fiber composites. Compression molding can be used for particulate composites with either thermoplastic or thermoset polymer matrix.

Metal matrix composites Typically, metal-matrix composites for biomedical applications consist of a metal alloy matrix and a particulate ceramic second phase. Such composites are investigated primarily for having the strength and stiffness provided by the metal matrix which are required for load-bearing implants. The particulate second phase is normally a bioactive bioceramic. The techniques for fabricating particulate metal-matrix composites can be classified on the basis of the temperature used in the manufacturing process: (1) liquid phase process, (2) solid state process, and (3) solid–liquid process. Liquid phase processes involve the incorporation of dispersing the particulate second phase in a molten matrix metal, followed by the solidification of the composite. Stir casting can be applied in the liquid phase process. In solid state processes, both phases in the composites are kept at the solid state but the high temperature and pressure exerted provide the energy needed to form bulk composites. Powder metallurgy can be used in the solid state process. Solid–liquid processes involve the mixing of the particulate second phase and metal matrix in a region of the phase diagram where both liquid and solid phases exist in the matrix. If the temperature in all these processes is too high, chemical reaction(s) between the particulate second phase and metal matrix can occur, which may be detrimental to composite performance.

Ceramic matrix composites Ceramic-matrix composites for biomedical applications generally consist of a bioceramic matrix and particulate, bioceramic second phase. These composites can be made using processing technologies for bioceramics, which have been presented in the previous section. Pressureless sintering, hot pressing, and HIPing. The important issue in the manufacture of ceramic matrix composites is the densification of composites during manufacture. Efforts are required to make fully or highly dense (zero porosity or nearzero porosity) ceramic products, which is difficult to achieve.

Fabrication of Fibrous Biomedical Composites Short-fiber composites Two commonly used fabrication processes for particulate polymer matrix composites, namely, extrusion and injection molding, can also be used for the manufacture of short-fiber composites with polymer matrices. However, due to high shear forces generated in these manufacturing processes, fiber breakage is often encountered. The short fibers are also oriented in the extrusion or injection direction in the final products, which results in anisotropy of the material. Compression molding does not encounter these two problems but its production rate is low.

Long-fiber composites Producing long-fiber composites requires specialized equipment. Filament winding is a common process for fabricating long/ continuous-fiber composites. A schematic illustration of this process is shown as Fig. 10A. In this process, the long fibers are pulled through a low viscosity resin bath for polymer impregnation. Then the fibers are precisely wound onto a rotating mandrel. Successive layers are laid on the mandrel until product specification is reached. Subsequently, the fibrous composite on the mandrel is cured, and the product is removed from the mandrel. Filament winding is a good and efficient technique for manufacturing parts with rotational symmetry such as cylinders and tubes. This process can control fiber orientation and fiber content. The highly aligned fibers provide the final products with high tensile strength in the fiber direction. Filament winding has been investigated to make medical products such as prosthetic hip stems, arterial grafts, intervertebral disks, intramedullary rods for fracture fixation and ligament prostheses. Pultrusion (Fig. 10B) is another fabrication technique for many long-fiber composites. During this process, the long fibers are also pulled through a bath of liquid resin for impregnation. The material then goes through a heated die, which has a cavity of constant cross-sectional area along most of its length, for curing. After being cured in the heated die, the solid, hot composite is cooled and cut into required size. This process can be used for both thermoset and thermoplastic polymer matrix composites. In

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(A)

(B)

Reinforcing fibers Heated die Mandrel

Cutter

Reinforcing fibers

Resin bath

Pulling system

Resin bath Fig. 10 Schematic diagrams showing two common technologies for the production of long-fiber composites: (A) filament winding, and (B) pultrusion.

the case of thermoset composites, the feedstock is long fibers and the bath is filled with liquid thermosetting resin. For thermoplastic composites, the feedstock consists of both long fibers and thermoplastic polymer. The feedstock is preheated to a temperature near or above the softening point of the polymer. It then enters a heated die for melting and forming the composite with designed crosssectional shape and size. Pultrusion is widely used for the mass production of parts with a constant cross-section such as rods, pipes, beams, sheets, and tubes. A new technique that can be used for making long-fiber composite is electrospinning. Over the past 20 years, electrospinning is intensively investigated for fabricating porous, nanofibrous scaffolds for tissue engineering applications. With the proper modification of the basic electrospinning technique, composite scaffolds can be produced via electrospinning. Electrospun nanofibers themselves can actually be used as reinforcing fibers in non-porous long-fiber composites. There are reports on research in this direction.

Fabrication of Laminated Structures Laminated structures are composites composed of 2D sheets or panels which are stacked and bonded together with the orientation of the high-strength direction (the fiber direction) varying with each successive layer. The general fabrication process is to use autoclaves/ovens to apply heat and pressure for curing stacked 2D sheets or panels into final or near-final products. Various methods have been developed to manufacture laminated structures, such as vacuum bag-autoclave and resin transfer molding. Vacuum bag-autoclave is a process for fabricating high-performance laminates such as fiber-reinforced epoxy. In this process, a prepreg is first produced with a fundamental structure of a thin sheet. This sheet consists of a resin matrix embedded with uniaxially oriented long fibers. For some matrix material, such as epoxy, the matrix in the thin sheet of is partially cured. Pieces of the thin sheets are cut out and placed on top of each other according to the composite design. The stacked prepreg sheets are moved to a shaped tool to form a laminate. Those sheets can be placed in different directions to achieve various strength, stiffness and other properties. Afterwards, the laminate is applied with a vacuum through a vacuum pump for vacuum-bagging and the entrapped air in the laminate is removed. The laminate enclosed in the vacuum bag is placed in an autoclave for curing. When the curing process is completed, the product is removed from the autoclave and taken out from the vacuum bag for other operations.

Fabrication of Porous Composite Scaffolds for Tissue Engineering Porous polymer-, metal- or ceramic-matrix composites (the so-called “scaffolds”) are made and studied for potential tissue engineering applications. Polymer matrix composite scaffolds are far more common that composite scaffolds based on metals or ceramics. The techniques for producing polymer scaffolds, such as solvent casting and porogen leaching, phase separation, electrospinning and 3D printing, can be suitably modified for fabricating polymer matrix composite scaffolds that contain bioactive bioceramic particles. However, with the addition of bioceramic particles (micro- or nano-size) in the polymer solution for scaffold fabrication, technical difficulties may be encountered, such as homogeneous distribution of bioceramic particles in composite scaffolds and sufficient amounts of bioceramic particles in the scaffolds. Nevertheless, composite scaffolds hold great promise for the regeneration of certain types of human body tissues and new and better techniques should be developed for making novel nanocomposite scaffolds.

Manufacture of Nano-Biomaterials Fabrication of Nano-Structured Biomaterials Advances in nanoscience and nanotechnology have now made a huge impact in the biomedical field. Nano-biomaterials, either nano-structured biomaterials or nano-sized biomaterials, are increasingly investigated owing to their distinctive advantages, and sometime unmatched advantages, for biomedical applications. Nano-biomaterials can be broadly defined as biomaterials that have typical structural features or physical sizes of at least one dimension within the range of 1–100 nm.

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Crystalline metals and ceramics made via conventional manufacturing technologies have grain sizes normally in the micrometer range. Various technologies are available to change the grain size for achieving targeted mechanical properties but the grain size cannot be reduced to the nanometer range. With new nanotechnology, bulk materials with nano-sized grains can now be made, which provide outstanding mechanical properties. For example, nano-structured HA has been produced for bone tissue repair. For producing nano-structured HA, HA nanoparticles, instead of HA microparticles, are firstly synthesized. The HA nanoparticles are compacted to form the greenbody which are subsequently sintered in a carefully controlled process. In this HA sintering, grain growth is avoided and hence sintered dense HA has only nano-sized grains. This nano-structured HA possess much high strength than dense HA made through the conventional route. Furthermore, nano-structured HA exhibits much improved biological performance than conventional dense HA. For some Ti alloys and Mg alloys, there are reports that new technology can help to decrease their grain size to the sub-micron range, which significantly improves their mechanical properties. Another important area for developing nano-structured materials is nanocomposites. R&D in nanocomposites has progressed very fast over the past decade. Nanomaterials such as HA nanoparticles, metal nanoparticles, carbon nanotubes (CNT), and graphene have been incorporated into polymers for various applications. These nanocomposites can be either non-porous or porous. The fabrication techniques for nanocomposites are mostly based on manufacturing technologies which have been described in previous sections. Obviously, there are particular problems in dealing with nanomaterials for making nanocomposites. Extensive investigations are required to overcome obstacles in nanocomposite manufacture.

Fabrication of Nano-Sized Biomaterials Nano-sized biomaterials (nanoparticles, nanofibers, etc.) have shown great potential in various areas including drug delivery, cancer diagnosis and treatment, and tissue regeneration. These nano-sized biomaterials can have different shapes and geometries (nanospheres, nanorods, nanotubes, etc.) and can be inorganic (e.g., HA), metallic (e.g., Au), polymeric (e.g., PLGA) or composite/hybrid. There are already a great variety of techniques for synthesizing or fabricating nano-sized materials or biomaterials. It is expected that there will be more new technologies for producing nano-sized materials. For controlled drug delivery, polymer nanoparticles (polymeric micelles, nanocapsules and nanospheres) as drug nanocarrier are fabricated either from an already-synthesized polymer or by polymerization of monomer units. Drugs can be loaded into/ onto polymer nanoparticles by either physical entrapment or chemical conjugation and later released in vivo for providing the therapeutic function. Some biomolecules such as growth factors can also be incorporated in polymer nanoparticles for other functions. For example, using the double emulsion technique, spherical polymer particles can be produced and a sustained release of incorporated biomolecules can be obtained. Nano-sized bioceramic particles, particularly HA and TCP nanoparticles, with high surface areas can be used for different applications. In general, HA nanoparticles can be made using a number of techniques. Conventional wet chemical synthesis with subsequent oven-drying or freeze-drying can produce HA nanoparticles but with many technical problems. HA nanoparticles with very small sizes and a narrow size range can be successfully synthesized using the “water-in-oil” emulsion technique. Nano-sized metals have shown high potential as diagnostic or therapeutic tools. Magnetic iron oxide nanoparticles can be used for magnetic resonance imaging (MRI imaging), drug-delivery, and hyperthermia. They can be synthesized with controlled size, size distribution and shape by using several different methods, including co-precipitation from of iron salts in aqueous media, microemulsion, and thermal decomposition and high temperature reaction. Gold nanoparticles have also received great attention owing to their unique properties such as localized surface plasmon resonance, which leads to their multiple applications such as labeling, sensing, imaging and photothermal therapy. There are a number of ways to synthesize gold nanoparticles and the simplest ways is the reduction of metal salt precursors with a reducing agent (e.g., chitosan) under controlled conditions in either organic solvents or water. However, more sophisticated ways need to be used to produce gold nanoparticles with desired shape and structure (highly branched nanoparticle, nanorod, nano-cage, etc.). CNTs have gained popularity in research and they can be potentially used in the biomedical field. CNTs can be composed of a single tube or concentric cylinders of carbon, termed single-walled carbon nanotubes (SWNTs) or multi-walled carbon nanotubes (MWNTs), respectively. CNTs are generally prepared using three different methods: arc-discharge, laser ablation, and chemical vapor deposition. Due to their high strength and unique electrical and physical properties, CNTs may find applications in cell labeling and tracking, chemical and biological sensing, and bioactive agent delivery. Another important area for developing nano-size biomaterials is the fabrication of nanofibers. Nanofibers can be used to construct tissue engineering scaffolds which mimic the nanofibrous extracellular matrix (ECM) of human body tissues. Nanofibrous scaffolds can serve as temporary ECM for cells and influence cell behaviour. Various methods have been developed for producing nanofibers of polymers (natural or synthetic) and typical techniques include self-assembly, phase separation, electrospinning and newly developed 3D printing technologies, as illustrated in Fig. 11. Electrospinning is described in detail in another article of this encyclopedia.

Surface Modification Techniques for Biomaterials Surface properties are crucial for biomaterials since they are closely associated with the reaction of the host to implanted biomaterials. For many biomedical applications, surface modification of biomaterials is frequently required. Surface modification is

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(A) Self-assembly

(B) Phase separation

(2) Formation of a micelle in the initial phase

(1) Single peptide

Temperature change

Bioactive head Hydrophilic domain Hydrophobic tail (3) A cylindrical micelle

(C) Electrospinning

(D) 3D printing

Pressure

V Fig. 11

Technologies for producing nanofibrous biomaterials.

a process that changes the composition and structure of the surface of bulk materials to obtain improved surface properties while maintaining the mechanical properties of the bulk material. Surface modification can be applied to biomaterials to improve cell adhesion, cell proliferation, ECM-secretion function of cells, tissue compatibility, blood compatibility, corrosion resistance, and many other characteristics/behavior/properties for the material. Therefore, many surface modification techniques are developed for all types of biomaterials.

Surface Modification Techniques for Biomedical Polymers When biomedical polymers interact with the human body environment, a process similar to the interactions between cells and their environment takes place on the polymer surface. Such a process is initiated by protein adsorption on the polymer surface. Interactions between cells and the material then occur on this newly formed protein layer. With the understanding of this process, two general surface modification strategies are often adopted: 1. modulating the material surface so that the adsorbed proteins can preserve their normal bioactivity, and 2. immobilizing desired biomolecules directly on the polymer surface to incur specific cellular response. Another strategy, micro- or nano-patterning is also often used in surface engineering of biomedical polymers. In the first strategy, many surface properties can be adjusted to improve biomedical polymers. These properties include surface roughness or surface micro- or nano-features after patterning, surface stiffness, surface charge, chemical composition, hydrophilicity/hydrophobicity, and other properties. Among them, hydrophilicity/hydrophobicity has been widely studied to improve adsorption or desorption of proteins. Generally, surface with moderate wettability, instead of extreme hydrophilicity/hydrophobicity, can adsorb a proper amount of proteins and maintain the natural bioactivity of proteins. Since most biomedical polymers are hydrophobic, several technologies have been investigated to increase the hydrophilicity of these polymers. Plasma treatment of polymer surface is one of the widely used methods. It uses various reaction gases, such as air, NH3, CO2, SO2 or other gases, to introduce polarized groups such as amino, hydroxyl, carboxyl, and sulfate groups on polymer surfaces. Grafting copolymerization of hydrophilic polymers onto biomaterial surface is another popular method. The initiation of copolymerization (introduction of initiators) can be achieved using different approaches, such as laser treatment, ozone oxidation, electron beam or g-ray irradiation, and oxidation by cerium (IV). Among these initiation methods, Ce(IV) induced grafting polymerization do not require irradiation source, plasma or ozone generator, making it a convenient method. Photo-oxidization is another method that can introduce peroxide groups onto polymer surface without specialized equipment. Simply immersing materials in hydrogen peroxide solutions under UV irradiation also achieves biomaterial surface modification.

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In the protein immobilization strategy, two types of proteins can be employed: 1. adhesive proteins derived from ECM, which include fibronectin, collagen, laminin, vitronectin and others. Cell adhesion and attachment can be improved through immobilization of this type of proteins. 2. growth factors, such as epidermal growth factor (EGF) and bone morphogenetic protein (BMP). Cell proliferation, differentiation, and other cell behaviors can be modulated by these growth factors. Proteins can be either covalently or physically immobilized on biomedical polymers. For covalent immobilization, reactive groups such as amino, carboxyl and hydroxyl groups should be initially introduced as coupling sites. Grafting poly(acrylic acid) or poly(methacrylic acid) is a widely used method to generate carboxyl groups. Water soluble carbodiimide is then applied to activate the carboxyl groups. Finally, the activated carboxyl groups react with the amino groups of the target proteins, resulting in protein immobilization. Hydrolysis and aminolysis can be applied to yield reactive groups for polyesters. An example of the aminolysis method is the treating of ester-containing polymers in diamino compounds solution, with subsequent use of glutaraldehyde as coupling reagent to graft proteins. One potential drawback of covalent immobilization is the natural conformation of the grafted proteins may undergo certain changes. Therefore, physical protein immobilization techniques are developed. Layer-by-layer (LBL), grafting and coating method (covalent grafting a protein layer, followed by physical coating of proteins), and entrapment of biomacromolecules through surface swelling with subsequent solvent transfer are examples of physical immobilization technologies.

Surface Modification Techniques for Metallic Biomaterials Metallic biomaterials are usually covered by a metal oxide film, which plays a vital role in both corrosion resistance and tissue compatibility. When a metallic biomaterial is in contact with living tissues, proteins are immediately adsorbed on the material surface. The adsorption of proteins can affect material corrosion, subsequent cells adhesion, and other properties/events. Protein adsorption, cells adhesion, corrosion resistance, and other properties determine the tissue compatibility of implanted metallic biomaterials. To make metallic biomaterials suitable for the targeted applications, surface modification has been widely used to improve wear resistance, corrosion resistance, antibacterial property, and tissue compatibility of materials. Many surface modification techniques have been developed for metallic biomaterials. These technologies can be classified into two major categories: dry processes and hydro-processes. Dry processes usually require sophisticated equipment to conduct modifications of the material surface, while hydro-processes share a basic process of immersion or electrolysis in aqueous solutions without the need to use sophisticated equipment or incur high costs. Ion beam technology is a representative for dry processes. The surface modification can be categorized into two types: thin film formation, and surface-modified layer formation. Thin film formation has the advantage of controlling the film composition but lacks strong adhesion to the substrate. This technology is usually used for forming apatite or other type of films to improve wear resistance, corrosion resistance, and osteoconductivity. The other route, surface-modified layer formation, cannot control the composition of the surface layer as easily as thin film formation technologies. However, it has the advantage of the formation of a fracture-resistant interface between the surface layer and substrate. Surface-modified layers with nitrogen and calcium ions implanted can improve wear resistance and osteoconductivity of metallic biomaterials. Another example of dry processes is the industrialized process of plasma spraying of HA coating on orthopedic or dental implants. Hydro-processes are performed in aqueous solutions. By controlling certain parameters in hydro-processes, such as composition and pH of the aqueous solution, current density of electrolysis and potential gain due to electrolysis, specific properties of resultant surface layers can be achieved. Examples of hydro-processes include apatite formation on titanium through electrochemical treatment, formation of surface-modified layer for bone implants produced by immersion in various solutions (such as alkaline solution, hydrogen peroxide solution, and calcium-containing solution; some of them need thermal treatment at same time), titanium oxide layer formation, modification with biomolecules or polymers, etc.

Biomimetic Deposition In general, biomimetic deposition is a solution-based method conducted in an environment that mimics the human body condition. In most cases, such body-like environment is provided by a simulated body fluid (SBF) at 37 C. The temperature, pH, and other parameters of the conditions for biomimetic deposition are carefully controlled to simulate the body environment. Traditional surface modification techniques, such as plasma spraying, ion beaming technologies, and hydrothermal-electrochemical treatment are usually conducted under non-physiological and extreme conditions. Bioactive agents are entrapped on the surface merely by the formed coatings through these techniques, resulting in their rapid release upon exposure to a physiological environment. To overcome this disadvantage, surface modification of implants is conducted under the biomimetic condition, such as immersing implants in a simulated body fluid at approximately 37 C and a physiological pH. Apatite (or other compound) can then be formed on the metal surface. With biomimetic deposition, metals exhibit enhanced biocompatibility and bioactivity, and proteins can be not only deposited on the surface but also incorporated into the coating. The proteins are thus released gradually instead of in a single burst. Biomimetic deposition generally consists of two steps: initial nucleation, and formation and growth of the coating. Various deposition methods have been developed for metals, as well as for bioactive glasses, glass-ceramics and polymers. A typical case of biomimetic deposition for metals is the coating of a calcium phosphate (Ca–P) layer on the surface of titanium alloys as

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bone implants. In this case, a titanium alloy implant is immersed in an SBF for Ca–P nucleation on the metal surface. After the nuclei have reached the critical size, crystal growth starts, gradually forming the final coating. The apatite coating achieved by biomimetic deposition can make implanted materials osteoconductive, leading to improved bone tissue repair by metal implants. Polymeric biomaterials have also been well investigated for biomimetic deposition, and a few coating techniques are developed, such as alternate soaking process. These techniques can be applied to produce osteoconductive polymeric scaffolds for bone tissue engineering.

Summary Many manufacturing technologies for metals, polymers, ceramics and composites employed in other industries can still be used in the biomedical industry for fabricating metallic biomaterials, biopolymers, bioceramics and biomedical composites, as well as their products. However, care needs to be taken when using these technologies due to requirements and restrictions for biomedical products. Additionally, many other manufacturing technologies are investigated and developed specifically for biomaterials manufacture. In using any of these fabrication techniques, general principles in materials science and engineering and in materials and product manufacture must be followed. Good manufacturing practice must also be adopted. With the increasing demand for newer and better biomaterials for medical applications, it is expected that many novel manufacturing technologies will appear. Existing manufacturing technologies can also be modified and/or combined for producing new biomaterials to meet clinical requirements.

Acknowledgments Min Wang thanks Nanyang Technological University, Hong Kong Polytechnic University, The University of Hong Kong, UK’s Engineering and Physical Sciences Research Council, Singapore’s Ministry of Education, Hong Kong Research Grants Council and the National Natural Science Foundation of China for research funding.

Further Reading Ducheyne, P., Healy, K., Hutmacher, D. W., & Grainger, D. W. (Eds.). (2011). Comprehensive biomaterials. Amsterdam: Elsevier. Lanza, R. P., et al. (Eds.). (2013). Principles of tissue engineering (4th edn.). Burlington, MA: Academic Press. Ratner, B. D., et al. (Eds.). (2013). Biomaterials science: An introduction to materials in medicine (3rd edn.). Amsterdam: Elsevier. Wang, M. (2004). Bioactive materials and processing. In D. L. Shi (Ed.), Biomaterials and tissue engineering (pp. 1–82). Berlin: Springer.

Materials and Their Biomedical Applications Min Wang, The University of Hong Kong, Pokfulam, Hong Kong Bin Duan, University of Nebraska Medical Center, Omaha, NE, United States © 2019 Elsevier Inc. All rights reserved.

Introduction Materials Classification Primary Bonds and Secondary Bonds Structure of Materials Crystalline Solids Common crystal structures Crystalline defects Polycrystalline structures Anisotropy of materials Amorphous Materials Glasses Metallic glasses Polymers Monomer and chain structures Amorphous polymers Semi-crystalline polymers Homopolymers and copolymers Hydrogels Composite Materials Membranes and Thin Films Natural Materials Processes for Materials Solidification of Metals and Phase Transformation Recovery, Recrystallization and Grain Growth of Metals Polymer Crystallization Glass Transition of Amorphous Polymers Diffusion in Solids Processing–Structure–Property Relationships for Materials An Introduction to Biomedical Materials History of Biomaterials Interdisciplinary Nature of Biomaterials Development Human Body Environment Biocompatibility of Materials In Vitro Assessment of Biomaterials In Vivo Evaluation of Biomaterials Ex Vivo Evaluation of Biomaterials Standards for Biomaterials Government Regulations Biomaterials in Action Materials in Orthopedics Materials in Wound Dressings Materials in Dentistry Materials in Ophthalmology Materials for Cardiovascular Devices Materials in Drug Delivery and Controlled Release Materials in Reconstructive Plastic Surgery and Cosmetic Surgery Materials in Bioelectrodes and Biosensors Looking Into the Future Cancer Theranostics Bioprinting Organs-on-Chips Acknowledgments Further Reading

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Glossary Biomaterial A biomaterial is a non-viable material used in a medical device, intended to interact with biological systems. Biocompatibility Biocompatibility is the ability of a material to perform with an appropriate host response in a specific application. Ex vivo Ex vivo is Latin, meaning “out of the living.” Ex vivo experiments refer to investigations or measurements conducted in or on tissue from an organism in an external environment with minimal alteration of natural conditions. In vitro In vitro is Latin, meaning “in the glass.” In vitro experiments are performed with biological species (biological molecules, cells, microorganisms, etc.) outside their normal biological environment. They are generally conducted in labware such as test tubes, flasks and Petri dishes. In vivo In vivo is Latin, meaning “within the living.” In vivo studies are those in which the effects are investigated inside the whole, living organisms including animals and humans.

Introduction The history of mankind is characterized by the materials used and the naming of ages of civilizations effectively emphasizes the importance of materialsdthe stone age, bronze age, iron age, cement age, steel age, silicon age and now new materials age. Looking at our society as it is now and as it will be in the future, every field in the engineering world requires new materials, with the biomedical field being no exception. Novel biomaterials and structures are needed for improving the performance of existing medical devices and for developing new devices to tackle currently insurmountable medical problems. Modern biomaterials science and engineering rests on two pillars: materials science and engineering, and biological and clinical sciences. A good understanding of fundamentals of materials science and engineering paves the way for the successful development of new biomaterials.

Materials Classification For whichever the engineering field, materials are classified into four categories: metals, polymers, ceramics, and composite materials. Metals are strong and ductile; polymers are light, easily processable and flexible; ceramics are hard and brittle. When the property of metals, polymers or ceramics are not good enough for the targeted application or new property is needed for the application, a composite material is designed and developed. Needlessly to say that Nature, the best designer and maker of materials, makes natural composites such as wood and nacre. Metals, polymers, ceramics and composites can be made into different forms (particles, fibers, films and membranes, non-porous bodies, porous bodies, etc.) for different functions. The structures made of these materials can be as small as having nanometer dimensions and as large as structures covering kilometers.

Primary Bonds and Secondary Bonds Properties (mechanical, electrical, thermal, acoustical, optical, etc.) of metals, polymers and ceramics differ greatly. What makes them different lies in the interatomic forces that bind the atoms together in these materials. Atoms form bonds because the compounds formed by them are more stable than the alternative arrangements of isolated atoms. For all materials, there are three types of primary interatomic bonds: ionic bonding, covalent bonding, and metallic bonding. Molecules are formed when atoms are bonded together by strong primary bonds. Polymer long chains have covalent bonds, metals possess metallic bonds and in ceramics, the bonds can be either ionic, covalent or a combination of both. While the primary bonds in materials are strong bonds with high bonding energies, there are secondary bonds which are weak physical bonds such as hydrogen bonding and exist between virtually all atoms or molecules. Hydrogen bonding forms between some molecules that have hydrogen atoms as one of the constituents. Hydrogen bonding is important in the biological systems and hence in the biomedical field.

Structure of Materials Crystalline Solids The physical structure of solid materials highly depends on the arrangements of atoms, ions or molecules that form the material. Crystalline solids are materials in which atoms are arranged in a repeating 3D pattern over large atomic distances (i.e., long-range order). Examples of crystalline solids include metals and ceramics. Crystalline solids are in contrast to amorphous/noncrystalline materials, whose atoms or ions are not arranged in long-range repeating 3D patterns. Liquid water is a representative material of this group. The repeated 3D arrangement of atoms in crystalline solids is called space-lattice, which means a 3D array of points coinciding with atom positions. The geometrical configuration of a space lattice can be described through a unit cell. A unit cell is normally the

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smallest group of atoms that form a repetitive pattern of the crystalline solid and maintains the characteristics of the overall crystal. In a perfect crystal, all the atom positions can be generated by translations of the unit cell at integral distances along each of its edges. Three lattice vectors a, b, and c is used to describe the geometrical features of the unit cell. The lattice constants of the unit cell include the axial lengths (a, b, and c) and the interaxial angles (a, b, and g). Fig. 1 exhibits a unit cell with the lattice vectors and constants.

Common crystal structures There are 14 different unit cells for all crystalline materials that are used today, representing 14 primary crystal structures. Among them, body-centered cubic (BCC), face-centered cubic (FCC), and hexagonal close-packed (HCP) are the most common structures in metals. Fig. 2 shows schematic drawings (unit cells) of these three crystal structures. In FCC and BCC which are both cubic, there is only one axial length a, In HCP, there are two axial lengths: a and c. FCC and HCP are the densest unit cells. They have the highest “atomic packing factor (APF).” APF is used to calculate the percentage of space occupied by atoms in a unit cell, assuming the atoms are spherical. In the BCC structure, a central atom is surrounded by eight atoms which are located at all eight corners. The center and corner atoms touch one another along the cube diagonals. Each BCC unit cell includes two atoms: the central atom is a complete one, and each corner atom is shared by eight unit cells. The APF of BCC crystals is 0.68. FCC is also cubic, with atoms existing at the corners and centers of all the cube faces. An FCC unit cell has the equivalent of four atoms, with each corner atom accounting for a 1/8 atom and each face center atom accounting for a 1/2 atom. The APF of FCC crystals is 0.74. Therefore, FCC crystals are more closely packed than BCC crystals. HCP structure is also important but not often encountered. Each HCP unit cell has six atoms. The APF of HCP is the same as that of FCC, that is, 0.74. A value of 0.74 for APF signifies the closest packing possible of “spherical atoms.”

Crystalline defects In reality, perfect crystalline solids do not exist. Various types of defects are present in crystals and considerably affect their physical and mechanical properties. The classification of crystalline defects is made according to their geometry and shape. The four types,

z

c

χ

β

a

α

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y α

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β x Fig. 1

A unit cell with x, y, z axes, showing lattice vectors and constants.

Fig. 2

Three common crystal structures (from left to right: BCC, FCC, and HCP).

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from zero dimension to three dimensions, of crystalline imperfections are: (1) point defects, (2) linear defects, (3) interfacial defects, and (4) bulk or volume defects. Vacancies, the simplest point defects, are the absence of atoms at the regular lattice points and can be generated during solidification as a result of local disturbances. Another type of point defects is interstitials, which means a small atom can occupy an interstitial site between surrounding regularly-sited atoms in a crystal. Both vacancies and interstitials are important for diffusion in crystalline solids. Linear defects are also known as dislocations. Dislocations can be generated during material solidification. Other processes (vacancy condensation, plastic deformation, etc.) can also create dislocations. There are two types of dislocations: edge dislocation and screw dislocation (Fig. 3). The combination of these two types is called mixed dislocation. An edge dislocation can be considered as the insertion of an additional half-plane of atoms in an otherwise perfect crystal. A screw dislocation can be created by applying shear stresses with opposite directions to perfect crystalline structure. Interfacial defects, or “planar defects,” include external surfaces, grain boundaries, and stacking faults. Grain boundaries and stacking faults are two major types of interfacial defects. Grain boundaries exist in polycrystalline materials (“Grains” are individual crystals in polycrystalline materials). They are surface defects that separate grains with different crystallographic orientations. This region is about two to five atomic diameters in width and is a region of atomic mismatch between adjacent grains. The atomic packing in grain boundaries is lower than within the grains. Grain boundaries also have some atoms in strained positions that raise the energy of the grain boundary region. Grain boundaries can be generated during solidification of metals when different nuclei grow simultaneously to form different grains and meet each other. Stacking faults mainly occur in HCP and FCC structures. HCP and FCC crystals are formed by stacking of atomic planes (e.g., A, B, and C planes): ABABAB. for HCP structures and ABCABCABC. for FCC structures. During the growth of these structures, one or more stacking planes may be missing, resulting in stacking faults. Bulk or volume defects include pores, cracks, and foreign inclusions in crystalline materials. Pores can occur when a cluster of point defects combine together to form a three-dimensional imperfection. This type of defects has a tremendous effect on the performance of materials.

Polycrystalline structures Polycrystalline materials are solids that consist of many small crystals (the “grains”). The grains are separated by grain boundaries and normally have random crystallographic orientations. The size of the grains may vary from nanometers to millimeters. During the solidification of polycrystalline materials, small nuclei initially form at different positions of the liquid with random crystallographic orientations. These nuclei grow into larger crystals by absorption of atoms in surrounding liquid. Eventually, the crystals impinge on one another forming a granular or polycrystalline structure. In such a structure, a region grown out from the nucleus with the same crystal orientation is called a grain. Grains are separated by grain boundaries, the interfaces across which the crystal orientation suddenly changes. The atoms are packed loosely in the region of grain boundaries, making them mechanically and chemically unstable. Thus cracks and corrosion occur more frequently at grain boundaries.

Anisotropy of materials Anisotropy is the directionality of properties, which implies different values of the same property in different directions. On the contrary, isotropy is the situation that properties are independent of directions. For single crystals, physical properties, such as elastic modulus and electrical conductivity, of many substances are anisotropic. These properties vary by the crystallographic direction in which the measurements are taken. In many polycrystalline materials, the orientations of individual grains are entirely random. Therefore, even though each grain is anisotropic, most polycrystalline materials are isotropic.

Fig. 3 Edge dislocation (left) and screw dislocation (right). (For the edge dislocation, the red dotted line refers to the linear defect. In the right figure, the red arrows indicate the forces that cause the screw dislocation, which occurs on the blue plane.)

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Amorphous Materials In contrast to crystalline solids, the atomic structures of amorphous materials lack long-range order because of factors that inhibit the formation of a periodic arrangement of atoms. Atoms in amorphous materials are sited on random positions instead of distinct spatial positions like the atoms in crystals. Fig. 4 provides schematic drawings for crystalline materials and amorphous materials. Most polymers, glasses and some metals (“metallic glasses”) are amorphous materials.

Glasses Glasses are amorphous solids that are usually transparent. They are mainly used for technological and decorative applications. The most common type of glasses is silicate glass for windows, which generally consists of the chemical compound silica and has the oldest history for glass applications. Other types of glasses include fiber glasses, network glasses, organic glasses, etc. In the viscous liquid state of glasses, the molecules have limited mobility. The formation of the crystalline structure is thus inhibited and the amorphous structure dominates after solidification.

Metallic glasses Metallic glasses are a type of metals that are formed as non-crystalline solids under specific conditions. Compared to normal glasses, metals have high mobility in the molten state and are very difficult to form amorphous structures. Metal alloys with a high percentage of semi-metals, such as Si and B, can form metallic glasses through hyper-quenching. The cooling rate is usually between 105 and 109 degrees celsius per second, which is high enough so that the atoms do not have enough time to form crystalline structures. Metallic glasses are ductile with exceptional mechanical strength. They also possess excellent anticorrosion properties since grain boundaries do not exist in the amorphous structures.

Polymers A polymer is a macromolecule that contains many chemically bonded subunits. Natural polymers, such as rubber and silk, have been used by humans for centuries. Various natural polymers including DNA, proteins, enzymes and cellulose play vital roles in biological processes of living creatures. From last century, many synthetic polymers have been invented and widely utilized in many industries and daily life due to their broad range of properties.

Monomer and chain structures A polymer can be synthesized through chemical combinations of many small molecules, which is termed as monomers. For a monomer, the number of its available sites for bonding with other monomers under specific polymerization conditions is termed as functionality. A monomer must have at least two active sites so that it can bond with two other monomers to form polymer chains. Bifunctional monomers can only enter into two linkages with other monomers since each of them only has two sites available. Polyfunctional monomers can be linked together as nonlinear structures. According to the chain structure, polymers can be categorized into four groupsdlinear, branched, cross-linked, and network polymers. If the monomers are linked together in a linear manner, the resulting structure is called a linear polymer. For branched polymers, the monomers are joined together in a branched manner. Cross-linked polymers are formed when the monomer units are linked in multiple chains and have interconnections between chains. Network polymers are the cases that cross-linked polymers possess sufficient interconnections between chains. The chain structures of polymers are related to the functionality of monomers. The combination of bifunctional monomers produces a linear polymer. Polyfunctional monomers can be linked together to form nonlinear polymers, include branched, cross-linked, and network polymers.

Amorphous polymers The structure of polymers can also be categorized as amorphous and semi-crystalline. Most polymers have non-crystalline (“amorphous”) structures. Their chains are randomly entangled like noodles in a large scale. The physical entanglement of the chains with

Fig. 4

Schematic drawings for crystalline materials (left, crystalline silica) and amorphous materials (right, amorphous polymers).

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the dipole secondary bonds between chains enhances the strength of amorphous polymers. For branched polymers, their side branches can create loose packing of the chains, leading to amorphous structures.

Semi-crystalline polymers In the crystalline parts of polymers, individual polymer chains are folded and packed in ordered arrangement. However, an entirely crystalline structure cannot be achieved for polymers since there are always amorphous region in the long polymer chains. A semicrystalline polymer consists of many small crystalline regions interspersed with amorphous regions. Each crystalline region possesses a particular alignment. Semi-crystalline polymers generally have enhanced mechanical properties, increased fatigue strength and distinctive thermal behaviors as compared to amorphous polymers.

Homopolymers and copolymers Polymers that are formed by single repeating units are termed homopolymers, while polymers made up of two or more different repeating units are called copolymers. The sequence of a homopolymer molecular chain is like AAAAA. (“A” represents the single repeating unit). For copolymers, four types of arrangement exist: random, alternating, block, and graft. For a random copolymer, different monomers are randomly arranged in the molecular chains, such as ABBAAABABBB. (“A” and “B” represent different types of repeating units). In an alternating copolymer, the monomers are arranged alternatively, like ABABABAB.. Block copolymers are the molecules that different monomers are arranged in long blocks of the same type, like AAAAABBBBAAABBBBB.. Graft copolymers are the cases that chains consisting of one type of monomer are grafted on the long chain of another type of monomer.

Hydrogels Hydrogels, as a group of polymeric materials, are semi-solid, cross-linked macromolecular networks made from hydrophilic polymers. Because of the hydrophilic functional groups attached to the polymer backbone, the three-dimensional networks of hydrogels can absorb and retain significant amounts of water in the cross-linked structures. The resistance of hydrogels to dissolution arises from the cross-links between network chains. The crosslinks in hydrogels can be formed by physical cohesion forces, ionic bonds or covalent bonds. Hydrogels can be classified as homogeneous or heterogeneous according to their network structure. Homogeneous hydrogels possess an isotropic (random distribution) network structure, with relatively mobile chains and pores in the network. Heterogeneous hydrogels exhibit an anisotropic network structure caused by the strong interpolymer interaction.

Composite Materials Composite materials are solids that consist of two or more chemically distinct phases which are separated by interface(s). In most cases, composites consist of two phasesdthe matrix phase (metal, polymer or ceramic), which is the continuous phase that provides the overall structure, and the reinforcing phase in the form of particulates or fibers. Composites can be categorized by the form of the reinforcing phase. There are three major forms of reinforcing phases: particle-reinforced, fiber-reinforced, and structural-reinforced. In particle-reinforced composites, the particles (reinforcing phase) are dispersed in the matrix. For fiber-reinforced composites, the reinforcing phase has the geometry of a fiber. Structural-reinforced composites include structures such as laminates. With the matrix and reinforcing phase, composites can combine the desirable properties from both constituents to meet specific application requirements.

Membranes and Thin Films Membranes are often used for purification and separation applications. They are often developed to possess high permeability and sufficient selectivity while matching the process conditions such as temperature and pressure. Therefore, membranes often have certain amounts of pores with specific pore sizes. Common membrane materials include polymers, inorganics (e.g., nanoporous silica), and metal-organic framework (MOF). Each type membrane has its own structural features. For example, silica membranes can be mesoporous with high uniformity. In biomedical field, membranes have great potential for biosensing, biosorting, immunoisolation, and drug delivery. Thin films are layers of materials (metals, polymers, and ceramics) with a thickness ranging from nanometers to micrometers. In many applications including biomedical applications, coating a thin film on the substrate is highly important. A stack of thin films is termed a multilayer and multilayered structures can provide distinctive properties. Thin films are vital for many industries since they can provide excellent properties of the materials to fulfill application requirements.

Natural Materials Numerous materials from the Nature have been used by human beings since the stone age and many of them (stone, wood, cotton, etc.) still play important roles nowadays. Structures of natural materials vary tremendously, from crystalline structures like the diamond to composites like wool. Among the natural materials, some are attractive for biomedical applications. A good example is silks, which possess high mechanical strength and good biocompatibility. Silk materials are produced by different species with unique properties. For silk made by silkworms, two main proteins, sericin and fibroin, form the main body of the silk. Fibroin

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is the structural center of the silk, which is surrounded by sericin. The polymer chains in silk materials are bonded by both crosslinks and mechanical adhesion. Another natural material which is as biomaterials is collagen, a group of structural proteins that widely exists in human and animals. Collagens are used in medical treatments for humans and have been intensively studied for tissue engineering applications.

Processes for Materials Solidification of Metals and Phase Transformation In industrial applications, metals are usually melted and then cast into a mold to form products of designed shapes. The solidification of metals is thus crucial for metal processing. Two major steps are included in the solidification process: (1) nucleation, which is the formation of stable nuclei in the molten metals, and (2) growth of nuclei into crystals to form the granular structure. In the nucleation process, the formation of stable nuclei in liquid metals can be either homogeneous or heterogeneous. In homogeneous nucleation, metal itself provides the atoms for nuclei formation. When the temperature of a pure molten metal is below a solidification point, many slow-moving atoms bond together to form numerous homogeneous nuclei. Heterogeneous nucleation can occur on the nucleating agent, such as insoluble impurities, surfaces of the container of liquid metal and other materials in a liquid. The nucleating agent must be wetted by the molten metal. Compared with heterogeneous nucleation, homogeneous nucleation requires a large amount of undercooling, which may not be achievable in the industrial casting process. After the formation of stable nuclei in liquid metal, nuclei begin to grow into crystals with different orientations. During the solidification process, the crystals grow and join together to form grains and grain boundaries in the final solidified metal. With more nuclei in the liquid metal, the average grain size will be smaller. Most engineering metal products have small grain sizes, which is beneficial for mechanical strength and uniformity. In general, two major types of grain structure can be produced without using grain refiners: equiaxed grains and columnar grains. Equiaxed grains form when crystals in the liquid metal grow equally in all directions during solidification, while columnar grains are long, thin, coarse grains that form in the presence of a steep temperature gradient during slow solidification. Phase is a homogeneous portion with uniform physical and chemical properties in a system such as a metal alloy. If two or more phases exist in a given system, there will be a boundary separating the phases, and each phase has its specific characteristics. For example, steam, water, and ice in a container are considered to be three phases since they are physically different. Phase diagram is the chart that exhibits conditions, such as pressure and temperature, under which distinct phases coexist at equilibrium. A general phase diagram for this three-phase system is shown in Fig. 5. Phase transformations happen when phase boundaries (the red curves) are crossed due to the changes of conditions. The arrow in Fig. 5 is an example of phase transformationdice melts into water. Phase transformation occurs during processing of materials and affects greatly their properties. In metal processing, control of phase transformation can produce distinct materials to achieve desired properties. The driven force of phase transformation is the difference of free enthalpy between two phases. Free enthalpy, or Gibbs free energy, is a function of the enthalpy (internal energy of the system) and entropy (randomness of the atoms/molecules). The temperature for a phase transformation to occur is the thermodynamic equilibrium temperature shown in the phase diagram. Phase transformations frequently start by the nucleation process. The fluctuations of large amplitude of the structure/composition can lead to nucleus formation. With the growth of nucleus, atoms/molecules will be absorbed to the nucleus at the interfaces between transforming phases, just like crystal growth during the solidification of metals.

Recovery, Recrystallization and Grain Growth of Metals

Pressure

Cold working, such as rolling, forging, extrusion and other metal forming processes that are conducted under the “cold” condition, strengthens metals through plastic deformation. The cold-worked metal has many dislocations and other defects. Metals that have

Liquid (Water) Solid (Ice)

Gas(Steam)

Temperature Fig. 5

Schematic phase diagram for the ice-water-steam phase system.

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undergone cold working possess superior mechanical strength but are much less ductile. To enhance the ductility of cold-worked metals, a process called annealing can be conducted for the metals. During annealing, the cold-worked metal is heated to a sufficiently high temperature with adequate time, and three steps of changes occur: recovery, recrystallization, and grain growth. A cold-worked metal has higher dislocation density and internal energy than the original one. When the cold-worked metal is heated in the recovery temperature range, sufficient thermal energy is supplied to the dislocations to increase atomic diffusion. As a result, some dislocations are annihilated and some other dislocations move and rearrange themselves into states with lower energy. The internal stress of the metal is thus partially relieved. The recovery temperature range is just below the recrystallization temperature range. After recovery, the mechanical strength of the metal is reduced slightly, while the ductility is often greatly increased. Recrystallization is the process that new strain-free grains form and grow in the metal after the recovery process. The newly formed grains have low dislocation densities and are characteristic of conditions before cold working. The driving force of the nucleation of strain-free grains is the difference of internal energy between the strained and unstrained material. The progress of recrystallization depends on both temperature and time. The mechanical strength of metals decreases notably during recrystallization. Grain growth starts after recrystallization. If the temperature is high enough, the newly formed strain-free grain continue to grow. With a sufficiently long time and proper temperature, the new grains grow and completely replace the previous ones, resulting in an entire reduction in the internal energy. The internal energy is associated with the area of grain boundaries, which decrease when the grains increase in size. The reduction of internal energy is the driving force for grain growth. Large grains grow at the expense of small ones which shrink and eventually disappear. The Hall–Petch equation is applied to correlate the average grain size with yield strength of polycrystalline materials.

Polymer Crystallization Crystallization in polymers is different from that in metals and ceramics. Polymer crystallization is a process which transforms an amorphous, crystallizable polymer into a semi-crystalline material that its molecular chains are partially aligned. The folded and aligned polymer chains constitute an area called lamellae. Fig. 6 shows schematically structures of amorphous and semicrystalline polymers. Polymer crystallization can be conducted through different processes, such as solidification from the melt and solvent evaporation. The solidification of molten polymers starts by nucleation in which some chains in the polymer become parallel. Like the solidification of metals, nucleation for semi-crystalline polymers also has two mechanisms: homogeneous, and heterogeneous. Due to heat motion, some polymer chains occur parallel and result in homogeneous nucleation. Heterogeneous nucleation is initiated by impurities or additives. After nucleation, crystal growth occurs at the temperature between the crystalline melting temperature Tm and the glass transition temperature Tg. At this step, more folded chain segments are added on the nuclei. When the temperature gradient is high enough, the direction of crystal growth correlates with the steepest gradient. If the polymer has an isotropic and static temperature distribution, the crystal (lamellae) grows radially and forms a large aggregated called spherulite. Solvent evaporation can also crystallize polymers. Generally, the polymer is dissolved in a solvent to form a dilute solution, where the polymer chains are separated from each other, and the polymer chains can fold to form single-chain crystals. When the solvent evaporate, the concentration of the solution increases and interactions between polymer chains occur to form semicrystalline polymers. The solvent evaporation process may generate polymers with the highest degree of crystallinity.

Glass Transition of Amorphous Polymers Glass transition, or glass-liquid transition, is the transition of a hard and brittle glassy state into a softer and rubbery state for amorphous and semi-crystalline polymers. Glass transition occurs at a narrow temperature range termed as glass transition temperature

lamellae Fig. 6

Schematic drawings for structures of amorphous polymers (left) and semi-crystalline polymers (right).

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Specific volume

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Temperature

Tg

Tm

Fig. 7 Specific volume-temperature curves of amorphous and semicrystalline polymers. (The blue zone is the region of the glassy state. Yellow zone represent the range of rubbery state. Red zone is the liquid region.)

Tg, which is usually lower than the crystalline melting temperature Tm. For semi-crystalline polymers, the glass transition only occurs at the amorphous regions, while this process happens in the entire structure of amorphous polymers. The glass transition can be illustrated by a typical specific volume-temperature curve as shown in Fig. 7. The green line refers to the curve of an amorphous polymer, while the black line represents a semi-crystalline polymer. In the blue zone, the amorphous polymer is heated at a low temperature. Here, the polymer expands at a constant rate with an increase in temperature. When reaching Tg, the volume expansion rate increase instantly to a higher level, meaning the hard and brittle glassy state changes to the soft and rubbery state in the yellow zone. Upon further heating, the amorphous polymer gradually transforms to the liquid state in the red zone. For the semi-crystalline polymer upon heating, the change at Tg is less intense since glass transition does not happen in the crystalline regions. In the yellow zone, the semi-crystalline polymer is in a state that crystals dispersed in a rubbery amorphous matrix. At Tm, the volume of the semi-crystalline polymer expands drastically due to melting of the crystalline regions in the polymer.

Diffusion in Solids The phenomenon in which matter is transported through matter is defined as diffusion. Diffusion is achieved by atomic motion. Atomic movements in gases and liquids are relatively easy and rapid. In solids, such motion is restrained because of bonding of atoms to the equilibrium position. Thermal vibrations in solids can enhance atomic mobility and thus allow some atoms to move from high concentration areas to low concentration areas. Solid-state reactions are usually related to diffusion. The recrystallization of cold-worked metals is an example of diffusion. The diffusion in a crystalline solid is the stepwise migration of atoms from lattice sites to lattice sites. Vacancy diffusion and interstitial diffusion are the two major mechanisms of diffusion in a crystalline lattice. Vacancy diffusion, also termed as substitutional diffusion, is a mechanism by which atoms with sufficient thermal energy move from original sites to vacancies or other crystal defects in the lattice. Interstitial diffusion is the mechanism by which atoms migrate from interstitial positions to neighboring empty ones. In this mechanism, the moving atoms must have sufficiently small sizes as compared to the matrix atoms. Atoms such as hydrogen, carbon, oxygen, and nitrogen can diffuse in metals via this mechanism. Steady-state and non-steady-state diffusions are often involved in the study of diffusion. Steady-state diffusion is the situation that the diffusion flux and concentration gradient do not vary with time. On the contrary, in non-steady-state diffusion, the diffusion flux and concentration gradient change with time. Steady-state diffusion is only suitable for specific situations that a highpressure non-reacting gas diffuses through a metal foil to a region where the pressure of this gas is low. Non-steady-state diffusion works for most practical diffusion situations. Steady-state diffusion and non-steady-state diffusion are dealt with by Fick’s first law and Fick’s second law of diffusion, respectively.

Processing–Structure–Property Relationships for Materials Processing, structure, and property are three essential components in materials science and engineering. The relationships among these components play a vital role in the design and application of materials. Basically, how a material is processed result in its specific structure, while the structure of a material leads to distinct properties. A linear progression can be made from processing to structure, then to the properties, resulting finally in the specific performance and applications of a material. If a specific property is needed for the application, the structure of the material must be suitably tuned, and hence the appropriate processing route must be selected and processing parameters must be controlled. This article has mainly discussed structures of materials. Other article in this encyclopedia will discuss manufacture and properties of materials in the context of biomedical applications.

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An Introduction to Biomedical Materials Many materials, including metals, polymers, ceramics and composites, are now used for biomedical applications and hence they are called “biomedical materials.” Some of these materials were not intended as biomedical materials originally. But their use in the biomedical applications has been successful and therefore they extended their service into the biomedical field. In the middle of the last century, modern biomaterials science and engineering began. Since then, many materials are designed and have been developed specifically for biomedical applications.

History of Biomaterials The concept of biomaterials is a relatively new idea and term, although biomaterials have actually been used for a very long time. A biomaterial is a material that is capable of being introduced into the body or living tissue without inducing a harmful biological response. Today, they are extremely widely used in the healthcare field for many different purposes, ranging from sutures to permanent implants. For many years, people were using biomaterials to help solve medical problems without realizing what a biomaterial was and why they worked or did not work. For a long time, there were no medical device companies producing biomaterials, besides some externally used materials, and materials were not being made for the purpose of using them as biomaterials. This means that there were no regulations or standards to follow when using materials for use in the human body, and there were mixed results in the safety and efficacy of the materials. The earliest used biomaterials were made from basic materials found in nature that were easily usable. There were sutures made of linen or catgut, as well as evidence of using seashells as dental implants. Some groups of people also used the jaws of large ants to hold wounds closed in lieu of sutures. As technology and the processing of materials advanced, so too did the biomaterials used in medical procedures. The use of metals as biomaterials allowed for stronger devices, such as metal sutures and a variety of implants, including dental implants. The longer these metals were used as biomaterials, people realized that there were problems with some of them, and the concept of biocompatibility was born. Many metals were tested and some were kept for use as biomaterials and some were discarded due to their adverse effects on the body. Eventually, plastics were developed, which allowed for a wider variety of properties for biomaterials from which to make medical devices. This leads to an advancement in the types of devices that could be created from the biomaterials. The polymeric biomaterials are tested for their biocompatibility, now that it is standard for testing. There are constantly new biomaterials, like hydrogel for example, being created and tested for different applications. Hydrogels are polymers that can swell but not dissolve in the water, which brings more new ideas to the world of biomaterials. The 1980s and 1990s were the golden era of ceramic biomaterials with many inventions and innovations on bioceramics. Biodegradable bioceramics are found to be useful for hard tissue regeneration in this century when tissue engineering and regenerative medicine have attracted great attention. Composite biomaterials play no lesser role in the biomedical field. They have found many biomedical applications, ranging from biosensors, bioseparation, drug delivery devices, to human tissue repair and regeneration.

Interdisciplinary Nature of Biomaterials Development Developing new biomaterials is a complicated process that requires a variety of skills and knowledge. To create and use new biomaterials, knowledge is needed in chemistry, biology, engineering, medicine, and other related areas. Due to this fact, it unlikely that one person has obtained all of the knowledge and skills necessary to develop a new biomaterial. This leads to people working together to develop, test, and use biomaterials. Chemistry and materials science and engineering knowledge is used in the development of new types of biomaterials and in optimizing existing materials. It helps with knowing how materials react in the presence of other materials and bodily substances. Biological and medical knowledge is key in understanding the reaction of the body to the biomaterials. It also helps in being able to determine what biological processes can be avoided and which to use to the materials advantage. Engineering skills are used to utilize biomaterials into devices or processes that can be used to solve a problem. These skills can be used to determine the properties of the biomaterials and how those properties are suited for certain biomedical applications. Experience in medicine enable people to know what types of biomaterials and devices are needed to make their jobs easier and to make clinical outcomes better. They can lead to looking at new ways to do things. The people with medical skills also have the ability to test the new biomaterials and devices made from them and are in charge of using them most of the time. The fact that so many different pieces are put into making a biomaterial, and a device from the biomaterial, makes a need for developing them to be interdisciplinary. Different sets of knowledge contribute to the development at different stages, and some contribute throughout the process. If people with all of these sets of knowledge can work together seamlessly, then the development of a biomaterial and its associated devices can run smoothly and quickly and could turn into a quality product.

Human Body Environment The human body is a complex system and environment. It has many different types of tissues, cells, proteins, growth factors, and other biological components. To add to the complexity, each tissue has a unique extracellular matrix (ECM) that affects the local environment of cells and other biological components. There are many different types of proteins in the body, all of which can adsorb, desorb, denature, and increase or decrease in activity in the presence of different biological materials and chemicals. Blood

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contains proteins that are reactive and can adhere to foreign bodies and injury sites and cause clotting and inflammatory responses. Blood also contains cells that have the job of fighting and removing foreign bodies. Along with the biological components of the body, the body parts feel mechanical forces. These mostly come from body motion and blood flow, and can affect the function and wear of parts of the body. The goal of a working biomaterial is to not induce any unwanted reaction from any components of the body, whether it is toxicity or unwanted wear and tear.

Biocompatibility of Materials For a biomaterial to be safe and effective, it must be biocompatible. If a material is biocompatible, it means that a material does not cause any harmful effects when it is in contact with the body. It also means that the material should perform its specific function without any unplanned effects. In general, the harmful effects that should be avoided are producing or supporting toxic substances, irritation due to movement, rubbing, or improper mechanical forces, and improper host reactions to the biomaterial. There are biomaterials that are said to be bioactive. Not only do these biomaterials not cause an adverse reaction, but they actually cause wanted reactions that may be used to enhance the positive healing responses to the material. Bioabsorbable materials degrade over time into safe substances in the body, allowing healing to occur naturally in the implantation site. A biomaterial that is bioactive, bioabsorbable, or both can have enhanced biocompatibility. A biomaterial must be viewed as a piece of a device. A biomaterial by itself may cause a different reaction than it does in the context of a full device. The goal of a biocompatible material is to have the minimized amount of inflammation possible. The initial reaction from the body when an implant, or any foreign body, is introduced is to undergo a foreign body reaction, which is a form of non-specific inflammation. If the foreign body is recognized as a foreign substance, then inflammation occurs as macrophages attempt to break down and remove the substance. Even if it does not occur instantly, if particles break off of the biomaterial, this can occur. In some cases cells that respond in inflammation can kill tissue around the biomaterial. Reducing inflammatory reactions reduces the risk of damage. Another important factor in biocompatibility is for the material to be non-toxic. This means that the material itself, or its products, should not harm or kill cells or tissues. There are multiple types of toxicity that all must be avoided. These include genotoxicity, carcinogenicity, reproductive toxicity, and cytotoxicity. Genotoxicity is the tendency of a chemical to mutate the genes of a cell, carcinogenicity is the tendency to cause a cell to become cancerous, reproductive toxicity is the tendency to cause death of reproductive cells, and finally, cytotoxicity is the tendency to kill living cells. The chemicals released by materials can cause any of these forms of toxicity. Some materials are susceptible to bacterial growth and may cause an infection. A material cannot be considered biocompatible if it has any form of toxicity or infection. When a material comes in contact with blood, it can cause thrombosis, embolization, and the consumption of platelets and coagulation factors. The biomaterials developed so far are never as resistant to thrombosis as natural endothelium in the body. The materials that are exposed to blood and the vascular system for an extended period of time are especially susceptible to thrombosis. Thrombosis can cause major problems in the function of the biomaterials and for the health of the person who received the implant. For a material to be considered biocompatible, it must not cause any harm to the cells or tissue while performing as intended in the body.

In Vitro Assessment of Biomaterials The goal of in vitro assessment of biomaterials is to mimic how the biomaterials and cells react to each other, as they would in the body, as closely as possible so that the biocompatibility of the material can be determined and also cell behaviors be assessed. The focus of the tests is to identify the chemical components of the materials that might be release in vivo, and determine if those chemicals are toxic to cells. Assays are used to determine the genotoxicity, carcinogenicity, reproductive toxicity, and cytotoxicity. All of these can occur in multiple different ways, so they must be assessed thoroughly. There are many standardized methods for the in vitro testing of biomaterials. Tests are designed to assess the characteristics of chemicals that may be released by the material under development during its time in the body. This includes how the material reacts to cells and fluids, as well as how the cells and fluids react to the material. A toxicology risk assessment is performed on the material to identify hazardous chemicals coming from it, predict the potential exposure to the chemicals, determine the dose-response relationship of the chemicals, and then use all of the information to characterize the potential risk to patient from the material. The test is done with a worst case scenario in mind for the amount of chemicals a patient could be exposed to, in order to be sure that the material is safe for less than the worst case. The tests for genotoxicity test for gene mutations in bacteria, as well as chromosomal damage in mammalian cells. If either one is positive for damage, then an in vivo test most likely is needed. Since all genotoxic chemicals are carcinogenic, a test for carcinogens should be performed before any testing on live animals is done. The carcinogenicity tests check for change in cell morphology, anchorage-independent replication, and discorded colony growth, which are indicators of malignant cells that have been formed. Reproductive toxicity testing does not need to be performed unless the biomaterial is able to come in contact with an area of the body or cells that are associated with reproduction. In vitro cytotoxicity tests determine whether or not a chemical from a material cause death to cells or disruption of essential process. Cytotoxic chemicals can do this by changing the environment around the cells so that they do not work or by altering a specific part of the cell that disrupts the proper function. Cytotoxicity can also lead to chronic inflammation. There is no single test that can fully characterize the toxicities for all chemicals and hence multiple must be performed to assess the full extent of the material toxicity.

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Due to the components and overarching importance of blood in the body, it is important to test the interaction between the biomaterial and blood in vitro. The material should be tested for whether or not it causes thrombosis, how it interacts with platelets, whether it damage red blood cells, or if it stimulates an immune response. The assessment models are designed to mimic the conditions inside the body. This means that the constituents of the test fluid should have the coagulation factors, correct temperature, flow, and other components that make the blood function in the body. Getting an assay to mimic clinical conditions as closely as possible is important in the assessment of a material for any type of test. There are needs to develop more in vitro toxicity tests to ensure that biomaterials are safe and effective without having to spend a lot of time and money on testing. The tests need to be developed to mimic as closely as possible the in vivo environment, so that, hopefully, an in vivo test would not be needed. It might not be completely realistic due to complexity differences between being in a body and outside of one, but being able to test materials in vitro with accuracy and completely free of any in vivo tests would be the ultimate goal.

In Vivo Evaluation of Biomaterials In vivo evaluation of biomaterials is centered on testing the biocompatibility of the biomaterial in a complex biological environment and in many cases, determining the efficacy of the medical device made of the material. The evaluation of biomaterials in vivo gives a better idea of how a biomaterial will react with the human body than in vitro tests. This is due to the complex nature of the body environment. The in vivo tests help to determine whether or not a medical device and its associated biomaterials perform as intended without any safety risks. The government regulatory bodies that have control over the use and testing of medical devices issue protocols, guidelines, and standards that should be used for the in vivo assessment of medical devices. The in vivo tests are different depending on what the intended use of the medical device is. They can also be used to test the individual biomaterials and their reactions or the device as a whole. The tests are based on the intended use of the device, and devices are grouped into different categories. These include the area of tissue contact, which includes surface contact, external communicating devices, and implant devices, as well as the duration that the device is in contact with the tissue, which includes limited (less than 24 h), prolonged (24 h–30 days), and permanent (more than 30 days). Biocompatibility tests in vivo include tests for sensitization reactions, irritation, intracutaneous reactivity, toxicity in all forms, blood compatibility, carcinogenicity, degradation, and other immune responses. Choosing the animal for tests is important when determining the safety and efficacy of a biomaterial or device. Some animals are more closely related to humans in certain areas than in others, which means that using that animal for one test might work well, but for another test it could mean nothing. For example, using a mouse model for tests with bone could work well to replicate humans to an extent, but using a mouse for vascular tests would not translate well at all. Selecting the right animal model is crucial in determining the safety and efficacy of the biomaterials. Certain knowledge of animal biology and the similarities and differences between humans and animals is needed for determining which model(s) should be used for which purpose(s). The important part of testing biomaterials in vivo is to match the human tissue as closely as possible in order to accurately and correctly model the interaction of the biomaterial in the human body.

Ex Vivo Evaluation of Biomaterials Ex vivo evaluation of biomaterials allows the testing to be done without having to have a large number of animals. It is done by taking tissue from the animal to test against the biomaterial outside of the body. An ex vivo test can be done using a shunt to test the blood interaction of the biomaterial. In ex vivo tests, the physical, chemical, mechanical and other properties of the tissue and biomaterials that have interacted with them after being implanted for a time or during the interaction are analyzed. Ex vivo evaluations of biomaterials are another way for ensuring the safety and efficacy of biomaterials.

Standards for Biomaterials Standards for biomaterials are needed so that there is consistency in biomaterials R&D and in quality of biomaterials is maintained around the world. Without the standards, there would be major problems with the manufacturing, quality control and use of the biomaterials. The standards developed for a material are used by manufacturers, users, researchers, etc. to consistently define and use a material every time. They specify what the chemical, mechanical, physical, and electrical properties of a material should be. There are standards developed for material specifications, material uses, material testing, biocompatibility, and other important standards. Material standards allow multiple manufacturers to make the same product, as well as letting the user know exactly what they are getting every time. Most standards are consensus standards, meaning they are developed by using popular opinion from a committee/community. A test method standard is developed to specify the conditions of the test, type of and how many test subjects, and what data is available to be analyzed from the test. These test method standards can be used to test a material by anyone after the standard has been developed and approved. Getting a standard approved can be a long process. It starts with a need for a standard and a group of people is formed and tasked to decide how a test should be done. They then send out materials to be tested to multiple laboratories and write the draft standard. They review information from multiple sources and finally test and produce the document. From beginning to end, it can take about 3–5 years to produce a standard for a material or test.

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The main organization in the United States that makes standards for biomaterials is the American Society for Testing and Materials (ASTM). Some other standardization bodies are the Association for the Advancement of Medical Instrumentation (AAMI), which deals with medical electronics, sterilization techniques, vascular and cardiac valve materials, the American National Standards Institute (ANSI), which deals with the reviewing and acceptance of standards documents, and the American Dental Association (ADA), which deals with dental materials standards. The ANSI also stays in contact with the other standards organizations, as well as the International Standards Organization (ISO) in order to help produce standards for international use. Other countries have standards organizations similar to the ones in the United States that also work with ISO. ISO gathers the standards from all of the standards organizations to create international standards. The international standards make it easier for companies to manufacture materials and devices that can be sold and used internationally. The standards organizations make their approved standards public so that they are available for use by everyone who needs them.

Government Regulations Government regulations on medical devices, including biomaterials, are important in that they provide a way to ensure the safety and efficacy of devices by providing a legally regulated path of approval. In the United States, the Food and Drug Administration (FDA) was given the regulatory powder of medical devices in 1976. Along with the regulatory authorities in other countries, the FDA regulates biomaterials on the basis of the risk associated with the intended use of the biomaterial. It is tasked to make sure that the device or biomaterial is safe and functions as intended in the United States. This means that the biomaterial could be regulated one way for one use and a different way for another use. Since the regulations are based on the risk associated with use of the biomaterial, they are grouped into three main classes in the United States. Class I medical devices and biomaterials are determined to be low risk. These devices and materials are mostly external devices and are not intended to support life and failure of the device would not cause any real harm. The Class I medical devices and biomaterials require the lowest regulatory control, as the risk of harm from them is very low. This allows them to be quickly and easily passed through the regulatory process. Examples of Class I devices are bandage and dental floss. Class II medical devices are found to pose moderate risk. These devices and materials are generally in contact with the body for a short period of time or detect internal components of the body. They are devices and materials that are safe for the function they are designed for and will not cause any major injury if failure occurs. Some of these devices are external and some are internally used. Due the moderate risk factors and intended uses, Class II medical devices and biomaterials require more strict regulations than Class I medical devices, with more stringent reviews of the technical documentation. Examples of Class II devices are gastroscope and magnetic resonance imaging (MRI) equipment. Class III medical devices and biomaterials are determined to be high risk. These devices and materials are used in extremely invasive procedures and life-sustaining devices. They can even be permanently implanted into the body. The failure of Class III medical devices can cause major injury or even death. For this reason, they are subject to the most scrutiny for approval. They must perform exactly the way they are specified without failure and that having them in the body will not harm the body. To do this, Class III devices are tested in animals and clinical trials to be completely sure of their safety and efficacy. Examples of Class III devices are heart valve, pacemaker and hip implant. As expected, the higher the risk, the higher amount of scrutiny is given to the devices before they pass the regulatory process. There are development and manufacturing regulations as well as premarket entry requirements. The development and manufacturing requirements are mostly standard throughout the world but the premarket rules can vary depending on the country. The easiest way for a medical device or biomaterial to get approved is to effectively compare it to one that has previously been approved. If they are determined to be similar enough, the regulatory process may move quickly for the new device or biomaterial. If not, tests and other work may be needed, which slows down the process. If a manufacturer wants to change a product slightly, this may cause a need for a new review of the device or material. It is up to the manufacturers to prove the safety and efficacy of the product that they are producing in order to pass the regulations. They must provide full documentation for the medical device or biomaterial for gaining regulatory approval. The regulatory process is an ongoing and ever changing system with reviews and adjustments going on all the time. Getting through the regulatory process can take a long time, which makes new medical devices and biomaterials take a long time to get to clinical use.

Biomaterials in Action Materials in Orthopedics In orthopedics, biomaterials are generally chosen for their strength or for mimicking the structure and properties of bone. They are also wanted for promoting the mineralization of tissue around the implants, which calls for bioactive materials. Many times, this can mean the use of metals or ceramics. Calcium phosphate ceramics, particularly synthetic hydroxyapatite (HA), closely resemble bone apatite and have been developed for bone tissue repair. They allow better bone growth in areas surrounding the ceramic implant than any other material such as orthopedic metals. In some applications, there can be metal implants coated with a bioactive ceramic material for the strength of metals and for the bioactivity (i.e., osteoconductivity) of calcium phosphates. Metals for use in orthopedics include titanium and its alloys, cobalt’chromium alloys, and 316L stainless steel. Metals are usually used for implants that require mechanical strength such as hip implants, bone screws, nails, pins, and fixation plates. Popular bioceramics are bioactive glasses, glass-ceramics and calcium phosphates such as HA and b-tricalcium phosphate (b-TCP). Materials

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such as silicon can be added to HA to enhance the biological property and promote bone growth. Bioactive bioceramics can be used as coatings for metals and as bone defect fillers. Two “bio-stable” polymers, ultra-high molecular weight polyethylene (UHMWPE) and poly(methyl methacrylate) (PMMA), are commonly used in hip joint replacements. UHMWPE is used for acetabulum cup while PMMA is used to cement metal stem of the hip prosthesis to bone. Commonly used biodegradable polymers include poly(ε-caprolactone) (PCL), polylactide (PLA), polyglycolide (PGA), poly(2-hydroxyethylmethacrylate) (PHEMA), collagen and hyaluronic acid. These polymers are used for a variety of applications based on their properties. They can be used to repair bone fractures, as substitutes for bone, ligament and cartilage repair, sutures, fixation plates, etc. that will degrade over time. Bioactive composite, such as HA reinforced high-density polyethylene (HDPE), are also developed for bone tissue repair.

Materials in Wound Dressings There are a variety of wounds, ranging from cuts and scrapes to ulcers and burns. This means wound dressings must be able to perform different functions, depending on the type of wound. In general, wound dressings must stop the bleeding, keep the wound clean and disinfected, absorb excess liquid and promote the proper gas and temperature levels in order to start and promote the body’s healing process. The biomaterials used must also be able to be formed into sheets easily in order to function properly. There are passive biomaterials that are used to just cover the wound and keep it clean. Traditional passive wound dressings are made of woven fabrics and foams that assist healing by absorbing excess liquid and blocking outside sources of contamination. Some wound dressings are able to promote the healing of wounds, allowing them to heal quicker and better, instead of just covering them. Interactive dressings allow for the movement of substances such as water and oxygen to pass through them without allowing contaminants to enter. These types of dressings are made of materials that can be made into thin films. Another type of wound dressing that promotes healing is termed bioactive. These bioactives are made of materials that come from the body and are known to aid in the healing processes. By bringing these materials to the wounds, the healing process can be sped up. These natural biomaterials are hemostatic agents. Combining multiple types of biomaterials into one wound dressing can allow it to effective cover multiple functions needed by a certain type of wound. Some common synthetic biomaterials used in wound dressings include PLA, PGA, PCL, poly(lactic-co-glycolic acid) (PLGA), poly(ethylene glycol) (PEG), polyesters, and polycarbonates. These biomaterials are able to be formed into sheets and sponges for covering wounds. Some commonly used natural polymers are proteins such as fibrinogen, thrombin, collagen, gelatin and albumin and polysaccharides such as chitin, chitosan and cellulose. These materials have been able to be made into sheets, sponges, gels, powders, and liquids. These hemostatic materials help in regenerating new tissue in the wound due to the fact they are naturally present in wound healing or support the growth without inflammation. Cotton is commonly used in bandages and gauze. Some polymers such as hyaluronic acid and methacrylates can be used in hydrogels that support wound healing. A combination of different biomaterials can be helpful, depending on the need for healing and how each material interacts with the wound.

Materials in Dentistry Biomaterials used in dentistry are very similar to those used in orthopedics, since teeth are similar to bone and are anchored in bone. Dental biomaterials include metals, polymers, ceramics, and composites. Using these materials is to either prevent or fix problems, and different from the application of most other biomaterials, some of dental biomaterials are visible and hence matching the color of the surrounding tissue can be important. Metals are often used as anchors in dental implants. They provide the strength needed to hold the implant in the bone under the stresses that the mouth goes through. The metals include titanium and its alloys, cobalt-based alloys and stainless steel. Dental crowns and bridges can also be made of metal. Gold crowns are common. Dental amalgam, which is composed of mercury, copper, tin, zinc, and silver, is used as a filler material for cavities after tooth decay. Bioceramics are common dental biomaterials. In some cases, they can be used as the anchors of implants as well. Some bioceramics, for example, medical grade alumina, have the high strength needed in the mouth and have low thermal conductivity, while exhibiting a color similar to natural teeth. Calcium phosphates including HA are also commonly used. HA can be coated on metal implants to promote osteointegration with bone. Polymers are used mostly as cements or fillers in dentistry. Cements or fillers start out as liquids and/or solid powders and are hardened to hold two solids together or fill holes. Some harden on their own after mixing, while others need to be hardened, for example, by UV light. They are usually made by mixing solid and liquid components. There are cements made of zinc phosphate, zinc polycarboxylate, glass ionomers such as calcium and aluminum silicate, resins such as urethanes, and other types that can be used for specific purposes. All dental biomaterials must survive the harsh and fluctuating conditions of the mouth. They are hard and stable material that work together in multiple ways. The metals, ceramics, polymers and composites work together and allow for integration with the host tissue.

Materials in Ophthalmology Biomaterials used in ophthalmology are contact lenses, optical implants such as artificial corneas, intraocular implants, glaucoma drainage tubes, scleral buckling materials, adhesives for repair of perforations and ulcers, etc. They are generally soft polymers,

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considering the eyes are made of soft and delicate tissues. Some of the main biomaterials used are PMMA, silicones, and hydrophilic polymers in the form of hydrogels. PMMA is used for hard contact lenses but it does not have great oxygen permeability. Hence it can cause problems if used for too long. It is also used for artificial cornea and intraocular lens. This hard polymer allows the artificial corneas and lenses to retain their shape and not to break down easily. The clear material allows for good optics as well. Silicones are used in the forms of a hard rubber, soft rubber, sponge, and liquid with high viscosity. Silicone rubber is used in some cases for contact lenses. It has high oxygen permeability but can cause the formation of deposits due to its hydrophobicity. Soft silicone rubber and sponge are used in retinal detachment surgeries as scleral buckling materials.

Materials for Cardiovascular Devices The cardiovascular system needs a variety of biomaterials. Biomaterials for heart and blood vessels have different requirements. A common requirement for all biomaterials for the vascular system is that they must not react or cause a reaction in the presence of blood since they are in constant contact with blood for their entire lifespan. The materials also need to withstand mechanical stresses because of the heart pumping and constant blood flow. Metals, carbons, ceramics, and polymers are common biomaterials for cardiovascular applications. Metals have high structural integrity. They are used in structural parts in heart valves, pacemakers, and stents. Some of the main metals for these purposes are stainless steel, cobalt’chromium alloys and titanium alloys. Nickel’titanium shape memory alloy (SMA, e.g., Nitinol) is used in vascular stents owing to its shape memory. Metals are also used in electrodes and wires in pacemakers. Platinum alloys are used in electrodes, and stainless steel and tantalum are used in sensors and wires braids. Carbons, especially graphite, are used in pyrolytic coatings of heart valve components. They are thromboresistant, have high lubricity and are resistant to wear. Sapphires have been used in high speed blood pumps to reduce the amount of friction during rotation. Polymers such as polytetrafluorethylene (PTFE) and polyester are used as small grafts and sutures. They help to repair tissue in blood vessels. Silicone has been used in artificial heart valves as tissues and cells generally do not attach to it. Biological materials are generally used as coatings for other biomaterials. All of the biomaterials used in the cardiovascular system must not cause thrombogenesis, must resist physical wear, and must not break down over time.

Materials in Drug Delivery and Controlled Release The goals of controlled drug delivery are to control the duration of release and the amount of drug released, to deliver the drug to a specific part of the body, to get through tissue barriers, and to get through cellular barriers. Biomaterials used in drug delivery must have right chemical compositions. Drugs can be delivered through the digestive systems, respiratory system, through skin or muscles, intravenously, or through the respiratory system. Each delivery method needs a specific type of device or biomaterial to deliver the drug effectively to the targeted area. The drug delivery methods are meant to control the amount of drug in the system and where the drug is delivered to and even to overcome cellular barriers. Good delivery vehicles deliver the drug with an appropriate amount to the proper location by releasing the drug or breaking down themselves under the conditions of the targeted sites. Controlled release means controlling the rate at which a drug is released into the system, as opposed to having the drug delivered into the system in full at one point in time. The main mechanisms for controlled release are diffusion, chemical reaction, and solvent activation and transport. The diffusion mechanism uses a polymeric material that creates a reservoir for the drug, or a material in which the drug can be uniformly distributed throughout. In the delivery through chemical reaction, the biomaterial making the delivery vehicle degrades in the presence of water or other agent. Solvent delivery uses a material that can swell in the presence of water to release the drug that has been locked in the capsule or by osmotic effect. Polymers are the main biomaterials for drug delivery and controlled release purposes. On the other hand, ceramics such as mesoporous silica can also be used for controlled delivery. Hydrogels are good for diffusion-based controlled release because they swell when exposed to water or other biological fluids. This allows for drugs to diffuse out of the expanded gels. They can also be used for solvent activation. Some hydrogels can respond to environmental factors such as ionization. They can increase the amount of swelling due to responses to the environment, which can allow the drug to be released more quickly or slowly. Specific locations in the body can be targeted by creating a hydrogel that binds to a specific tissue.

Materials in Reconstructive Plastic Surgery and Cosmetic Surgery Reconstructive plastic surgery has the goal of repairing defects for the sake of the damaged tissue appearing normal and hopefully regaining its function. Cosmetic surgery has the goal of changing the appearance of a bodily feature. Since they are closely related to each other and have the hope of creating structures that will hold shape and last in the body, they can use similar biomaterials to change and fix the structure of body parts. The biomaterials used in reconstructive surgery are intended to provide mechanically stable structures of the tissue while allowing native tissue to use it as a scaffold to regenerate its proper structure. Some of the materials are meant to degrade over time, while some integrate with the body. Much attention is paid to wound healing so as to prevent the appearance of scars after surgery.

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Most of the biomaterials for plastic surgery should be mechanically strong and chemically stable. There is also the reason for the materials to have some bioactivity. If they are bioactive, they can integrate with the surrounding tissue, which means a better functioning tissue repair that will heal to look normal. For support materials, PTFE, PMMA and titanium have been used but they do not integrate well with the tissue due to low bioactivity. Therefore, for soft tissue, a biomaterial called acellular dermal matrix (ADM) was developed. ADM integrates well with native tissue and allows for the healing process to occur through tissue regeneration rather than scar tissue formation. This makes for better healing and less noticeable scarring. For hard tissue reconstruction as in craniofacial reconstruction, calcium phosphates (HA and TCP) are often used. The bioactivity of these materials allows for good integration with the bone. Some polymers are also used for facial defect repair. They include silicones, PGA, PLGA, PCL, and PMMA. These are used for structures without being bioactive. Growth factors are commonly added for enhancing new tissue formation.

Materials in Bioelectrodes and Biosensors Bioelectrodes are devices that send an electrical signal to the body. Biosensors are devices that detect certain biochemical signals and convert them to electrical signals that can be measured. Biosensors are made of a transducer with a biologically active molecule. The biologically active molecule responds to the biochemical signal and relays that response to the transducer. Bioelectrodes and sensors can be on the body surface or implanted. Implanted electrodes and sensors have many more needs for compatibility than the surface ones. Bioelectrodes are commonly made of metals of high electrical conductivity. A common metal is stainless steel due to its low reactivity when in contact with the body. In some cases, copper, gold or platinum is used. Copper needs to be coated in the body as it reacts to the biochemical conditions in the body. Biosensors are made of a transducer with a biologically responsive material. The biologically active material responds to a certain biochemical factor which it wants to detect. This reaction in turn causes a slight change in the material, which also affects a piezoelectric material that makes up the transducer. The slight change in the materials causes a small electrical signal to be transmitted and measured. This is how the device measures the amount of a substance that is present. The transducers can be made of pH- or ion-selective electrodes, thermistors, optical fibers, or piezoelectric crystals. The biologically active material is what really makes the biosensor function. Materials that have been used as the biologically active materials include enzymes, antibodies, DNA, organelles, microorganisms, cells, etc. Each of them can be selected or modified to react specifically to certain biochemical signals. Purified enzymes are commonly chosen because they are specific in their reaction to the factors of interest. Biosensor materials must be able withstand bodily factors, such as temperature, pH, movement and chemically reactive substances while still functioning. There is a new trend in using nanomaterials as biosensors. They include graphene, carbon nanotubes, nanowires, quantum dots and nanocomposites. They have structures that allow them to be modified to detect certain biological factors. Their sensitivity, response time, stability and specificity make them promising materials for future biosensors.

Looking Into the Future Cancer Theranostics Cancer theranostics combine cancer diagnosis and therapy. Diagnostic and therapeutic substances can be incorporated in and released from theranostics. The release can be stimulated by internal or external stimuli. Internal stimuli come from within the body and include pH, redox potential, oxidative stress, enzymatic presence, temperature, etc. External stimuli come from outside of the body and include light, ultrasound, magnetic field, etc. Using biomaterials that can respond to these factors have advantages. Nanoparticles can be used to form theranostics. The nanoparticles can deliver diagnostic materials, such as dyes, that can be used to image the target tumors. They can also deliver cancer drug to cancer cells. Metal nanoparticles are major biomaterials for cancer theranostics. They include gold, silver, quantum dots, iron oxide, and other nanoparticles. There are magnetic metal nanoparticles that can be tracked in the body and manipulated. Some nanoparticles can react to external stimuli, such as ultrasound, to kill the cancer cells that they are attached to. There is a long list of polymers and other biomaterials that can be used in certain forms to react to stimuli and release their cargo. These include modified PEG, PCL, PLGA, polyesters, and many others. The development of novel nanoparticles and nanostructures will allow for better and more effective cancer theranostics in the future.

Bioprinting Bioprinting uses additive manufacturing with biomaterials or biologics to create complex three-dimensional structures for regenerative medicine through the layer-by-layer additive processes. 3D bioprinting is able to create complex structures based on computer designs. It can also create objects with personal anatomical structures on the basis of computer files of medical imaging. It therefore uses a wide range of biomaterials for different types of applications. With a wide range of desired structures and materials to use, there are a number of types of bioprinters. The three main types of bioprinting are based on inkjet, laser/light, or extrusion. Inkjet bioprinting uses multiple mechanisms, such as thermal, piezoelectric actuator, laser-induced forward transfer, and pneumatic pressure, to deposit tiny droplets of the biomaterial onto a substrate. These tiny droplets eventually build up into the desired shape. Stereolithography/projection bioprinting uses lasers or other light sources such as UV light to project a 2D image on a photopolymerizable material. During the process, a stage lowers and the next image is projected, and the layers are polymerized on top of each

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other, creating a 3D structure. Extrusion-based bioprinting uses mechanical forces, such as air pressure or a motor, to extrude material through a nozzle in a specific pattern. The biomaterials used must be able to be extruded but still able to hold their shape once printed. Each bioprinting technique has advantages and limitations and use different materials. The biomaterials used in bioprinting are called bioinks. Bioinks are biocompatible materials that can contain cells and other biological factors. For a bioink to work, it must be able to be bonded to the layers below and above and to be able to hold its shape. The bioinks can have a wide variety of properties, depending on the desired characteristics of the human body tissue it is meant to mimic. Bioinks are made of biomaterials that are either curable or are mechanically tough materials that harden on their own after printing and soft biomaterials. The soft biomaterials are capable of supporting cell growth, whereas the hard materials generally are not suitable for this. Hence the harder materials are usually used as support structures while the softer biomaterials support the cells. The hard biomaterials are generally curable and self-hardening polymers, while the soft bioinks are synthetic or natural polymers that are often formed into hydrogels. Some synthetic polymers used in bioprinting are PEG, and its acrylated version and PCL. PCL is a hard polymer with a melting temperature suited for bioprinting. Some natural polymers that are commonly used for bioprinting are hyaluronic acid and its methacrylated version, collagen, gelatin, and alginate. The soft natural and synthetic polymers can be made into hydrogels for bioprinting, which also allows them to support cellular life and biologics. The fact that they can support cells makes them particularly valuable for 3D bioprinting for tissue regeneration. Each of these biomaterials has different mechanical and biological properties and it is important to use the right biomaterial for a specific purpose. The development of new bioinks and optimizing current ones with cells and biologics will help to realize the potential of bioprinting in tissue engineering and regenerative medicine.

Organs-on-Chips An organ-on-a-chip is a micro-scale system used for mimicking the human body environment. The goal for organ-on-a-chip is to develop human tissue models for disease modeling and drug testing. They use microfluidics, along with cells, to imitate the physiological and mechanical conditions experienced in the body. They can control the movement and behavior of materials and cells by using channels, chambers, membranes, etc. The devices are manufactured using soft lithography and BioMEMS (BioMicroElectroMechanical Systems), which allow for the micro scale details to be properly produced. These fabrication techniques allow the use of different materials, including thermoplastics and thermoset polymers. Most of the materials used to create organ-on-a-chip devices need to be optically clear for viewing and imaging purposes, although whether they are stiff or flexible depends on the use of the device. The materials must also have the right chemistry and reactivity so as to not improperly affect the system. Glass and silicone have been used as materials for microfluidic devices. A commonly used soft, synthetic polymer is polydimethylsiloxane (PDMS). It is optically clear, easy to stretch and easy to fabricate and has high oxygen permeability. Organ-on-a-chip systems that need to be mechanically stable can use thermoplastics such as polystyrene. They are stiff materials with stable surface chemistries. Other synthetic polymers used in making organ-on-a-chip systems are PMMA and polycarbonate. Natural materials, such as collagen, in the form of hydrogels have been used in organon-a-chip systems to assist cell organization. In some cases, biodegradable materials are desired as scaffolds in the system. Materials such as PLGA and polydioxanone (PDO) are thus used. A major requirement for the materials used in organ-on-a-chip systems is that they need to be able to be manufactured with small details. A major technique for manufacture is soft lithography, which normally uses PDMS as the material for chips. Hot embossing and injection molding are also used to make devices from thermoplastics. 2D printing now appears promising for constructing organ-on-a-chip systems.

Acknowledgments The authors thank members of their respective research group in the The University of Hong Kong and University of Nebraska Medical Center for assistance in writing this book chapter. Min Wang thanks The University of Hong Kong, Hong Kong Research Grants Council and the National Natural Science Foundation of China and Bin Duan thanks University of Nebraska and American Heart Association for funding their research in biomaterials and tissue engineering.

Further Reading Agrawal, P., Soni, S., Mittal, G., & Bhatnagar, A. (2014). Role of polymeric biomaterials as wound healing agents. The International Journal of Lower Extremity Wounds, 13, 180–190. Callister, W. D., Jr., & Rethwisch, D. G. (2014). Materials science and engineeringdAn introduction (9th edn.). Hoboken, NJ: John Wiley & Sons. D’Souza, S. F. (2001). Immobilization and stabilization of biomaterials for biosensor applications. Applied Biochemistry and Biotechnology, 96, 225–238. Ebewele, R. O. (2000). Polymer science and technology. Boca Raton, FL: CRC Press. Hench, L. L. (Ed.). (2013). An introduction to bioceramics (2nd edn.). Singapore: World Scientific. Inamdar, N. K., & Borenstein, J. T. (2011). Microfluidic cell culture models for tissue engineering. Current Opinion in Biotechnology, 22, 681–689. Kim, J. J., & Gregory, G. R. D. (2012). Applications of biomaterials in plastic surgery. Clinics in Plastic Surgery, 39, 359–376. Kumar, S., Ahlawat, W., Kumar, R., & Dilbaghi, N. (2015). Graphene, carbon nanotubes, zinc oxide and gold as elite nanomaterials for fabrication of biosensors for healthcare. Biosensors and Bioelectronics, 70, 498–503.

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Langer, R., & Peppas, N. A. (2003). Advances in biomaterials, drug delivery, and bionanotechnology. AIChE Journal, 49, 2990–3006. Lanza, R. P., et al. (Eds.). (2013). Principles of Tissue Engineering (4th edn.). Burlington, MA: Academic Press. Mayet, N., Choonara, Y. E., Kumar, P., et al. (2014). A comprehensive review of advanced polymeric wound healing systems. Journal of Pharmaceutical Sciences, 103, 2211–2230. Molli, R. G., Lombardi, A. V., Jr., Berend, K. R., Adams, J. B., & Sneller, M. A. (2012). A short tapered stem reduces intraoperative complications in primary total hip arthroplasty. Clinical Orthopaedics and Related Research, 470, 450–461. Navarro, M., Michiardi, A., Castaño, O., & Planell, J. A. (2008). Biomaterials in orthopaedics. Journal of the Royal Society Interface, 5, 1137–1158. Ratner, B. D., et al. (Eds.). (2013). Biomaterials science: An introduction to materials in medicine (3rd edn.). Amsterdam: Elsevier. Refojo, M. F. (1982). Current status of biomaterials in ophthalmology. Survey of Ophthalmology, 26, 257–265. Skardal, A., & Atala, A. (2014). Biomaterials for integration with 3-D bioprinting. Annals of Biomedical Engineering, 43, 730–746. Teo, A. J. T., Mishra, A., Park, I., et al. (2016). Polymeric biomaterials for medical implants and devices. ACS Biomaterials Science & Engineering, 2, 454–472. Von Recum, A. (Ed.). (1998). Handbook of biomaterials evaluation: Scientific, technical, and clinical testing of implant materials (2nd edn.). Boca Raton, FL: CRC Press. Wang, Y., Shim, M. S., Levinson, N. S., Sung, H. W., & Xia, Y. (2014). Stimuli-responsive materials for controlled release of theranostic agents. Advanced Functional Materials, 24(27), 4206–4220.

Nano-Biomaterials and their Applications Mian Wang, Northeastern University, Boston, MA, United States Thomas J Webster, Northeastern University, Boston, MA, United States; and Wenzhou Medical University, Wenzhou, China © 2019 Elsevier Inc. All rights reserved.

Introduction Nano-Hydroxyapatite (nHA) Advantage of the nHA Nanostructure nHA Synthesis Biomedical Applications of HA Composites Drug delivery Tissue engineering Carbon Nanotubes (CNTs) Advantages and Disadvantages of CNTs Biomedical Applications (Fabrications and Improvements) Metal Nanomaterials Tissue Regeneration of Metal Nanomaterials Drug Delivery Using Metal Nanomaterials Polymer Hydrogels Drug Encapsulation and Delivery Tissue Engineering Protein Based Nanomaterials Drug Delivery Tissue Engineering Antibacterial Applications Conclusions References Further Reading

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Introduction The utilization of biomaterials in medicine is not new. The first applications of biomaterials happened in ancient Egypt and Greece, where they used biomaterials such as corals, shells, ivory, stone, wood, and some metals like gold and silver as implants for the skeletal system (Piao et al., 2008). The Chinese were the first to use a dental amalgam to repair decayed teeth. More and more biomaterial applications have emerged since then. Recent studies have focused on the use of small scale biomaterials into a growing number of biomedical applications (Kang et al., 2014). The discovery and manipulation of innovative nanomaterials, such as nano-hydroxyapatite, carbon nanotubes (CNTs), metallic nanomaterials, protein-based nanomaterials, and other nanostructured materials are becoming increasingly common in biomedical applications in general. Nanomaterials for different pharmaceutical applications, like tissue engineering, drug or DNA delivery, drug formulations, dentistry, and medicine have all been made. In this article, we review some known methods for obtaining nanomaterials with multiple properties for numerous biomedical applications. We then demonstrate the multifunctionalization of nanomaterials and the improvement in medicine from those modifications, as well as emerging applications.

Nano-Hydroxyapatite (nHA) Biomimetic nanohydroxyapatite [nHA, Ca10(PO4)6(OH)2] has received much attention in bone regeneration, owing to their similarity in chemical composition to that of natural bone mineral. Bone is a complex composite with a porous structure consisting of inorganic and organic components. The inorganic part is mainly hydroxyapatite, which is the hydroxyl end member of the complex apatite group. Therefore, nHA plays an essential role in bone defects and related diseases, which can significantly affect the quality of life. In the US alone, the number of orthopedic procedures has more than doubled, from 138,700 in 2000–310,800 in 2010 (Thein-Han and Misra, 2009). The number of hip replacements performed in the US has continuously increased, and the procedure has become more common in younger people. Traditional orthopedic implants use various metals such as titanium and its alloys. However, these conventional implants face a lot of problems, such as corrosion, mechanical issues, and biocompatibility issues.

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To achieve better bone cell adhesion and osseointegration, hydroxyapatite has been considered as the primary material which can create bioactive and biomimetic interfaces with traditional implants. Another reason that nHA has been widely used in bone regeneration is that nHA has a high surface to volume ratio and high surface energy giving nHA an ability to serve as a reinforcement in biomaterial composites. Studies have revealed that nHA biomaterials possess enhance bioactivity and mechanical properties to conventional, or micron structured, hydroxyapatite. These reasons allow nHA to form a promising osteoconductive bond between native bone and implant replacements.

Advantage of the nHA Nanostructure Natural hydroxyapatite is usually found in the form of Ca10(PO4)6(OH)2 with a Ca/P ratio of 1.67. Although other ratios of calcium phosphates exist, the ratio of 1.67 is the most thermodynamically stable and hence is the most used calcium phosphate material in bone regeneration studies. nHA with fine crystals apparently improved mechanical properties of composites when compared with microHA or noncrystalline HA. In natural bone, nHA plays a vital role in reinforcing mechanical strength. Besides, type I collagen fibers serves as the extracellular matrix providing support for cell adhesion and proliferation (Shen et al., 2016). Therefore, two components of bone enable its strong properties and allow for new bone formation. nHA has been proven to have a positive impact on cellular behavior such as osteoblast cell adhesion, proliferation, and bone formation when compared with large sized hydroxyapatite (micro and macro) (Shen et al., 2016). In addition, studies have indicated the importance of nHA in in vivo studies, indicating constant bone formation when using nHA. Webster et al. published a series of reports that the surface roughness of implants created by adding nHA had a positive effect on cell proliferation and osteogenic differentiation (Wang et al., 2017). During this process, cellular markers of cell adhesion, proliferation, and differentiation, such as alkaline phosphatase synthesis and calcium deposition, were observed in greater amounts on a nanostructured hydroxyapatite based surface than on amorphous and micro hydroxyapatite surface. Protein adsorption on nHA, which leads to initial cell attachment, was also found to be higher than standard hydroxyapatite according to those studies. The same group also indicated that whereby tartrate-resistant acid phosphatase synthesis (TRAP, a phenotypic marker for osteoclasts which express this enzyme) and resorption pitting, which are primary activities of osteoclasts was significantly greater for cells exposed to nanostructured surfaces (compared to a hydroxyapatite sample with a surface area greater than 100 nm).

nHA Synthesis It is possible to synthesize hydroxyapatite with a wide variety of methods, such as wet chemistry, hydrothermal, mechanical–chemical synthesis, plasma spraying, and sol-gel. Generally, there are two main synthesis ways: wet chemistry and high-temperature methods. For this section, we will introduce two typical methods, wet chemistry and hydrothermal methods. Wet chemistry is a precipitation method which depends on a solution environment of calcium and phosphorus ions to precipitate onto an exposed surface. It is a wet method that enables calcium phosphate precipitation. The advantage of the wet chemistry method is that it is cheaper and a more accessible process when compared with other hydroxyapatite synthesis methods. However, the disadvantage of this technology also exists in this traditional approach. The size of HA particles in the wet chemical method usually varies significantly, or are not evenly dispersed. To improve dispersion, multiple methods have been developed. For example, ammonium hydroxide is added to control pH at a slow rate under stirring. The precipitation process is slow and crystallinity is controlled by a temperature change. Generally, the higher the temperature the better the crystallinity. An alternative method is by adjusting the pH value of the solution that the samples are exposed to, resulting in the control of the calcium phosphate phase and morphology. For instance, amorphous calcium phosphates are considered to form from pH 5 to 12. If the pH range is narrowed to 5.5–7, hydroxyapatite will develop best with an even morphology. Another method for forming hydroxyapatite uses an ionic solution which mimics human blood plasma. Simulated body fluid was developed by Kokubo and Takadama (2008) and was used for hydroxyapatite formation. The process is thought to occur through dissolution, renucleation, crystal growth and finally transformation to hydroxyapatite. According to this method, hydroxyapatite forms in a very uniform and well-dispersed way. The hydrothermal method of hydroxyapatite synthesis is required for high pressure, high-temperature processes to form a crystalline phase. Hydroxyapatite particles are generated from the reaction of calcium carbonate and diammonium hydrogen phosphate at a temperature range from 200 to 300
C and pressures around 1–2 kbar. Particle sizes synthesized by this method are about 200 nm to 1 mm in a needle-like morphology. Temperature changes in a small range (30
C) will not influence particle sizes too much according to Jiang et al. studies (Chen et al., 2016).

Biomedical Applications of HA Composites Because of structural advantages, high surface to volume ratios, unique physical and chemical properties, hydroxyapatite has been used for a multiple of biomedical applications. In addition, the surface properties of nHA make it easy to modify, possess excellent biocompatibility and noninflammatory responses, which are all important for numerous biomedical applications. Biocompatibility and nontoxic materials are essential requirements for nanocarriers used in drug delivery biomedical applications. To be a qualified nanocarrier, an easily modified surface with covalently immobilized ligands can provide controlled release of proteins or drugs. There are two main groups of using hydroxyapatite as nanocarriers: one is the conjugation of a drug or protein

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with a covalent linkage, while the other is with physical interactions. Covalent linkage can be realized by conjugating amino or hydroxyl functional groups on the surface with linker groups such as iodoacetyls, maleimides, and the bifunctional linker pyridyl disulfide. One of the advantages of those linkages is a sustainable release of drugs or proteins. On the other hand, physical interactions, such as hydrophobic/hydrophilic and electrostatic, can also lead to the coupling of drug molecules with the surfaces of hydroxyapatite (Pàmies and Stoddart, 2013). However, a cationic polymer which is usually used as a physical linker may bring toxicity, which is a big challenge for its utilization in physical interactions. To overcome such constraints, the modification of polymer linkers has been investigated to reduce cytotoxicity and enhance cellular uptake.

Drug delivery Hydroxyapatite has a powerful ability to absorb proteins on its surface. One of the positive aspects of nHA as a drug carrier molecule is its strong ability to adsorb proteins on the surface. Studies have demonstrated the various proteins such as bovine serum albumin (BSA), hemoglobin, and cytochrome were used with HA for multiple purposes. In addition, studies have also revealed the release of BSA from an HA carrier at different pH conditions. There are two stages of HA-conjugated protein release: the initial stage is a fast release via desorption followed by a slow release stage because of crystal dissolution. Chitosan and HA have an ability to form a stable partnership as a drug delivery carrier due to the surface charge properties of those two items. A negatively charged HA core plays a vital role in attracting positively charged chitosan. One study from Venkatesan et al. showed that HA-chitosan nanoparticles were used to sustain the release of Celecoxib for antitumor research (Venkatesan et al., 2016). Less than 15% of drugs were released by a burst effect of the adsorbed drugs, while over 10% of the drug remained after sustained released for 2 weeks. More interesting applications of HA as a drug delivery carrier have been done. Targeted drug delivery and targeted drug therapy are able to transport the designed drug to the affected sites of a body for a better therapeutic effect. A lot of nHA nanoparticles are designed for targeted delivery via conjugation with iron oxides. Investigations have also been performed on magnetic nHA nanoparticles based on other dopants. Tran et al. mentioned that HA-coated iron oxide nanoparticles were able to increase osteoblast functions (Tran et al., 2010). They synthesized HA coated Fe3O4 nanorods which had a saturation magnetization (Ms) of  35 emu/g revealing good osteoblast activity when cultured with osteoblasts.

Tissue engineering The performance of HA materials in different applications mainly rely on their chemical composition, particle size, crystal morphologies, and aggregation. Since they have similar morphology and functions to natural bone composition, HA materials, especially nHA, have been widely used in bone regeneration, as well as other tissue regeneration studies. Bioceramics or biopolymer composites are the best approach to make the artificial bone material with the required properties. One of the most widely used materials is chitosan and its composite biomaterials for orthopedic tissue engineering. Chitosan is able to be processed into numerous forms and used as bone graft substitutes, such as nanofibers, microspheres, membranes, and 3D scaffolds. nHA/chitosan composites are widely studied for bone graft implants and have proven to be excellent in bone tissue engineering, since it has excellent biodegradable, pore-forming ability, and antibacterial properties. The addition of HA to chitosan will further enhance mechanical properties and be able to mimic the natural structure of bone. Wang et al. reported on a composite of HA/chitosan, which was prepared by a freeze-drying method, and used for bone regeneration (Wang et al., 2014). Nanocomposite scaffolds with 20% wt HA had the ideal porous structure with a pore size ranging from 100 to 500 mm. Bone cell attachment and proliferation on the scaffold indicated that the nHA/chitosan is nontoxic and has good cytocompatibility. Collagen is the main component of bone that possesses a fibrous structure with a diameter of 50–500 nm. Numbers studies on nHA/collagen have been undertaken for bone tissue engineering. Collagen has been reported to have several kinds of negatively or positively charged groups as well as uncharged but polar groups on its surface. Based on those groups, charged sites on the surface play an important role to impact the nucleation of the HA crystals on collagen membranes through chemical interactions. nHA/collagen composite scaffolds displayed homogeneous interconnected macroporous structures which have proper mechanical properties to natural bone. PLA is another widely used synthetic polymer in bone tissue engineering because of its biodegradability and excellent biocompatibility. nHA/PLA composites prepared by phase separation methods showed excellent osteoconductive, osteoinductive and mechanical properties. The porosity of scaffolds, prepared by a solid–liquid phase separation method, was up to 85% with a pore diameter of around 64–175 mm. Among those, the size of HA is not only related to the synthesis methods, but also has a crucial influence on cell response. The Webster group has demonstrated that it is beneficial if the surface roughness of synthetic hydroxyapatite is in the nanoscale. In particular, a nanoscale topography appears to have a positive effect on cell proliferation and osteoblastic differentiation. A nano-roughened hydroxyapatite surface was superior to a less roughened one for positive bone cell responses and therefore better osseointegration. More specifically, osteoblast adhesion, proliferation and cell differentiation markers, such as alkaline phosphatase synthesis and calcium deposition, were significantly greater on nanoscale topographical surfaces than on a standard HA sample surface with a roughness greater than 100 nm. Wang et al. fabricated nanoscale, microscale, and amorphous HA/chitosan scaffolds via a freeze-drying method used for a bone regeneration study (Wang et al., 2014). Nanosized HA was proven to impact cell attachment, proliferation, and differentiation for bone generation, when compared to microscale and amorphous phases. The reasons behind such improvements are still largely under investigation. However, a common hypothesis is that protein adsorption such as vitronectin, fibronectin, and bioactivity on surfaces or particles with nanoscale features is different from that on conventional microscaled surfaces.

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Carbon Nanotubes (CNTs) Carbon nanotubes (CNTs) or carbon-based nanotubular structures have received much attention since the 1990s. They are hollow cylinders formed by one (single-walled CNTs, SWCNTs) or several (multiwalled CNTs, MWCNTs) layers of graphene. CNTs can be fabricated via a variety of methods, including chemical vapor deposition (CVD) and arc-discharge, and they have garnered high interest in many different fields due to their unique mechanical, chemical, and electrical properties (Fig. 1).

Advantages and Disadvantages of CNTs Progress in nanomedicine and using nanomaterials in biomedical applications have lately allowed us to develop multifunctional nanosystems consisting of disease diagnosis, targeting, and treatment, all in one. Currently available nano drug delivery systems include nanoparticles, liposomes, dendrimers, carbon nanotubes, nanoshells, and quantum dots. The primary purpose for the development and application of those nanocarriers are for a better therapeutic response as well as low cytotoxicity. Among all nanomaterials, carbon nanotubes are receiving more and more attention for more effective and safe drug delivery. The advantages of carbon nanotubes include high cell membrane permeability, excellent electrical, optical, thermal and mechanical properties, ultra-lightweight and ultra-high surface area, photoluminescence property, photoacoustic effect, and pH-dependent unloading of loaded materials. Meanwhile, there are some disadvantages of carbon nanotubes in biomedical applications, such as nonbiodegradable, aggregation and bundling phenomenon, some toxicity, accumulation in the liver, and that they are insoluble in organic and inorganic solvents.

Biomedical Applications (Fabrications and Improvements) The high aspect ratio of CNTs is a feature to apply CNTs for diverse biomedical applications. They have captured a lot of attention as other nanoscale materials because of their nanometric structure and remarkable list of excellent properties that encourage their exploitation for promising applications. Significant progress has been made to overcome many limitations such as aqueous solubility and cytotoxicity for better applications in biomedical fields. Functionalized CNTs have been widely used in imaging, targeting drug delivery, gene delivery, biosensing, gene delivery, and tissue engineering. The following section demonstrates the possible potential of CNTs in imaging, targeted delivery, tissue engineering, and diagnostics and their future potential for biomedical applications. To overcome barriers like cytotoxicity and dispersibility, functionalization of CNTs provides a possible solution. Functionalization of CNTs can not only enhance their solubility but also provide more functional sites to conjugate genes, drugs, ligands, and agents. For example, PEG-conjugated CNTs helped to impede opsonization in vivo, as well as prevented reticular endothelial system (RES) uptake. PEGylation is the most effective method to enhance in vivo properties of CNTs, and low cytotoxicity. Covalent functionalization of CNTs showed better dispersion. Before covalent functionalization, a carboxylic acid group on the CNTs surface must be activated using reagents like thionyl chloride, oxalyl chloride or N-hydroxysuccinimide in order to obtain highly reactive intermediate groups for stable covalent bonding. CNTs offer many benefits in various applications like tissue engineering, targeted drug delivery, imaging and diagnosis, and photothermal therapy. The bone grafting process requires a bone mimetic scaffold that is associated with the natural bone healing process. An improved healing process may be possible with CNTs assisting new bone formation. CNTs have been investigated to

Fig. 1 Representation of using a SWNT as a nanocarrier for targeted drug delivery of Doxorubicin (Dox) and a schematic of SWNT transportation into a cell. Reprinted from Meng, L., Zhang, X., Lu, Q., Fei, Z. and Dyson, P. J. (2012). Single walled carbon nanotubes as drug delivery vehicles: Targeting doxorubicin to tumors. Biomaterials 33, 1689–1698 (Copyright © 2012, with permission from Elsevier).

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replace collagen in scaffolds for bone growth. CNTs are also used for gene therapy and stem cell differentiation, since CNTs have sufficient contractility for muscle tissue regeneration. PEGylation of CNTs makes them stealth to prevent white blood cells from an inflammatory response on CNTs, allowing them to circulate in the bloodstream for a sustained duration of time. Surface engineered CNTs have emerged as a nano delivery carrier and imaging for disease treatment and health monitoring. Meanwhile, due to extravagant mechanical properties, CNTs could also be embedded into materials as a nanocomposite for tissue engineering applications. The elasticity, mechanical properties, conductivity, and toughness are significantly promoted by the addition of CNTs. CNTs are not only able to facilitate structural reinforcement of a composite, but also enhance cellular performance. CNTs have the capability to absorb light in the NIR region, which results in state jumping of CNTs. During this process, heat generated by a photothermal process will be released in the targeted site. The photothermal agents in an in vivo tumor could induce thermal destruction of cancer cells with sufficient CNT concentrations. Targeted CNTs have been able to spare cells from damage. CNTs have also been used as a contrast agent in imaging and recognition of cancerous cells owing to their remarkable optical properties. Fluorescent agent conjugated CNTs are utilized as a radio-opaque substance providing an image of the desired in vivo organs. De La Zerda et al. revealed the potential use of RGD peptide conjugated SWCNTs in photoacoustic imaging (De la Zerda et al., 2008). Welsher et al. investigated the use of antibody conjugated SWCNTs for probing cell surface receptors such as NIR fluorescent labels (Welsher et al., 2008). CNTs could enhance cell specificity by conjugating antibodies to CNTs, resulting in better internalizing within mammalian cells. Meanwhile, CNTs have not only enhanced drug delivery, but also reduced cytotoxicity which has been found in an Amphotericin nanotube study. In addition, an Amphotericin nanotube study indicated increased antifungal efficacy and reduced toxicity to mammalian cells when compared to amphotericin antifungal study without nanotubes. Upon conjugation with peptides, it may be possible to be used as a vaccine delivery since they have excellent structures for vaccine delivery. With the advancements in molecular dynamic simulations, as well as unique structure, CNTs have potential use as a gene delivery tool. The ability of CNTs to transport DNA across cell membranes is used in gene therapy. During DNA therapy process, CNTs could conjugate with DNA or can load DNA inside tubes. It has been found that gene therapy with b-galactosidase marker gene nanotubes showed higher expression compared to the transfer of naked DNA. Another application of CNT surfaces is the incorporation with carboxylic and ammonium groups which makes them more suitable for the delivery of peptides, nucleic acids, and other molecules. Traditional CNTs are waterinsoluble forms which result in high in vitro toxicity compared with PEGlated CNTs which could increase water dispersible. It has been proved that the extent of toxicity decreases with special functionalization. However, functionalization of CNTs has a negative impact on the elimination of nanotubes. SWNTs have a high renal uptake. Whereas modest liver uptake as compared to singlewalled nanotubes with conjugation to a monoclonal antibody has higher liver uptake and lower renal uptake. In addition, it was also reported that carbon nanotubes, except acetylated ones, when conjugated with peptides produced a higher immunological reaction when compared to free peptides. Therefore, owing to the excellent properties of CNTs such as high surface area to volume aspect ratio, high mechanical strength, high electrical conductivity, ultra-lightweight, and high thermal conductivity, multiple advantages over other drug delivery carriers exist for CNTs.

Metal Nanomaterials Traditional grafting techniques for tissue regeneration are being replaced by tissue engineering approaches of using 3D biomimetic materials. Of these biomaterials, metals have the highest mechanical strength, as well as very good performance in biocompatible properties. For example, lots of metal nanoparticles play an essential role in bone formation and regeneration, since they have an inherent ability to promote osseointegration and osteoconductivity. Besides, a lot of metal nanoparticles could possess antimicrobial activity. This section will review multiple metal nanoparticles, such as Zn, Ti, Zr, B, Sr, Mg, Ag, and Cu along with their significant applications in the biomedical field, including tissue engineering, drug delivery, and antimicrobial studies.

Tissue Regeneration of Metal Nanomaterials Metal nanoparticles have been used in biomedical applications for many years due to many attractive properties, especially for use in dental implant and orthopedic surgery. Their great strength, ductility, and durability provide excellent properties as a bone substitution. Several studies have shown that the incorporation of HA nanoparticles and metallic nanoparticles such as titanium and iron oxide could increase collagen and calcium deposition by osteoblast cells, as well as promoting stem cell differentiation and bone formation. It has been hypothesized that the nanostructural topographical properties generated by the incorporation of nanomaterials play a critical role in the improved biological performance of nanocomposites rather than the chemistry. Indeed, natural tissue is also a nanostructured material, consisting of collagen fibrils and proteins with dimensions in the 100 nm size or less. Although metal nanoparticles showed excellent biomedical applications in multiple aspects, metal ions release may have toxicity and induce some complications in the clinics. Numerous studies have been completed in vitro to study the cytotoxicity of multiple metal nanoparticles. Stainless steel contains more than one kind of metal element and may have more potential safety problems. Some alloys may generate genotoxicity of medical implants, such as TiAlV and CoCrMo. The Cu, Zn, Mg, Ag, Al may have some exceptional properties as essential elements in the human body. For instance, Zn can improve DNA synthesis, enzyme activity, and hormonal activity. On the other hand, these metal elements also can lead to cytotoxicity when only used in high doses. Different metal ions have different mechanisms resulting in cytotoxicity.

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Drug Delivery Using Metal Nanomaterials Recent studies have focused on antiinfection studies of metallic nanoparticles. Infection-induced by bacteria has existed over the past several decades, even though sterilization technology and surgical equipment cleaning has improved. Antibiotics have been used for many years for antibacterial purposes both in vivo and in vitro, as well as sterilization of biomedical equipment. Although sterilization techniques significantly reduce the chance of infection, new technology is still needed to improve resisting bacterial colonization. Passive strategies including surface modification and drug delivery of antibiotics have been proven to reduce the chance of bacterial infection. One of the most direct approaches for improving the efficacy of implant materials is to deliver antibiotics at the implant from a surface coating. Another current strategy for the development of improved antibacterial biomaterials is via surface modification, such as changing surface roughness and decorating with nanomaterials. Researchers filled nanoparticles like silver (Ag), titanium (Ti), zinc (Zn), zinc oxide (ZnO) and zirconia (ZrO2) to further decrease bacterial adhesion of implant materials. The results suggested that nanoparticle combined coatings significantly improved the antibacterial performance relative to conventional Ti and Ti nanotube implant materials (Fig. 2). Puckett et al. determined decreased adhesion and growth of Staphylococcus aureus, Staphylococcus epidermidis, and Pseudomonas aeruginosa on nanorough Ti generated by electron beam evaporation and nanotubular and nanotextured Ti produced by two anodization methods compared to traditional Ti. Silver(Ag) is known for its broad-spectrum antibiotic activity against Gram-negative and Gram-positive bacteria, fungi, chlamydia, mycoplasma and certain viruses, especially antibiotic-resistant strains. There are many implant surface modifications using Ag as the antibacterial agent. At proper doses, it is possible to fabricate coatings with long-term antibacterial characteristic by introducing and controlling Ag release. Zinc (Zn) has been reported to act broadly against a wide range of bacterial adhesion. Notably, the antibacterial properties of Ag have been known for centuries. The chemical nature of silver affords antibacterial activity in multiple ways. Karlsson et al. showed that CuO particles were much more toxic than the Cu ions since CuO particles had more potential to cause DNA damage and oxidative lesions (Karlsson et al., 2008).

Polymer Hydrogels Many improvements to implants and cell interfaces have been made which could be pivotal in translating current polymer or hydrogel coating technologies into the nanoscale. It will benefit long-term implants as well as have a positive effect on an antiinflammatory body response and poor tissue integration. Towards this end, a lot of effort has been made to look into implant-cell interactions and to investigate the possibilities for controlling cell fate, such as adhesion, proliferation, and migration through the nanoscale as well as managed drug delivery.

Drug Encapsulation and Delivery In addition to traditional implants, nanosurface coatings also hold significant benefits over conventional metallic or inorganic semiconductor electrodes as functional molecules, such as extracellular matrix (ECM) molecules or antiinflammatory molecules, can be easily incorporated via simple adsorption, entrapment, covalent binding or the doping process to encourage neural attachment and to reduce inflammatory responses. In recent years, it has been recognized that cells within all tissues are surrounded by a three-dimensional matrix called the ECM, which provides structural support for cellular constituents and initiates critical biochemical and biomechanical dialogues between

Fig. 2 Surface SEM images of NT-TiO2, NT-Ag, and NT-Sr-Ags (Upper row); Cross-sectional SEM images of the samples (Bottom row). Reprinted from Chen, Y. et al. (2017). Antibacterial, osteogenic, and angiogenic activities of SrTiO3 nanotubes embedded with Ag2O nanoparticles. Materials Science and Engineering: C 75, 1049–1058 (Copyright © 2017, with permission from Elsevier).

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various cellular components that eventually regulate cell proliferation, differentiation, migration, and apoptosis. One of the essential strategies for understanding and controlling processes that determine cell-material interfacing is to recreate the microenvironment cells experience in native biological settings. In the context of neural tissues, the intricate 3D structure of neural tissues also plays a huge role in determining neural functions. Motivated by these advances, there’s a trend in converting 2D CP surfaces into functional 3D structures with the intent to further promote tissue integration. In general, a 3D construct of conducting polymers can be created via two primary methods: CPs can be deposited onto nonconducting 3D–structured materials, such as electrospun fibers or 3D–printed scaffold, or a 3D construct can be fabricated directly with a CP material. Although it is still at its early stage, results indicate that porosity, morphology, mechanical properties and mechanical stability in these 3D scaffolding could be tunable; excellent neuronal cell adhesion, proliferation, and alignment can be achieved; and antiinflammatory drugs can be precisely released upon electrical stimulation. Taken together, this represents a promising way to construct biomimetic and biologically active coatings, which will facilitate the development of stable long-term implants with better tissue integration and reduced foreign body responses.

Tissue Engineering One of the most heavily investigated set of molecules to improve the biological responses of nanorough coatings is the ECM due to their well-known role in regulating cell attachment, proliferation, differentiation, migration, and apoptosis. It has been demonstrated that scaffolds containing small, degradable polymer beads that release nerve growth factors (NGF) to mimic the chemical microenvironment of developing tissue were found to improve the viability of fetal neural cells transplanted into rat brains, and a later study further illustrated that this improvement in biological responses is in a dose-dependent manner. Additionally, the discovery of adhesive domains derived from ECM proteins, notably the amino acid sequence Arg-Gly-Asp (RGD), has enabled the development of synthetic nanomaterials that can modulate cell adhesion. Inspired by these advances, the incorporation of ECM molecules with a nanosurface and their subsequent release has been one of the leading strategies to improve the biological responses of nanoscaled coatings. For example, oligopeptides, such as RGD and YIGSR, have been introduced into PEDOT coatings to form a nanoscale surface, and has been shown to increase cell adhesion in vitro, to be able to establish strong connections with the neuronal structure and thereby improve recordings at the coated sites. Results showed that cell survival, attachment, extension, and differentiation were indeed improved on a nanoscaled surface, suggesting the potential to help establish stable nanomaterials and cell interfaces.

Protein Based Nanomaterials Peptide amphiphiles (PAs) are another nanomaterial which receives more and more attention for biomedical applications. PA is composed of a hydrophobic portion and a hydrophilic portion which meditate self-assembled supramolecular structures. Due to its biocompatibility and biodegradability, a number of PA and its functionalized derivates has promising applications in drug delivery, tissue engineering, and antibacterial studies.

Drug Delivery Due to excellent biodegradability, biodegradability, and efficient cell membrane permeability, PAs have exceptional advantages over liposomes and polymeric drug carriers. PAs are designed to target tumor cells via enhancing the permeability and retention (EPR) effect of tumors. Self-assembled PA would be encapsulated by a tumor cell by being passively engulfed by tumors. Chen et al. developed a novel PA of 15 nm micelles loaded doxorubicin (DOX) with a high loading efficiency, long-circulation and in vivo stability (Chen et al., 2011) (Fig. 3). In addition, studies demonstrated that nanoparticles modified with cationic charged peptide sequences conjugated with RGD showed excellent selectivity of the tumor. Chang et al. investigated amphiphilic peptide nanoparticles (APNPs) with arginine-rich and RGD peptide sequences in their hydrophilic head groups (Chang et al., 2015). APNPs encapsulated with curcumin showed excellent antitumor selectivity. APNPs were able to self-assemble into 15–20 nm diameter nanospheres in water, but disassemble at a pH below 4, which can be beneficial for pH-sensitive drug carriers. With pHsensitive properties, self-assembly at an acidic environment allows curcumin to be encapsulated, and when the pH went back to pH 7, the APNPs could reassemble and release curcumin. The curcumin-loaded APNPs were easily encapsulated into MG-63 osteosarcoma cells compared with normal human osteoblast cells within 2 hours. Low water solubility plays a vital role in the process of internalization of drug loaded nanoparticles into a cell membrane. In the first 24 h, curcumin-loaded APNPs decreased 85% of the MG-63 osteosarcoma cell viability, which was remarkably lower than osteoblast cell viability. This novel nanoparticle indicated that the APNP peptide sequence was selective towards osteosarcoma cells.

Tissue Engineering Self-assembled PA nanomaterials have promising potential for tissue regeneration. The ECM is a microenvironment containing insoluble macromolecules and soluble bioactive factors, which can regulate the behavior and functions of cells. Based on a firm understanding of cellular interactions with the ECM, peptide sequences are capable of targeting cell receptors to promote cell

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Fig. 3 A schematic of micelles consisted of amphiphilic triple-helix peptide-PEG conjugates for Dox delivery: (A) The short peptide with PEG chains attached to the middle and the C-terminus; (B) The neighbored amphiphiles formed into subunits and further self-assembled into micelles (C); (D) The TEM image of the self-assembled micelles. Reprinted from Chen, J. X. et al. (2011). Construction of surfactant-like tetra-tail amphiphilic peptide with RGD ligand for encapsulation of porphyrin for photodynamic therapy. Biomaterials 32, 1678–1684 (Copyright © 2015, with permission from Elsevier); Chen, Y. et al. (2017). Antibacterial, osteogenic, and angiogenic activities of SrTiO3 nanotubes embedded with Ag2O nanoparticles. Materials Science and Engineering: C 75, 1049–1058.

functions. PAs modified with bioactive sequences can activate cell receptors and facilitate targeted cell adhesion, proliferation, and activation of endogenous regeneration mechanisms of diseased tissue. A self-assembled PA system with a BMP-2 binding sequence, NH2-TSPHVPYGGS-COOH, was designed and studied by Lee et al. (2013). The osteogenesis effects of the PA nanofibers loaded with BMP-2 was evaluated by the level of alkaline phosphatase (ALP) activity, which is an indicator for osteoblast differentiation from C2C12 premyoblasts. In addition, Hartgerink et al. designed a nanostructured scaffold formed by PA nanofibers, which contained a phosphorylated serine residue that can interact strongly with calcium ions (Hartgerink et al., 2002). Those nanofiber matrices showed excellent bone cell adhesion, proliferation, and bone formation. The nanostructured PA systems have also been revealed to have a promising potential in angiogenesis promotion. The selfassembled nanostructured PA system can keep a low growth factor level, such as vascular endothelial growth factor (VEGF) and fibroblast growth factor-2 (FGF-2), which is a key factor in restoring blood flow during tissue regeneration and wound healing. Conjugating with growth factors onto the self-assembled nanostructures of PAs can stabilize their biological conformation and enhance binding strength for interactions with cell receptors. Those nanofibers had a strong affinity with loaded molecules, which can increase local retention in the microenvironment. For example, the FGF-2 loaded PA heparin-binding nanofibers displayed controlled-release over 10 days with only 57.1% of the growth factor released. The heparin-binding nanofibers combined with growth factors significantly promoted rat cornea neovascularization in vivo.

Antibacterial Applications Antimicrobial peptides (AMPs) are usually short peptides with a 2–9 positive net charge, which increased the interaction between bacterial membrane and peptides. As a result, amphipathic conformations could stabilize the peptide into secondary structures such as a-helix, b-sheet and loop structures are important to AMP activity. The cationic domain can induce the interaction with a membrane surface while the hydrophobic domain would then drive the peptide insertion into the hydrocarbon chain membrane core. A difference from traditional antibiotics, AMPs bind and penetrate into bacterial cell membranes, causing membrane disruption, membrane lysis and eventually cell death. More importantly, cationic AMPs can disrupt cell membranes especially for bacteria membranes which consist of a large number of anionic phospholipids. However, mammalian cell membranes which consist of more zwitterionic phospholipids, have a low association ability of interaction with cationic AMPs. Therefore, a conjugation with hydrophobic domains would enhance the activity of amphiphilic AMPs. For example, self-assembly of amphiphilic peptides would promote the interactions with bacterial cell membranes by increasing the local cationic charge. Cell-penetrating TAT peptides and six arginine residues (R6) were shown to have a strong antimicrobial property. Compared with conventional antibiotics, these peptide-based nanoparticles were superior to penicillin G in killing Gram-positive Bacillus subtilis, and had a lower hemolytic effects than amphotericin B. AMPs have attracted more and more attention because of this selectivity between bacteria and mammalian cells. Some antimicrobial peptides are already in clinical and commercial applications. For these peptides, it is possible to differentiate between those structural features that contribute to the specificity of initial membrane binding of the bacterial membrane. Multiple parameters would optimize the design of novel AMPs. Those parameters that should be taken into consideration include production costs, toxicity against normal healthy cells, and degradation rate. Recombinant DNA techniques could make biosynthesis feasible but the peptides are usually damaged by the bacteria used to produce them. Although there were early hopes that bacteria would not easily develop resistance to AMPs, it is clear that some strains of bacteria already have.

Conclusions Owing to the success of the rapid development of technologies for the synthesis of multiple nanomaterials during the past decades, investigators currently have at their disposal an enormous diversity of available nanomaterials with required parameters with

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respect of size, shape, structure, electrical and strength properties. Moreover, the question that is now on the agenda is the biomedical applications of those nanomaterials. From the standpoint of biomedical applications, much significance was held by the development of efficient technologies for nanomaterials belonging to various classes. This article reviewed multiple nanomaterials including nano-hydroxyapatite, CNTs, metallic nanomaterials, protein-based nanomaterials, and nanostructured polymer hydrogels, as well as multiple biomedical applications, such as drug or DNA delivery, tissue engineering, bioimaging, and diagnosis. Nevertheless, the results here describe the way to use nanomaterials in many biomedical applications and more complex systems; the challenge is now to transfer nanomaterial technology to in vivo applications, and to develop next-generation devices able to impact biomedical performance in a controlled fashion efficiently.

References Chang, R., Sun, L., & Webster, T. J. (2015). Selective inhibition of MG-63 osteosarcoma cell proliferation induced by curcumin-loaded self-assembled arginine-rich-rgd nanospheres. International Journal of Nanomedicine, 10, 3351–3365. Chen, J. X., et al. (2011). Construction of surfactant-like tetra-tail amphiphilic peptide with RGD ligand for encapsulation of porphyrin for photodynamic therapy. Biomaterials, 32, 1678–1684. Chen, Z., et al. (2016). Osteoimmunomodulation for the development of advanced bone biomaterials. Materials Today, 19, 304–321. De la Zerda, A., et al. (2008). Carbon nanotubes as photoacoustic molecular imaging agents in living mice. Nature Nanotechnology, 3, 557–562. Hartgerink, J. D., Beniash, E., & Stupp, S. I. (2002). Peptide-amphiphile nanofibers: A versatile scaffold for the preparation of self-assembling materials. Proceedings of the National Academy of Sciences, 99, 5133–5138. Kang, I. K., Ito, Y., & Kwon, O. H. (2014). Nano/microfabrication of biomaterials. BioMed Research International, 2014. Karlsson, H. L., Cronholm, P., Gustafsson, J., & Möller, L. (2008). Copper oxide nanoparticles are highly toxic: A comparison between metal oxide nanoparticles and carbon nanotubes. Chemical Research in Toxicology, 21, 1726–1732. Kokubo, T., & Takadama, H. (2008). Handbook of Biomineralization: Biological Aspects and Structure Formation (vol. 3, pp. 97–109). Lee, S. S., et al. (2013). Bone regeneration with low dose BMP-2 amplified by biomimetic supramolecular nanofibers within collagen scaffolds. Biomaterials, 34, 452–459. Pàmies, P., & Stoddart, A. (2013). Materials for drug delivery. Nature Materials, 12, 957. Piao, Y., Burns, A., Kim, J., Wiesner, U., & Hyeon, T. (2008). Designed fabrication of silica-based nanostructured particle systems for nanomedicine applications. Advanced Functional Materials, 18, 3745–3758. Shen, X., et al. (2016). Sequential and sustained release of SDF-1 and BMP-2 from silk fibroin-nanohydroxyapatite scaffold for the enhancement of bone regeneration. Biomaterials, 106, 205–216. Thein-Han, W. W., & Misra, R. D. K. (2009). Biomimetic chitosan-nanohydroxyapatite composite scaffolds for bone tissue engineering. Acta Biomaterialia, 5, 1182–1197. Tran, N., Pareta, R., Taylor, E., & Webster, T. J. (2010). Iron oxide nanoparticles: Novel drug delivery materials for treating bone diseases. Thermec 2009 Suppl., 89–91, 411. Venkatesan, J., Jayakumar, R., Anil, S., & Kim, S. K. (2016). Nanocomposites for musculoskeletal tissue regeneration (pp. 161–174). https://doi.org/10.1016/B978-1-78242-4529.00007-8. Wang, M., et al. (2014). Design of biomimetic and bioactive cold plasma-modified nanostructured scaffolds for enhanced osteogenic differentiation of bone marrow-derived mesenchymal stem cells. Tissue Engineering. Part A, 20, 1060–1071. Wang, M., Geilich, B. M., Keidar, M., & Webster, T. J. (2017). Killing malignant melanoma cells with protoporphyrin IX-loaded polymersome-mediated photodynamic therapy and cold atmospheric plasma. International Journal of Nanomedicine, 12, 4117–4127. Welsher, K., Liu, Z., Daranciang, D., & Dai, H. (2008). Selective probing and imaging of cells with single walled carbon nanotubes as near-infrared fluorescent molecules. Nano Letters, 8, 586–590.

Further Reading Chen, Y., et al. (2017). Antibacterial, osteogenic, and angiogenic activities of SrTiO3nanotubes embedded with Ag2O nanoparticles. Materials Science and Engineering: C, 75, 1049–1058. Meng, L., Zhang, X., Lu, Q., Fei, Z., & Dyson, P. J. (2012). Single walled carbon nanotubes as drug delivery vehicles: Targeting doxorubicin to tumors. Biomaterials, 33, 1689–1698.

Natural Biopolymers for Biomedical Applications Natalia Davidenko, Ruth Cameron, and Serena Best, University of Cambridge, Cambridge, United Kingdom © 2019 Elsevier Inc. All rights reserved.

Introduction Definition, Classification, Sources Requirements, Advantages and Disadvantages of Natural Polymers Demand and Areas of Application Properties and Applications Polysaccharides Glucans Seaweeds-derived polysaccharides Chitin-chitosan Mammalian glycosaminoglycans Proteins Combination of Polysaccharides and Proteins Examples of Commercially Available TE Biomedical Implants and Wound Healing Products Concluding Remarks Further Reading

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Glossary Antiangiogenic properties The ability of substance to stop the growth of tumors and progression of cancers by limiting the pathologic formation of new blood vessels (angiogenesis). Anticoagulant activity The ability of substances to prevent or reduce coagulation of blood, prolonging the clotting time. Bioactivity A beneficial or adverse effect of a substance on living tissue (organism). Biocompatibility The ability of a material to perform with an appropriate host response in a specific situation. Biodegradability The capacity of being decomposed by biological agents, especially bacteria. Hemostatic activity An act to arrest bleeding or hemorrhage. Immune response The reaction of the cells and fluids of the body to the presence of a substance which is not recognized as a constituent of the body itself. Immunogenicity The ability to provoke an immune response in the body of a human or animal. Self-assembly A process in which a disordered system of preexisting components forms an organized structure or pattern as a consequence of specific, local interactions among the components themselves, without external direction. Thrombogenicity The tendency of a material in contact with the blood to produce a thrombus, or clot.

Introduction Definition, Classification, Sources Biological structures are a source of inspiration for solving challenges in material science and medicine. Nature has developed, with comparatively few base substances, a range of materials with remarkable diversity of functional properties and biological activities. The term “natural” has always been used to refer to something that is present in- or produced by nature (rather than being artificial or man-made) and is often synonymous with something which is good, pure or healthy. Throughout history, humans have relied on natural biological materials such as wool, leather, silk, and cellulose. Natural biopolymers may be defined as biological macromolecules that are produced by living organisms and that can be used in/as biomedical materials. In general, they are composed of repeating structural units which are linked together to form long chains or branched networks. The simple building blocks are called monomers, while the more complicated building units are sometimes referred to as “repeat units.” There are three main classes of natural biopolymers, categorized according to the monomeric units used and the structure of the biopolymer formed: (1) Proteins (include polypeptides) which are macromolecules composed of a-amino acids; (2) Polysaccharides, which are linear or branch-bonded polymeric carbohydrate structures; (3) Polynucleotides (nuclei acids DNA and RNA), which are polymers composed of 13 or more nucleotide structural units.

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Natural polymers, such as nucleic acids and proteins, carry and manipulate essential biological information, while others, like polysaccharides,dthat is nature’s family of sugarsdprovide fuel for cell activity and serve as structural elements in living systems. Biopolymers, playing a central role in the natural world, can come from many forms of life and at any point or level of the Evolutionary Tree. Interestingly, the form of lifedbe it simple or highly organized- does not define their usefulness in the biomedical area. The natural sources of biopolymers include (1) microorganisms, (2) plants, or (3) animal tissues and can originate from aerial, terrestrial, or marine living organisms. For example, microorganisms including bacteria, smuts, nests, yeasts, molds, fungi, and many other forms of what is considered to be a primitive life can be very valuable to provide an enormous variety of polymeric biomolecules with excellent structural and biochemical characteristics. These biopolymers include polysaccharides (cellulose, dextran, chitin, hyaluronic acid, etc.) and proteins (silk, keratin, etc.). Plants have also, over time, been an extremely valuable and renewable source of both polysaccharides (cellulose, starch, alginate, carrageenan, etc.) and proteins (soy, zein, wheat gluten, etc.). Animals, whether highly developed or poorly developed, whether they live on land, in the sea, or in the air can provide natural structures with strong potential in biomedical field. Examples of these structures include glycosaminoglycans (chitin, heparin, hyaluronic acid, etc.), proteoglycans and proteins (collagen, elastin, gelatin, heparin, etc.) as well as deoxyribonucleic acid, the genetic material originated from all living creatures.

Requirements, Advantages and Disadvantages of Natural Polymers Any substance to be used in- or as a biomedical material, needs to fulfill some important requirements which derive from the definition of a biomaterial as “a substance (other than a drug) synthetic or natural in origin, that can be used as a system or as a part of a system that treats, augments, or replaces any tissue, organ or function of the body.” As such, an ideal biomaterial is one that is biocompatible, biodegradable and nontoxic or immunogenic. It is also extremely convenient that this material can be extracted or processed at reasonable cost and in different forms, and that it provides a possibility to be functionalized, if required, with bioactive chemicals. Remarkably, many naturally derived biopolymers are inherently biodegradable, biocompatible and nontoxic. In addition, some of them carry specific protein binding sites and other biochemical signals that may assist in tissue healing, regeneration and integration. These properties make them favorable contenders for biomedical applications in different fields of medicine. However, natural biopolymers, as many other natural materials, can be associated with several problems. These include evoking immunogenic reaction after being implanted; the batch to batch variability in molecular structure associated especially with animal sourcing, and the tendency to denature or decompose at temperatures below their melting point. Another major problem yet to be overcome with natural starting materials is their propensity for calcification (the abnormal deposition of calcium in a body tissue) and they are also prom to bio-deterioration. Some important advantages and disadvantages of natural biopolymers are summarized in Table 1. It should be stressed that current advances in extraction, purification and characterization of proteins and polysaccharides from different natural sources provide effective solutions to some of these drawbacks and contribute to the development of a diversity of new biomaterial platforms.

Demand and Areas of Application The use of natural products as biomaterials is not a new area of science. The field is extremely multidisciplinary encompassing elements of medicine, biology, chemistry, tissue engineering, and materials science. The demand for biocompatible, biodegradable, and functional structures is constantly increasing and this is a tendency which is expected to continue in the future. For example, in the US the demand is currently growing at a compound annual rate of 6.2% reaching $7.12 billion in 2018. Broadly, the main area of applications of natural biopolymers include (1) implantable devices, (2) wound management products, (3) controlled drug delivery systems, and (4) tissue engineering scaffolds. Reviews have appeared recently describing the structure, properties, classification and potential use of natural biopolymers in different biomedical fields including orthopedics, ophthalmology, dentistry, soft and cardiovascular tissue regenerations, among others. Some commercially available biopolymers are relatively expensive but their application in specialized biomedical fields can justify their costs. Improvements in technologies for their extraction and processing have expanded the use and incorporation of natural biopolymers as biomaterials; especially as drug-delivery vehicles and tissue

Table 1

Advantages and disadvantages of natural biopolymers

Advantages

Disadvantages

➢Biocompatible ➢Nontoxic, do not elicit a foreign body response ➢Possess biological functionality on both macroscopic and molecular levels ➢Biodegradable via natural enzymes ➢Degradation kinetics can be adjusted by modification including crosslinking ➢Can be obtained at reasonable cost

➢May elicit immunological reaction ➢High natural lot-to lot variability ➢Structurally more complex than synthetic materials ➢Technological manipulation may be more elaborate ➢Propensity for calcification

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engineering scaffolds. These and other applications of the natural biopolymer-based materials in biomedicine are described in following sections. A comprehensive overview of the most important members of each family of natural biopolymer (according to classification) is given and includes proteins and polysaccharides. Polynucleotides are mainly employed in molecular biology principally for diagnosis and monitoring of hereditary diseases, in forensic science and in parentage testing for the analysis of genetic fingerprints for DNA profiling, etc. These areas are not of direct interest in the biomaterial field and, as such, are out of scope of this article.

Properties and Applications Polysaccharides Polysaccharides are families of natural sugar polymers that serve two main functions in living systems: for energy storage and as extracellular structural elements of plant or animals. Polysaccharides comprise 75% of all organic material on earth. They are also known as glycans, the term which is defined by IUPAC as synonym to polysaccharide, meaning “compounds consisting of a large number of monosaccharides linked by glycosidic bonds.” The exact placement of the linkage can vary, and the orientation of the associating functional groups may result in a- and b-glycosidic bonds. Polysaccharides may be linear (cellulose, chitin and hyaluronan) or branched (starch, dextran) and may also differ in ring size (furanose or pyranose), absolute configuration (D- or L-) and in the chemical identity of substitutes (hydroxyl, sulfate, carboxylate, amine, etc.). Examples of polysaccharides relevant to the biomaterial fields include cellulose (plant and bacterial), starch, dextran, alginate, carrageenan, chitin/chitosan, pullulan, agar, pectin, hyaluronic acid (hyaluronan), chondroitin sulfate, and heparin. Some of these polysaccharides and their structural units are given in Figs. 1 and 2. The numbering in the polysaccharide nomenclature indicates the location of the linking carbon in the ring. Glycans are mainly of two types: those made up of one type of monomer unit (homo-polysaccharides) or those composed by two or more types of monosaccharide residues (hetero-polysaccharides). Most homo-polysaccharides store fuel (starch) or are extracellular organizational elements (cellulose and chitin) while hetero-polysaccharides (hyaluronan and chondroitin sulfate) typically provide cellular support for microorganisms and animal tissues.

Fig. 1

Polysaccharides: Structures of glucans and seaweed-glycans.

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Fig. 2

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Polysaccharides: Structures of glycosaminoglycans (GAGs).

The term glycan should not be confused with that of glucan which specifically defines a polysaccharide comprising D-glucose monomers meaning that not all glycans are glucans. For example, cellulose, composed of b-1,4-linked D-glucose is glucan (Fig. 1) while chitin, comprised by b-1,4-linked N-acetyl-D-glucosamine monomeric unit (Fig. 2), is glycan, but not glucan. Polysaccharides, as the other classes of biopolymers, may be derived from different natural sources including microorganisms, plants and animals. They may be classified a nonmammalian and mammalian according to/depending on their origin. The most abundant nonmammalian polysaccharides include cellulose, starch, chitin, alginate and dextran which can be extracted and purified using relatively simple procedures originating large quantities of material at low cost. This economic advantage coupled with their low immunogenicity makes these natural polymers attractive for biomedical applications. Isolation and purification of mammalian polysaccharides, such as chondroitin sulfate, hyaluronan and heparin, are more difficult than those of the nonmammalian polysaccharides, but they exhibit important biological functionality, including specific binding with multiple proteins, which has motivated interest in their use in biomedical application. Many naturally-derived polysaccharides exhibit properties which make them ideally suited for bio encapsulation technology and tissue engineering. These properties include a unique gelation ability, high water binding capacity, biodegradability and similarity to extracellular matrices (ECM) as well as the possibility to be easily modified and processed into adequate forms (micro /nanoparticles, films, hydrogels, and 3D scaffolds). Unfortunately, there are concerns about their purity and pathogen content especially when extracted from sources such as algae, insects, bacterial expression, and/or mammalian tissues. Different approaches have been developed to address these concerns; some of them are mentioned in this article. Polysaccharides are polyelectrolytes, macromolecules that possess a relatively large number of functional groups which either are charged, or under suitable conditions can become charged. In fact, their polyelectrolyte nature is one of the most important characteristics of this class of biopolymers. The interaction between oppositely charged polyelectrolytes (by reversible electrostatic and dipole–dipole bonding or/and hydrogen and hydrophobic bond formation) leads to development of polyelectrolyte complexes (PEC). The great advantage of these complexes is that they can generally be formed by a simple procedure consisting of simultaneous mixing of the corresponding polyelectrolytes in solution without need for any additional chemicals as covalent crosslinkers. As such these complexes are generally nontoxic, well-tolerated and biocompatible; very important characteristics for their use as/in biomaterials. Furthermore, the stability and other important features of these PECs may be adjusted and modified in many ways, for example, by changing the density of charges, degree of ionization, pH of reaction medium, concentration of polyelectrolytes, distribution of ionic groups, molecular weight, mixing ratio, order of reacting polyelectrolytes and drying process. Recently, polysaccharides-based PECs have been extensively investigated for biomedical applications, such as drug encapsulation and controlled delivery, DNA-binding, enzyme immobilization, tissue engineering and biosensor. A brief description of the structure, properties and areas of biomedical applications of selected nonmammalian and mammalian polysaccharides, relevant to the biomedical field, are reviewed as follows under specific headings. This selection includes: (a) Glucans: cellulose (linear b-glucan from both plant and bacterial sources), starch (branched a-glucan of plant origin), and dextran (branched a-glucan of bacterial origin).

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(b) Seaweeds-derived polysaccharides: alginate and carrageenan, both of nonmammalian origin and of anionic nature but with different functional groups in the saccharide rings. (c) Chitosan: a unique positively charged nonmammalian glycosaminoglycan (GAG) with important characteristics for biomedical use. (d) Mammalian glycosaminoglycans: hyaluronan, chondroitin sulfate, and heparin, all highly used in numerous biomedical applications.

Glucans Cellulose, starch and dextran are all composed of D-glucose monomers but differ in the length of chains; the type of the linking units and in the degree of branching; which define their diverse properties and applications. Cellulose may be of plant or bacterial origin. Both possess the same molecular formula, composed of repeated D-glucose building blocks (Fig. 1). They are linear homo-polymers characterized by their high hydrophilicity, biodegradability and broad chemical modifying capacity. Plant cellulose is the most abundant organic compound on Earth used for more than 150 years as a chemical raw material. The bacterial form may be produced by certain bacteria (Acetobacter xylinum), algae and fungi. Due to differences in their origin, these celluloses significantly differ in their properties and architectural characteristics. Plant cellulose generally forms a tough, mesh-like bulkwork structure and contains impurities such as hemicellulose or lignin, while microbial cellulose, with ultrafine micro-fibril architecture, is more chemically pure, more crystalline and more porous. The bacterial form has higher swelling capacity and hydrophilicity as well as a greater tensile strength than its plant-derived counterpart. Furthermore, bacterial cellulose can be produced on a variety of substrates and can be grown to virtually any shape due to the high moldability during formation. Cellulose, independent of its origin, is degraded by microbial enzymes. Animal cells cannot cleave the b-(1 / 4)-bond between the two glucose moieties, thus, cellulose degradation in tissues takes place by a slow nonenzymatic hydrolysis and therefore cellulose can be regarded as an almost stable matrix. Both types of cellulose and their derivatives are well-tolerated by most tissues and cells and as such cellulose-based nontoxic biocompatible materials offer numerous possibilities in medical fields. For example, cellulose sponges have been used widely in experimental surgery as they do not affect the healing process, but act as a chemoattractant inducing cells involved in the repair process to migrate towards them. Cellulose-based biomaterials are also used as coatings for drugs, supports for immobilized enzymes, artificial kidney membranes, in wound care and as implant materials and scaffolds in tissue engineering. Starch, a branched a-glucan, is one of the most abundant and cheap polysaccharides. Natural starch occurs in a granular form, and it is a principal carbohydrate storage product of plants. It is found in a variety of botanic sources such as cereals (maize, wheat, rice, cassava, etc.) and tuber plants (e.g., potatoes). The content of amylose and amylopectin in starch is largely dependent on the source. Most often, starch consists of about 30% amylose, a linear a-(1–4) glucan, and 70% amylopectin, a highly branched version. Starch is a highly biodegradable polymer as it can be broken down into its constituent sugars by enzymes (known as amylases) which are found in plants and in animals, including human saliva and pancreas secretions. Starch is insoluble in cold water, but it is very hygroscopic and binds water reversibly. It is biodegradable, inexpensive, and as such can play an important role in the medical field. The unique physicochemical and functional characteristics of starches such as swelling power, solubility, and capacity of gel formation, rheological characteristics, mechanical behavior and enzymatic digestibility make them potentially useful for a wide variety of biomedical applications ranging from bone cements, drug carriers for controlled release to tissue engineered cellular supports. Dextran is a nonmammalian glucan. This polysaccharide is expressed by certain lactic acid bacteria, the best-known being Leuconostoc mesenteroides and Streptococcus mutans. The structure of dextran consists of linear (1 / 6)-a-D-glucose, with branches extending mainly from (1 / 3) and occasionally from (1 / 4) or (1 / 2) positions. Dextran is highly water-soluble and easily functionalized through its reactive hydroxyl chemistries. Characterization of many types of dextran have indicated that branching, average molecular weight and molecular weight distributions can vary widely depending on the conditions and strain of bacteria used for expression. In the early 1940s, dextran was investigated as a blood plasma replacement and has since become of interest as a biodegradable and biocompatible material. Dextran biodegradation occurs through natural enzymatic splitting of saccharide bonds by dextran-1, 6-glucosidase found in spleen, liver, lungs, kidneys, brain, and muscle tissue as well as by dextranases expressed by bacteria in the colon. Dextran lacks nonspecific cell binding and resists protein adsorption, characteristics which have increased interest in its use as a biomaterial. Like other nonmammalian glucans, the relatively low cost and availability of dextran as well as its hydroxyl functionality for chemical modification has enlarged the utilization of dextran in the field of polysaccharide-based biomaterials. For example, dextran has recently been investigated as an alternative to polyethylene glycol for low protein-binding and cell-resistant coatings on biomaterial surfaces. It has been shown that the multivalent properties of dextran are advantageous when high-density surface immobilization of biologically active molecules to low protein-binding surface coatings is desired.

Seaweeds-derived polysaccharides Seaweed-originating nonmammalian polysaccharides have been used extensively over decades for biomedical and biological applications including tissue engineering, drug delivery, wound healing and biosensor being alginate and carrageenan among the most widely employed resources. Alginates represent a whole family of linear copolymers containing blocks of alternating b-D-mannuronic acid and a-L-guluronic acid residues with (1 / 4)-linkages, displaying carboxylic acid functionality at the C5 residue. Alginates are hydrophilic non

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mammalian polysaccharides commercially extracted from marine brown algae (Phaeophyceae). They can also be produced by soil bacteria such as Azotobacter and Pseudomonas. Depending on source and processing, alginates have broad distributions in molecular weights (from 10 to 1000 kDa) and chemical structures. Bacterial alginate has more defined chemical composition and physical characteristics than its seaweed-derived equivalent. In the unmodified form, alginates are not enzymatically degradable in mammals and do not support cellular adhesion. Nevertheless, alginates have been extensively investigated for many biomedical applications, due to their biocompatibility, low toxicity, relatively low cost, and mild gelation by forming ionic complexes with divalent cations such as calcium. Over decades, sodium alginate has been indispensable part of dental practice as a cost-effective and simple impression material. More recently, alginates have been used for coencapsulation of proteins and growth factors, such as vascular endothelial growth factor. The experimental procedure is very simple and consists of the simultaneous addition of alginate and protein solutions into calcium chloride solutions. Alginate, in form of hydrogels has been particularly attractive in wound healing, drug delivery, and tissue engineering applications. Alginate-based wound dressings maintain a physiologically moist microenvironment, minimize bacterial infection and facilitate wound healing. The lack of cellular adhesion has been addressed by covalent association to its structure of different cellular adhesive peptides and proteins. As negatively charged polyelectrolyte, alginate can interact electrostatically with the positively charged macromolecules, for example chitosan, forming PECs. These PECs possess many improved physicochemical properties and advanced biomedical applications, especially for controlled drug delivery, in comparison with alginate calcium gels. These alginate-chitosan delivery systems, in the form of nanoparticles, microspheres and hydrogel beads, have been used to encapsulate small drugs, including anticancer, ocular, asthma and pulmonary drugs, antiinflammatory medications, etc. In form of membranes and films alginate-base PECs have been employed to encapsulate biological macromolecules (proteins, peptides, and nucleotides) for implantable drug delivery systems. The term carrageenan defines a naturally occurring biopolymer family of linear sulfated polysaccharides, extracted from red marine algae (Rhodophyta). The main chain of the carrageenan molecule is composed of D-galactose and D-anhydrogalactose units linked by glycosidic bonds. The properties of carrageenan are defined by its structural characteristics. Depending upon sources and the extraction methods, they are generally of three types: kappa (k), iota (ι), and lambda (l), mainly differing in number of the sulfate groups per disaccharide repeating unit and the presence of the 3,6-anhydro bridges. As a result, these carrageenans show different linear charge density and solubility, both of which increase with sulfate group content (k < ι < l). Among the three types of carrageenan, only k- and ι-varieties possess gelling ability, with the former being firm and rigid while the latter being soft and elastic. Different strategies have been applied to exploit carrageenans biocompatibility and attractive physicochemical properties including special gelling mechanisms, strong water absorption capacity and abundant functional groups. Due to the presence of the ester sulfate groups, carrageenan molecule is highly negatively charged and has the ability to interact with positively charged molecules forming PECs. Carrageenan incorporation into drug delivery matrices increases drug loading and drug solubility, enabling the release of orally administrated medications in zero-order or in a significantly prolonged period. Other features of carrageenan-based formulations, such as pH-sensitivity and adhesive properties, contribute to its broad application in pH / temperature-sensitive delivery systems in response to physiological environments. Some promising results have been achieved using carrageenan-containing tissue engineering devices due to their capacity to induce important osteogenic, adipogenic, and chondrogenic differentiation in stem cells. However, it has been shown that carrageenan may provoke some adverse effects on the biological system such as unwanted immune responses and inhibition of blood coagulation. As such there is a pressing demand for a careful and comprehensive analysis of all carrageenan properties for a better development and a safer use of this biopolymer in biology and medicine.

Chitin-chitosan Chitin, a homopolymer of b (1 / 4) linked N-acetyl-D-glucosamine, is the second most abundant biopolymer in nature. It is commonly found in the exoskeletons of crustacean and insects as well as in the cell walls of fungi. Because of its insolubility in aqueous media, most chitin applications are based on its deacetylated form, chitosan, which occurs rarely in nature, but can be efficiently obtained by extensive deacetylation of chitin. The resultant chitosan consists of b (1 / 4) linked glucosamine and N-acetylglucosamine, randomly or block distributed throughout the biopolymer chain, depending on the preparation method and source material. N-deacetylation not only increases the aqueous solubility of the polymer but also provides reactive primary amines for further chemical modifications. The microstructure of chitosan and its solubility are influenced strongly by the characteristics of the deacetylation procedure: chitosan is soluble when obtained under homogeneous conditions and insoluble when heterogeneous deacetylation is carried out. The molecular weight and degree of deacetylation are important parameters determining chitosan properties and applications. Chitosan itself possesses hemostatic, antimicrobial, antiviral and antitumoral activities. Over recent decades it has been one of the most popular naturally-derived biopolymers in biomedical field due to its biocompatibility (approved for human use), excellent biodegradability (degrades by human enzymes, such as lysozyme and lipase), low toxicity, wound healing properties, strong bactericidal effect as well as abundant availability and low production cost. Many chitosan-based formulations have been investigated for wound-dressing materials, drug encapsulation and drug-delivery devises. As a unique positively charged polysaccharide, chitosan can form PECs by electrostatic interactions with many negatively charged natural biopolymers, for example anionic nonmammalian (alginate, carrageenan, cellulose, etc.) and mammalian (hyaluronan, chondroitin sulfate, etc.) polysaccharides; essential components of ECM (such as proteoglycans and proteins) and other negatively-charged biomolecules (growth factors, nucleic acids, cytokines, etc.). Biomaterials fabricated from chitosan-based polyelectrolyte complexes in forms including nanoparticles, microspheres, beads, tablets, gels, films, and membranes, are especially popular in drug encapsulation applications. Recently, increasing attention has been paid to chitosan-based biomaterials in tissue-engineering approaches

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particularly in epithelial and soft tissues repair and regeneration. Both types of tissues play an important role in supporting anatomical structures and physiological functions of living organisms. Chitosan-based matrices have been investigated, for example, to repair body surface linings, for regeneration of connective tissue, and to support nerve connection and vascular growth. It has been reported that chitosan may be beneficial in regulating many tissue regeneration functions including preservation of cellular phenotype, binding and enhancement of bioactive factors. Chitosan also assists in controlling gene expression and in synthesis and deposition of tissue-specific extracellular matrix. Although chitosan itself is nonadhesive to cells it can be efficiently modified by attaching to its structure of cell-adhesive polymers such as peptides, proteins and mammalian GAGs. This modification results in providing important cell-recognition sites to otherwise inert chitosan-based formulations. The choice of chitosan as a base component for many tissue supports is determined by the fact that its biological, physical and chemical properties can be modified, controlled and engineered by multiple ways and under mild conditions. The novel use of chitosan-based scaffolds in promoting regeneration of various tissues demonstrates that these materials may serve as promising substrates for a great number of future applications.

Mammalian glycosaminoglycans Mammalian glycosaminoglycans are complex polysaccharides that play important roles in cell growth, differentiation, morphogenesis and migration. They can be classified as nonsulfated, such as hyaluronic acid (HA), and sulfated, such as heparin, chondroitin sulfate, dermatan sulfate, and keratin sulfate. Many GAGs, (except from hyaluronic acid) are generally attached to a central protein forming proteoglycans. Due to their polyelectrolyte nature, GAGs possess high water absorption capacity. Their chemical versatility and biological performance make them attractive building blocks for great variety of biomaterials aimed especially at tissue engineering applications. Hyaluronic acid (HA), also known as hyaluronan or hyaluronate, is a natural, nontoxic, biocompatible and biodegradable mammalian polysaccharide composed by 2-acetamide-2-deoxy-a-D-glucose and b-D-gluconic acid residues linked by alternate (1 / 3) and (1 / 4) glycoside bonding. Hyaluronan, a unique nonsulfated mammalian glycosaminoglycan, is found naturally in the extracellular matrix of skin, cartilage, vitreous body of the eye, and other body tissue. It also occurs as an extracellular polysaccharide in a variety of bacteria. HA has a high capacity of lubrication, water sorption, and water retention, and influences several cellular functions such as cell migration, adhesion, and proliferation. For biomedical applications it is produced mainly via microbial fermentation, to reduce the risk of crosspieces viruses, infection, and contamination. HA clinical uses include dermal filling, viscosupplementation in deteriorated joints and ophthalmic surgical aids. For such applications, HA has the advantage of being biocompatible and safe, inducing minimal foreign body reaction. Owing to its important physiological role in tissue repair, HA is also used for wound healing. HA can be chemically modified in two different ways: crosslinking or conjugation. Due to the high density of negative charges, hyaluronan can form PECs with different cationic biopolymers among which chitosan is one of the most explored. HA-based PECs in forms including nano- and micro-particles, films, membranes, hydrogels and porous matrices have been investigated for drug delivery systems and tissue engineering. Hyaluronan itself has been indicated to impact cell–cell and cell–substrate interactions, to aid in the organization of proteoglycans and to promote angiogenesis. This range of properties together with its high lubrication capacity and low cytotoxicity have motivated the widespread use of HA in ophthalmic surgery, arthritis treatment, in vocal fold repair, wound repair and healing, antiinflammatory materials, drug delivery, and tissue engineering scaffolds, especially for soft tissue repair and regeneration. Chondroitin sulfate (CS), is a sulfated mammalian polysaccharide, widely distributed as one of the most physiologically important GAGs in ECMs and at cell surfaces. It is mainly found attached to proteins as proteoglycan in connective tissues, working as a structural component, or on cell surface and basement membranes, functioning as a receptor. CS is composed of repeating disaccharide units of D-glucuronic acid and N-acetyl galactosamine sulfate at either 4- or 6-positions. Commercially available CS is obtained via extraction and purification from sources including shark and whale cartilage, and bovine or porcine tissues. CS is a polyelectrolyte with strong negative charges and as such it could efficiently interact with proteins and different positively charged polyelectrolytes (for example chitosan), forming PECs and other associates. This provides multiple ways for modification and tailoring of physicochemical, mechanical and biological properties of CS-based biomaterials. CS, as an ECM component, plays important roles in modulating the stability, activity, release, and spatial localization of growth factors. It also contributes to wound healing, tissue morphogenesis, hemostasis and cell division in vivo. Its biodegradability, biocompatibility, high versatility and availability make CS a promising candidate for biomedical applications. The ability of CS to induce the synthesis of cartilagespecific markers, to inhibit cartilage destruction processes and to stimulate the cartilage formation explains its wide use in cartilage engineering platforms. Other TE applications of CS include repair and regeneration of damaged bone, skin and neural tissues. In spite of its potential as biomaterial there are some concerns associated with the quality differences and properties variability of different samples of chondroitin sulfate due the fact that it is extracted frim living organisms. Recent technological advancements in extraction and purification methods have significantly contributed towards overcoming these potential drawbacks, increasing clinical utility of CS-base formulations. Heparin and heparin sulfate are both sulfated heterogeneous linear mammalian glycosaminoglycans composed by a or b (1 / 4) linked uronic acids (90% a-L-iduronic acid, 10% b-D-glucuronic acid) and a-D-glucosamine residues, containing a heterogeneous mixture of carboxylic acids, hydroxyls, sulfates, and amine functional groups. Heparin is only produced in mast cells, where it is cleaved from the core protein (serglycin), while heparin sulfate is found in most tissues where it remains attached to the core protein forming proteoglycan. Both play an important role in many biological processes. Commercial production of heparin from mammallian tissues such as porcine or bovine intestinal mucosa involves complicated extraction and purification processes which add to

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heterogeneity in chemical structure and molecular weight of naturally-derived heparin. The most important characteristics of heparin, for biomedical use, include its anticoagulant activity and ability to bind and stabilize growth factors preventing them from denaturating and increasing the affinity of complex to cell receptors. As such, heparin integration to biomaterials was originally aimed at reducing the thrombogenicity of materials in contact with blood. Heparin also exhibits antiangiogenic and apoptotic effects, characteristics which favored its use as a component in tumor-targeted delivery systems. The high negative charge of heparin promotes its ionic interactions with proteins, such as growth factors, proteases, and chemokines, and other natural biomolecules. For practical use heparin is often conjugated to naturally-derived hydrogels to provide sustained release of the anticoagulant, small drugs and protein-based systems. Recent biomedical applications include heparin-containing nanoparticles and scaffolds for tissue engineering, where heparin is generally associated with other polysaccharides (e.g., hyaluronan, chitosan) or proteins, including collagen, gelatin, fibrin, etc. Heparin-based biomaterials in form of nanoparticles and scaffolds have been employed as cell carriers in therapeutics for cancer treatments and in tissue engineering applications. Heparin-functionalized nanocarriers stimulate cell differentiation and promote therapeutic efficacy of drugs by increasing the cellular uptake of the delivered molecules. Due to its biocompatibility, low toxicity, benign gelation conditions and advantageous biological activities together with its relatively low cost heparin incorporation in biomaterials is constantly growing and extending to new emerging applications.

Proteins Proteins form a vast variety of natural tissues and organs from which they can be extracted and processed. Some of the most familiar protein materials include wool, leather, silk, gelatin, and collagen. Proteins are large biological macromolecules composed of one or more long chains of a-amino acids. Once linked in the protein chain, an individual amino acid is called a residue, and the linked series of carbon, nitrogen, and oxygen atoms are known as the main chain or protein backbone. A linear chain of amino acid residues is called a polypeptide. Fig. 3 shows the structures of a-amino acid, the formation of the peptide bond in polypeptide chain and the three-dimensional representation of protein structure. Proteins differ from one another primarily in their sequence of a-amino acids, which is dictated by the nucleotide sequence of their genes. They are in fact unique biopolymers where the position of the exact amino acid in the chain is determined by a specific template (DNA). Like other biological macromolecules (polysaccharides and nucleic acids), proteins are essential parts of living organisms where they play a vital role in cell signaling, immune responses, cell adhesion, and the cell cycle. Some of them are enzymes that catalyze biochemical reactions and are essential to metabolism; the others have structural or mechanical functions, conferring stiffness and rigidity to biological components. Most structural proteins such as collagen and elastin, for example, can be derived from ECM of mammalian connective tissue and cartilage where they are crucial components. Other important structural protein, keratin, is found in hard- or filamentous structures such as hair, nails, feathers, hooves, and some animal shells. The great advantage of proteins is that they are structurally identical or very similar to macromolecular substances which the biological environment is prepared to recognize and deal with metabolically. As a result, the problems of toxicity and simulation of chronic inflammatory reaction are suppressed or highly reduced. Over decades proteins have been viewed as potential resources for many biomedical platforms owing to their intrinsic ability to perform very specific biochemical, mechanical and structural roles. ECM-derived structures, in particular, have become an obvious

Fig. 3 Structures of Proteins. Peptide chain is from https://commons.wikimedia.org/wiki/File:Peptide_bond.png by Webridge. Protein structure is from https://commons.wikimedia.org/wiki/File:Proteinviews-1tim.png by Opabinia regalis.

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starting point in the development of devices in different tissue engineering approaches due to their similarity in composition and properties to native tissue. Most of them possess a defined, three-dimensional microstructure that supports cell-guided tissue formation providing appropriate binding motifs for cell attachment, proliferation and maintenance of phenotype. As such, the use of collagen, albumin, gelatin, silk fibroin and keratin in the naturally-derived biomaterials has been increasing. Due to the importance of proteins in the biomedical field, selected members of this class of natural biopolymers (namely, silk, keratin, elastin, and collagen) are reviewed as follows under specific heading. Particular focus is directed towards collagen as this is the most significant and widely used component in tissue engineered devices and other biomedical materials. Further details on its origin, type, and structural diversity are provided and the results of our research group in developing collagen-based formulations, especially for tissue engineering applications are presented. Silk is a natural fibrous protein produced by Lepidoptera larvae such as silkworms, spiders, scorpions, mites, and flies. Silk fibers are composed of at least two main proteins: sericin (globular protein) and fibroin (fibrous protein). This biopolymer has always been a material of great fascination. Spiders can process silk protein into a material that has a tensile strength 16 times greater than that of nylon and a very high degree of elasticity. Silk without sericin displays higher stability than others. Silk fibroin shows excellent physical and chemical properties and for decades it has been used as a surgical suture material. As a biomaterial, silk can be processed in different forms including fibers, films, membranes, gels, sponges, powders, and scaffolds. Due to its low toxicity, biocompatibility, tunable biodegradation rate and remarkable mechanical characteristics silk, has been popular choice for a broad range of tissue engineering applications. The unique mechanical properties of silk fibroin together with its ability to support the differentiation of mesenchymal stem cells along the osteogenic lineage; predicates its use especially for bone tissue engineering. Silk fibroin can be also combined synergistically with other biomaterials to form composites and can be chemically modified. As such silk fibroin scaffolds can be tailored to specific applications including vascular prostheses, structural implants, nets, and sutures among others. The term “keratin” was initially referred to the broad category of insoluble proteins originated from hair, wool, horns, hooves and nails. Mammalian keratins can be classified into two distinct groups, “hard” and “soft” keratins, based on their structure, function and regulation. By their origin they can be divided in epithelial and hair keratins both containing nonhelical domains and central alpha-helical domain. Hair keratins are richer in cysteine residues in their nonhelical domains and thus form tougher and more durable structures. Advances in the extraction, purification, and characterization of keratin proteins from hair and wool fibers have led to the development of numerous keratin based biomaterial platforms. Like many naturally-derived biomolecules, keratins have intrinsic biological activity and biocompatibility. In addition, extracted keratins are capable of forming self-assembled structures that regulate cellular recognition and behavior. These qualities have directed its use in forms including powders, films, gels, coatings, fibers, foams and scaffolds in wound healing, drug delivery and tissue engineering devices. Elastin is an ECM protein that is known to provide elasticity to tissues and organs and allows many tissues in the body to resume their shape after stretching or contracting. It is most abundant in connective tissues and other body structures where elasticity is of major importance. Keratin comprises up to 70% of the dry weight in elastic ligaments, about 50% of large arteries, 30% of lung, and 2%–4% of skin tissues. The insoluble elastin is one of the most stable proteins with a half-life of about 70 years. Tropoelastin, the precursor protein of elastin, and elastin-like peptides possess an important ability to self-assemble under physiological conditions. Biomaterials based upon elastin and elastin-derived molecules are increasingly investigated for their application in tissue engineering. This interest is fueled by the remarkable properties of this structural protein, such as elasticity, self-assembly, long-term stability, and biological activity. Elastin can be applied in biomaterials in various forms, including insoluble elastin fibers, hydrolyzed soluble elastin, recombinant tropoelastin (or tropoelastin fragments) and as block copolymers of elastin. Furthermore, repeated elastin-like sequences can be produced by synthetic or recombinant means and incorporated to biomaterials. Used alone or in combination with other biopolymers, elastin-containing biomaterials have been suggested for applications including skin substitutes, vascular grafts, heart valves and elastic cartilage. Collagen is the most abundant of all proteins found in mammals, typically accounting for more than 30% of total body protein mass. It forms a significant part of connective tissues such as bone, tendons, ligaments, blood vessels, nerves and skin. Collagen, being a principal structural protein present in all vertebrates, comprises a family of genetically distinct molecules with a common triple helix configuration of three polypeptide subunits, known as a-chains (Fig. 4). These triple helices form a molecule of tropocollagen, the basic building block of collagen. Tropocollagen molecules associate in a staggered fashion to produce collagen fibrils, which are strengthened and stabilized mainly by enzymatically and nonenzymatically catalyzed covalent crosslinks. The extent of these crosslinks is age-dependent and tissue-specific. To date 28 types of collagen have been identified and described. Variations in collagen type are due to differences in the assembly of the polypeptide subunits, lengths of the helix, interruptions and terminations of the collagenous helical domains. The best known and the most abundant are fibrillar collagens I, II and III, each containing different affinity cell-recognition motifs that support cellular activity, mainly through interaction with cell-surface integrins a1b1, a2b1, a10b1, and a11b1. In biological systems all these collagens have important functions: Col I, a major ECM component, accomplishes both structural and cell adhesive roles in many vital organs and tissues while Col II is the chief element in articular cartilage (approximately 60% of the dry weight of this tissue) and Col III is an important component of reticular fibers, where it is commonly found alongside Col I. For biomedical purposes, for decades, collagen, especially type I, in forms including gels, films, nanoparticles, microspheres, membranes and scaffolds has been the principal biomaterial of choice. Its use is based on characteristics such biocompatibility, biodegradability, low immunogenicity as well as on its ability to form strong fibers with high tensile strength. In addition, collagen displays numerous triple-helical cell binding motifs, GFOGER being a major and most important ligand of collagen I.

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Fig. 4 Chemical structure of collagen (showing Gly-Hyp-Pro repeats, left), and the triple-helical arrangement of collagen peptides (right). Reproduced from Goodsell, D. RCSB Protein Data Bank. 2000. www.rcsb.org/pdb/101/motm.do?momID¼4.

In tissue engineering applications, collagen is often used in combination with other biomolecules such as proteins and polysaccharides. Collagen-containing scaffolds are biologically active and typically promote cell adhesion and growth. Furthermore, they are biodegradable and, over time, allow host cells to produce their own ECM and replace the degraded scaffold. Along with these attractive physical and biological attributes collagen, especially type I, is abundant, easy to purify at reasonable cost and to modify. Furthermore, it can be easily processed in different forms and shapes which additionally broaden its use in tissue engineering constructs, wound management products and drug delivery strategies. For example, in the form of nanoparticles collagenous formulations have been employed for gene delivery; as mini-pellets and tablets for protein release, in the form of thin films as shields in ophthalmology and as sponges for burn and wound healing products. Collagen gels, in combination with liposomes, have been successfully used as a controlling material for transdermal delivery and as matrices for cell culture systems. However, fabricating biomaterial devices from naturally-derived collagens with suitable stability and reproducible structures presents some challenges. One of these concerns is the poor mechanical properties of collagenous matrices for TE applications, which limits their use in different biological environments and especially in load bearing situations (e.g., orthopedic). It should be stressed, that in its native state collagen possesses multiple inter- and intramolecular cross-links, which provide strength and durability to their matrices. However, extraction and the purifying methods applied to process it into useful forms markedly reduce collagen cross-linking density with a consequent impact on material properties and biological response. To overcome this limitation, collagen-based scaffolds are usually crosslinked by different chemical (aldehydes, isocyanates, carbodiimides, etc.) and physical (ultraviolet irradiation, dehydrothermal treatments, etc.) procedures. Among them, the chemical treatment based on carbodiimide hydrochloride (EDC) in the presence of succinimide constitutes one of the most successful methods for stabilization of collagenous biomaterials in TE applications. This treatment is highly efficient, nontoxic (cross-linkers do not take part in linkage), and the resultant by-products can be easily removed by washing. However, EDC-promoted bonding presents a significant drawback as it uses free primary amino groups (on lysine residues) and carboxylate anions (on glutamate or aspartate residues) for crosslinking. Many of these amino acid side chains, for example the acidic E of GFOGER, are essential components of cell binding sites on collagen-type materials and so EDC-promoted crosslinking can impinge on bioactivity of collagenous formulations. This should be addressed for the successful use of collagen-based scaffolds as templates for tissue regeneration. Over recent years, intensive research in our Center has been directed to development nano-and macro-particulate systems for controlled drug-delivery, and for production of tailor-made three-dimensional collagen-based scaffolds. The latter 3D matrices have been mainly designed for (1) in vivo use as 3D engineered environments for regenerative medicine and (2) as models of natural tissue for in vitro applications. A controlled freeze drying experimental approach has been employed to obtained highly porous, interconnected sponges of collagen (mainly type I) alone or in combination with glycosaminoglycans (chondroitin sulfate, hyaluronic acid, heparin, and chitosan) and other proteins (gelatin, elastin, fibrinogen, and peptides). These scaffolds have been directed mainly to repair and replacement of bone, tendon, cartilage, ligament, skin, nerve and myocardial tissues. Recent in vitro applications have included cell therapies to aid in the culture of megakaryocytes for in vitro production of platelets from adult stem cells and in vitro models of mammary glands to assist in drug screening for breast cancer. These in vitro models of mammary glands have been developed to reduce the use of animals in testing and to assist in the development of personalized medicine initiatives through tailored assays of breast biopsy material. To achieve a desirable biological performance from 3D engineered matrices, several important parameters have been finely tuned. These include the nature and availability of cell binding ligands, the material (swelling profiles, degradation rates, etc.) and the mechanical properties of these scaffolds as well as their morphology and spatial characteristics. To accomplish this goal, the scientific base of freeze-drying technology has been broadened, different treatments collagenous formulations have been applied and appropriate biomolecules have been incorporated into collagenous matrices. By applying compositional, positional and mechanical controls at the developing stage we have obtained collagen-based cellular supports with varying local biochemical, spatial and mechanical environments, closely matching those of natural tissues. Furthermore, 3D structures with controlled and

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complex pore orientation that closely mimic many normal multioriented tissue arrangements have been produced recently. Fig. 5 shows some examples of different scaffold morphologies achieved by a combination of freeze-drying technique with molding technology, which has been designed to induce uniaxial or multidirectional temperature gradients in collagen slurries during scaffold fabrication stage. The attachment to collagen-based scaffolds of novel peptides designed to control and guide cell behavior has opened avenues for successful restoration of their biological activity after crosslinking treatment. A range of characterization methodologies have been developed to assess the contribution of the important structural determinants of the biological performance of collagen-based materials. These include the control of porosity, interconnectivity, percolation diameter, pore size and morphology together with adjustment of swelling properties, dissolution characteristics, structural stability as well as mechanical control at both a local and macroscopic level to provide mechanical cues for cell activity. The research is continuing to achieve the ultimate goal of producing systems with structural and functional properties of native tissues for novel emerging approaches.

Combination of Polysaccharides and Proteins The integration and extensive use of natural polymers as biomaterials can be significantly expanded by combination of different biopolymers in the same formulation. Typically, each class of biopolymers employed in the fabrication of biomaterials has specific advantages and disadvantages. The development of composite materials comprised of different components, frequently produces a desired synergistic effect, so enhancing the material biological performance and practical use. As mentioned previously, while describing design and properties of different biopolymers, the addition to biologically inactive nonmammalian polysaccharides of GAGs or proteins (peptides) enables a biological recognition to be introduced to otherwise benign polymeric systems. As such, the combination of nonmammalian polysaccharides with bioactive polymers has led to the design of many robust systems while maintaining the low cost and low immunogenicity of the material with specific receptor responses and cell stimulation. Combination of proteins and peptides with the natural polysaccharides can be achieved via versatile and efficient protocols to provide biologically active composite materials that can be recognized by enzymes or cells. The precise amino-acid sequence and structural conformations of these peptides and proteins dictate high affinity binding constants and efficient assembly compared to those of polysaccharides alone. This leads to more specific control over biomaterial design and properties. Taken together, the

Fig. 5 Examples of different morphologies of collagen scaffolds. Degree and type of anisotropy in the microstructures of the scaffolds were achieved by imposing temperature gradients, uniaxial or multidirectional, during the phase of crystallization of water in collagen suspensions, using molding technology. Images from Cambridge Centre for Medical Materials, University of Cambridge, UK.

Table 2

Examples of commercially available biomedical implants based on naturally-derived biopolymers Product

Composition

Use/Form

Skin

Apligraf, Organogenesis

Lower layer of human fibroblasts and bovine collagen, upper layer of keratinocytes Allogenic fibroblasts and human collagen with additional layer of keratinocytes Porous bovine collagen crosslinked with chondroitin-6-sulfate with upper layer of silicon Granulated bovine collagen crosslinked with chondroitin-6sulfate ECM protein (amelogenins) inpropylene glycol alginate carrier Human fibroblasts on a laser-microperforated membrane of benzyl hyaluronate Human keratinocytes on alaser-micro perforated membrane of benzyl hyaluronate Bovine type I collagen sponges Bovine type I collagen Porous foam of bovine type I collagen Composite of human demineralized bone matrix with type I collagen Composite of nanocrystalline hydroxyapatite and porcine collagen and dextran co-polymer Human mineralized bone matrix in porcine gelatin carrier Composite of highly purified natural bone mineral combined with native collagen Hyaluronic acid (Hylan GF-20 and Hylan B) from chicken combs Rat-tail type I collagen matrix Bovine type I collagen with hyaluronic acid and GAGs HYAFF (esterified derivative of hyaluronic acid) scaffold Type I collagen membrane Bilayer collagen matrix Gelfoam porcine gelatin foam sponges Bovine type I collagen nerve conduits

Leg ulcers/Sheet

ICX-SKN, Intercytex Integra dermal regeneration template, Integra lifesciences Integra flowable wound matrix, Integra lifesciences Xelma, Molnlycke Hyalograft 3-DTM (fidia advanced biopolymers) Laserskin™ (fidia advanced biopolymers) Bone

INFUSE bone graft, medtronic OP-1, Stryker Vitoss Scaffold FOAM, Orthovita Bioset IC, pioneer surgical FortrOss, pioneer surgical

Cartilage

Blood vessel Nerve

Regenafil, regeneration technologies/exatech Orthoss® Collagen and Geistlich Bio-Oss® Collagen, Geistlich Pharma AG Synvisc, genzyme CaReS, arthro kinetics Menaflex, regenbiologics Hyalograft C autograft, Fidia advanced biopolymers MACI, genzyme unreg Chondro-Gide®, Geistlich pharma AG VascuGel, pervasis NeuraGen, integra

Burns and acute wounds/Sheet Burns/Sheet Ulcers/Gel Epidermal skin substitute/Membrane Epidermal skin substitute/Membrane Spinal fusion/Solid Bone injury/Paste Bone injury/Foam Bone injury/Paste Bone injury/Paste Bone injury/Paste Orthopedic bone regeneration/Porous block Synovial fluid replacement/Gel Articular cartilage injury/3D disc Meniscus cartilage injury Articular cartilage injury/Mesh Articular cartilage injury/Sheet Cartilage regeneration/Porous block Vessel reconstruction/Tubular Nerve injury/Tubular (Continued)

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Examples of commercially available biomedical implants based on naturally-derived biopolymersdcont'd

Tissue

Product

Composition

Use/Form

Pancreas

Islet sheet, Cerco medical

Diabetes mellitus/Sheet

Amcyte/ReNeuron Novocell Durepair® Dura regeneration matrix, medtronic Duraform dural graft implant, depuy Biodesign® Dural graft, cook biomedical Duramatrix, Collagen Matrix Inc. Biodesign fistula plug, cook medical

Alginate sheet with encapsulated pancreatic islets (pig or human) Alginate/poly-L-lysine encapsulated human pancreatic islets Alginate/PEG encapsulated human pancreatic islets 3D matrix of type I and type III collagen from fetal bovine tissue Collagen- based biocompatible graft Collagen-based scaffold Matrix from highly purified type I collagen Collagen scaffold

Permacol surgical implant (Covidien) Neuragen nerve guide, Integra): Collagen tendon sheet, Rotation Medical Inc. GraftJacket® regenerative tissue matrix, wright medical

Cross-linked porcine collagen for enhanced durability Bovine type I collagen and GAG conduit and collagen matrix Resorbable type I collagen matrix Noncross-linked human dermal collagen scaffold

Extracel, glycosan biosystems

Crosslinkable bacterial hyaluronic acid, bovine and porcine gelatin, porcine heparin and PEG derivatives Comprised primarily of collagen type I Collagen membrane

Hernia supports/Mesh Peripheral nerve conduit/Conduit, porous matrix Tendon protector/Matrix Support, protection, and reinforcement of tendon and ligamentous tissue/Scaffold Examples: cartilage, bone, vocal fold TE/Gel, sheet, tubular Soft tissue repair/Mesh Oral tissue regeneration/Membrane

Collagen 3D matrix

Soft tissue regeneration/Spongy or compact blocks

Dura mater

Enter cutaneous fistula Abdominal wall Peripheral nerve Tendon/ligament Various

Miromatrix biological mesh, miromatrix medical Inc. Geistlich Bio-Gide® and geistlich BioeOss pen, geistlich pharma AG Geistlich Mucograft® and geistlich Mucograft®Seal, Geistlich Pharma AG

Diabetes mellitus/Beads Diabetes mellitus/Beads Dural draft substitute/3D matrix Dural draft substitute Dural substitute Dural substitute/Membrane matrix Fistula filler/Scaffold

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development of polysaccharide–protein composites has permitted the production of increasingly diverse and useful materials with tailored biological responses and chemical and mechanical properties to better mimic the properties of the natural extracellular matrix. This has led to important new classes of environmentally sensitive materials with response to the biological environment and stimuli, that can further increase the efficacy of materials as tissue replacements.

Examples of Commercially Available TE Biomedical Implants and Wound Healing Products The history of the development and marketing of TE biomedical implants has shown that a medically effective product is not in itself sufficient to ensure commercial success. Apart from the obvious scientific, medical and human benefits, for it to reach the clinic it needs to bring a clear financial reward to those who invested in translating a new technology to the clinic. As such, for material to be able to compete successfully in the market a combination of clinical performance, marketing and cost-effectiveness should be effectively achieved. This explains why so many promising formulations have not been converted into commercially offered products. Nevertheless, there are a growing number of TE implantable biomaterials, based on naturally-derived biopolymers, which are already approved and commercially available as therapies for soft and hard tissues. These tissues include skin, cartilage, bone, blood vessel, pancreas, nerve, abdominal wall, tendon, and ligament. A nonexhaustive set of examples of biomedical implants, based on naturally-derived biopolymers that are currently on the market is presented in Table 2. With the advancement in technology, more than 3000 products have been developed to treat different types of wounds by targeting various aspects of the healing process. Based on the cause and type of wound, numerous products are available in the market, among them being bioactive dressings, a modern generation of wound healing materials. They are generally composed of naturally-derived biopolymers such as collagen, hyaluronic acid, chitosan, alginate, and elastin, and are available in different forms including gels, pads, particles, pastes, powders, sheets, or solutions (more than 900 products). Some examples of the commercially existing wound healing products, based on naturally-derived biopolymers, are shown in Table 3.

Table 3

A nonexhaustive set of examples of commercially available wound dressing products based on natural biopolymers

Principal components

Product and manufacturer

Form

Alginate (calcium/sodium)

AlgiDERM (Bard), AlgiSite(Smith & Nephew, Inc.), Algosteril (Johnson & Johnson), CarraSorb H (Carrington), CURASORB and CURASORB Zinc (Kendall), Dermacea (Sherwood-Davis & Geck) FyBron (B. Braun), Gentell (Gentell), Hyperion Advanced (Hyperion), Alginate Dressing (Medical, Inc.), KALTOSTAT (ConvaTec), KALGINATE (DeRoyal), Maxorb (Medline), PolyMem (Ferris Mfg.), Restore (Hollister), SORBSAN (Dow Hickam), SeaSorb (Coloplast Sween Corp.), Tegagen HG and Tegagen HI (3 M Health), etc. Granuflex™ (Conva Tec), Comfeel™ (QuickCompany), Tegasorb™ (AliMed), WALKER® (Compeed), etc. HemCon (Medical Technologies, Inc.), HemCon Bandage (Cath Lab Digest, SP Services UK Ltd., Chitoderm® plus (Trusetal, POmedic, etc.) etc Biofill® (Cellulose Solution LLC.), Bioprocess® and XCell® (Medline Industries, Inc.), etc. Exm. Hyiodine (H&R Healthcare), IDRA® hydrogel (Trusetal), Hyalomatrix KC Wound Dressing (Laserskin), etc. BIOPAD™(Angelini Pharma, Inc.), Helix3® Bioactive Collagen (Amerx Health Care Corp.), Stimulen™ Collagen Powder (Southwest Technologies, Inc.), Cutimed® Epiona (BSN medical, Inc.), CellerateRX® Gel and CellerateRX® Powder (Wound Care Innovations, LLC), DeRoyal C-Pro® 3D (DeRoyal), Gentell Collagen (Gentell Wound and Skin Care), Puracol® Plus MicroScaffold™ Collagen (Medline Industries, Inc.), Simpurity™ Collagen Pad and Simpurity™ Collagen Powder (Safe n’ Simple), Stimulen™ Collagen Gel (Southwest Technologies, Inc.), SkinTemp™ II Dressings Human (BioSciences, Inc.), Triple Helix Collagen Dressing (MPM Medical, Inc.), etc. FIBRACOL™ Plus (KCIdAn Acelity Company), etc. DermaCol™ (DermaRite Industries, LLC), ColActive® Plus (Covalon Technologies, Ltd.),etc. BIOSTEP* Collagen Matrix (Smith & Nephew, Inc.) PROMOGRAN® Matrix (KCIdAn Acelity Company),etc. Endoform Dermal Template™ (Hollister IncorporateddWound Care)

Films, patches, membranes, foams

Carboxymethyl cellulose, gelatin and pectin Chitosan Bacterial cellulose Hyaluronic acid Collagen

Collagen, alginate Collagen, alginate, cellulose and EDTA Collagen, matrix metalloproteinases Collagen, cellulose Collagen, ECM proteins

Thin films, sheets Films, sheets Membrane, films Viscose gel, pads Sponges, thin films, sheets, powders, sprays, gels, ropes, porous pads

Sheets 3D pads 3D pads Pads, sheets Sheets

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Concluding Remarks A variety of biomaterials has been developed from biopolymers using approaches inspired by nature. While major advances in the use of natural biopolymers have taken place with increasing numbers of products entering the market and into clinical trials, significant research is still required. Currently increasing number of technologies are entering the clinical and commercial fields with more ongoing research dedicated to the development of sophisticated biomimetic biomaterials with added levels of complexity and stimuli-responsive properties. New and expanding research in this area demonstrates just how multidisciplinary the field of biomaterials has become and while the challenges are huge, the opportunities for improving human health in a whole variety of areas are immense. Advances in these areas will facilitate the development of new generations of biopolymer-based platforms with controlled function, offering expanded options for guiding cellular behavior for applications in tissue replacement and other areas. Recent technological progresses and growing knowledge in biochemistry, molecular biology, and bioengineering have added to the optimization of biological performance of biopolymer-based formulations, expanding significantly, the areas available for emerging applications.

Further Reading Cen, L., Liu, W., Cui, L., Zhang, W., & Cao, Y. (2008). Collagen tissue engineering: Development of novel biomaterials and applications. Pediatric Research, 63, 492–496. Daamen, W. F., Veerkamp, J. H., van Hest, J. C. M., & van Kuppevelt, T. H. (2007). Elastin as a biomaterial for tissue engineering. Biomaterials, 28, 4378–4398. Davidenko, N., Schuster, C. F., Bax, D. V., et al. (2015). Control of crosslinking for tailoring collagen-based scaffolds stability and mechanics. Acta Biomaterialia, 25, 131–142. Klemm, D., Heublein, B., Fink, H.-P., & Angew, A. B. (2005). Cellulose: Fascinating biopolymer and sustainable raw material. Angewandte Chemie International Edition, 44, 3358–3393. Kogan, G., Soltes, L., Stern, R., & Gemeiner, P. (2007). Hyaluronic acid: A natural biopolymer with a broad range of biomedical and industrial applications. Biotechnology Letters, 29, 17–25. Koha, L.-D., Cheng, Y., Tenga, C.-P., et al. (2015). Structures, mechanical properties and applications of silk fibroin materials. Progress in Polymer Science, 46, 86–110. Kwoni, H. J., & Han, Y. (2016). Chondroitin sulfate based biomaterials for tissue engineering. Turkish Journal of Biology, 40, 290–299. Lee, K. Y., & Mooney, D. J. (2012). Alginate: Properties and biomedical applications. Progress in Polymer Science, 37, 106–126. Liang, Y., & Kiick, K. L. (2014). Heparin-functionalized polymeric biomaterials in tissue engineering and drug delivery applications. Acta Biomaterialia, 10, 1588–1600. Liu, J., Zhan, X., Wan, J., Wang, Y., & Wang, C. (2015). Review for carrageenan-based pharmaceutical biomaterials: Favourable physical features versus adverse biological effects. Carbohydrate Polymers, 121, 27–36. Rouse, J. G., & Van Dyke, M. E. (2010). A review of keratin-based biomaterials for biomedical applications. Materials, 3, 999–1014. Yang, T.-L. (2011). Chitin-based materials in tissue engineering: Applications in soft tissue and epithelial organ. International Journal of Molecular Sciences, 12, 1936–1963. Baldwin, A. D., & Kiick, K. (2010). Polysaccharide-modified synthetic polymeric biomaterials. Biopolymers, 94, 128–140. Coimbra, P., Ferreira, P., Alves, P., & Gil, M. H. (2014). 1. Polysaccharide-based polyelectrolyte complexes and polyelectrolyte multilayers for biomedical applications. Carbohydrates Applications in Medicine, 2014, 1–29. Dash, M., Chiellini, F., Ottenbrite, R. M., & Chiellini, E. (2011). ChitosandA versatile semi-synthetic polymer in biomedical applications. Progress in Polymer Science, 36, 981–1014. Gagner, J. E., Kim, W., & Chaikof, E. L. (2014). Designing protein-based biomaterials for medical applications. Acta Biomaterialia, 10, 1542–1557. Ige, O. O., Umoru, L. E., & Aribo, S. (2012). Natural products: A minefield of biomaterials. ISRN Materials Science, 2012, 1–20. https://doi.org/10.5402/2012/983062. Lee, C. H., Singla, A., & Lee, Y. (2001). Biomedical applications of collagen. International Journal of Pharmaceutics, 221, 1–22. Kaplan, D. I. (Ed.). (1998). Biopolymers from renewable resources. Verlag Berlin Heidelberg New York: Springer. Luo, Y., & Wang, Q. (2014). Recent development of chitosan-based polyelectrolyte complexes with natural polysaccharides for drug delivery. International Journal of Biological Macromolecules, 64, 353–367. Reis, R. (Ed.). (2008). Natural-based polymers for biomedical applications. Cambridge England: Woodhead Publishing Limited. Ricard-Blum, S. (2011). The collagen family. Cold Spring Harbor Perspectives in Biology, 3, 1–19.

Polymeric Coatings and Their Fabrication for Medical Devices Dimitrios A Lamprou, University of Kent, Canterbury, United Kingdom Nikolaos Scoutaris, Steven A Ross, and Dionysios Douroumis, University of Greenwich, Greenwich, United Kingdom © 2019 Elsevier Inc. All rights reserved.

Introduction Applications Antimicrobial Drug-Eluting Coating Stents Microneedles Coatings for Osseointegration Coating to Improve the Mechanical Properties Technology Spray Coating Pulsed Laser Deposition Chemical Vapor Deposition Sputter Coating Inkjet Printing Immobilization of Molecules Onto Implant’s Surface Layer by Layer Coating (LbL) Conclusion References Further Reading

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Introduction According to U.S. Food Drug Administration (FDA), medical implants are devices or tissues that are placed inside or on the surface of the body. Many implants are prosthetics, intended to replace missing body parts. Other implants deliver medication, monitor body functions, or provide support to organs and tissues (U.S. food and Drug Administration, 2015). The implant surface is critical for determining the integration of the implant with the human body. This can range from enzyme inhibition, whether specific or unspecific, to how antibodies and albumin may act when stored in glass containers, the biocompatibility of hip replacements or how blood proteins may react with stents. These interactions can be either useful (improved biocompatibility) or detrimental (thrombosis and toxicity) so it is important to understand how and why they occur. Appropriate coatings are advantageous and often essential for the acceptance and smooth functioning of the implant. The coating on medical devices can help to reduce friction in the body to improve the placement of the implant and also minimize irritation and inflammation. It can improve the biocompatibility inhibiting the formation of scar tissue surrounding implanted devices, the chance of infection related to the implanted device, and encouraging the growth of tissues to help the healing process. Applying a coating to a device that is placed in the body is a very critical process. The coating must be uniform, covering the complete surface which often consists of a complex structure and eliminating the bringing across the structure. Nowadays, numerous technologies have been developed in order to achieve a thin coating on medical devices. These include spray coating to deposit a thin film, physical vapor deposition (PVD) to transfer a solid source to a surface film, chemical vapor deposition (CVD) for a surface reaction to create film, surface polymerization to create a film from a monomer vapor, and inkjet placement of coating via impingement of tiny droplets.

Applications Antimicrobial All implanted medical devices, from transient, easily inserted, and retrieved contact lenses, urinary catheters, and endotracheal tubes, to more permanently surgically implanted cardiac valves, embolic coils, vascular grafts, hip, knee and shoulder joints, pacemakers, coronary stents, and plastic surgery augmentation devices suffer from recognized risks of “device-related” or “implantassociated” infection (Wu and Grainger, 2006). In order to reduce the incidence of device-related infections two main strategies have been implemented: first, antiadhesive biomaterials using physicochemical surface modification methods such as nondrug containing coating and second direct incorporation of drugs into or onto the medical device either immobilized or released (Wu and Grainger, 2006).

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There have been several antimicrobial surfaces described in the literature, including antibiotic coatings such as cefazolin minocycline–rifampin, teicoplanin, and vancomycin which have been tested on biomaterials. Alternative antimicrobial coatings such as salicylic acid, quaternary ammonium compounds, and chlorhexidine have also been trialed. In this concept, polymers which are used as carriers for the antibiotics are deposited on implant’s surface. Harris et al. (2006) used various polymers including (poly(D,L-lactide) (PDLLA), politerefate (PTF), calcium phosphate/anodic plasma-chemical treatment (CaP/APC), polyurethane (PU), and polyvinylpyrollidone (PVP) on titanium surfaces with and without chlorhexidine diacetate (CHA) to investigate their cytocompatibility for Staphylococcus aureus, Staphylococcus epidermidis, and human telomerase reverse transcriptase (hTERT) fibroblasts. The study showed that the release kinetics varied from slow (over 200 h) to burst release: PDLLA > PTF > PU > CaP/ APC ¼ PVP. This study showed that PDLLA and PTF have the best potential as coatings on implants for drug delivery, as they were cytocompatible to hTERT fibroblasts, eluted CHA effectively, and passed mechanical testing. The actual release kinetics of PDLLA and PTF are important, as the amount of CHA present should remain above the minimal inhibitory concentration value for a limited time before disappearing completely. Silver antimicrobial agents have also been utilized as an alternative strategy for reducing bacterial adhesion and preventing biofilm formation. Silver (Ag) is a potent bactericide and has attracted increasing attention due to a host of additional benefits such as a broad antibacterial spectrum including antibiotic-resistant bacteria, noncytotoxicity at proper doses, satisfactory stability, and a smaller possibility to develop resistant strains. However, the use of silver must be undertaken with caution, since a concentration-dependent toxicity has been demonstrated. Panácek et al. (2006) assessed the suitability of a mouse spermatogonial stem cell line as an in vitro model to evaluate the toxicity of silver resulting that concentrations of silver nanoparticles between 5 and 10 mg/mL induced necrosis or apoptosis of mouse spermatogonial stem cells. Apart from silver, other promising metals for antimicrobial coating are Copper (Cu) and Zinc (Zn). Besides conventional antibiotics, different types of alternative antimicrobial coatings for medical devices have been proposed including antimicrobial peptides (AMP) and polymers. Polymers have been introduced in order to overcome problems associated with the low-molecular-weight antimicrobial agents, such as toxicity to the environment and short-term antimicrobial ability, where antimicrobial functional groups can be introduced into polymer molecules. The use of antimicrobial polymers offers promise for enhancing the efficacy of some existing antimicrobial agents and minimizing the environmental problems accompanying conventional antimicrobial agents, by reducing the residual toxicity of the agents, increasing their efficiency and selectivity, and prolonging the agents lifetime. For instance, Hoffman et al. applied poly(ethylene glycol-stat-propylene glycol) prepolymers (Star PEG). Taken together, for Star-PEG-covered substrates demonstrated a profound reduction of various blood–biomaterial interactions compared to noncoated substrates. A number of polymers that have been applied to influence the amount and/or conformation of adsorbed proteins, preventing bacterial adhesion and biofilm formation, such as poly(hydroxyethylmethacrylate), poly(methacrylic acid), and polyurethanes. Recently, antimicrobial peptides (AMP) have been applied to the implant’s surface in order to prevent colonization of biomaterials as they display long-term stability, even though the sterilization process, and have a low cytotoxic profile. Because of their physicochemical characteristics, which involve their highly cationic character and tendency to adopt amphipathic structures, AMPs have the tendency to associate with negatively charged microbial surfaces and membranes. These peptides offer several attractive advantages: they exhibit bactericidal, fungicidal, viricidal, and tumoricidal properties, they act at a very low concentration, and they are less likely to promote bacterial resistance. Hence, AMPs’ coating on implants has been developed to reduce bacterial strains of Gram-positive (S. aureus) and Gram-negative (Pseudomonas aeruginosa) bacteria.

Drug-Eluting Coating Stents Stents are small expandable tubes used to treat coronary heart diseases. A stent is implanted using a procedure called percutaneous coronary intervention (PCI), also known as coronary angioplasty. PCI restores blood flow through narrow or blocked arteries. A stent helps support the inner wall of the artery in the months or years after PCI. Biocompatibility of the stents materials is important to minimize inflammation and to allow endothelial cell growth to continue as normal. The first generation of stents was made of bare metal. Although bare-metal stents almost eliminated the risk of the artery collapsing, they only modestly reduced the risk of restenosis. About 25% of all coronary arteries treated with bare-metal stents would close up again, usually within about 6 months. Stents are normally made of a metal like tantalum or an alloy such as 316L stainless steel. While these materials typically show reasonable biocompatibility, coating stents with polymer or ceramic layers has the potential to improve the properties further. Therefore, new platforms for drug-eluting stents have been developed in order to eliminate the risk of restenosis. Coatings can either be passive providing an improved tissue interface or active to elute drugs which can control neointimal growth, restenosis, and inflammation which would narrow the lumen and may give rise to dangerous thrombi. There are currently stents are made of biodegradable polymers for drug elution. Polymeric coating of metallic stents also reduces the risk of toxicity from metal released through corrosion. Drug-eluting stents consist of the metallic platform, a polymer, and the antiproliferative agent. In drug-eluting stents, the polymer is used as a reservoir which controls the drug release, whereas the drug is used to inhibit the endothelial cells to proliferate inhibiting restenosis. Employing a spray-coating technique Bege et al. (2012) coated rotating stents with a poly(ethylene carbonate) (PEC) solution of 0.8% (w/w) in dichloromethane with or without paclitaxel. Other drugs that were used for stent coating are rapamycin, tacrolimus, everolimus, zotarolimus, and biolimus. Moreover, a wide range of polymers can be permanent, such as poly-n-butyl methacrylate (PBMA) and polyethylene-co-vinyl acetate (PEVA), which is used in cypherÒ stents developed by

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Cordis. Most modern polymer coatings are bioerodible and degraded by hydrolysis and are eventually metabolized to water and carbon dioxide, examples include polyglycolic acid, poly-L-lactid acid grades, and PLGA-PEG block polymers. Drugs and polymers can be deposited on stent’s surface by using various techniques including dip coating, spray coating, inkjet printing, and layer by layer.

Microneedles Microneedles is a novel approach for transdermal delivery of drug substances, vaccines or macromolecules which cannot be administered orally due to their poor absorption or enzymatic degradation in the gastrointestinal tract and liver. The aim of microneedles is to create large pathways of microscope dimensions using an array of microscopic needles attached to a metal or polymer base. Regarding metal microneedles, drug/excipient compositions are coated both sides of each needle and delivered transdermally by piercing the skin. Currently, microneedles are coated using the dip-coating approaches. A dip-coating process typically involves dipping and withdrawing an object from a coating solution, after which a continuous liquid film adheres and dries on the object’s surface, leaving behind a uniform coating. However, conventional dip-coating methods have difficulty controlling the material deposition of specified sections of micron-dimensioned structures, especially when those structures are closely spaced. McGrath et al. (2011) spray coated microneedle patches using a conventional film-coating process and investigated two film polymeric-coating materials, hydroxypropylmethylcellulose and carboxymethylcellulose; these could be potentially loaded for intradermal drug and vaccine delivery. However, the main disadvantage of the latter technology is the loss of material. Therefore, recently inkjet printing technology has been applied to deposit single droplets of the coating solution directly onto the microneedles.

Coatings for Osseointegration Titanium (Ti) and its alloys are the most commonly used metallic materials for medical implants in orthopedic and dental applications, due to their low density, high strength, nontoxicity, and excellent corrosion resistance. However, there have been reports on inflammatory reaction around these implants as a result of the creation of an avascular fibrous tissue that encapsulated the implants. Therefore, a hydroxyapatite layer is deposited on the metal alloy to assist the osseointegration of these implants with surrounding tissues. Synthetic hydroxyapatite [Ca10(PO4)6(OH)2, HA] has been extensively used in a number of dental, orthopedic, and other medical applications, due to its similarity in chemical composition and crystal structure with bone mineral. It has been shown that hydroxyapatite stimulates the expression of bone-related mRNAs and proteins in osteoblasts. Several techniques have been used to create the HA coating on metallic implants, such as plasma spraying process, thermal spraying, sputter coating, pulsed laser ablation, dip coating, sol–gel, electrophoretic deposition (EPD), ion-beam-assisted-deposition, and hot isostatic pressing. A number of approaches toward improving the integration rates of HA with bone have included the incorporation of biological entities such as growth factors, proteins and cells, into the HA implant (Mankani et al., 2006). Other strategies to improve the osteoinduction of embedded biomolecules are the coating on the surface of the implant with biomolecules. Large proteins or glycosaminoglycans such as collagen and chondroitin sulfate provide a biomimetic coating on the surface of an implant which can improve integration once implanted in the body. In order to manage efficient loading of biomolecules onto the implant surface and to modulate the release of such molecules in a controlled manner which in turn can promote osseointegration hydrogel coatings, layer-by-layer (LBL) coating techniques have been developed. In hydrogel coatings, orthopedic implants are immersed in hydrogel solutions that contain biomolecules of interest. The implant is then removed and air dried to allow adsorption of molecules onto the surface of the implant. Given its ease of application, hydrogel coatings have been applied to coat various orthopedic implants with a broad range of biomolecules including growth factors, viruses, and peptides. Hydrogels are appealing candidates for the development of scaffolds for mineralization. The intrinsic elasticity and water retention ability of synthetic hydrogels resemble those of collagen, and their porosity may be controlled by various techniques. Another important feature of hydrogels is that they can be assembled in a three-dimensional form, displaying multiple functional domains through copolymerization of different monomers. Hence, proteins regulating mineral growth, and biological epitopes such as the tripeptide RGD that promote cellular adhesion have been applied to collagen hydrogel and poly(ethylene glycol) hydrogels. Other hydrogel coatings involve the utilization 2-hydroxyethylmethacrylate (HEMA) and HEMA copolymerized with a phosphate-containing monomer, 2-methacryloyloxyethyl phosphate (MOEP), where P(HEMA–MOEP) coating was to entrap and release rhVEGF which showed its potential to enhance implant site calcification and angiogenesis. Moreover, coating metal surfaces with suitably functionalized/modified polymer films capable of improving implant osteointegration and at the same time inhibiting metal substrate corrosion is another useful strategy. Such as polymers include poly(bis(trifluoroethoxy)phosphazene), pyrrole, poly(methyl methacrylate), and poly(acrylic acid). In LBL approach, the adsorption of biomolecules onto the implants’ surface occurs by introducing a surface charge to the implant surface by plasma etching or modifying with molecules such as heparin. The charge is achieved by dipping implants repeatedly in polyelectrolyte solutions with opposite charges. The LBL technique has been used for the deposition of growth factors on a variety of implant surfaces. Polycaprolactone/b-tricalcium phosphate (PCL/b-TCP) scaffolds coated with bone morphogenetic protein 2 (BMP-2) via an LBL approach led to enhanced ectopic bone formation in the rat hindlimb tissue.

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Coating to Improve the Mechanical Properties The most common coating material which is currently used to improve the mechanical properties of the implant is the diamondlike carbon (DLC) coating. DLC is an amorphous carbon biomaterial that in recent years has demonstrated its coating potential due to its high hardness, low frictional coefficient, high wear and corrosion resistance, chemical inertness, high electrical resistivity, infrared-transparency, high refractive index, and excellent smoothness. All these properties match well with the criteria of a safe biomaterial for applications in orthopedic, cardiovascular, contact lenses, or dentistry. Also, it has been shown that DLC has better biocompatibility and wear resistance than stainless steel, titanium, and Ti alloys. Moreover, when DLC coatings are applied to implant’s abutment’s screws, they demonstrate more resistance to an applied force than without the coating. The most common method used for coating DLC is plasma-enhanced CVD (PECVD). In addition to PECVD-coating process, other technologies which were used to impart the implant with desirable coating properties include pulsed layer deposition, sputtering, bipolar type plasma-based ion implantation and deposition (PBII&D), and more rarely cathodic arc deposition. In the PBII&D processing, the glow discharge plasma is formed near the target surface by applying a positive pulse voltage directly to the target, and the plasma is omnidirectionally implanted into (or deposited on) the target surface by a subsequent negative high pulse voltage.

Technology Spray Coating Spray coating is a technology that has been successfully used for various purposes such as for drug release, antibacterial, and osseointegration coating by depositing various active agents such as polymers and drugs on implantable medical devices, including catheters, stents, drug-eluting balloons, pacemakers, heart valves, sensors, surgical implant coatings, and orthopedic implant coatings. There are various technologies which can be characterized as spray coating. Among the most prevalent is the ultrasonic spray nozzle. According to this, the coating solution is atomized into a fine mist spray using high frequency sound vibrations. Piezoelectric transducers convert the electrical input into mechanical energy in the form of vibrations, which create capillary waves in the liquid when introduced into the nozzle. For the liquid to atomize, the vibrational amplitude of the atomizing surface must be carefully controlled. Below the critical amplitude, the energy is insufficient to produce atomized drops. If the amplitude is excessively high, the liquid is ripped apart, and large “chunks” of fluid are ejected, a condition known as cavitation. Only within a narrow band of input power is the amplitude ideal for producing the nozzle’s characteristic fine, low velocity mist. Ultrasonic spray nozzle has been extensively used to coat stents. The operating requirements in this application call for flow rates on the order of 20–100 mL/min, spray diameters in the range of 0.5–2 mm, and tiny median drop diameters. In this method, the stent is placed on a mandrel that is attached to a rotating shaft. The nozzle is mounted above the stent. The sprayed liquid consists of the polymer/drug system dissolved in a suitable organic solvent. Typically, high vapor pressure solvents are used, so that drying occurs quickly. Repeated traverses, coupled with low flow rates, produce the best coatings and maximum material transfer efficiency. By varying rotational speed, the distance of the spray from the stent, and the number of traverses a process can be optimized. Specifically, poly (D,L-lactic-co-glycolic acid) (PLGA) was used as a drug carrier to generate two types of stents loaded with different concentrations of sirolimus, and curcumin-eluting PLGA coatings were fabricated on the surface of 316L stainless steel stents by an ultrasonic spray method. An interesting study from Choi et al. showed that sirolimus (SRL) release from the biodegradable (PLGA) matrix is affected by the organic solvents due to the interaction between SRL and PLGA and the different surface topography (Choi et al., 2014). Another spray coating technique is the use of aerosol. For aerosol generation, gas is passed through a loose powder contained in a chamber, thereby producing a fluidized bed. Driven by a pressure difference the aerosolized particles are transported from the aerosol chamber through a nozzle to the deposition chamber. The particles in the focused jet collide with the substrate at high speed. Hahn et al. used aerosol deposition on AZ31 Mg alloy samples, sprayed in a vacuum, with a specially mixed HA–chitosan powder at room temperature. The nozzle was placed vertically opposite the sample, which allows rotation in the x- and y-directions, but not in the z-direction (Hahn et al., 2011). Moreover, HA and 4-hexylresorcinol (4-HR), a well-known antiseptic, were successfully coated onto a titanium surface. A spray-coating technique which is widely used for coating of orthopedic implants with hydroxyapatite are the thermal spray techniques. These are coating processes in which melted (or heated) materials are sprayed onto a surface. The “feedstock” (coating precursor) is heated by electrical (plasma or arc) or chemical means (combustion flame). Another technique is called high velocity oxygen fuel coatings (HVOF). In HVOF spraying, fuel and oxygen are pressed into a combustion chamber in a continuous flow, producing a jet of combustion products at extremely high speed. Powder particles injected into this gas stream are accelerated to a very high velocity. Fusion is obtained by the kinetic impact of the coating particles, rather than by their increased temperature. This process is carried out in an ambient atmosphere. HVOF has been applied in order to coat implants with HA with titanium addition. The titanium was found to improve the Young’s modulus, fracture toughness, and shear strength of HVOF-sprayed HA-based coatings. However, the increase of titanium content from 10 to 20 vol% induces a small decrease in Young’s modulus. Moreover, the influence of particle temperatures and velocities has been investigated, and it was observed that they have linear correlation with the phase composition, Vickers microhardness, and the residual stress. One of the most promising suspension spray technology is the high-velocity suspension flame spraying (HVSFS) process developed at the Institute for Manufacturing Technologies of Ceramic Components and Composites. HVSFS is a new approach for

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spraying micron, submicron, and nanoparticles with hypersonic speed by feeding a suspension directly into the combustion chamber of a HVOF torch. The aim in mind is to achieve dense coatings with an improved microstructuredprobably reaching the nanoscale, from which superior physical properties are expected. Compared to the alternative approach, that is, using agglomerated (nano- and micron-sized) powders, direct spraying of suspensions shows much higher flexibility in combining and processing different materials and is far less expensive. Several suspensions consisting of an organic solvent and a solid phase composed of a micron or a nanopowder have been prepared and HVSFS sprayed, examples include suspensions containing oxide nanopowders of titanium oxide (n-TiO2), chromium(III) oxide (n-Cr2O3), yttrium-stabilized zirconia (n-YSZ), and n-hydroxyapatite (n-HA). Different process parameters such as gas flow, spray distance, air-fuel ratio, and electric arc current can affect the coating properties. Hence Gadow et al. (2010) prepared thermally sprayed hydroxyapatite coatings by using three different thermal coating technologies HVOF, APS, and HVSFS. They found that HVSFS HA coatings have the most refined microstructure regarding the microstructure of the deposited coatings. Also, the main effects of the different processes on the adhesive behavior were reviewed. The pull-off performance of the HVSFS–HA coatings strongly depends on the dispersive medium and the oxygen flow rate, for HVOF–HA coatings, it is the spray distance and for the APS–HA coatings, the electric arc current is the most influencing parameter. Moreover, APS has been applied in order to produce amorphous HAP. It was found that three factors strongly influence the formation of the amorphous phase; dehydroxylation of the molten particle during flight, the cooling rate as it impinges onto the metal substrate, and the substrate temperature. Crystalline regions were identified as unmelted particles and elongated recrystallized areas. Amorphous phase regions vary throughout the coating but are more commonly found at the coating-substrate interface, that is, the areas decrease toward the surface of the coating.

Pulsed Laser Deposition Pulsed laser deposition (PLD) is a PVD process, carried out in a vacuum system, which shares some process characteristics shared with molecular beam epitaxy and some with sputter deposition. In PLD, a pulsed laser is focused onto a target of the material to be deposited. For sufficient high laser energy density, each laser pulse vaporizes or ablates a small amount of the material creating a plasma plume. The ablated material is ejected from the target in a highly forward-directed plume. The ablation plume provides the material flux for film growth. The film growth process depends greatly on the parameters of the condensing plasma fluxes (density, spatial distribution, particle energy, ionization degree, etc.), as well as on the thermodynamic parameters of the substrate surface, such as temperature, density of adsorption sites, activation energy of surface desorption and diffusion, etc. PLD is one of the most promising coating technique for hydroxyapatite providing excellent quality, and high performance coatings. Pure crystalline hydroxyapatite (HA) films with thicknesses of ranging from 0.5 to 5 mm have been deposited on titanium substrate using the PLD technique (Fig. 1). Experimental results indicate that the structure and properties of the PLD-HA films varied with deposition parameters. Torrisi and Setola (1993) have obtained a smooth, uniform PLD-HA film on a titanium substrate at 400 C. When deposited on a room temperature substrate, however, the coating had a granular-uniform mixed morphology. Jelinek et al. (1995) have also obtained a uniform laser-deposited HA film on a titanium substrate at intermediate temperatures (200–400 C). When deposited on a higher temperature (600–760 C) substrate, the coating roughness increased and inhomogeneity and buckling were observed. Moreover, Zeng and Lacefield (2000) showed that the Ca/P ratios of HA powder and targets were larger than 1.67. Ca/P of HA ratios of coatings varied with the deposition conditions. Specifically, coatings deposited at room temperature under vacuum had higher Ca/P ratios than the HA target material. Coating occurred at room temperature but with the presence of an ambient gas had a Ca/P ratio closer to that of the target, whereas coating at elevated temperature with an

Fig. 1 SEM micrograph of the cross section of a PLD coating, showing a dense HA film. Reproduced with permission from Dinda, G.P., Shin, J. and Mazumder, J. (2009). Pulsed laser deposition of hydroxyapatite thin films on Ti–6Al–4V: Effect of heat treatment on structure and properties. Acta Biomaterialia 5(5), 1821–1830.

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ambient gas had a Ca/P ratio that was in between the Ca/P ratios for specimens deposited under the other two conditions. Cotell et al. (1992) have reported that the amorphous PLD-HA films deposited at room temperature were demonstrated calcium deficiency, displaying Ca/P ratios between 1.2 and 1.6. However, when deposited at 400–700 C crystalline HA films with Ca/P ratios ranging from 1.6 to 3.6 were obtained. Although the substrate temperature was perhaps the most critical parameter, variations in Ca/P ratio of as much as 60% were noted, even when deposited at the same substrate temperature.

Chemical Vapor Deposition In CVD, there is no universal equipment, but each piece of CVD equipment is individually tailored for specific coating materials, substrate geometry, etc., whether it is used for research and development or commercial production. In general, the CVD equipment consists of three main components: the chemical vapor precursor supply system, a CVD reactor, and the effluent gas handling system. The role of the chemical vapor precursor system is to generate vapor precursors and then deliver to the reactor. A CVD reactor consists of a reaction chamber equipped with a load lock for the transport and placement of the substrate into the chamber, a substrate holder, and a heating system with temperature control. The primary function of the CVD reactor is used to heat the substrate to the deposition temperature. The CVD reactor can either be a hot-wall or cold-wall reactor. A hot-wall reactor uses a heated furnace into which the substrates are placed for indirect heating, whereas in the cold wall reactor, only the substrate is heated and the wall of the reactor is cold. Finally, the effluent gas handling system consists of a neutralizing part for the exhaust gases, and/or a vacuum system to provide the required reduced pressure for CVD processing, that performs at low pressure or high vacuum during deposition. The main function of the effluent gas handling system is to remove the hazardous by-product and the toxic unreacted precursor safely. The unreacted precursors and corrosive by-products such as HCl are neutralized or trapped using a liquid nitrogen trap to prevent these gases from entering the rotary or diffusion pump, which can cause damage to the pump. In terms of medical implants, CVD is suitable for titanium nitride, titanium carbonitride, and aluminium oxide, although many other materials may be deposited by CVD as well. The main parameters of CVD that affect the quality of coating are temperature, pressure, reactant gas concentration, and total gas flow, which require accurate control and monitoring. Specifically, the temperature at which the coating is deposited is a critical factor, as it controls both the thermodynamics and the kinetics of the coating process. The deposition temperature must be achieved and maintained in order for the reaction to occur on the substrate and not in the gas phase, and with an appropriate microstructure. Moreover, CVD processes are carried out from atmospheric pressure to high vacuum. At atmospheric pressure, the growth processes are often considered to be “transport controlled.” The main disadvantage of CVD is that the heat input can result in damage to temperature-sensitive substrates and so alternative forms of energy input have been developed which allow deposition at lower temperatures. One way of reducing growth temperatures is to use plasma-assisted or PECVD. With this technique, deposition can occur at very low temperatures, even close to ambient, since electrical energy rather than thermal energy is used to initiate homogeneous reactions for the production of chemically active ions and radicals that can participate in heterogeneous reactions, which, in turn, lead to layer formation on the substrate. In PECVD processes, deposition is achieved by introducing reactant gases between parallel electrodesda grounded electrode and an RF-energized electrode. The capacitive coupling between the electrodes excites the reactant gases into a plasma, which induces a chemical reaction resulting in the reaction product being deposited on the substrate. The substrate, which is placed on the grounded electrode, is typically heated to 250–350 C, depending on the specific film requirements. In comparison, CVD requires 600–800 C. PECVD has been applied extensively to deposit DLC films producing a smooth surface morphology exhibiting high hardness and elastic modulus (200 GPa) and good adhesion to Ti–6Al–4V substrate. Also, other biocompatible coatings for implants involve the deposition of SiOx, TiOx, and SiO–TiO mixed oxide films and TiN and TiAlN coated on Ti Surface to enhance the osteoblast’s proliferation (Fig. 2). Finally, multilayers with extension to layer thicknesses in the nanometer range of TiN/Ti–B–N multilayers have been deposited in an industrially sized deposition chamber to optimize mechanical and tribological properties.

Sputter Coating In sputter coating, a gaseous plasma is created for the purpose of ion bombardment of the source material known as the target. The source material is eroded by the arriving ions via energy transfer and is ejected in the form of neutral particlesdeither individual atoms, clusters of atoms, or molecules. As these neutral particles are ejected, they will travel in a straight line unless they come into contact with something, such as other particles or a nearby surface. If a “substrate” is placed in the path of these ejected particles, it will be coated with a thin film of the source material. Sputter coating has been used extensively to coat hydroxyapatite and other CaP materials. However, it has been reported that it would produce coatings whose chemistry was different upon deposition, than the bulk target. Specifically, the surface Ca/P ratio for the sputtered coatings has been reported to range from 1.6 to 2.6. Normally, the coating coming from sputtering technique is uniform and dense, where the thickness can be as small as 10 nm depending on the deposition time. Furthermore, it was found that CaP deposition can improve the osseointegration even in osteoporotic conditions (Fig. 3). Also, it has been applied to deposit zinc oxide to control and prevent the spreading and persistence of bacterial infections. EPD is a technique that exploits the movement of charged particles in suspension in the presence of electric field. This electric field enables the consolidation of particles onto any shaped substrate. Two groups of parameters determine the characteristics of this process: (i) those related to the suspension and (ii) those related to the process including the physical parameters such as the

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Fig. 2 Scanning electron microscopy (SEM) images of MG63 cells cultured on (A) Ti, (B) TiN, and (C) TiAlN-coated surfaces after 24 h of cultures. Reproduced with permission from Jang, H.W., et al. (2011). Surface characteristics and osteoblast cell response on TiN- and TiAlN-coated Ti implant. Biomedical Engineering Letters 1(2), 99–107.

Fig. 3 Histological images present bone features around coated and noncoated implants in osteoporotic vs. healthy rats. In osteoporosis (A and B) more bone shows in direct contact with coated implants but disconnected from the preexisting bone compared to healthy (C and D). Bone covering noncoated implant surface is thin and sparse, and a layer of fibrous tissues could be always detected (Scale bar ¼ 500 mm at 10 objective magnification). Reproduced with permission from Alghamdi, H.S., et al. (2013). Osteogenicity of titanium implants coated with calcium phosphate or collagen type-I in osteoporotic rats. Biomaterials 34(15), 3747–3757.

electrical nature of the electrodes and the electrical conditions (voltage/intensity relationship, deposition time, etc.). The main parameters that affect the coating quality are the suspension properties (which are related to particle size), dielectric constant of the liquid, conductivity and viscosity of the suspension, zeta potential, and stability of suspension. Other parameters involve the deposition time, the applied voltage, concentration of solid in suspension, and the conductivity of substrate. Nowadays, even though plasma spraying is currently the only commercial process for fabricating HA coatings on metallic implants. EPD has been suggested to deposit HA onto implants. Usually, the electrodeposition is carried out from a bath containing low concentrations of Ca(NO3)2 and NH4H2PO4 at pH 6 by cathodic polarization. The advantages of EDP include control over the composition and structure of the coating, and the ability to coat irregular surfaces efficiently.

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EPD has been applied chitosan/silk fibroin to deposit composite coatings onto titanium substrates which were recommended for protein stability and cell viability. The silk fibroin content in the coatings increased proportionally with the increase of the silk fibroin in the electrophoretic solutions, whereas the shear and tensile bond strength of the coatings to titanium substrates increased with the increasing silk fibroin content. The in vitro biological tests indicated that chitosan/silk fibroin composite coatings had better cellular affinity than pure chitosan coatings. Moreover, colloidal suspensions of lysozyme and silver nanoparticles were electrophoretically deposited onto the surface of stainless steel surgical blades and needles. This antimicrobial coating on medical instruments combines the bacteriolytic activity of lysozyme and the biocidal properties of silver nanoparticles. Also, EPD has been applied to coat stents with curcumin-loaded PLGA nanoparticles producing uniform surface morphology.

Inkjet Printing Inkjet printing encompasses a set of material deposition technologies which are dispersed in a single, or in a system of solvents and are ejected onto a substrate through a nozzle. The advantages of inkjet printing compared to other conventional coating techniques are that it is an entirely automated process, provides accurate and precise control over the droplet size which in turn minimizes material losses, allows the design of various deposition patterns, and reduces the contamination of the deposited material. Inkjet printing can operate either in continuous (CIJ) or drop on demand (DOD) modes. In CIJ mode the printer is operated in a continuous stream mode at high speeds which results in low printing quality, and it is mainly used for high-speed graphical applications. When using DOD mode, the droplets of a solution are jetted only according to the requirements and are perfectly located onto a predetermined position. Depending on the technology, inkjet printing is categorized into thermal, piezoelectric, acoustic, or electrostatic inkjet. An excellent paper which describes these technologies was published in 1998 (Le, 1998). Thermal ink-jetting (TIJ) and piezoelectric ink-jetting (PIJ) are the predominant technologies, whereas acoustic and electrostatic ink-jetting are still in the development stages. In thermal inkjet printing, the ink is heated, by means of a thin film resistor of heater which is activated by an electrical pulse and thus creating a rapidly expanding vapor bubble which in turn ejects the ink from the nozzle. Conversely, the piezoelectric dispensers rely on the deformation of a piezoelectric material. Specifically, when a voltage is applied to the piezo element, it changes its shape to a preordained direction, causing a sudden volume change which creates pressure waves resulting in a drop being ejected from the orifice. Most of the research nowadays rely on PIJ technique due to its flexibility regarding the ink compatibility and the higher range in the operation frequency over the TIJ. In thermal inkjet printing, the degradation of the compounds due to the applied heat is possible. However, this has not been proved as the heat lasts for a few milliseconds and hence it does not affect the stability of the compounds. The most important properties of the liquid dispensers that implement a piezoelectric response are viscosity, surface tension, and density. These fluid properties influence the drop formation mechanism and subsequent the drop size at a given voltage. Also, these properties can provide information about the impact of the droplet into the substrate. Hence, Jang et al. (2009) investigated the printability of the fluids by applying the inverse (Z) of the Ohnesorge number (Oh) which relates to the viscosity, surface tension, and density of the fluid. They have experimentally defined the printable range as 4 < Z < 14 by considering characteristics such as single droplet formability, positional accuracy, and maximum allowable jetting frequency. This range of Z corresponds to low viscosity fluid. Yang et al. (2006) showed that a smaller value of surface tension represents a weaker cohesive which leads to a slender liquid ligament with a relatively longer break-up length. The shapes of the liquid and tail droplet tend to be round for the cases of high surface tension. It also takes more time for contraction of a spherical droplet of the liquid with lower surface tension. The predominant factor that determines the size and volume of the droplet is the size of the dispenser’s nozzle. For a given dispenser, the volume of the droplet is directly dependent on the voltage and the duration of the signal that is applied. Hence, the droplet velocity and volume are found to show a linear relation with driving voltage but show a more complicated and periodic behavior with changing frequency and pulse width. Reis et al. (2005) showed that droplet velocity exhibits a maximum as a function of pulse width, which remains unchanged when driving voltage amplitude increased. This factor can be proved critical to control the deposited amount. Also, due to the low surface tension of organic solvents, the voltage increase can result to air bubble formation inside the dispenser. Finally, during the printing process, droplet satellites can be formed which are undesirable as it can create nonprecise and nonhomogenous coating. Many methods are being developed to eliminate these droplet satellites which are related to the wave propagating in the dispenser and the amplitudes and pulse durations (Tsai and Hwang, 2008). Due to the controllable and reproducible nature of the droplets in the jet stream and the ability to direct the stream to exact locations on the device surfaces, inkjet printing technology has shown its potency to applications where high-resolution coating is required. Tarcha et al. (2007) programmed target deliveries of 100 mg of the drug, a typical dose for a small stent, into cuvettes, which demonstrated standard deviation of 0.6 mg per stent. According to the coating pattern, the dispenser moves longitudinally across the stent and is then positioned in such a way that the stream of reagent droplets is tangent to the cylinder swept out by the stent as it is rotated. Jetting on coated uncut stent tubes exhibited 100% capture efficiency with a (137  1.8 mg), while continuous jetting on actual stents presented efficiencies up to 91% with coefficients of variation as low as 2%. Also, Scoutaris et al. (2015) implemented inkjet printing technology to apply biodegradable and biocompatible polymeric coatings of PDLLA with the antiproliferative drugs, simvastatin and paclitaxel, on coronary metal stents. Finally, Gill and Prausnitz (2007) developed a new coating methodology based on inkjet printing whereby metal microneedles are coated with varies molecules such as calcein, an anticancer drug, and insulin protein. In terms of medical implants, inkjet printing has been applied to create rifampicin as antibiotic and calcium eluting as a means of preventing the formation of biofilm colonies and facilitating osteogenic cell development on orthopedic implant surfaces. The micropatterns consisted of a periodic array of circular dots.

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Fig. 4 Schematic diagram of BMP-2 protein attachment to TI surface by initial introducing of (11-hydroxyundecyl)phosphonic acid. Reproduced with permission from Adden, N. et al. (2006). Phosphonic acid monolayers for binding of bioactive molecules to titanium surfaces. Langmuir 22(19), 8197–8204.

Immobilization of Molecules Onto Implant’s Surface Generally, active pharmaceutical ingredients are deposited onto the implant’s surface. However, this approach failed to achieve delivery of a sustained and efficient dosage over a relatively prolonged period of time. Vancomycin has been successfully attached to titanium and proven to be bactericidal to S. aureus and S. epidermidis. Nevertheless, the immobilization of molecules on implant’s surface is extremely beneficial in the case of depositing AMP. In general, the surfaces of metals are derivatized into reactive groups, then biomolecules are conjugated to the surface by reacting with these groups. One of the most prevalent methods is the silanization of the titanium implant resulting into a terminal-maleimide or into aldehyde groups where RGD-containing peptides proteins such alkaline phosphatase and albumin were bound on the surface, respectively. Immobilization of Arg-Gly-Asp (RGD) containing peptides has received significant interest because RGD is the essential sequence mediating cell adhesion in many extracellular matrix proteins. Other strategies involve functionalized PLL-g-PEG polymer adsorbed as a monolayer on a negatively charged metal oxide surface (e.g., titanium oxide). The PLL backbone is positively charged at pH 7.4 due to its amine groups, and the PEG chains are exposed to the fluid as a brush. The peptide sequence arginine–glycine–aspartic acid (RGD) is covalently attached to a fraction of the PEG chains. Another strategy to attach RGD peptide involves the dopamine’s polymerization onto implant’s surface and subsequently immobilized of the RGD sequence. Also, BMP-2 protein has been successfully attached to titanium by introducing of monolayers of (11-hydroxyundecyl)phosphonic acid and (12-carboxydodecyl)phosphonic acid molecules produced by a simple dipping process. The terminal functional groups on these monolayers were activated (carbonyldiimidazole for hydroxyl groups and N-hydroxysuccinimide for carboxyl groups) to bind amine-containing molecules (Fig. 4). In a recent study, BMP-2 was encapsulated in chitosan coatings on functionalized Ti surfaces. Slow release of the protein enhanced the osteoinductivity of the implant.

Layer by Layer Coating (LbL) LBL coating involves the dipping of surfaces repeatedly in polyelectrolyte solutions with opposite charges. LBL has several applications on the biomaterial area as it enables the nanometer level control of the composition of a thin film, and the generation of highly sophisticated, tailor-made coating compositions. The LBL technique has been used for the deposition of growth factors on a variety of implant surfaces. Hence, BMP-2, which are capable of directing the host tissue response to create bone from native progenitor cells, has been coated on PCL/b-TCP scaffolds. In another work, osteogenic rhBMP-2 (recombinant human bone morphogenetic protein-2) and angiogenic rhVEGF165 (recombinant human vascular endothelial growth factor) in different ratios in a degradable [poly(b-amino ester)/polyanion/growth factor/polyanion] LBL tetralayer repeat architecture were promoted the differentiation cascade in MC3T3-E1S4 preosteoblasts, while rhVEGF165 upregulated HUVEC proliferation and accelerated closure of a scratch in HUVEC cell cultures in a dose-dependent manner (Shah et al., 2011). To improve the surface biocompatibility of titanium films, chitosan (Chi) and gelatin (Gel) were used leading to the formation of multilayers on the titanium thin film surfaces. The film growth was initialized by deposition of one layer of positively charged poly(ethyleneimine) (PEI). Then the thin film was formed by the alternate deposition of negatively charged Gel and positively charged Chi utilizing electrostatic interactions. The results showed that the cells’ proliferation and viability had been increased due to the coating (Fig. 5). Moreover, chitosan combined with heparin which is oppositely charged polymer electrolyte form a stent coating through a LBL self-assembly method in order to promote reendothelialization.

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Fig. 5 Focal laser scanning microscopy images of osteoblasts adhered to (A) original titanium and (B) LBL-modified (PEI/Gel/(Chi/Gel)3) titanium surfaces after 24 h incubation. Initial cell seeding density was 5000 cells/cm2. Bar ¼ 20 mm. Osteoblasts attached to LBL-modified titanium film displayed more bundles (arrow) of actin microfilaments than that of the control sample. Reproduced with permission from Cai, K., et al. (2005). Polysaccharide-protein surface modification of titanium via a layer-by-layer technique: Characterization and cell behaviour aspects. Biomaterials 26(30), 5960–5971.

Finally, PLGA microneedles were coated by using a spray LBL technique with DNA encapsulated lipid-modified PLGA nanoparticles. PLGA microneedle arrays were immersed in solutions of polycationic protamine sulfate and polyanionic poly(4-styrene sulfate). Additional multilayers were then deposited through alternating deposition of polycationic polymer-1 and polyanionic plasmid DNA or PLGA NP to give pDNA or PLGA NP-coated arrays, respectively. Polyelectrolyte LBL thin films can be assembled with nanometer scale control over spatial architecture and morphology, which can control the release of multiple types of biomolecules, but this approach is not widely adopted. The reason for this, according to Goodman et al. (2013), is that a few hundred layers are required to avoid a burst release of the biomolecules, rendering the LBL method labor intensive, costly, and may lead to batch-to-batch variability. Also, LBL-coating process is typically performed using acidic solution for efficient loading. However, this approach is not considered to be biomolecule friendly.

Conclusion The performance of implantable medical devices is dictated by the surface of which it is implanted. Therefore, the implant’s surface has to be modified in order to avoid any inflammatory response during implantation, for local drug delivery and to improve the biocompatibility. Moreover, the coating must be homogenous without cracks covering most of the surface. Modern technologies have been developed for the purpose of combating this problem. These techniques involve solvent-based deposition methods such as spray coating and inkjet printing, plasma-coating techniques such as PECVD and HVOF and techniques which take advantage the chemistry of the surface including the LBL method. Each of these technique offers a unique approach to deposit materials on implant’s surface rendering the implantation of medical devices a low-risk clinical operation.

References Bege, N., et al. (2012). Drug eluting stents based on poly(ethylene carbonate): Optimization of the stent coating process. European Journal of Pharmaceutics and Biopharmaceutics, 80(3), 562–570. Choi, J., et al. (2014). Effect of solvent on drug release and a spray-coated matrix of a Sirolimus-eluting stent coated with poly(lactic-co-glycolic acid). Langmuir, 30(33), 10098– 10106. Cotell, C. M., et al. (1992). Pulsed laser deposition of hydroxylapatite thin films on Ti–6Al–4V. Journal of Applied Biomaterials, 3(2), 87–93. Gadow, R., Killinger, A., & Stiegler, N. (2010). Hydroxyapatite coatings for biomedical applications deposited by different thermal spray techniques. Surface and Coatings Technology, 205(4), 1157–1164. Gill, H. S., & Prausnitz, M. R. (2007). Coated microneedles for transdermal delivery. Journal of Controlled Release, 117(2), 227–237. Goodman, S. B., et al. (2013). The future of biologic coatings for orthopaedic implants. Biomaterials, 34(13), 3174–3183. Hahn, B.-D., et al. (2011). Aerosol deposition of hydroxyapatite–chitosan composite coatings on biodegradable magnesium alloy. Surface and Coatings Technology, 205(8), 3112–3118. Harris, L. G., et al. (2006). Bacteria and cell cytocompatibility studies on coated medical grade titanium surfaces. Journal of Biomedical Materials Research. Part A, 78A(1), 50–58. Jang, D., Kim, D., & Moon, J. (2009). Influence of fluid physical properties on ink-jet printability. Langmuir, 25(5), 2629–2635. Jelinek, M., et al. (1995). Effect of processing parameters on the properties of hydroxylapatite films grown by pulsed laser deposition. Thin Solid Films, 257(1), 125–129. Le, H. P. (1998). Progress and trends in ink-jet printing technology. Journal of Imaging Science and Technology, 42(1), 49–62.

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Mankani, M. H., et al. (2006). Canine cranial reconstruction using autologous bone marrow stromal cells. The American Journal of Pathology, 168(2), 542–550. McGrath, M. G., et al. (2011). Determination of parameters for successful spray coating of silicon microneedle arrays. International Journal of Pharmaceutics, 415(1–2), 140–149. Panácek, A., et al. (2006). Silver colloid nanoparticles: Synthesis, characterization, and their antibacterial activity. The Journal of Physical Chemistry B, 110(33), 16248–16253. Reis, N., Ainsley, C., & Derby, B. (2005). Ink-jet delivery of particle suspensions by piezoelectric droplet ejectors. Journal of Applied Physics, 97(9). https://doi.org/10.1063/ 1.1888026. Scoutaris, N., et al. (2015). Development and biological evaluation of inkjet printed drug coatings on intravascular stent. Molecular Pharmaceutics, 13(1), 125–133. Shah, N. J., et al. (2011). Tunable dual growth factor delivery from polyelectrolyte multilayer films. Biomaterials, 32(26), 6183–6193. Tarcha, P. J., et al. (2007). The application of ink-jet technology for the coating and loading of drug-eluting stents. Annals of Biomedical Engineering, 35(10), 1791–1799. Torrisi, L., & Setola, R. (1993). Thermally assisted hydroxyapatite obtained by pulsed-laser deposition on titanium substrates. Thin Solid Films, 227(1), 32–36. Tsai, M.-H., & Hwang, W.-S. (2008). Effects of pulse voltage on the droplet formation of alcohol and ethylene glycol in a piezoelectric inkjet printing process with bipolar pulse. Materials Transactions, 49(2), 331–338. U.S. food and Drug Administration, Implants and prosthetics. 2015; Available from: http://www.fda.gov/MedicalDevices/ProductsandMedicalProcedures/ImplantsandProsthetics/. Wu, P., & Grainger, D. W. (2006). Drug/device combinations for local drug therapies and infection prophylaxis. Biomaterials, 27(11), 2450–2467. Yang, A., Cheng, C., & Lin, C. (2006). Investigation of droplet-ejection characteristics for a piezoelectric inkjet printhead. Proceedings of the Institution of Mechanical Engineers, Part C: Journal of Mechanical Engineering Science, 220(4), 435–445. Zeng, H., & Lacefield, W. R. (2000). XPS, EDX and FTIR analysis of pulsed laser deposited calcium phosphate bioceramic coatings: The effects of various process parameters. Biomaterials, 21(1), 23–30.

Further Reading Adden, N., et al. (2006). Phosphonic acid monolayers for binding of bioactive molecules to titanium surfaces. Langmuir, 22(19), 8197–8204. Alghamdi, H. S., et al. (2013). Osteogenicity of titanium implants coated with calcium phosphate or collagen type-I in osteoporotic rats. Biomaterials, 34(15), 3747–3757. Cai, K., et al. (2005). Polysaccharide-protein surface modification of titanium via a layer-by-layer technique: Characterization and cell behaviour aspects. Biomaterials, 26(30), 5960–5971. Dinda, G. P., Shin, J., & Mazumder, J. (2009). Pulsed laser deposition of hydroxyapatite thin films on Ti–6Al–4V: Effect of heat treatment on structure and properties. Acta Biomaterialia, 5(5), 1821–1830. Jang, H. W., et al. (2011). Surface characteristics and osteoblast cell response on TiN- and -coated Ti implant. Biomedical Engineering Letters, 1(2), 99–107.

Porous Biomaterials and Scaffolds for Tissue Engineering Liliana Liverani, University of Erlangen-Nuremberg, Erlangen, Germany Vincenzo Guarino, National Research Council of Italy, Naples, Italy Vincenzo La Carrubba, University of Palermo, Palermo, Italy Aldo R Boccaccini, University of Erlangen-Nuremberg, Erlangen, Germany © 2019 Elsevier Inc. All rights reserved.

Tissue Engineering and Scaffolds Requirements Porous Scaffolds Fabrication Techniques Electrospun Scaffolds Principle/Mechanism at the Basis of the Technique Advantages and Limitations of the Techniques In Vitro/In Vivo/Clinical Applications Polymeric Foams and Scaffolds Principle/Mechanism at the Basis of the Technique Advantages and Limitations of the Technique Polymeric foams produced by TIPS Polymeric membranes produced by DIPS In Vitro Applications Inorganic Scaffolds Inorganic Porous Scaffold Fabrication Techniques Scaffolds Fabricated by Additive Manufacturing Principle/Mechanism at the Basis of the Technique Advantages and Limitations of the Technique Biomaterials printing Tissue and organ printing In Vitro/In Vivo Clinical Applications Integration of Different Scaffolds Fabrication Techniques Conclusions and Future Perspectives Acknowledgments References

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Tissue Engineering and Scaffolds Requirements Tissue engineering is an emerging field proposing a novel approach for repairing, improving, and restoring tissue and organ function. Its definition started in the 1990s and evolved through the last decades (Langer and Vacanti, 1993; Vacanti and Langer, 1999; O’Brien, 2011). Tissue engineering approach aims to overcome all the limits presented by the current treatments for tissue replacement, like autografts, homografts, or xenografts. In fact, this approach is based on the use of functional three-dimensional scaffolds fabricated by using biodegradable biomaterials, having the function to support, promote, and enhance the new tissue formation. Scaffolds could be fabricated by using polymeric, ceramic, metallic, and composite materials suitable for different tissue target of the regeneration and should be characterized by the properties briefly reported as follows. Biocompatibility is a property related to the host tissue response, and it was introduced as fundamental properties of biomaterials in the 1940s regarding implantable medical devices. The first definitions considered biocompatible less reactive or inert materials, while the evolution of the definition implies a suitable and appropriate host response, related to the specific anatomical site. Recently, in the last years, a definition of biocompatibility specifically addressed to scaffolds was stated by Williams (2008), and it refers to the scaffold’s ability to “support the appropriate cellular activity, including the facilitation of molecular and mechanical signaling systems, in order to optimize tissue regeneration, without eliciting any undesirable local or systemic responses in the eventual host” (Williams, 2008). Biodegradability is another pillar of scaffold characteristics; in fact, the kinetic of the scaffold degradation should be synergic with the new tissue formation, allowing an initial support from mechanical and morphological point of view, suitable for cell adhesion, migration, and colonization of the scaffolds, supported eventually also by biomolecules release. Degradation rate and progressively decrease in mechanical properties are strictly related to the selection of the appropriate materials, combined also with the scaffold surface functionalization with biomolecules or growth factors that enhance the new tissue formation (Hutmacher, 2001). Scaffold mechanical properties play a pivotal role in an effective scaffold design and should be tailored according to the tissue target of the regeneration. In fact, the scaffolds should retain the desired shape, facilitating the implant in vivo and should have similar or comparable mechanical strength of the healthy native tissue surrounding the implant. Scaffold mechanical and structural integrity should not be lost before the complete neotissue formation (Karande et al., 2014).

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Scaffold surface properties are fundamental to improve and enhance cell response, adhesion, and migration. The functionalization of scaffold surface and the related release of biomolecules could be enhanced by appropriate surface scaffold morphologies, like fibrous structure (Karande et al., 2014). Porosity and pore interconnectivity, finalized to cell migration, oxygen and nutrient permeation and transport, and metabolite removal to improve tissue vascularization and new tissue formation, are also fundamental for the regeneration of all the tissues. Appropriate scaffold porosity and pore interconnectivity should be compromised with decreased or suboptimal scaffold mechanical properties and stability (Loh and Choong, 2013). In particular, recently, many research works have been focused on the fabrication of scaffolds with controlled three-dimensional architecture to enhance the biomimetic approach including the fabrication of scaffolds able to mimic the native healthy tissue in terms not only of composition but also in particular of morphology, threedimensional structure, and therefore cell response (Hollister, 2005). Usually, referring to scaffold pore size, it is not commonly applied the classification introduced by IUPAC (Everett, 1972), which defined the pores exceeding 50 nm as macropores, the one not exceeding 2 nm as micropores, and the intermediate size between 2 and 50 nm as mesopores. Considering the average cell dimensions and that the pore size needed to fulfill the scaffold requirements explained earlier, this classification seems to not report an appropriate range of pore size suitable for tissue engineering applications. In order to ensure all these functional requirements, the selection of the appropriate biomaterials is fundamental. In particular, it is possible to fabricate polymeric, ceramic, metallic, and composite scaffolds. Fabrication techniques and scaffold properties regarding polymeric, ceramic, and their composite scaffolds are analyzed more in details in the next paragraphs. Metallic scaffolds are not suitable for applications in soft tissue regeneration, and they are mainly use for hard tissue regeneration as bone tissue engineering, in particular for load-bearing areas, because of metal fatigue resistance and high compressive strength. There are some limits and disadvantages in the use of metals for scaffold fabrication in particular for the obtainment of controlled internal scaffold architecture with conventional processes. In particular, the limited number of biodegradable metals (mostly used are magnesium (Mg) (Yazdimamaghani et al., 2017), iron (Fe) (Mohd Daud et al., 2014), and their alloys) and the metal degradation rate in the human body are pivotal parameters to take into account during the scaffold design and fabrication (Wang et al., 2016). In vitro and in vivo characterizations, by using animal models, are fundamental to assess cell–scaffold interactions and their capability to enhance tissue regeneration (Peric et al., 2015).

Porous Scaffolds Fabrication Techniques As reported in the previous paragraph, porosity is a key scaffold characteristic. In particular, the pore size and interconnectivity play a pivotal role in triggering cell–scaffold interactions. Several scaffold fabrication techniques could be used for the obtainment of porous scaffolds; the main characteristics of these techniques are summarized in Table 1. Salt-leaching technique is based on the use of a porogen material (i.e., salt crystals), which is added to the polymeric solution and subsequently removed by dissolution of the porogen in a proper solvent. The pore size is related to the size of the porogen used (Aramwit et al., 2015). In the gas foaming process, the gas is used as porogen for the formation of porous structure (Mikos and Temenoff, 2000), and it is possible to remove the leaching step during the fabrication process. As a drawback, it is not possible to control the pore size and interconnectivity, but it is possible to combine this technique with other scaffold fabrication techniques to improve the control on these two key parameters (Joshi et al., 2015). The scaffold fabrication techniques based on the phase separation, electrospinning, additive manufacturing, and foam replica method processes are described in details in the following paragraphs. In the freeze-drying technique, the formation of pores is based on the sublimation of the frozen water into the gas phase. The porosity and pore size are strictly related to the process parameters (like gradient of temperature and time of the process) and solution parameters (like water-to-polymer ratio and viscosity of the solution) (Zhang et al., 2014). There are different techniques for the measurement and the evaluation of scaffold porosity and pore size. The total porosity, expressed in percentage, is often calculated starting from the bulk density of the scaffold, and it is related to the amount of pores presented in the scaffolds. Physical characterization could be used for the assessment of the porosity as gravimetric method, mercury porosimetry, gas adsorption analysis, liquid displacement method, etc. Also, imaging techniques (i.e., scanning electron microscopy analysis and microcomputed tomography imaging) could be exploited for the evaluation of scaffold porosity and pore size (Loh and Choong, 2013).

Electrospun Scaffolds Principle/Mechanism at the Basis of the Technique The electrospinning technique is based on the application of a high electric potential between two electrodes of opposite polarities. This high voltage is able to overcome the surface tension inside a polymeric solution, or melt, allowing the complete evaporation of the solvent and the formation of fibrous structure. This technique is widely used for the fabrication of nano- and microfibrous scaffolds, resembling the morphology of the extracellular matrix (ECM) (Teo et al., 2011). The standard setup is composed of a high-voltage generator, a syringe pump able to control the solution flow rate, a syringe with a metallic needle, and a grounded fiber collector. There are three categories of parameters affecting the process and the properties of the obtained electrospun fibers: parameters related to the solution properties (i.e., polymer concentration, solvent system,

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Summary of the characteristics of the different scaffold fabrication techniques

Scaffold fabrication technique Advantages Salt leaching

- Pore size related to porogen particle size - Versatile in combination with other scaffold fabrication techniques, in particular to improve the pore interconnectivity

Gas foaming

- Avoid the leaching step in the scaffold fabrication process - Reduced use of solvents

Phase separation

Avoid the leaching step in the scaffold fabrication process

Freeze drying

High temperature and leaching step not required during the process

Additive manufacturing

Control of the geometry, pore size, and interconnectivity

Electrospinning

Control over fiber diameter and related openings between the fibers

Foam replica method

Control over the pore size and pore interconnectivity, related to the selected sacrificial foam

Disadvantages

Range of pore size (mm)

Applications (tissue target of regeneration)

- Not control on pore interconnectivity - Low mechanical properties respect to the native tissues (not suitable for hard tissue regeneration) - Possible residual of porogen or toxic solvents not suitable for biomedical applications - Not control on pore size and on pore interconnectivity - Low mechanical properties respect to the native tissues (not suitable for hard tissue regeneration) - Addition of organic solvents could inhibits the incorporation of biomolecules during the process - Small pore size not suitable for all the application - Low mechanical properties respect to the native tissues (not suitable for hard tissue regeneration) - Longtime processing - Low mechanical properties respect to the native tissues (not suitable for hard tissue regeneration) - Constrains in the process of some polymer/biomaterials

Pore size related to porogen particle size (usually in the range of 100–500)

Many fields of applications, in particular related to other scaffold techniques, like wound healing (Aramwit et al., 2015), bone tissue engineering (Sadiasa et al., 2014)

Pore size up to 100

Many fields of applications, in particular related to other scaffold techniques, like skin tissue engineering (Poursamar et al., 2016)

1–100

Bone tissue engineering (Rezabeigi et al., 2016), vascular tissue engineering (Pavia et al., 2013)

50–200

Cartilage tissue engineering (Vishwanath et al., 2016); osteochondral tissue engineering (Levingstone et al., 2016)

100–500

Bone tissue engineering (Wang et al., 2016), cartilage tissue engineering (Jessop et al., 2016), organ printing (Murr, 2016) Mainly soft tissue engineering applications and interface tissue engineering if combined with other scaffold fabrication techniques (O’Connor and McGuinness, 2016) Bone tissue engineering (Miguez-Pacheco et al., 2016; Baino et al., 2016)

- Limited mechanical properties - Required the integration with other techniques for the fabrication of bulk 3-D scaffolds - High values of porosity but reduced mechanical properties

From 1 to 500

Pore size related to the foam (usually in the range of 100–600)

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Table 1

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conductivity, viscosity and type (solution vs. melt electrospinning) of the solution, its volatility, dielectric effect, and surface tension), electrospinning process parameters (i.e., applied voltage, feed rate of polymeric solution, its temperature, capillary-tocollector distance, effect and type of collector, needle diameter, and configuration of the nozzle), and environmental properties (like temperature and relative humidity) (Putti et al., 2015). The fabrication of porous electrospun scaffolds and in particular the obtainment of macroporosities in the electrospun mats has been the focus of many researches in the last years (Rnjak-Kovacina and Weiss, 2011). In particular, besides the combination with other scaffold fabrication techniques, it is not often possible to achieve a pore size and pore interconnectivity appropriate for allowing cell infiltration inside the scaffold structure (Rnjak-Kovacina and Weiss, 2011).

Advantages and Limitations of the Techniques The electrospun mats have a fibrillar structure resembling the morphology of the native extracellular matrix (ECM), having the advantage of representing a suitable example of biomimetic approach in scaffold fabrication (Woods and Flanagan, 2014), enhanced also by the possibility to obtain aligned fibers resembling the native tissue anisotropy. In fact, this technique is versatile because it allows the use and the processability of a huge number of natural and synthetic polymers and their blends, the possibility to select different type of solvents (Shenoy et al., 2005), with particular focus in the recent research works on the use of benign solvents and on green electrospinning (Sun et al., 2010). In particular, the use of polymeric blends of natural and synthetic polymers could be interesting to trigger the polymer degradation rate, as showed in Fig. 1, in which electrospun mats obtained from a blend of poly (ε-caprolactone) (PCL) and zein showed fiber structure degradation after 1 day and 1 week of immersion in phosphatebuffered saline (PBS) solution, demonstrating an increased degradation rate respect to the neat PCL. The control over the fiber degradation rate is relevant to improve cell adhesion and enhance the new tissue growth and also for the release of drugs or particles embedded in the polymeric solution. It is also possible to obtain composite electrospun fibers by adding a different phase in the polymer matrix; in particular, the most commonly used have been hydroxyapatite particles (Liverani et al., 2014), bioactive glasses particles (Liverani and Boccaccini, 2016; Gönen et al., 2016; Lepry et al., 2016), and ß-tricalcium phosphate particles (Erisken et al., 2008) relevant and suitable for bone tissue engineering applications (Liverani et al., 2016). Particularly, relevant results were obtained in terms of preservation of bioactive glass particle bioactivity, even if embedded in polymeric electrospun fibers (Liverani and Boccaccini, 2016; Gönen et al., 2016; Lepry et al., 2016). It is also possible to obtain different fiber functionalizations with biomolecules, like growth factors (Li et al., 2006; Fu et al., 2008b). As reported in the previous section, the obtainment of an appropriate pore size and effective scaffold pore interconnectivity in the electrospun mats could be challenging, and different techniques have been used to increase the pore size, like the increase of the fiber diameter, the use of different fiber collectors, wet electrospinning, the use of a sacrificial polymer to add in blend before the electrospinning, the addition of surfactant in the solution before the electrospinning, cryogenic electrospinning, and the combination with salt leaching and gas foaming (Rnjak-Kovacina and Weiss, 2011).

In Vitro/In Vivo/Clinical Applications Many studies have been focused on the investigation of cell–electrospun fiber interactions, in particular for the evaluation of the effect of fiber alignment (Jin et al., 2017). Those studies about the applications of 3-D structure and gradient morphologies are particularly relevant, in particular for bone (Obata et al., 2013), cartilage (Chen et al., 2016b), and vascular tissue engineering (Ju et al., 2010). Even if in vitro tests play a fundamental role in the evaluation of scaffold performance in terms of cell response and new tissue formation, in vivo studies represent the gold standard in the preclinical step before human clinical trials. During the design of an in vivo experiment is pivotal to select the proper animal model considering the tissue target of the regeneration (i.e., in the case of vascular or bone tissue (Rocco et al., 2014; Peric et al., 2015)). For what concerns the clinical applications, there are several electrospun products already available on the market, like coronary balloon expandable stent system (Papyrus, Biotronik AG and Bioweb, Zeus Inc.), vascular access graft (AVflo, Nicast Ltd.), dural patch (ReDura, Medprin Regenerative Medical Technologies Co. Ltd.), and synthetic bone (Rebossis, Ortho ReBirth Co. Ltd.).

Fig. 1 SEM micrographs of electrospun mats obtained started from a blend of PCL and zein: as-spun (A) after 1 day of immersion in PBS solution (B) and after 1 week of immersion in PBS solution (electrospun mats fabricated and characterized by the authors, unpublished data).

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Polymeric Foams and Scaffolds Principle/Mechanism at the Basis of the Technique The largest part of porous structure is produced by techniques based on liquid–liquid phase separation of polymer solutions (van de Witte et al., 1996a), which still seem to be the most promising, flexible, and tunable route for the preparation of 3-D interconnected polymeric porous structures, owing to its versatility and ease of operation and to the vast latitude of achievable morphologies. It has been shown by several attempts in the last few years that the thermodynamics of phase separation plays a crucial role in determining the final structure (Matsuyama et al., 2000). Depending on the use of a thermal or diffusive driving force, the technique is often called thermally or diffusion-induced phase separation (TIPS or DIPS, respectively) (Akki et al., 1999). Along decades, many research groups have studied the thermodynamic and kinetic properties of polymer solutions (Akki et al., 1999; Han, 2000), in order to achieve a better control in membrane and foam production processes. The most widely model system studied is a ternary solution composed of a polymer, a solvent, and a nonsolvent, which has the function of promoting phase separation. The typical mechanism of phase separation of polymer solutions is liquid–liquid demixing leading to a polymer-rich and a polymer-lean phase (Han, 2000), according to a binodal demixing or to a spinodal decomposition or to a combination of both. The former, characterized by nucleation and growth, takes place inside the metastable region (between the binodal and the spinodal curve), while the second takes place inside the spinodal region (unstable). The morphology developed during a phase separation process is therefore crucially affected by the particular path followed in the thermodynamic diagram. If the prevailing mechanism is nucleation and growth, a porous structure with poorly interconnected cells is obtained, while spinodal decomposition gives rise to an interconnected network (Gong et al., 2006). The growth and coalescence of the polymer-lean phase leads to a porous structure according to phase separation path, which depends upon the utilization of a binary system (polymer/solvent) and/or of a ternary system (polymer/solvent/nonsolvent). After polymer solidification and solvent removal from the demixed solution, the space that was occupied by the solvent will create voids (pores). By changing the polymer concentration, the solvent, the cooling rate, and the final temperature, it was then possible to vary the phase separation path leading to scaffolds exhibiting different micro- and macrostructures and final properties (Ho et al., 2004). Among the polymers employed for preparing scaffolds via phase separation routes, poly-L-lactide (PLLA) occupies a predominant role. With this respect, van de Witte et al. (1996b) studied the influence of solid–liquid demixing, liquid–liquid demixing, and vitrification on PLLA membrane morphology obtained via immersion precipitation. Chloroform and methanol were used as solvent and nonsolvent, respectively. In a different study, the same research group applied the relations between phase diagram and membrane morphology for the ternary PLLA–chloroform–methanol system to PLLA–dioxane–methanol, PLLA–dioxane–water, and other ternary systems (van de Witte et al., 1996c).

Advantages and Limitations of the Technique Polymeric foams produced by TIPS Thermally induced phase separation (TIPS) is a widely used technique to prepare scaffolds for tissue engineering purposes (Gong et al., 2006) and porous membranes (Tanaka and Lloyd, 2004). It involves cooling a homogeneous polymeric solution to a temperature where the single-phase system becomes thermodynamically unstable and spontaneously separates into a polymer-rich and a polymer-lean phase. Generally, the system used is a ternary solution of a polymer, a solvent, and a nonsolvent. Thermodynamic diagrams provide information about a possible phase separation process, depending on temperature, pressure, and composition. Moreover, composition of separated phases and nucleating phase(s) is determined by thermodynamics. The dense, microporous, macroporous nucleating phase, in its turn, determines the type of microstructure obtained. Therefore, knowing the phase diagram of a given system can aid scientists in choosing the appropriate processing conditions, to obtain the desired microstructure. Nevertheless, the construction of a phase diagram for a ternary polymeric system is not straightforward. Experiments for the determination of tie lines are time-consuming. Moreover, as commercial polymers are polydisperse, phase diagrams lose their deterministic accountability. A more straightforward way to get experimental information about phase separation is the determination of cloudpoint curves (Barton and McHugh, 1999); however, this method is not fully quantitative. A more precise way to carry out turbidity measurement is to adopt light transmission detection methods (Bulte et al., 1996). Besides thermodynamic properties, even kinetic features must be considered in the design of a TIPS protocol. Phase separation kinetics depends on temperature, composition, cooling rate, and demixing time. Each parameter affects the final morphology of porous structure (Nunes and Inoue, 1996). An overview of the morphology exhibited by the foams allows one to infer that a multilevel hierarchical structure sets in, like often it occurs in complex processing procedures inducing structure gradients, for example, injection molding of polymers. As a matter of fact, at least three hierarchically ordered structural levels might be recognized (Pavia et al., 2008): a “macroporosity” level, with pores characterized by dimensions up to 10 mm, depending upon the experimental conditions (viz., demixing temperature and time); an “interconnection” level, whose effectiveness relies on the number of interconnecting channels per pore (with a characteristic size of the order of ca 10 mm); and a “microporosity” level, appearing when very long residence times in the metastable region are applied. The foams prepared at lower a demixing temperatures (25 C) present a constant pore size independent of demixing time, whereas at higher demixing temperatures (30 and 35 C), a classical sigmoidal shape is observed, with a final plateau level depending upon the demixing temperature (in the range of 40–100 mm). Pore size results can be interpreted on the basis of a balance between nucleation and growth processes. A slight variant of the aforementioned approach consists in the use of an innovative experimental apparatus able to imposedvia Peltier cellsddifferent thermal histories on two sample

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surfaces, thus giving the opportunity to produce nonhomogeneous foams. Although the underlying physical principles are the same regarding phase separation in homogeneously cooled polymer solutions, this technique allows the fabrication of foams with pore size monotonously varying along the thickness within a single sample, thus obtaining a material with features not accessible via a “classical” TIPS approach (Mannella et al., 2015). This type of scaffold structure has promising tissue engineering applications wherever a hierarchical architecture involving morphological variations in space is required, for example, in bone tissue repair.

Polymeric membranes produced by DIPS The production of polymeric membranes via diffusion-induced phase separation (DIPS) has been widely studied and applied for a number of model systems. The main steps of membrane formation through DIPS protocol are the casting of a polymer solution on a support and the immersion into a coagulation bath, usually composed of a nonsolvent. The mutual exchange of solvent and nonsolvent will modify the casted solution composition, inducing phase separation when the equilibrium curve is crossed. The liquid–liquid phase separation is responsible for the formation of a porous membrane, and it can occur via nucleation and growth and/or spinodal decomposition mechanisms. Local composition determines the demixing mechanism and the amounts of separated phases, according to its location in phase diagram. As a general rule, the immersion in a coagulation bath composed only of nonsolvent will produce a porous membrane but with a dense skin; this was related to a rapid increase of polymer concentration at the interface, owing to a fast outward solvent diffusion, which induces crystallization. This fact has been inferred by mass transfer modeling, providing an estimate of concentration profiles inside the polymeric film (Reuvers et al., 1987). Several experimental studies regarding to the formation of PLLA membranes by immersion precipitation have been reported in literature (Zoppi et al., 1999), where the effect of different solvents and nonsolvents, casting and coagulation bath concentration, and temperature were investigated to control the membrane morphology. However, some points, related to the external surface morphology and the influence of drying step, need to be further investigated and clarified. As a matter of fact, membranes prepared via DIPS ordinarily show a completely closed and nonporous external surface, owing to a local polymer concentration increase (Bulte, 1996). This fact reduces sensibly the overall membrane permeability, thus affecting the choice of potential applications. A number of solutions to this limitation were investigated, for example, employing a coagulation bath composed of both nonsolvent and solvent (Wijmans et al., 1983) or multiple coagulation baths (Yang et al., 2007) adopting a sacrificial layer approach (Li et al., 2008) or wetting both surfaces with a modified spinneret (He et al., 2003). Furthermore, the influence of desiccation step, that is, of solvent removal on the morphology of external surface needs to be accounted for. In typical DIPS practice, the membrane is washed or dried immediately after the immersion step; however, researchers often neglect the effects of this last operation on the membrane microstructure. As a matter of fact, during solvent removal, the system composition will vary; thus, an influence on membrane morphology can be reasonably expected (Montesanto et al., 2015).

In Vitro Applications Porous and biodegradable PLLA scaffolds for vascular tissue engineering applications, with a vessel-like shape, can be produced by a two-step experimental protocol, including a dip coating of a viscous polymer solution around a nylon fiber and a diffusioninduced phase separation (DIPS) by immersion into an antisolvent bath (Pavia et al., 2013). The first step (dip-coating bath, PLLA/dioxane solution) allows the production of the tubular structure of the vascular graft, whereas the second (DIPS) is responsible for the interesting level of porosity and for the surface microporosity as potentially ensuring nutrient and metabolites exchange during blood flow. The as-prepared scaffold can be utilized either as vascular grafts or embedded into a porous scaffold for the regeneration of an injured tissue as a pseudoperipheral circulatory system, with the aim of promoting a rapid vascularization of the whole tissue-engineered construct (Pavia et al., 2013). A cell culture inside the vessel-like scaffolds was carried out, by using endothelial cells (EC), which are the solely components of capillary bed and the first to form during embryonic development. After 21 days, the internal lumen of the scaffold is totally covered by EC, which have organized themselves into a well-differentiated vessel structure. Cells shown to form stable cell–cell interactions and spindle membrane protrusion, characteristic of mesenchymal endothelial phenotype, were not observed. These results suggest that the scaffold produced could be usefully employed in vascular tissue engineering applications (Pavia et al., 2013). In another research study, mouse mesoangioblasts (A6) were seeded onto bidimensional matrices within three-dimensional porous scaffolds of poly (L-lactic acid) (PLLA) prepared according to the protocol described in (Carfì-Pavia et al., 2009), in the presence or absence of a type I collagen coating. Results show that PLLA films allow direct cell adhesion and represent an optimal support for cell growth (Carfì-Pavia et al., 2009).

Inorganic Scaffolds Inorganic biomaterials, like hydroxyapatite, other calcium phosphate phases, apatite–wollastonite glass ceramics, and bioactive glasses, have been commonly used for tissue engineering applications, in particular for bone-tissue-related applications (Bose et al., 2012), because of the similarity with the mineral phase of native bone, but they could also find other applications in contact with soft tissues (Miguez-Pacheco et al., 2015). The use of the aforementioned inorganic biomaterials is suitable for tissue engineering applications because of their bioactivity, defined as the property of the materials, like bioactive glasses and bioactive glass ceramic, to form interfacial bonding with tissues (Cao and Hench, 1996). In healthy native hard tissue, the mechanism of apatite biomineralization occurs as a self-remodeling process leaded by bone cells and proteins, but it also possible to detect mineralization

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on the surface on bioactive inorganic materials, like crystalline calcium phosphate and bioactive glasses, acting as an interface for the integration with tissues (Kim et al., 2004). The use of bioactive glasses is also valuable for tissue engineering applications because of their angiogenic, osteogenic, and antibacterial properties (Hoppe et al., 2011; Jones et al., 2006). Despite these properties, these inorganic biomaterials showed also low mechanical strength and low fracture toughness, limiting their applications in loadbearing tissue regeneration (Gerhardt and Boccaccini, 2010).

Inorganic Porous Scaffold Fabrication Techniques Different scaffold fabrication techniques could be used for the production of inorganic porous scaffolds, like foam replica method, foaming techniques, freeze casting, and additive manufacturing. It is also possible to incorporate the inorganic particles in a polymeric matrix and subsequently process it to fabricate porous composite scaffolds by using other techniques, as electrospinning or TIPS. A brief description of these techniques is reported in this section. The foam replica method, used for scaffold fabrication since 2006 (Chen et al., 2006) is based on the use of a sacrificial porous template, like polyurethane foam, for the obtainment of a positive replica of this template. In details, the bioactive glass particles were dispersed in a slurry with a binder, that is, polyvinyl alcohol (PVA); then, the template foam is immersed in the slurry and coated with it; during a heat treatment, the sacrificial foam is burned out, and the bioactive glass is sintered at high temperature, an example of the scaffold morphology that it is possible to obtain is reported in Fig. 2. The application of this technique is not limited to a specific bioactive glass composition, allowing the fabrication of scaffolds with different type of bioactive glasses (Chen et al., 2006; Fu et al., 2008a) and also enhancing the scaffold functionalities, by using ion-doped bioactive glasses (Hoppe et al., 2013). The obtained scaffolds usually could reach total porosity values up to 80%–90%; that value plays a crucial role in inducing in vivo vascularization, as reported by Arkudas et al. (2013), but the obtainment of this high porosity implies scaffolds’ low mechanical properties. In order to reinforce the scaffold structure, different solutions were proposed, as the use of different type of sacrificial templates with different porosities, in order to compromise a decrease the total porosity value and an increase in the mechanical properties (Boccardi et al., 2015), or it is possible to perform polymeric coating to enhance the scaffold mechanical properties, for example, by using gelatin (Desimone et al., 2013) or blend of natural and synthetic polymers (Fereshteh et al., 2015). Other porous scaffold fabrication processes are the foaming techniques. These approaches are based on the formation of pore and based on the introduction of gas (Jones and Hench, 2003), with the limit of the lack of control in the pore size or in the use of surfactants inside a bioactive glass slurry, before the aging and the sintering process (Sepulveda et al., 2002). Few research works are focused on the use of freeze drying for the fabrication of inorganic porous scaffolds; in fact, this technique is consolidated for the fabrication of polymeric or composite (polymeric scaffolds with incorporated inorganic phase). The use of this technique has still the limit in the pore size that it will be possible to obtain; in fact, it is not possible to obtain pore bigger than 100 mm (Mallick, 2009; Song et al., 2006). Interesting and promising results were obtained from in vivo studies about inorganic porous scaffolds (Fu et al., 2010; Livingston et al., 2002), even if there are not still available clinical applications of these porous scaffolds.

Scaffolds Fabricated by Additive Manufacturing Principle/Mechanism at the Basis of the Technique In the last years, additive manufacturing (AM) technologies have imposed themselves as the most appropriate method for the production of ordered 3-D structures to be successfully used as instructive scaffolds for tissue and organ regeneration. They are

Fig. 2 SEM micrographs of 45S5 bioactive glass-based porous scaffold (A) and higher magnification of the wall of the pore (B) (porous scaffolds fabricated and characterized from the authors, unpublished data).

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generally based on layer-by-layer fabrication strategies that allow manufacturing solid platforms with complex shapes and microstructures from three-dimensional (3-D) model data, showing high degree of automation, good accuracy, and reproducibility. Therefore, they allow finely controlling morphological patterns to exert all the structural functionalities required to support viability of cells within the 3-D printed structure. In the last 10 years, several studies have widely demonstrated the enormous potential of AM technologies to design tailor-made scaffolds to properly guide cell activities for the regeneration of different kinds of soft and hard tissues (Billiet et al., 2012). Strong efforts have been spent to improve current technology based on the AM principia to further enhance the control of morphological features (i.e., pore size distribution, pore volume, and pore interconnectivity, anisotropy) or to implement less invasive processing routes able to more easily manipulate recently synthesized biomaterials with biological functionalities (i.e., modified proteins and polysaccharides, smart polymers, and new bioactive hydrogels) (Forbes and Rosenthal, 2014). Three-dimensional platforms fabricated by properly adapted AM technologies are also successfully emerging as 3-D in vitro models able to bridge traditional cell culture and in vivo modeling; indeed, they have been used to predict relevant aspects of in vivo behavior, only traditionally assessed by animal implants and/or human trials (Khademhosseini et al., 2006). Besides, it is well-known that traditionally used in vitro 2-D models show significant limitations to recapitulate the complex tissue microenvironment due to the innate tendency of cells to mimic in vivo behavior when grown in 3-D conditions (Fischbach et al., 2007). In this context, novel approaches based on 3-D bioprinting can combine main advantages of consolidated rapid prototyping (RP) techniques with innovative biofunctionalization strategies, thus providing much more physiologically relevant information about organogenesis, disease progression, and molecular release onto specific targets.

Advantages and Limitations of the Technique To date, a large variety of 3-D printing-based techniques have been implemented to design 3-D scaffolds for the replacement of tissues and organs. Each technique presents benefits and disadvantages in terms of feasibility, material processability, strut resolution, and productivity. A first classification may be attempted based on the working principle, that is, inkjet, laser-assisted techniques, electrically assisted induced atomization, pneumatic or screw extrusion, and stereolithographic techniques. Other classifications of AM technologies mainly refer to the implementation of custom-made setups to realize complex geometries during the printing process (Onuh and Yusuf, 1999) with the support of tailor-made dispensing devices (i.e., laser-, nozzle-, or printerbased systems) (Billiet et al., 2012). As for biomedical applications, an alternative but useful classification is between AM and cell-friendly AM technologies depending upon (a) the processing of biomaterials in the form of 3-D bioinspired scaffolds or (b) the creation of 3-D tissue/organ analogues by the combination of biomaterials with living cells (i.e., bioinks), as summarized in Fig. 3.

Biomaterials printing In the class of AM technologies to process biomaterials, it may be included traditional 3-D printing technology used to create 3-D ordered structures made of ceramic, metal, plastic, and polymers (synthetic or natural ones) via layer-by-layer approach. This process is commonly guided by 3-D modeling softwaredthat is, computer-aided design (CAD) or computer tomography (CT) scan imagesdable to physically reproduce the CAD/CAM sketch with the support of automatic printing equipment. Their working principle may drastically vary as a function of selected materials and desired structural/functional features by the application of different driving forces, that is, UV or laser light, electrostatic, piezoelectric, pressure, thermal, thermodynamic, or magnetic ones. Among them, most commonly used technologies to fabricate layer-by-layer structures are based on ultraviolet (UV) light photopolymerization. They include stereolithography (SLA) (Skoog et al., 2014), digital light processing (DLP) (Wallace et al., 2014), and selective laser sintering (SLS) (Ko et al., 2007) that reproduce complex designs by fast processing and high resolution. These methods allow fabricating 3-D scaffolds by hardening a photopolymer resin under the controlled exposure to UV light or another similar power source. In these cases, much attention is required for the selection of cytocompatible photoinitiators to minimize damaging effects on cell membrane, protein, and nucleic acids, ascribable to the formation of free radicals that may potentially restrict their use in tissue engineering applications (Hutmacher et al., 2004). In tissue engineering, another relevant benefit of photopolymerization mainly concerns the suitability to fabricate hydrogel-like scaffolds. In the past, all the main photopolymerization techniques generally worked by a bottom-up scheme; it means that the object was built from a fabrication support just below the resin surface. Subsequent layers were cured on the top of the previous layers by irradiation from earlier (Liska et al., 2007). Only recently, alternative approaches based on top-down setups are emerging with enormous success (Melchels et al., 2010). Focusing on the fabrication of polymer or composite scaffolds, the most commonly used technologies are based on extrusion processes such as fused deposition modeling (FDM) (Drummer et al., 2012). In this case, thermoplastic polymers are extruded from the 3-D printer head under melting conditions until forming thin filaments that can create 3-D architectures by adding layer by layer (Guarino et al., 2012). FDM allows for an accurate control over scaffold morphological properties, despite their use is restricted in the case of polymer melt with relatively high viscosity that can be hardly extruded through small diameter nozzle. Independently, upon the intrinsic viscosity of melt polymer solutions, fiber diameter cannot achieve average size down to 100 mm. Hence, similar technologies have been properly revisited to process polymer solution or slurry (i.e., 3-D colloidal inkjet printing) in order to drastically scale down strut resolution, that is, about 0.5–10 mm as average diameter. However, the generation of clinically relevant bioinspired systems from such colloidal inks shows relevant problems in terms of regulatory perspective, which drastically limit their use for medical applications (Smay et al., 2002).

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Biomaterial printing

• •

Selective laser sintering (SLS) 3D printing (3DP)

• •

Stereolithography (SLA) Digital light processing (DLP)

• • •

Fused deposition modeling (FDM) Precise extrusion deposition (PED) Melt electrospinning writing (MEW)

• • •

Pressure assisted mycrosyringe (PAM) Low temperature assisted manufacturing (LDM) Robocasting

Tissue/Organ printing

Powder based

Light based

• •

Laser assisted bioprinting (LaBP) Biological Laser Printing (BioLP)

Melt based

• •

Bioextrusion Bioplotter

Solution based

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Thermal/piezoelectric ink-jet Electrofluidodynamic ink-jet printing (EIP)

AM Technologies Fig. 3

Classification of AM technologies for biomaterials and cell printing.

Tissue and organ printing Recent discoveries in medical and materials science have enabled for the development of new AM technologies to design biocompatible materials in concert with cells to form complex 3-D functional biohybrid tissues that better address strong need of regenerative medicine to ex novo fabricate tissues and organs suitable for transplantation. Compared with biomaterial printing strategies, this approach involves additional complexities, such as the choice of materials, cell types, growth and differentiation factors, and technical improvements related to the sensitivities of living cells and the construction of tissues (Murphy and Atala, 2014). To date, biological AM technologies have already been used for the generation and transplantation of several tissues (i.e., bone, vascular grafts, trachea, myocardium, skin, and cartilage). More recently, they are forcefully emerging also for the development of highthroughput 3-D bioprinted tissue models for research, drug discovery, and toxicology (Oskui et al., 2016). The working principle is based on a layer-by-layer precise positioning of biological materials, biochemicals, and living cells, with spatial control of the placement of functional components, into a three-dimensional arrangement. There are several AM approaches to combine cells and biomaterials into hierarchically ordered 3-D constructs with biological and mechanical properties suitable for clinical restoration of tissue functions. They may provide (a) the manufacturing of identical replica of cellular and extracellular components of tissues/organs (Ingber et al., 2006), (b) the reproduction of biological tissues by self-assembling embryonic organs as a guide for mature organ development (Kasza et al., 2007), or (c) the fabrication of mini-tissue building blocks as smallest structural and functional components to be assembled in macrotissues with biologically inspired design and organization (Mironov et al., 2009). In particular, it is possible to roughly distinguish between bioprinting systems able to deposit a continuous material bead (dispensing systems) or multiple materials in short interrupted or defined spaces (dropping systems) to form the 3-D structure, as reported in Fig. 4. The former ones commonly involve the use of microextrusion or inkjet bioprinters properly set to process biological matter at safe body conditions (i.e., temperature). As for microextrusion technologies, working principle is generally based on robotically controlled extrusion of materials by a properly designed head that allows dispensing biological matter in the form of continuous beads rather than liquid droplets. Small beads of bioslurry or bioink are deposited layer by layer in two dimensions onto the stage, under specific instructions imparted by a CAD–CAM system. A myriad of bioinks are compatible with microextrusion printers,

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Fig. 4

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Schematic summary of AM technologiesdexploiting specific drawing forcesdfor 3-D cell printing.

including hydrogels, biocompatible copolymers, and cell spheroids (Peltola et al., 2008). The most common methods to extrude biological materials via 3-D bioprinting are driven by pneumatic or mechanical (piston or screw) forces (Cohen et al., 2006). Analogously, inkjet printing technologies may allow directly printing spatially well-organized patterns by drop-to-drop deposition of cells and biomaterial slurries for the realization of bioinspired microstructures variously encoded with bioactive molecules to address specific cell behaviors. Among AM technologies, inkjet techniques are the most commonly used in 3-D tissue/organ printing for their increasing resolution, precision, and speed in comparison with other ones. The working principle is based on the application of thermal and/or piezoelectric forces able to form droplets with controlled sizes via actuation mechanism (Boland et al., 2006). More in general, the most part of current printing technologies are constrained by several limitations related to (a) the ink viscositydup to 10 cP excessive force may be required to eject dropsd(b) clogging of small size nozzles, and (c) the generation of pattern smaller than the nozzle size. More recently, many researchers are focusing on electrohydrodynamic inkjet printing (EIP) based on the administration of electric field to polymeric solution or slurry. The use of electric field as driving force to draw micropatterns by physical additive mass flow is promoting the development of a new set of bioprinting techniques to form continuous patterning, drop-on-demand printing, and thin-film deposition (i.e., electrospray) as a function of the ejection mode of liquid from the nozzle (Jayasinghe et al., 2006).

In Vitro/In Vivo Clinical Applications Despite the fact that a certain number of AM processes and 3-D scaffold products have already been approved by regulatory bodies for (routine) clinical use, an insufficient number of studies have been deeply characterized though extensive in vitro and in vivo studies and obtained results referred to a restricted number of biomaterials. Indeed, there is a still limited availability of materials really compatible with AM processes. Indeed, traditionally used biomaterials cannot be processed by AM techniques, while the bestperforming materials in AM machines, in terms of accuracy and functionality, are often not biocompatible or do not exhibit the required biodegradation behavior for the regeneration of specific tissues. The development of novel biomaterials or bioinks is one of the most important challenges for the future of AM technologies in biomedical field. Future outlook of advanced AM technologies still remains the design and fabrication of products processed by a rational printing of cells and biomacromolecules derived from native extracellular matrix (ECM), in combination with unexplored biomaterials (i.e., bioinks and powders), in order

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to generate in vitro and/or in vivo tissue analogue structures. In this case, the choice of bioink may be further limited by stringent conditions related to the combination of printing materials with cells. Recent approaches based on dropping strategies (i.e., inkjet or laser-assisted bioprinting) are brightly overcoming these limitation by encapsulating cells in biodegradable hydrogels to mimic a tissue-like environment (Bertassoni et al., 2014). In this case, the peculiar characteristics of hydrogels can protect inner cells from the shear force generated during the printing process, maintaining cell biofunctions (i.e., self-renewal ability and multilineage differentiation potency in the case of stem cells) without penalizing basic key features of AM technologies (i.e., high resolution and accuracy in single-cell deposition). Moreover, multiple bioinks and cell types can be also distributed within 3-D printed scaffolds to guide tissue generation processes, thus enabling regeneration of more complex tissue structures (Rutz et al., 2015). By newly implemented AM technologies, it will be possible to design and rapidly fabricate mechanically stable, functional, human-scale tissues such as the mandible, calvarial bone, cartilage, and skeletal muscle by simultaneously plotting cell-laden material with biodegradable porous structures to assure an efficient transport of nutrient transport and the required exchange of metabolites.

Integration of Different Scaffolds Fabrication Techniques The integration of different scaffold fabrication techniques allows to overcome the limitations of the single techniques, obtaining scaffolds with enhanced properties in terms of morphology, compositional and porosity gradients, functionalization with biomolecules, and appropriate mechanical properties related to the tissue target of regeneration. Usually, this combination of different techniques results in particularly suitable for interface tissue engineering applications, aiming to repair or regenerate the functions of diseased or damaged zones at the interface of different tissue types (also called “interface tissues”) (Seidi et al., 2011). In this section, few examples of synergic integration of different techniques are reported. Mi et al. (Mi et al., 2016) reported the successful integration of electrospinning technique and TIPS. Interesting results were reported by Biswas et al. (2017), reporting the integration of mechanical foaming process with TIPS to control pore formation inside scaffolds. In order to achieve highly performance porous structures using an automated stage, melt electrospinning writing (MEW) based on the combination of electrospinning with largerscale AM is recently emerging as the most interesting AM technology to realize ordered 3-D architectures to be used as scaffolds for hard tissue engineering applications. This technique is based on the interaction of melt polymers by electrostatic forces (i.e., melt electrospinning) and is supported by the use of automated machines able to mechanically control spinneret translation along x–yaxis, in order to realize fiber dispensing systems to directly write polymers in the form of 3-D scaffolds with controllable architectures and patterns (Brown et al., 2011). MEWs allow for the deposition of polymeric filaments with average diameters down to 10 mm, by a fine and synergic control of extrusion rate and spinnerets/collectors speeds, until to form a reproducible threedimensional (3-D) lattice with characteristic dimensions suitable for cell and tissue colonization (Farrugia et al., 2013). Using this approach, scaffold architectures can be readily created without the need for organic solvents, combining specific benefits of melt extrusion-based AM methods and electrospinning in a unique way (Dalton et al., 2013). Accordingly, some authors also attempted to combine solution and melt electrospinning to reproduce a random nanofiber texturing onto surfaces of ordered filament meshes for the fabrication of a hybrid vascular graft with interesting results in terms of cell adhesion and structural recognition (Centola et al., 2010). Moreover, through an accurate manipulation of the melt electrospinning process parameters (i.e., spinneret diameter, voltage, and collector distance), different struts with peculiar fiber sizes can be processed by using a large set of biodegradable and/or bioactive materials, for the fabrication of tailor-made scaffolds for tissue engineering use. To date, poly(ε-caprolactone) (PCL) is the most frequently processed polymer by MEW due to its low melting pointdabout 63 Cdin combination with well-known biodegradation properties and acceptable in vivo host response (Brown et al., 2012). Recently, it has been readily copolymerized with other cyclic monomers such as lactide and trimethylene carbonate to generate new substrates with different biodegradation properties and biological response. More recently, photo-cross-linkable poly(L-lactide-co-ε-caprolactone-co-acryloyl carbonate) copolymers (poly(LLA-ε-CL-AC)) have been processed by MEW followed by UV irradiation at room temperature to realize 3-D scaffolds without relevant alteration of surface structure but a significant improvement of mechanical stiffness under dynamic loads to be successfully used in bone and connective tissue regeneration (Chen et al., 2016a). For instance, melt electrospinning writing (MEW)drising from the combination of FDM and electrospinningdhas demonstrated to be a computer-aided electrofluidodynamic printing technique able to rationally design and fabricate fibrous scaffolds with thinner fiber sizesddown to 50 mm as diameter. MEWs have been successfully used to process thermoplastics like poly(ε-caprolactone) or photo-cross-linkable biopolymer poly(L-lactide-co-ε-caprolactone-co-acryloyl carbonate) to obtain 3-D scaffolds with improved mechanical properties for hard tissue engineering (TE) applications. Besides, melt electrospinning allows bridging the gap between nanoscale production methods with insufficient fiber deposition control (i.e., electrospinning) and resolution-limited AM methodologies. Several studies have demonstrated that large meshes generated by MEW may more efficiently contribute to 3-D cell invasiveness, with relevant improvements respect to electrospun scaffolds, frequently perceived by cells as 2-D structures with limited cell penetration (Simonet et al., 2007) or, respect to FDM processed scaffolds with not efficient cell-to-cell and cell-to-material interactions, due to too much larger pores (> 100 mm). Moreover, polymers approved by Food and Drug Administration (FDA) such as PCL can be processed through basic MEW in their pure form without the use of toxic solvents (Hutmacher and Dalton, 2011), thus minimizing postproduction costs. In terms of functional graded scaffolds, a recent and innovative approach is based on the “4-D bioprinting,” considering the variable time as a fourth dimension, relevant for the scaffold property evaluation and functionalities. In fact, in this approach, the fact that living cells colonized the scaffolds and dynamically interact with it induces an evolution in the scaffold properties (Gao et al., 2016).

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Conclusions and Future Perspectives The main challenge for the regeneration, restoration, or replacement of defective or injured functional living organs and tissues is to design 3-D scaffolds with specific functions tailor-made for specific targets. In particular, they should provide internal pathways for cell attachment and migration, promoting the exchange of specific growth factors and waste products and preserving their structural/mechanical properties during ex novo tissue formation. In order to achieve these functions, they have to present a highly interconnected porous structure able to create a friendly microenvironment for cells from chemical and morphological point of view. To reach this goal, many scaffold fabrication techniques could be exploited and also combined, as reported in this contribution. In the next future, it could be possible to design innovative implantable devices with more efficiently control of drug administration for innovative therapeutic uses.

Acknowledgments Liliana Liverani acknowledges funding from the European Union’s Horizon 2020 research and innovation program under the Marie SkłodowskaCurie grant agreement No. 657264.

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Preparation and Properties of Coatings and Thin Films on Metal Implants Zhong Li and Khiam Aik Khor, Nanyang Technological University, Singapore © 2019 Elsevier Inc. All rights reserved.

Introduction Preparation of Coatings and Thin Films on Metal Implants Commonly Used Metallic Implant Materials Physical Methods for Preparing Coatings and Thin Films Thermal spraying Glow discharge plasma Ion implantation and deposition Physical vapor deposition Chemical Methods for Preparing Coatings and Thin Films Electrochemical methods Chemical vapor deposition Sol–gel processes Biomimetic modifications Properties of Coatings and Thin Films on Metal Implants Desirable Surface Functionalities of Metallic Implants Osteoconductive and Osteogenic Coatings Corrosion-Resistant Coatings Wear-Resistant Coatings Hemocompatible Coatings Antibacterial Coatings Concluding Remarks References

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Glossary Angioplasty Procedure with a balloon-tipped catheter to enlarge a narrowing in a blood vessel, especially a coronary artery. Bioactivity The specific biological response at the interface of a material that results in the formation of a bond, usually through a carbonated hydroxyapatite layer, between the tissues and the material. Biocompatibility The ability of a biomaterial to perform its desired function without eliciting any undesirable local or systemic effects in the recipient. Biodegradability/bioresorbability The disintegration of materials by biological means. Biomaterial A natural or synthetic substance that has been adapted and is used to direct, by control of interactions with components of living systems, the course of any therapeutic or diagnostic procedure. Hemocompatibility An implanted material’s compatibility with blood. Osseointegration The direct structural and functional connection between living bone and the surface of an implant. Osteoconductivity The supply of a surface along which bone migrates. Osteogenesis The development and formation of the bone. Osteoinductivity The process that induces osteogenesis by stimulating immature cells into preosteoblasts. Osteolysis An active resorption of bone matrix by osteoclasts, cells that function in the breakdown and resorption of bone tissue. Osteosynthesis A surgical procedure used to reduce and internally fix a bone fracture with implantable devices. Stress-shielding effect Reduction in the density of the bones around an implant site caused by the implant’s higher Young’s modulus that reduced mechanical loads on the bones. Thrombosis Formation or the presence of a blood clot within a blood vessel.

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Introduction The surgical use of metallic implants can be traced back to the 16th century (Gotman, 1997). However, it was not until the introduction of Lister’s aseptic surgery techniques in the 1860s that the steep increase in the use of metals and alloys in implant surgery was realized (Lister, 1867). Since then, metallic biomaterials have played a predominant role in orthopedic surgery. The unmet clinical need for human hard tissue repair has been the primary driving force for the development of metal implants. This unmet need gap is being widened by the world’s rapidly aging population. A wide range of materials, including metals, polymers, and ceramics, have been developed and used for reconstructive surgery of hard tissue. Among them, metallic biomaterials fall into a most attractive category because of their excellent mechanical properties such as high ultimate tensile strength, good damage tolerance, ductility, formability, and machinability. The early medical uses of metal implants were mainly in bone contacting and replacement applications, such as the internal fixation of fractured long bones. Their applications were later extended to dental implants and cardiovascular stents. More recently, increasing research interest has been drawn to the study on biodegradable magnesium-based implants (Hornberger et al., 2012). Metallic biomaterials are currently being utilized in a vast number of permanent (such as total joint replacements) and temporary implants (such as bone plates, pins, and screws). Corrosion behavior is a key consideration for the selection of metallic implant materials because it fundamentally determines the implants’ long-term biocompatibility. Furthermore, in load-bearing applications, they have to possess sufficient strength, fracture toughness, and wear resistance to sustain cyclic loading and present minimal implant-bone modulus mismatch to mitigate the stress-shielding effect. While the bulk properties are primarily responsible for the biomechanical requirements, the corrosion and wear resistance, which constitutes a major disadvantage of metal implants, is more dependent on the surface characteristics. The surface properties of metal implants play a vital role in their interactions with cells and are strongly correlated with their attachment to the surrounding tissue. Therefore, tuning the physiochemical niche through surface modification of metallic implants has long been employed to enhance their corrosion and wear resistance and hence biocompatibility and, at the same time, retain their meritorious bulk attributes. The surface modification of metal implants is also warranted by the relentless pursuit of improved functionality in various medical practices. Biocompatibility and functionality stand for two sets of properties that regulate the performance of an implant material. Most metallic biomaterials are classified as first-generation biomaterials, meaning they are bioinert (Hench and Polak, 2002). Therefore, most of them are surrounded by a fibrous foreign body capsule (FBC) and do not elicit active biological responses in the human body because of their inherent bioinertness. This may be considered acceptable when they are only sought to match the anatomical form of the diseased tissue and provide mechanical support and stabilization. Nevertheless, for an increasing number of biomedical applications, the implant materials are expected to evoke and elicit desired responses from the body. This has resulted in the growing popularity of the clinical use of second- and third-generation biomaterials, either in their bulk form or as coatings or thin films on a substrate. Second-generation biomaterials are either bioactive or bioresorbable (e.g., BioglassÒ, poly(lactic-co-glycolic acid)), and third-generation ones are both bioactive and bioresorbable (e.g., hydroxyapatite/collagen composite) (Murugan and Ramakrishna, 2005). The “economical” use of such materials by coating them on metallic implants can significantly diversify and augment the biological aspects of the implant surfaces, such as osseointegration and suppression of adverse effects (e.g., FBC formation, thrombogenicity, and implant infection) (Goodman et al., 2013). The ultimate goal is to facilitate the restoration of full tissue and organ function. Two general approaches for the surface modification of metal implants are (1) mechanically/physically altering the atoms in their existing forms and (2) fabricating a surface coating. Typical examples of the first approach are machining, grinding, polishing, grit blasting, and beading. They can be used to achieve specific roughness, topography, and surface energy. They are also frequently employed as pretreatment steps for subsequent coating deposition, mostly for the removal of surface contamination and strengthened interfacial bonding. This section focuses on the surface modification of metal implants through the fabrication of coatings and thin films. Emphasis will be attached to the creation and evaluation of the produced coatings in the context of intended clinical applications of the metal implants.

Preparation of Coatings and Thin Films on Metal Implants Commonly Used Metallic Implant Materials The selection of metallic biomaterials is essentially dependent on their corrosion resistance and toxicity. Commonly used metallic implant materials include stainless steel, titanium alloys, cobalt-based alloys, and magnesium alloys (Chen and Thouas, 2015). Stainless steel is a generic term referring to iron-based alloys with at least 10.5 wt% Cr. It is the minimum amount required for the formation of a protective passivation chromium oxide film. The types of stainless steel frequently used in medicine include 316, 316L, 420, and 440. Stainless steel has relatively low corrosion resistance compared with Ti- and Co-based alloys but incurs lower production cost and has higher availability. It is the predominant stem material in total hip replacements. It has also been widely used in temporary/transient implants such as fracture plates, bone nails, and screws. Their service life ranges from several months to several years. Major issues associated with stainless steel implants are corrosion-induced toxicity and device degradation, as well as allergic responses of surrounding tissues caused by wear debris.

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Compared with stainless steel, Ti- and Co-based alloys possess higher chemical stability and are more preferred in long-term and permanent implants. Commercially pure Ti and Ti6Al4V are among the most frequently used Ti-based implant materials. At room temperature, a 3–7 nm thick oxide film (mainly composed of TiO2) is generated on the surface of Ti-based materials when they are exposed to the atmosphere. The strong passivation and repassivation ability renders them highly corrosion-resistant. Besides, they have smaller density and Young’s modulus and larger specific strength than most metallic biomaterials. Ti and its alloys have been successfully used in permanent load-bearing implants such as total joint replacements and dental implants, in addition to various osteosynthesis applications (e.g., bone plates, nails, screws, and craniofacial implants). Outstanding issues facing such applications are mainly the insufficient bending fatigue strength and low wear and abarasion resistance of Ti materials (Fu et al., 1998). Ti and Ti alloys have also been frequently used in cardiovascular applications such as prosthetic heart valves, cardiovascular stents, and protective cases in pacemakers. One of the major challenges facing these applications is the formation of blood clots on device surfaces. Particularly, NiTi alloy is one of the most popular materials in angioplasty by virtue of its special shapememory effect (Cragg et al., 1983). This effect refers to the restoration of the predefined shape of a material by heating it. A number of alloys, such as Cu–Zn, Au–Cd, and NiTi alloys, exhibit shape-memory effect, but NiTi is the most attractive and has been most widely used. When utilized as a self-expandable stent, the implant is usually inserted as a thin wire; upon being triggered by the body temperature, it dilates and returns to its preset shape. Besides vascular stents, gastrointestinal stents made from NiTi alloy have also achieved commendable clinical success. It should be noted that the biocompatibility of NiTi alloy is still controversial because of the potential toxicity of released Ni ions. Cobalt alloys utilized in medical implants usually refer to CoCrMo alloys. The alloying element Cr imparts to them a high corrosion resistance, which is comparable with that of Ti materials. The wear resistance of Co-based alloys is superior to that of Ti alloys because of their higher hardness. CoCrMo alloys are therefore the most frequently used metallic materials in joint-bearing systems. By contrast, they are less attractive as stem materials because of the stress-shielding effect caused by their high Young’s modulus. The release of Co, Cr, and Ni ions from permanent implants is also of concern. Mg alloys are classified as third-generation biomaterials. Such materials are designed to provide temporary mechanical support and would be gradually degraded and eventually replaced by native host tissues. The complete degradation eliminates the need to remove them with a second procedure when their function is no longer required. Mg alloys are advantageous over many other metallic implant materials because their density and Young’s modulus are similar to those of human bone. Their major drawback is the poor corrosion resistance and resultant hydrogen evolution and alkaline pH shift. The tribological properties of Mg alloys also need to be improved to mitigate the adverse effects caused by wear particles. Mg alloys have been more widely used in cardiovascular stents than in orthopedic implants (Witte, 2010). To enhance the performance of the aforementioned metallic implant materials, a myriad of techniques have been employed to apply coatings and thin films on them. They can be generally classified into physical and chemical methods. Based on the production mechanism, the coatings can be broadly categorized into two types, conversion coatings and deposited coatings (Hornberger et al., 2012). While deposited coatings have a wider compositional range, in conversion coatings coating delamination is less of a problem. In terms of chemical composition, metal, ceramic, polymer/biomolecule, and inorganic–organic hybrid coatings have been produced onto metal implants.

Physical Methods for Preparing Coatings and Thin Films Physical processes exploited for coating preparation are characterized by the supply of high thermal, electric, and kinetic energy. This generic term does not necessarily exclude methods involving chemical reactions.

Thermal spraying In thermal spraying processes, particulate materials are injected into a high-temperature heat source where they are thermally melted and energetically propelled to a substrate on which the droplets flatten and stick. They are capable of depositing thick coatings (up to several millimeters) over a large area at high deposition rates. In addition, thermal spraying processes allow versatile control of the coatings’ microstructural and chemical characteristics and are suitable for substrates in complex shapes. Common thermal spraying processes include plasma spraying, high-velocity oxygen-fuel spraying (HVOF), flame spraying, wire arc spraying, cold spraying, and warm spraying, with the first two being most frequently used for modifying the surface of metal implants. Plasma refers to partially ionized gas comprising free radicals, ions, photons, and electrons. Plasma spraying processes make use of electric energy to ionize the gas and create plasma. The plasma jet can offer very high temperatures (typically 12,000 K at the core) and energy densities. Plasma spraying is thus widely employed to prepare ceramic coatings. There are two types of plasma spraying processes, atmospheric plasma spraying and vacuum plasma spraying. Fig. 1A presents the typical surface morphology of a HA coating deposited by atmospheric plasma spraying. The HVOF technique generates a supersonic flame through the combustion of fuel gas (e.g., hydrogen, propane, propylene, and acetylene) in oxygen under high pressure (Khor et al., 2003). The powder is axially fed into the jet by an inert carrier gas, heated, and accelerated in the flame, whose temperature and velocity are typically > 1000 m/s and < 3000 C. The properties and performance of thermally sprayed coatings depend strongly on the process parameters. These include power setting, atmosphere, gas combination and flow rate, powder feedstock, standoff distance, substrate temperature and angle, and cooling rate.

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Fig. 1 (A) Scanning electron microscopy (SEM) image of a plasma-sprayed HA coating on Ti6Al4V. (B) SEM image of commercially pure Ti treated by MAO in phosphoric acid. (C) Atomic force microscopy image of diamond-like carbon (DLC) on stainless steel prepared by chemical vapor deposition (CVD).

Glow discharge plasma Glow discharge plasma is formed when the voltage applied on a low-pressure gas exceeds its breakdown voltage and causes its ionization. The workpiece immersed in glow discharge plasma is bombarded by electrons and ions, leading to rearrangement and sputtering of the surface atoms. The two main glow discharge plasma sources are magnetron discharge and radio-frequency (RF) discharge. Glow discharge plasma is frequently used to remove surface contamination and increase surface energy of biomaterials. It has also been used to generate nitride, carbonitride, and oxynitride coatings on Ti alloys (Sobiecki et al., 2001).

Ion implantation and deposition Ion implantation and deposition processes employ ions with high energy and speed to bombard the substrate surface. It is one of the most widely and successfully used surface modification techniques for Ti alloys. It can be classified into two broad categories, conventional beamline ion implantation and plasma-immersion ion implantation (PIII). The former is a line-of-sight technique and thus not suitable for objects with complex geometry. This restriction can be circumvented by PIII, where the negatively biased workpiece is immersed in a plasma sheath and ions are accelerated normally to its surface. Plasma-immersion ion implantation and deposition (PIIID) is the current pinnacle of ion implantation technology. It is a hybrid process that combines ion implantation and deposition and can generate not only a thick coating obtainable in PIII but also an atomically intermixed layer between the substrate and the coating that leads to high bonding strength (Sun et al., 2011).

Physical vapor deposition Physical vapor deposition (PVD) refers to a variety of vacuum deposition methods. Physical processes such as sputtering and evaporation are used in PVD to generate a vapor, in the form of atoms, molecules, or ions, of the coating material supplied from a target. They are then transported to and deposited on the substrate surface, resulting in coating formation. In PVD processes, the substrate temperature is substantially lower than the melting temperature of the target material, making it feasible to coat temperaturesensitive materials. Examples of commonly used PVD processes include evaporative deposition, ion plating, pulsed laser deposition, and sputter deposition. Compared with evaporating, sputtering is more suitable for target materials that are difficult to deposit by evaporation, such as ceramics and refractory metals (Thian et al., 2006). In addition, coatings prepared by sputtering usually have a better bonding strength to the substrate than those deposited by evaporation.

Chemical Methods for Preparing Coatings and Thin Films Surface modification of metal implants by most chemical methods involves chemical reactions at the interface between the implant and a gas/liquid phase. Coatings produced by these processes can be conversion coatings or deposited coatings.

Electrochemical methods Anodic oxidation, or anodization, is a most widely used electrochemical method for the surface modification of Ti- and Mg-based metallic implants. In this process, the workpiece is immersed in an electrolyte and acts as the anode of an electric circuit. The DC electric field drives the migration of oxygen and metal ions in the electrolyte. The process can be regarded as the facilitation and enhancement of the naturally occurring oxidation process on the metal surface. The oxide film formed is harder and more corrosion- and wear-resistant than the substrate and has high adhesive strength. As a special type of anodization, microarc oxidation (MAO) occurs when the anodizing voltage applied is higher than the breakdown voltage of the oxide film. Microdischarges, intense plasma, and high local temperature and pressure are generated on the sample surface in the electrolyte. The representative surface morphology of MAO-treated Ti is shown in Fig. 1B.

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Chemical vapor deposition While PVD employs physical forces to deposit the coating, CVD uses chemical processes. The volatile chemical precursors are injected into a chamber, where they react with or decompose on the heated substrate to form a nonvolatile compound. The byproducts and unreacted precursor gases are exhausted out of the chamber by gas flow. CVD usually produces films with higher uniformity than PVD and is particularly suitable for elaborately shaped objects. Thin films with high purity can be deposited by CVD as impurities can be removed with ease from gaseous precursors by distillation.

Sol–gel processes A sol is a colloidal suspension of solid particles in a continuous liquid, while a gel contains an integrated solid network. The conversion of small molecules into solid materials is realized in sol–gel processes through chemical reactions. The earliest commercial application of sol–gel technology was in the preparation of coatings and thin films because of its versatility and simplicity. Solgel processes have been widely used to deposit thin ceramic coatings such as calcium phosphates (CaP), silica, and titania. One of the drawbacks is coating shrinkage upon drying, which may lead to fracture due to large internal stress. Sol–gel-derived coatings are deposited coatings.

Biomimetic modifications Naturally occurring biochemical processes offer a wealth of inspiration for biomaterial design. Biomimetic surfaces are developed by imitating properties or processes found in physiological systems to induce specific cell and tissue response. Biomimetic surface modifications are typically realized by creating nature-inspired structural and topographical features and through the immobilization/anchoring of proteins, peptides, growth factors, or certain polymers (Chen et al., 2006a).

Properties of Coatings and Thin Films on Metal Implants Desirable Surface Functionalities of Metallic Implants Generally, the surface of metallic implants is expected to be highly corrosion-resistant and biocompatible. The corrosion of metallic implants causes the release of metal ions that may be necessary in trace amounts but can turn toxic at high concentrations. This relates to the essence of their biocompatibility. The biocompatibility of 17 elemental metals has been investigated in a recent study (Zhang et al., 2017). It should be noted that the lower pH value of inflamed tissues, like those frequently found at surgical sites, expedites the corrosion process. In addition, the pH value and oxygen concentration, which strongly influence the corrosion behavior, differ in different parts of our body. Wear resistance is another important consideration for many metallic implants. Wear debris generated at the bearing surfaces attracts macrophages and tends to kill them subsequently. It causes severe acidification of the microenvironment and facilitates the corrosion process. This in turn leads to the formation of more debris and promotes third-body wear. Besides corrosion, cyclic stress can also complicate the wear process as it results in faster materials fatigue than static fixed loading. It is critical for load-bearing implants to achieve efficient tissue integration. Most metallic biomaterials are intrinsically bioinert, and hence, they are surrounded by a FBC in vivo. The strength of this intervening soft-tissue bonding is far from sufficient. By modifying the implant surfaces with a bioactive coating, rapid fixation, fast integration, and strong bonding can be realized. The surface of many implants is also required to possess antibacterial properties because the surrounding tissues at the implant site are susceptible to bacterial infection due to the compromised host immune ability. The purpose is to retard or prevent biofilm formation at the implant/tissue interface in the race between host cells and bacteria to colonize the implant surface. Antibacterial coatings on metallic implants have been actively researched and clinically used. Ideal coatings should be able to tackle various bacterial species and at the same time not hamper the tissue integration ability. The surface of blood-contacting implantable devices has to be hemocompatible. For example, the surface of coronary artery stents should ideally induce no platelet activation and adhesion, thereby minimizing the occurrence of thrombosis or restenosis (renarrowing of the arteries). Both organic and inorganic coatings have been developed to prevent blood coagulation. While organic coatings allow more versatile chemical functionalization, inorganic ones usually possess higher stability and durability.

Osteoconductive and Osteogenic Coatings Osteoconductive and osteogenic coatings facilitate implant anchorage. They are especially important for cementless implantation that is widely used on young patients. Direct chemical bonds can be formed between bioactive materials and living human tissues. As a typical bioactive ceramic from the family of CaP, hydroxyapatite (HA) is the most popular coating material on metallic implants by virtue of its close chemical resemblance to the inorganic part of human hard tissue. After HA-coated devices are implanted, the partial dissolution of the coating causes the saturation of the surrounding solution and the precipitation of a carbonated HA layer. Collagen is then incorporated in this layer, and bone remodeling occurs in the areas of load transfer. Generally, ceramic materials are chemically stable and corrosion-resistant but brittle. The insufficient fracture toughness prevents bulk HA from being used in major load-bearing applications. Ti and its alloys are frequently used substrates for HA coatings, because they have a lower density, stronger bonding strength to the coating, and closer thermal expansion coefficient to HA compared with many other metals.

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Fig. 2 Histological images of HA-coated steel K-wire. (A) Physiological new bone formation (arrow) starting from the endosteal part of the cortex (A) (magnification, 10 ). (B) Detailed histology of the marked region reveals deposition of uncalcified bone matrix (arrowhead) by a layer of osteoblasts (arrows) that subsequently become osteocytes (double arrows) after being embedded by mineralized bone in rabbit (objective, 100). Reproduced from Fig. 6 of Alt, V., Bitschnau, A., Österling, J., Sewing, A., Meyer, C., Kraus, R., Meissner, S.A., Wenisch, S., Domann, E. and Schnettler, R. (2006). The effects of combined gentamicin–hydroxyapatite coating for cementless joint prostheses on the reduction of infection rates in a rabbit infection prophylaxis model, Biomaterials 27, 4627–4634.

Thermal spraying is one of the most commonly used methods to deposit HA coatings. Plasma-sprayed HA coatings have been clinically used in orthopedics and dentistry since the 1980s (Geesink, 1989). Typical applications include dental implants, femoral stems, bone screws and pins, and total knee arthroplasty. Most short- and medium-term clinical studies have shown faster and stronger fixation of HA-coated bone plates, screws, pins, intramedullary nails, dental implants, and femoral stems. Fig. 2 shows the histological images of HA-based coatings for a cementless metallic prosthesis implanted in the intramedullary canal of rabbit tibia. The deposition of uncalcified bone matrix and its mineralization could be clearly observed 28 days after implantation (Alt et al., 2006). The major concerns about plasma-sprayed HA coatings are their degradation, resorption, and debonding in physiological environments. Coatings with low crystallinity usually suffer from coating debonding due to their fast dissolution, although they can bring about faster implant fixation (Cheang and Khor, 1996). On the other hand, coatings with higher crystallinity and stability would usually comprise more unmelted or partially melted particles (see Fig. 1A). They thus have weaker coating/substrate interfaces and interparticle cohesive strength. Another important reason for controlling coating crystallinity is that the fast dissolution of coatings with low crystallinity and the weak interparticle bonding (for many coatings with high crystallinity) can both lead to the generation of disintegrated HA debris. This may accelerate third-body wear and result in osteolysis. In biomedical applications, the required degree of crystallinity is 65%–70% to ensure satisfactory long-term performance of HA coatings. Post-spraying heat treatment can be utilized to increase the coatings’ crystallinity (Yu et al., 2003). CaP represent the most widely used osteoconductive and osteogenic coatings. To achieve higher bonding strength to the substrate, other bioactive ceramics such as calcium silicate and BioglassÒ have also been plasma-sprayed (Liu et al., 2002). However, such coatings also face the challenge of rapid coating degradation. To address these issues, increasing research interest has been drawn to the fabrication of gradient coating (e.g., HA and tricalcium phosphate) and HA-containing composite coatings (Li et al., 2002). Other methods that have been employed to prepare osteoconductive and osteogenic coatings on metallic implants include RF sputtering (Thian et al., 2006), sol–gel spin coating or dip coating (Wang et al., 2007), chemical treatment (e.g., H2O2, acid, and alkali) (Wang et al., 2003), anodic oxidation (Yang et al., 2004), and biomimetic procedures (Habibovic et al., 2002). To further assist tissue ingrowth and integration, growth factors, bioactive molecules, and DNA can be incorporated in such coatings (Bose and Tarafder, 2012).

Corrosion-Resistant Coatings While CaP coatings on Ti alloys are primarily intended for increased bioactivity and have been clinically used, such ceramics are deposited on Mg and its alloys mainly to improve their corrosion resistance, and currently no Mg-based medical implants are commercially available. Their fast and localized corrosion is the major limitation to their clinical applications (even as biodegradable implants). Therefore, the research work on the surface modification of Mg and its alloys has been focused on realizing their delayed and controlled in vivo degradation. Conversion coatings have been prepared by a variety of processes on Mg and its alloys. The natural oxide layer on Mg-based materials is vulnerable and provides very limited protection. This layer can be thickened by using thermal, hydrothermal, or alkali treatment (Lorenz et al., 2009). CaP coatings are commonly prepared on Mg and its alloys by immersion in solutions containing calcium and phosphate ions such as simulated body fluid (SBF) (Gray-Munro and Strong, 2009). However, satisfactory results have yet to be obtained due to crack formation and poor coating adhesion. Similar immersion procedures can also be used to fabricate fluoride-containing thin films.

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MAO is frequently and commercially utilized to enhance the corrosion and wear resistance of Mg-based materials through the formation of a stable, hard ceramic coating. Higher corrosion resistance and hence improved cell adhesion have been observed for a MAO-treated Mg–Ca alloy compared with the uncoated samples (Gu et al., 2011). By tuning the electrolyte composition, it is possible to produce coatings containing CaP and fluoride by MAO. In addition, MAO is widely used as a pretreatment for the preparation of polymeric deposited coatings (Zhang and Kang, 2014). Such composite coatings are expected to possess increased corrosion resistance and enhanced adhesion strength. Most deposited coatings on Mg and Mg alloys are fabricated by wet chemical procedures as their high chemical reactivity makes many highenergy physical methods inappropriate. Frequently used processes for preparing deposited coatings include electrodeposition, dipping, immersion, and sol–gel methods. Coatings produced by such processes can be not only corrosion-resistant but also highly functional. They have thus been actively investigated for their potential in drug delivery, immobilization of biomolecules (such as growth factors, enzymes, and peptide sequences), and enhanced osteointegration (Di Mario et al., 2004). Nevertheless, the primary focus is to obtain long-term uniform degradation at an appropriate rate. Diamond-like carbon, which was developed in the 1970s, has been successfully used to improve the corrosion and wear resistance, biocompatibility, and hemocompatibility of metallic implants. DLC is a hydrogenated amorphous carbon and contains various quantities of sp2 and sp3 carbon bonds. It possesses bioinertness, low friction coefficient, and higher hardness and stiffness than most ceramics. A variety of methods have been developed to deposit DLC coatings, such as CVD, ion beam deposition, PIIID, magnetron sputtering, pulsed laser ablation, and filtered cathodic arc deposition (Dearnaley and Arps, 2005). Fig. 1C presents the typical surface topography of a DLC coating prepared by CVD. Although DLC coating is able to effectively retard the corrosion of Mg and its alloys, the nondegradability and durability make it not suitable for biodegradable Mg-based implants. Nevertheless, DLC has been widely deposited on metallic implants designed for long-term applications and led to commendable outcomes. For instance, experimental results have shown that with a DLC coating, the corrosion rates of stainless steel, Ti6Al4V and CoCrMo in 10 wt% HCl solution, could be decreased by a factor of 10,000–15,000 (Lappalainen et al., 1998). While the corrosion behavior of metals can be effectively improved through surface modification, the intrinsic corrosion resistance of the bulk material plays a fundament role. For example, the purification of Mg and incorporation of suitable alloying elements can both substantially enhance its corrosion resistance.

Wear-Resistant Coatings The excessive wear of acetabulum is a major challenge facing total hip and knee replacements, where Ti alloys, CoCrMo alloys, and stainless steels are widely used. It leads to joint loosening and the generation of wear debris that can cause osteolysis. The wear resistance of Mg and its alloys also needs to be enhanced to use them in load-bearing medical implants. Such problems can be alleviated by depositing a wear-resistant coating with high hardness and a low friction coefficient. Oxide ceramics such as Al2O3, ZrO2, and TiO2 have excellent corrosion and wear resistance. They have been thermally sprayed to enhance the tribological properties of Ti6Al4V and stainless steel (Cizek et al., 2013; Chen et al., 2002). For the same purpose, thermal spraying has also been used to prepare metal-matrix or ceramic-matrix composite coatings (Khun et al., 2015). Another very successful approach is the use of ion implantation. For coatings and thin films prepared by ion implantation, surface delamination is seldom an issue because of the graded composition and obscure coating/substrate interface. Ions commonly implanted into metallic implant materials include oxygen, nitrogen, carbon, calcium, and sodium ions. Among them, nitrogen ions are frequently implanted to improve the wear, corrosion, and fatigue resistance of Ti and Ti alloys. It was found that the (Ti, O, N)/Ti composite coating prepared by PIIID could augment not only the mechanical properties and wear resistance but also the biocompatibility of NiTi shape-memory alloy (Sun et al., 2012). Similarly, nitrided and carbonitrided surface layers on Ti alpha alloy fabricated by glow discharge plasma showed enhanced wear and corrosion resistance and increased fatigue strength (Sobiecki et al., 2001). As mentioned earlier, DLC holds great promise as a wear-resistant coating material. Compared with pyrolytic carbon, DLC can be produced at lower temperatures and is thus more suitable for heat-sensitive substrates such as acetabular cups lined with ultrahigh-molecular-weight polyethylene (UHMWPE) or polytetrafluoroethylene. It has been found that the frictional behavior and wear resistance of a UHMWPE/CoCrMo sliding pair in SBF could be significantly improved by coating both sliding surfaces with DLC (Sheeja et al., 2005). It is worth mentioning that the surface roughness of the DLC coating strongly influences the wear rate of the counterface, especially in orthopedic applications with a sliding surface made from UHMWPE. It was pointed out earlier that the poor wear resistance is an obvious drawback for Mg and Mg alloys used as implant materials. Their tribological properties can also be enhanced by MAO. It was found in sliding wear tests that the mass loss of AZ91D Mg alloy in Hanks’ solution could be reduced by one-third after MAO treatment (Zhang et al., 2007).

Hemocompatible Coatings Besides biocompatibility, hemocompatibility is another crucial consideration for biomaterials used in blood-contacting implants. A wide range of surface modification strategies have been employed to achieve hemocompatible surfaces. Certain organic coatings can render device surfaces hemocompatible through biochemical reactions at a molecular level. However, they are inferior to inorganic coatings in terms of stability and durability and therefore have been much less used.

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DLC is a well-known hemocompatible coating material, which should be attributed to their bioinertness and smoothness (smooth DLC surface shown in Fig. 1C). Conventionally, artificial heart valves are made from low-temperature isotropic (LTI) carbon that risks fracture-induced failure. DLC shows comparably good hemocompatibility to LTI carbon, and the replacement of LTI carbon with DLC in the heart valves has been commercially successful. DLC has also been coated onto metallic cardiovascular stents. The results from many in vitro and in vivo tests proved that DLC could obviously suppress platelet adhesion, activation, and aggregation and decrease their area coverage on cardiovascular implants. Furthermore, the reduced metal ion release as a result of the DLC coating may further contribute to the device’s biocompatibility and hemocompatibility (Gutensohn et al., 2000). When DLC is deposited on needles and medical guidewires, it brings about an additional advantage as the low friction coefficient helps reduce the friction force. Another group of inorganic hemocompatible materials are ceramics such as titanium oxides and titanium nitrides. Implantation of oxygen ions has been used to enhance the hemocompatibility of Ti by thickening the surface oxide layer. Ti–O thin films manufactured by PIII exhibited superior thromboresistant properties than LTI carbon in both in vitro and in vivo experiments (Yang et al., 2002). The coimplantation of nitrogen and oxygen ions was carried out by PIIID, and the incorporation of nitrogen in the coating was found to be able to further improve its blood compatibility (Tsyganov et al., 2005). While the above techniques emphasize the passivation of a surface, there has been relentless research work targeting active inhibition of platelets. This includes the use of drug-eluting coatings and the growth of endothelial cells on implant surfaces. Endothelial cells form the interior lining of blood vessels, and their anticoagulant mechanisms are attracting increasing research interest (Werner et al., 2007).

Antibacterial Coatings Antibacterial coatings are deposited on device surfaces to mitigate implant-associated infection. They can be either passive or active, depending on whether antibacterial agents are locally delivered. Passive coatings can hinder the bacterial attachment and/or kill bacteria upon contact. The physical and chemical characteristics of the coating, such as surface roughness, wettability, and conductivity, have strong influences on the bacterial behavior (Li et al., 2016). By creating a crystalline anatase-enriched surface on Ti through anodization and heat treatment, the attachment of three bacterial strains could be markedly reduced, while the soft- and hard-tissue cellular response was not negatively influenced (Del Curto et al., 2005). Ti surfaces coated with certain polymers, such as poly(ethylene glycol) and poly(methacrylic acid), have also shown the ability to inhibit bacterial adhesion (Harris et al., 2004; Zhang et al., 2008). Passive coatings are preferable as they induce no bacterial resistance development, but their antibacterial efficacy is limited and varies for different bacterial strains. Active coatings release antibacterial agents, such as antibiotics, bioactive molecules, and inorganic antimicrobial agents, to the surrounding tissues. A number of antibiotics have been incorporated in bioceramic or biodegradable polymer coatings, such as gentamicin, amoxicillin, carbenicillin, cephalothin, cefamandole, vancomycin, and tobramycin (Stigter et al., 2004). Although they possess a broad antibacterial spectrum, it is difficult to achieve optimum release kinetics with minimum harmful effects on cellular functions and tissue integration. Besides, potential drug resistance of bacteria is another problem. The development of pathogen resistance is less of an issue for coatings containing nonantibiotic organic antibacterial agents (Tan et al., 2014). These include chloroxylenol, chlorhexidine, poly(hexamethylene biguanide), chitosan, and hyaluronic acid (Campbell et al., 2000). Again, it is challenging to realize controlled and sustained release of such molecules. Convincing in vivo evidence has yet to be obtained to unambiguously prove their cytocompatibility and efficacy. In addition, since antibiotics and organic antibacterial agents are both heat-sensitive, they cannot be loaded in the coatings through high-energy process routes such as thermal spraying. Active coatings can also be loaded with inorganic antimicrobial agents, among which Ag is the most commonly used. Silver possesses a broad antibacterial spectrum and high biocompatibility. It is not prone to antibiotic resistance and can be introduced by well-established methods such as PIIID and thermal spraying (Sanpo et al., 2009a). Compared with pure HA coating on Ti substrate, Ag-doped HA coating showed remarkably enhanced antibacterial ability without compromising the coating’s cytocompatibility (Chen et al., 2006b). Nevertheless, further research is needed to understand the possible long-term tissue toxicity and its exact antibacterial mechanisms. Other inorganic antibacterial agents include Cu, F, Ca, N, and Zn (Sanpo et al., 2009b). Like Ag, they can be introduced to a wide variety of biomaterials including ceramics, polymers, metals, and DLC. To realize widespread clinical applications of antibacterial coatings requires more information on their in vivo performance. Since host immune ability plays a paramount role in infection prevention, future studies may be directed to antimicrobial coatings with the capability to facilitate tissue integration and enhance host defense ability.

Concluding Remarks This section broadly presented the preparation techniques and properties of coatings and thin films on metallic implants. While the different coating characteristics were discussed separately, it is of vital importance to take into account the whole range of properties and functionalities holistically and comprehensively in coating design and manufacture. The clinical application of a coating technology for metallic implants can only become realistic when it possesses high reproducibility and satisfies the stringent

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requirements in the aforementioned interrelated aspects. This calls for effective communication between researchers, clinicians, and manufacturers. Furthermore, coatings on metallic implants not only present a new physiochemical niche but also alter the topography and mechanical properties. The latter two factors also need to be scrutinized as they can elicit significant influences on the surrounding tissues. To encourage natural interactions between the coating surface and host cells, one promising method is to mimic the exquisite topographic features of extracellular matrix. The highest level of biocompatibility is realized when the body cannot distinguish the implant material from its own. The concept of biomimetic fabrication of medical implants has thus been embraced by many researchers. In biological evaluations of deposited coatings and thin films, it is crucial to keep in mind that the in vitro conditions are usually far less aggressive and complicated, both mechanically and biochemically, than the in vivo environment. This warrants necessary in vivo tests to bridge the knowledge gaps in this aspect. While researchers have been persistently seeking to extend the functionality of coatings and thin films, it is critical to keep the complex coating-induced immune responses under control. On top of other desirable properties, an ideal coating should be able to enhance the host immune ability and promote the body’s self-healing. Finally, an increasing fraction of patients are outliving the service longevity of their implants. Therefore, extending metallic implants’ service life is a critical consideration for their coating design. Alternatively, biodegradable metallic implants, which may be produced from Mg alloys, can be developed to allow their complete adsorption and replacement by newly grown tissues.

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Titanium Alloys Mitsuo Niinomi, Tohoku University, Sendai, Japan; Osaka University, Osaka, Japan; Meijo University, Nagoya, Japan; and Nagoya University, Nagoya, Japan © 2019 Elsevier Inc. All rights reserved.

Introduction Nontoxic Ti Alloys Low Modulus Ti Alloys Low Young’s Modulus b-Type Ti Alloys With Changeable Young’s Modulus and for Reconstructive Implants Improvement of Mechanical Properties of Low Modulus b-Type Titanium Alloy for Biomedical Applications Improvement of Static Strength by Severe Cold Working Improvement of Dynamic Strength by Thermomechanical Treatment Improvement of Dynamic Strength Using O as Alloying Element Effect of Low Young’s Modulus on Stress Shielding Summary References Further Reading

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Introduction Biomaterials are grouped into biotolerant, bioinert, and bioactive materials from the pattern of osteogenesis as listed in Table 1 (Niwa, 2004). With regards to biocompatibility, biotolerant materials are superior to bioinert materials, which are in turn less than bioactive materials. Some representative metallic biomaterials are austenitic stainless steel (mainly SUS 316L), Vitallium (Co–Cr–Mo alloys), and titanium (Ti) and its alloys. As shown in Table 1, Ti and its alloys are classified as bioinert materials. On the other hand, SUS 316L and Co–Cr–Mo alloys are grouped into the biotolerant category. Therefore, the biocompatibility of Ti and its alloys is better than that of SUS 316L and Co–Cr–Mo alloys. Ti is regarded as a nontoxic and nonallergenic element based on cell viability characterization (Steinemann, 1980; Kawahara et al., 1963; Yamamoto et al., 1988) and rate of metallic allergy of pure metals (Uggowitzer et al., 1997; Niinomi, 2000). Ti and its alloys also present fewer imaging artifacts, which results in clear MRI (magnetic resonance imaging) images (Radzi et al., 2014). Furthermore, Ti alloys exhibit high corrosion resistance, high specific strength meaning a light weight and an high strength, and excellent balance of strength and ductility. For these reasons, Ti alloys are receiving significant attention as biomaterials for constructing implants that replace failed hard tissues as cortical bone (Fig. 1; Nakano, 2010). In the present article, from the view point of biomedical applications, review of the latest research on nontoxic Ti alloys, low Young’s modulus b-type Ti alloys, and low Young’s modulus b-type Ti alloys with changeable Young’s modulus and for reconstructive implants is presented. Further, improvements in the mechanical properties of low Young’s modulus b-type Ti alloys for biomedical applications and the effect of low Young’s modulus on stress shielding are described.

Nontoxic Ti Alloys The earliest used Ti-biomaterial was pure Ti. Lately, an alloy of Ti has also come into use: Ti–6Al–4V ELI (extra low interstitial) (mass %). There are four grades for pure Ti: grades 1 through 4, where the concentration of impurities such as nitrogen (N), iron (Fe), and oxygen (O) increases with increasing grade number. The strength of pure Ti increases with impurity concentration, while the ductility decreases. Pure Ti and Ti–6Al–4V ELI have been standardized as ASTM F-67-13 (ASTM International, 2016a) and F13613 (ASTM International, 2016b). Additionally, Ti–6Al–4V has been standardized as ASTM F1108-14 (ASTM International,

Table 1

Biocompatibility of various biomaterials judged by patterns of osteogenesis

Pattern of osteogenesis Intervend osteogenesis Contact osteogenesis Bonding osteogenesis

Encyclopedia of Biomedical Engineering, Volume 1

Biomaterials Stainless steel, Vitallium, PMMA (polymethyl methacrylate) Titanium, titanium alloys, carbon, alumina, zirconia, titania, TiN, Si3N4 Bioglass, ceravital, tricalcium phosphate, hydroxyapatite, A-W glass ceramic

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Biotolerant materials Bioinert materials Bioactive materials

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Artificial dental implants

Artificial shoulder joints Spinal fixation devices

Artificial elbow joints Artificial hip joints

Artificial finger joints

Bone fixation devices

Artificial knee joints Artificial ankle joints

Fig. 1

Various kinds of implants constructed with metallic biomaterials: any of them are made of titanium alloys.

2016c) for castings in surgical implants and as ASTM F1472-14 (ASTM International, 2016d) as a wrought alloy for the same application. Pure Ti and Ti–6L–4V ELI continue to be the primary Ti-based materials for constructing implants, although they have been diverted from general structural usage. However, since the discovery of the toxicity of vanadium (V) which is a constitutional element of Ti–6Al–4V ELI, V-free Ti alloys for biomedical applications such as Ti–5Al–2.5Fe (ISO, 1996), Ti–6Al–7Nb (ASTM International, 2016e), Ti–6Al–6Nb–1Ta (Niinomi, 2002), and Ti–6Al–2Nb–1Ta (JISC, 2002a) have been developed. Since then, Al, a constitutional element of Ti–6L–4V ELI, has also been pointed out as a harmful element for the human body, V- and Al-free Ti alloys such as Ti–15Sn–4Nb–2Ta–0.2Pd (Okazaki et al., 1996), Ti–15Zr–4Nb–2Ta–0.2Pd (JIS T 7401-5) (JISC, 2002b), and Ti–4.5Al–6Nb–2Fe–2Mo (Ueda et al., 2013). The latter has been modified super plasticity Ti–4.5Al–3V–2Fe–2Mo (SP-700) (Suzuki et al., 1999) for biomedical applications. Ti alloys for biomedical applications mentioned earlier are all (a þ b)-type Ti alloys except for pure Ti, which is an a-type Ti alloy.

Low Modulus Ti Alloys The development of Ti alloys for biomedical applications has been not only by the need for nontoxic elements but also nonallergenic elements, but for inhibiting stress shielding, which is the inhomogeneous stress transfer between the host bone and the implant that leads to bone resorption and poor bone remodeling. The high-risk elements for allergy are Ni, Cr, and Co. Based on the rate of allergy of pure metals reported in the field of dentistry, Ni is the most restricted allergenic element in metallic biomaterials. The nontoxic components of Ti alloys for biomedical applications can be identified using cell viability testing results. From this data, Nb, Ta, Mo, Zr, etc. have been found to be appropriate nontoxic and nonallergenic elements for alloying with Ti, judging from cell viability and rate of allergy of pure metals for designing Ti alloys for biomedical applications. Ti alloys are roughly grouped into a-, (a þ b)-, and b-type Ti alloys based on the constituent phases. The Young’s moduli of b-type Ti alloys, where the b-phase with the body-centered cubic (bcc) structure is predominant, are expected to be lower than those of a-type Ti alloys where the a-phase with the hexagonal close packed (hcp) structure is predominant, and (a þ b)-type Ti alloys where a- and b-phases are mixed because the atomic density of the bcc structure is less than that of the hcp structure. Typically, Ti alloys designed to be b-type Ti alloys, as b-type Ti alloys are usually composed of nontoxic and nonallergenic b stabilizing elements (Nb, Ta, Mo, etc.) and exhibit Young’s moduli close to that of the cortical bone (approximately 10–30 GPa (Niinomi and Boehler, 2015)). It is noted that the Young’s moduli of representative metallic biomaterials such as Ti–6Al–4V ELI, SUS 316L, and Co–Cr– Mo alloy are approximately 110, 180, and 210 GPa, respectively (Niinomi et al., 2012). A number of b-type titanium alloys with low Young’s modulus for biomedical applications have been developed thus far. Typically, these alloys contain a large amount of b stabilizing elements such as Nb, Mo, and/or Ta, and a small amount of other elements such as Zr, Sn, Hf, Si, and/or Fe (Niinomi et al., 2012). Fig. 2 (Niinomi et al., 2012) shows the Young’s moduli of various low Young’s modulus b-type titanium alloys developed for biomedical applications. The data shown includes samples subjected to various treatments and moduli measured by various

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Alloy, treatment, measuring method Ti-6Al-4V ELI (Mill anealed, Unclear) Ti-6Al-4V (Annealed, Unclear) Ti-6Al-4V (Annealed,Ultrasonic method) Ti-6Al-4V (Annealed, Nanoindenter) Ti-20Nb (WQ: water quenching, NQ: liquid nitrogen cooling, AC: air cooling, FC: furnace cooling, Ultrasonic method) Ti-13Nb-13Zr (ST, STA, Ultrasonic method) Ti-13Nb-13Zr (ST, STA, AC, WQ+50-75%CW:cold working, Three point bending test) Ti-27Nb-8Zr (ST: solution treatment, STA: aged after ST, Tensile test) Ti-29Nb-13Ta (ST, STA, Tensile test) Ti-22.5Nb-0.7Zr-2Ta-(0.5-2.5)O (ST, Tensile test) Single cryatal of Ti-29Nb-13Ta-4.6Zr ((100), ST+ zone melting, Electromagnetic acoustic resonance) Single cryatal of Ti-29Nb-13Ta-4.6Zr ((111)ST+ zone melting, Electromagnetic acoustic resonance) Ti-23Nb-0.7Ta-2.0Zr-1.2O (ST, Dynamic mechanical analysis) Ti-29Nb-13Ta-4.6Zr (ST, STA, CR: cold rolling, Resonance vibration method) Ti-29Nb-13Ta-4.6Zr (0.3 mm dia. Wire, Cold drawing, Dynamic super micro hardness tester) Ti-29Nb-13Ta-4.6Zr (1.0 mm dia. Wire, Cold drawing, Dynamic super micro hardness tester) Ti-29Nb-13Ta-4.6Zr (0.1-0.7)O (ST, Resonance vibration method) Ti-29Nb-13Ta-4.6Zr-(0-0.5)B (ST+ cold rolling, Resonance vibration method) Ti-29Nb-13Ta-4.6Zr-(0.05-1.0))Y2O3 (ST + cold rolling, Resonance vibration method) Ti-35Nb-(0-10)Ta-(0-10)Zr (ST, STA, Tensile test) Ti-35Nb-2Ta-3Zr (in TD direction, ST,Tensile test) Ti-35Nb-2Ta-3Zr (in 45 degree, ST,Tensile test) Ti-35Nb-2Ta-3Zr (in RD direction, ST,Tensile test) Ti-35.5Nb-7Zr-5Ta (TNZT) (ST,Transmission ultrasonic tequnique) Ti-(22-35.5)Nb-(5-22)Ta-(4-7.2)Zr (ST, Transmission ultrasonic tequnique) Single crystal Ti-36Nb-2Ta-3Zr-0.09O (ST + zone melting, Electromagnetic acoustic resonance) Single crystal Ti-36Nb-2Ta-3Zr-0.36O (ST + zone melting, Electromagnetic acoustic resonance) Single crystal Ti-36Nb-2Ta-3Zr-00.51O (ST + zone melting, Electromagnetic acoustic resonance) Ti-35Nb-4Sn ( ST to sever cold rolling, Tensile test) Ti-16Nb-13Ta-4Mo (ST, STA, Tensile test) Ti-29Nb-13Ta-4Mo (ST, STA, Tensile test) Ti-29Nb-13Ta-2Sn (ST, STA, Tensile test) Ti-29Nb-13Ta-4.6Sn (ST, STA, Tensile test) Ti-29Nb-13Ta-6Sn (ST, STA, Tensile test) Ti-28Nb-13Zr-0.5Fe(TNZF) (ST, STA, Tensile test) Ti-24Nb-4Zr-7.9Sn (Ti2448) (ST + cold rolled, Hot forged,Cold rolled, Annealed, Tensile test Resonance vibration method) Ti-7.5Mo (ST, Unclear) Ti-15Mo (ST, Tensile test) Ti-(15-18)Mo (ST, Resonance vibration method) Ti-12Mo-3Nb (ST, Nanoindenter) Ti-12Mo-5Ta (ST, Ultrasonic method) Ti-15Mo-5Zr-3Al (ST, STA, Ultrasonic method) Ti-15Mo-2.8Nb-0.2Si (21RX) (ST, Tensile and compression) Ti-16Nb-9.5Hf (Tiadyne) (STA, Unclear) Ti-12Mo-6Zr-2Fe (TMZF) (Solution annealed, Tensile test) Ti-(10-70)Ta (ST, Resonance vibtation method) Ti-50Ta (ST, STA, Resonance vibration method) Ti-30Zr-(1-3)Cr-(3-5)Mo (ST, Resonance vibration method) Ti-(10-14)Cr (ST, Resonance vibration method)

0 Fig. 2

100 Young’s modulus (GPa)

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Young’s moduli of representative b-type titanium alloys for biomedical applications with treatment and measuring method.

methods. The Young’s moduli of low Young’s modulus b-type titanium alloys that are subjected to solution treatment (ST) lie between 40 and 80 GPa. These alloys show a single b-phase. However, in the specific case, where the entire sample is a single crystal growth directions, the Young’s modulus is found to be as that of cortical bone. For example, a single crystal of the low Young’s modulus b-type titanium alloy Ti–29Nb–13Ta–4.6Zr (TNTZ), growing in the h100i direction exhibits a lower Young’s modulus (35 GPa) when compared with crystal growth in directions such as h111i and h110i (Tane et al., 2008). It is noted that tensile tests appear to measure slightly lower values of Young’s modulus than the free-resonance or ultrasonic methods, while the dynamic super microhardness test measures slightly higher values. The Young’s modulus of these alloys may be increased to 100 GPa and over using processing techniques such as by applying an aging treatment. Due to the high cost of elements such as Nb, Ta, Mo, and Zr, biomedical b-type titanium alloys with low Young’s moduli composed of low cost elements such as Fe, Cr, Mn, Sn, and Al have also been developed. Some examples of low Young’s moduli alloys with low cost elements are Ti–10Cr–Al (Hatanaka et al., 2010), Ti–Mn (Ikeda et al., 2009), Ti–Mn–Fe (Ikeda et al., 2012), Ti– Mn–Al (Ikeda et al., 2010), Ti–Cr–Al (Ikeda and Sugano, 2005), Ti–Sn–Cr (Ashida et al., 2012), Ti–Cr–Sn–Zr (Murayama and Sasaki, 2009), Ti–(Cr, Mn)–Sn (Kasano et al., 2010), and Ti–12Cr (Nakai et al., 2011).

Low Young’s Modulus b-Type Ti Alloys With Changeable Young’s Modulus and for Reconstructive Implants Low Young’s modulus Ti alloys are considered to be suitable for rods used in spinal fixation devices. The flexibility of these alloys is patient-friendly feature as it is expected to inhibit the damage to adjacent vertebra. However, during operations, the rods in spinal fixation devices are bent by surgeons to match the physiological curvature of the patient’s spine as possible as closely as possible (Steib et al., 2004). This necessitates that the return back of bent rod to its original shape, that is, springback, should be prevented. In summary, the materials used in spinal fixation rods must have low Young’s modulus measured across the entire rod and minimal springback. To satisfy both requirements, an increased Young’s modulus is required at the bent section of the rod. This demand can be satisfied if the secondary phase with high Young’s modulus can be induced by bending, that is, deformation. It is known that the deformation-induced a0 martensite and a00 martensite phases can decrease the Young’s modulus of Ti alloys, while the u-phase is known to significantly influence the mechanical properties of b-type Ti alloys, and likely to increase the Young’s modulus. Using this rationale, Ti–12Cr was developed (Nakai et al., 2011). Fig. 3 shows the change in Young’s moduli of b-type, alloys A, B, and C, and Ti–12Cr subjected to solution treatment before and after deformation by cold rolling to a 10% reduction (CR), which simulates

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α”

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TEM DF of Ti-12Cr -12Cr after deformation

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Fig. 3 Change in Young’s moduli of b-type alloy A, B, C and Ti–12Cr subjected to solution treatment (ST) indicating before deformation and cold rolling by a 10% reduction (CR) indicating after deformation.

Ratio of springback per unit load, R/(%N–1)

bending deformation of the spinal fixation rods. The deformation was confirmed using TEM diffraction patterns shown in this figure. For b-type alloy A, no change in Young’s modulus is recognized as no deformation-induced secondary phase is formed. For b-type alloys B and C, the Young’s moduli decrease after deformation due to the formation of deformation-induced a00 martensite and a0 martensite, respectively. For Ti–12Cr, the Young’s modulus increases after deformation due to the formation of the deformation-induced u-phase. Fig. 4 (Zhao et al., 2012a) shows springback of Ti–12Cr and TNTZ. It is noted that in TNTZ, no deformation-induced phase is formed. The springback of Ti–12Cr is much smaller than that of TNTZ. The enhancement in the Young’s modulus of the Ti–Cr-based alloys can be further increased by adding a small amount of oxygen (O) to prevent the formation of the athermal u-phase and by controlling the Cr content. As a result, Ti–10Cr–0.2 and Ti–9Cr–0.2 have been developed as Young’s modulus changeable b-type alloys for the rods in spinal fixation devices (Liu et al., 2015, 2016). Ti–17Mo (Zhao et al., 2012b), Ti–30Zr–7Mo (Zhao et al., 2011a), Ti–30Zr–5Cr (Zhao et al., 2011b), and Ti–30Zr–3Cr–3Mo (Zhao et al., 2011b) were also developed as Young’s modulus changeable b-type Ti alloys with low Young’s modulus. Among those, Ti–30Zr–7Mo, Ti–30Zr–5Cr, and Ti–30Zr–3Cr3Mo were developed for use in reconstructive implants. Reconstructive implants, which are removable implants constructed using Ti alloys, should (a) not tightly fuse with bone and (b) have good biocompatibility. This is because, in some cases, certain internal fixation devices implanted into the bone marrow (femoral, tibial, or humeral marrow) or those, which would otherwise be incorporated into the bone, need to be removed after surgery. These include the screws used for bone plates fixation and implants used for children, which might have to be removed to address local symptoms such as palpable hardware and wound dehiscence/exposure of hardware, or to enable athletes returning to contact sport (Kobayashi et al., 2007). In these cases, the assimilation of the removable internal fixators with the bone (because of the precipitation of calcium

Fig. 4

0.06 0.05 0.04 TiNTZ 0.03 0.02 0.01 0

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Spring-backs of Ti–12Cr and Ti–29Nb–13Ta–4.6ZrTNTZ.

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phosphate) might cause refracture of the bone during the fixator-removal operation. The Ti alloys used in these fixators contain a large amount of Zr because Zr has been reported to show the ability to prevent the precipitation of calcium phosphate (Kobayashi et al., 2007; Tsutsumi et al., 2009). Further, some titanium alloys with Zr contents exceeding 25 mass% have also been reported to prevent the formation of calcium phosphate, which is the main component of human bones as mentioned earlier (Tsutsumi et al., 2009).

Improvement of Mechanical Properties of Low Modulus b-Type Titanium Alloy for Biomedical Applications Fig. 5 shows the elastic modulus (Young’s modulus), tensile strength, and elongation of selected biomedical b-type titanium alloys with a single b-phase. The Young’s modulus and elongation tend to increase and decrease with increasing tensile strength. In b-type titanium alloys with a single b-phase and with relatively lower Young’s moduli such as Ti–35.3Nb–7.1Zr–5.1Ta and TNTZ, the strengths are found to be poor. For these kinds of alloys, an improvement in strength, while maintaining a low Young’s modulus and a good ductility, is required.

Improvement of Static Strength by Severe Cold Working Static strength such as tensile strength can be improved significantly by severe cold working techniques such as cold rolling (Niinomi, 2010), cold swaging (Niinomi et al., 2002a), and cold drawing (Niinomi et al., 2007), as well as by severe plastic deformation such as high-pressure torsion (Yilmazer et al., 2013). By these methods, the tensile strength of TNTZ is increased to values equal to or over that of Ti–6Al–4V ELI, while maintaining good ductility and a Young’s modulus equal to or lower than that of TNTZ subjected only to solution treatment (approximately 60 GPa).

Improvement of Dynamic Strength by Thermomechanical Treatment While static strength can be improved, it is difficult to increase dynamic strength such as fatigue strength by the earlier-mentioned severe cold working process. To improve the dynamic strength, introducing a secondary phase is required. A short aging treatment to introduce a small amount of u-phase in TNTZ, which significantly increases the strength and Young’s modulus of a b matrix is effective in also increasing fatigue strength to values similar to those of Ti–6Al–4V ELI, with good ductility and the Young’s modulus at the level of approximately 75 GPa (Nakai et al., 2012). Adding a small amount of ceramics such as TiB or Y2O3 is also effective in improving the fatigue strength to values similar to those of Ti–6Al–4V ELI, while maintaining the Young’s modulus as same as that of TNTZ subjected only to ST (approximately 60 GPa (Song et al., 2010)). Further improvement in fatigue strength can also be achieved by special thermomechanical treatment such as a direct aging treatment after severe cold working, However, by this method, the Young’s modulus becomes relatively high (approximately order of 80 GPa, which is approximately the average Young’s moduli of low Young’s modulus b-type Ti alloy with single b-phase (Narita et al., 2012)).

Fig. 5

Elastic modulus (Young’s modulus), tensile strength, and elongation of selected b-type titanium alloys for biomedical applications.

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Fig. 6 (Narita et al., 2012) shows the special thermomechanical processes (CRA (cold rolling and aging) and CWA (cold swaging and aging)) and general thermomechanical process (STA (solution treatment and aging)) applied to TNTZ. Fig. 7 (Narita et al., 2012) shows the fatigue strengths (S–N curves) of TNTZ subjected to CRA and CWA, and STA along with those of Ti–6Al–4V ELI (Ti-64). The fatigue strengths of TNTZ subjected to CRA and SWA are much greater than those of TNTZ subjected to STA and Ti-64. The fatigue strength of TNTZ subjected to SWA is a little greater than that of TNTZ subjected to CRA. Evidently, the fatigue strength can be improved significantly by special thermomechanical treatments such as SWA and CRA.

Improvement of Dynamic Strength Using O as Alloying Element O is in general regarded as an impurity in Ti and its alloys, but currently O is being recognized as an effective alloying element to improve mechanical properties of titanium alloys. Fig. 8 (Liu et al., 2017a) shows the room temperature true stress–strain curves and the corresponding work-hardening rates of TNTZ containing 0.14, 0.33, and 0.70 mass% O subjected to solution treatment (ST), which forms a single b-phase. These alloys are (A)

Solution treatment : 1063K, 3.6ks

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Aging treatments: 723K, 259.2 ks

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SWA Cold swaging (Area reduction ratio: 91%)

Bar βtr : β transus WQ :Water quenching Fig. 6

Aging treaments: 723K, 259.2 ks

WQ

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Special thermomechanical processes (CRA and CWA) and general thermomechanical process (STA) for TNTZ.

Maximum cyclic stress, σmax/ MPa

1200 1000 800 600 400 200 0 104

STA CRA SWA Ti-64 105 106 Number of cycles to failure, Nf

107

Fig. 7 Fatigue strength (S–N curves) of TNTZ subjected to special thermomechanical processes (CRA and CWA) and general thermomechanical process (STA) for TNTZ shown in Fig. 7.

Work-hardening rate, dσ/dε (MPa) True stress, σ (MPa)

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True strain, ε Fig. 8

Room temperature true stress–strain curves and corresponding work-hardening rates of 0.1ST, 0.3ST, and 0.7ST.

referred to as TNTZ-0.14ST, TNTZ-0.33ST, and TNTZ-0.70-ST, respectively. The tensile strength and 0.2% proof stress of TNTZ increases with increasing O content. On the other hand, however, the elongation decreases at an O content of 0.33 mass%, and then increases at an O content of 0.70 mass%. This trend with O content is contradictory to those conventionally reported. Their tensile strength and elongation of TNTZ-0.70ST can reach to approximately 1100 MPa and 20%, respectively. Both the tensile strength and elongation of TNTZ-0.70ST are greater than those of Ti–6Al–4V ELI. An obvious yielding point can be recognized in TNTZ-0.7ST, but such yielding cannot be observed in TNTZ-0.14ST and TNTZ-0.33ST. Double yielding, which is observed in the case of TNTZ-0.14ST, is not observed in TNTZ-0.7ST. The work-hardening rate curve obtained for TNTZ-0.7ST, which shows the highest work-hardening rate among the examined alloys, intersects with the true stress–strain curve at true strains higher than 20%. This suggests that this alloy exhibits the greatest work-hardening effect among all the alloys, and undergoes late necking under tension. The excellent strength and elongation of TNTZ-0.70ST have been attributed to the occurrence of slips with wavy or polyline morphologies corresponding to cross slips with multiple directions in the same grain (Liu et al., 2017a). The Young’s moduli of TNTZ-0.14, 0.33, and 0.70ST are shown in Fig. 9 (Liu et al., 2017a) along with that of Ti–6Al–4V ELI. The Young’s moduli of TNTZ-0.14ST and TNTZ-0.33ST are lower than 65 GPa. The Young’s modulus of TNTZ-0.70ST is lower than  75 GPa, which is much lower than that of Ti–6Al–4V ELI, which shows a Young’s modulus of approximately 100–110 GPa. The fatigue strength of TNTZ can be improved by adding a large amount of O as shown in Fig. 10 (Liu et al., 2017b). In this case, deformation-induced martensite is formed. The width of the deformation-induced martensite decreases with increasing O content

120 110 Young's modulus, E/GPa

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Young’s moduli of TNTZ-0.14O, TNTZ-0.33O, and TNTZ-0.70O subjected to hot rolling followed by solution treatment respectively.

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Maximum cyclic stress,σmax/MPa

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S–N curves of prepared 0.1O, 0.3O, 0.5O, and 0.7O alloys with fatigue limit range of Ti–6Al–4V ELI.

as shown in Fig. 11. Therefore, the increment in the fatigue strength of TNTZ with O content can be considered to be due to O solid solution strengthening and microstructural refinement by deformation-induced martensite (Liu et al., 2017b). It has been reported that a small amount of O addition also increases both the strength and elongation in (a þ b)-type Ti–4Cr as shown in Fig. 12 (Kanga et al., 2014), which is a plot of the nominal stress–strain curve of Ti–4Cr and Ti–4Cr–0.2O subjected to an aging treatment at 1023 K for 18.0 ks. In this case, a significant work hardening was observed due to twinning-induced plasticity (TWIP) in the b-phase.

Effect of Low Young’s Modulus on Stress Shielding It is important to evaluate the bonding ability and bone affinity, and the role of low Young’s modulus in preventing stress shielding and bone remodeling using animal testing.

Width of martensite, nm

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Nominal stress–strain curve in aged Ti–4Cr and Ti–4Cr–0.2O alloy aged at 1023 K for 18.0 ks.

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A study of the bonding ability and bone affinity of rods made of TNTZ subjected to aging treatment (TNTZAT) and solutiontreated TNTZ (TNTZST), as well as rods of pure Ti (CP-Ti), and Ti–6Al–4V ELI (Ti64), was carried out using these rods in canine mandibular implant beds of 18-month-old, male beagle dogs (Edamatsu et al., 2015). Fig. 13 (Edamatsu et al., 2015) shows the bone area ratio (BAR) and bone contact ratio (BCR) in the cancellous bone area 3 and 6 months after the implantation of TNTZAT, TNTZST, CP-Ti, and Ti64 rods. Both BAR and BCR of TNTZ are greater than those of CP-Ti and Ti64, indicating that the bonding ability and affinity of TNTZ are greater than those of CP-Ti and Ti64. Furthermore, the effectiveness of the low Young’s moduli of the implants in inhibiting stress shielding, namely inhibiting bone resorption and enhancing bone remodeling, were proved by evaluating bone fracture healing and bone remodeling caused by implanting intramedullary rods and bone fracture fixation plates made of TNTZ, Ti-6Al-4V ELI (Ti-64), and SUS 316L into a fracture model made in the tibia of Japanese white rabbits (Niinomi et al., 2002b; Sumitomo et al., 2008). In both cases, the low Young’s modulus TNTZ was proved to be much more effective in inhibiting stress shielding. Fig. 14 (Niinomi et al., 2006) shows the three-point bend strength (fracture load) of tibia of three Japanese white rabbits with and without implants of intramedullary rods made of TNTZ, Ti–6Al–4V ELI, and SUS 316L. The three-point bending fracture

(A)

(B)

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Fig. 13 (A) Bone area ratio (BAR) and (B) bone contact ratio (BCR) in the cancellous bone area at (A) 3 and (B) 6 months after implantation of TNTZAT, TNTZST, CP-Ti, and Ti64.

TNTZ (n = 3)

kgf

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SUS316L (n = 3)

80 +19.4%

60 –15.0% –0.2%

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–12.5%

+9.2% –9.1% died by +10.5% disease at 43 weeks

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Fig. 14

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Three-point bend strength: fracture load.

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Analysis of BVF obtained from histological image analysis with paired t-test in rabbit BVF

Number of rabbits

Ti2448

TiAlV

T1 T2 T3 T4 T5 Mean SD Paired t-test P

0.479 0.430 0.420 0.475 0.484 0.458 0.027 0.00019

0.329 0.296 0.327 0.335 0.342 0.326 0.016

(A)

(B)

(C)

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A B

B

Middle level A

B

C

A

A B

B

Fig. 15 CMRs (contact micro-radiographs) of cross sections of fracture models implanted with and without bone plates made of TNTZ at middle and distal level at 48 weeks after implantation: (A) cross section of fracture model, (B) parts of , of (A), namely high magnification CMR of branched parts of bones formed outer and inner sides of tibiae, and (C) cross sections of unimplanted tibiae.

strength of tibia of rabbits with implants of the intramedullary rods made of TNTZ is greater than those of Ti–6Al–4V ELI and SUS 316L, suggesting that the low Young’s modulus is effective in inhibiting bone absorption and promoting good bone remodeling. A animal testing similar to those mentioned earlier was carried out using intramedullary rods made of Ti–6Al–4V ELI and Ti– 24Nb–4Zr–7.9Sn (Ti2448 (TNZS)), which have a Young’s modulus of approximately 42 GPa. These two intramedullary rods were implanted into the tibiae of New Zealand white rabbits for 4 weeks. After this period, the rabbits were sacrificed to evaluate the bone volume fraction (BVF) around each intramedullary rod (Guo et al., 2009). The results are listed in Table 2 (Guo et al., 2009). BVF in the case of Ti2448 is greater than that of Ti–6Al–4V ELI (TiAlV). Therefore, here too it is found that low Young’ modulus is effective in enhancing the bone formation. Furthermore, Fig. 15 (Sumitomo et al., 2008) shows an increase in the diameter of the tibia and the double wall structure in the intramedullary bone tissue, which was observed only for the case of the bone plate made of TNTZ (Narita et al., 2012). In this figure, the inner wall bone structure indicates the original bone, that is, the remaining old bone, whereas the outer wall bone structure indicates the newly formed bone. This bone remodeling is the direct result of using a bone plate with a low Young’s modulus. Therefore, in implants made of the b-type Ti alloy with low Young’s modulus for biomedical applications, stress transfer between the bone and the implant becomes homogeneous. The prevention of bone resorption and the good bone remodeling are achieved, although this has been proved only at the stage of animal testing.

Summary Biomedical Ti alloys with new concept have been developed and are currently being developed. However, their biofunctionalities are still poor. To address this issue, surface modifications using biofunctional materials are performed on Ti alloys for biomedical applications. However, such biofunctionalization of Ti alloys, for example, smart Ti alloys that change according to the mechanical circumstances in the body should be further developed.

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Steinemann, S. G. (1980). Corrosion of surgical implantsdIn vivo and in vitro tests. In G. D. Winter, J. L. Leray, & K. de Groot (Eds.), Evaluation of biomaterials (pp. 1–34). New York: Wiley. Sumitomo, N., Noritake, K., Hattori, T., Morikawa, K., Niwa, S., Sato, K., & Niinomi, M. (2008). Experiment study on fracture fixation with low rigidity titanium alloy – Plate fixation of tibia fracture model in rabbit. Journal of Materials Science: Materials in Medicine, 19, 1581–1586. Suzuki, N., Iizumi, H., & Ogawa, A. (1999). Super plastic formability of Ti-4.5%Al-3%V-2%Fe-2%Mo alloy. Journal of Japan Institute of Light Metals, 49, 363–367. Tane, M., Akita, S., Nakano, T., Hagihara, K., Umakoshi, Y., Niinomi, M., & Nakajima, H. (2008). Peculiar elastic behavior of Ti-Nb-Ta-Zr single crystals. Acta Materialia, 56, 2856–2863. Tsutsumi, Y., Nishimura, D., Doi, H., Nomura, N., & Hanawa, T. (2009). Difference in surface reactions between titanium and zirconium in Hanks’ solution to elucidate mechanism of calcium phosphate formation on titanium using XPS and cathodic polarization. Materials Science and Engineering C, 29, 1702–1708. Ueda, K., Nakaoka, S., & Narushima, T. (2013). b-grain refinement of a þ b-type Ti–4.5Al–6Nb–2Fe–2Mo alloy by using rare-earth-oxide precipitates. Materials Transactions, 54, 161–168. Uggowitzer, P. J., Bähr, W.-F., & Speidel, M. O. (1997). Metal injection molding of nickel-free stainless steels. Advances in Powder Metallurgy and Particulate Materials, 3, 18.113– 18.121. Yamamoto, A., Honma, R., Sumita, M., & M. (1988). Cytotoxicity evaluation of 43 metal salts using murine fibroblasts and osteoblastic cells. Journal of Biomedical Materials Research Part A, 30, 331–340. Yilmazer, H., Niinomi, M., Nakai, M., Hieda, J., Todaka, Y., & Miyazaki, T. (2013). Mechanical properties of a medical b-type titanium alloy with specific microstructural evolution through high pressure torsion. Materials Science and Engineering C, 33, 2499–2507. Zhao, X. L., Niinomi, M., & Nakai, M. (2011a). Relationship between various deformation-induced products and mechanical properties in metastable Ti-30Zr-Mo alloys for biomedical applications. Journal of the Mechanical Behavior of Biomedical Materials, 4, 2009–2016. Zhao, X. L., Niinomi, M., Nakai, M., Miyamoto, G., & Furuhara, T. (2011b). Microstructures and mechanical properties of metastable Ti-30Zr-(Cr, Mo) alloys with changeable Young’s modulus for spinal fixation applications. Acta Biomaterialia, 7, 3230–3236. Zhao, X. F., Niinomi, M., Nakai, M., Hieda, J., Ishimoto, T., & Nakano, T. (2012a). Optimization of Cr content of metastable b-type Ti-Cr alloys with changeable Young’s modulus for spinal fixation applications. Acta Biomaterialia, 8, 2392–2400. Zhao, X. F., Niinomi, M., Nakai, M., & Hieda, J. (2012b). Beta-type Ti-Mo alloys with changeable Young’s modulus for spinal fixation applications. Acta Biomaterialia, 8, 1990–1997.

Further Reading Ambrosioand, L., & Tanner, E. (2012). Biomaterials for spinal surgery. Sawston, Cambridge, England: Woodhead Publishinig Ltd. Kuroda, D., Niinomi, M., Morinaga, M., Kato, Y., & Yashiro, T. (1998). Design and mechanical properties of new beta type titanium alloys for implant materials. Materials Science and Engineering A, 243, 244–249. Niinomi, M. (1998). Mechanical properties of biomedical titanium alloys. Materials Science and Engineering A, 243, 231–236. Niinomi, M. (2003a). Fatigue performance and cyto-toxicity of low rigidity titanium alloy Ti–29Nb–13Ta–4.6Zr. Biomaterials, 16, 2673–2683. Niinomi, M. (2003b). Recent research and development in titanium alloys for biomedical applications and healthcare goods. Science and Technology of Advanced Materials, 4, 445–454. Niinomi, M. (2008a). Mechanical biocompatibilities of titanium alloys for biomedical applications. Journal of the Mechanical Behavior of Biomedical Materials, 1, 30–42. Niinomi, M. (2008b). Biologically and mechanically biocompatible titanium alloys. Materials Transactions, 49, 2170–2178. Niinomi, M. (2010). Fatigue failure of metallic biomaterials, metals for biomedical devices. Cambridge, England: Woodhead Publishing Ltd. Niinomi, M. (2014). Biomedical implant devices fabricated from low Young’s modulus titanium alloys demonstrating high mechanical biocompatibility. Materials Matters, 9, 39–46. Niinomi, M., & Akahori, T. (2010). Improvement of the fatigue life of titanium alloys for biomedical devices through microstructural control. Expert Review of Medical Devices, 7, 481–488. Niinomi, M., Narushima, T., & Nakai, M. (2015a). Advances in metallic biomaterials, Part I: Tissues, materials and biological reactions. Wang, M. (series editor). In Springer Series in Biomaterials Science and Engineering (Vol. 3). Berlin, Heidelberg, Germany: Springer. Niinomi, M., Narushima, T., & Nakai, M. (2015b). Metallic biomaterials, Part II: Processing and applications. In Springer Series in Biomaterials Science and Engineering (Vol. 4). Berlin, Heidelberg, Germany: Springer. Wang, M. (series editor). Pignatello, R. (2011). Biomaterials science and engineering. Rijeka, Croatia: In Tech. Ramalingam, M., & Kobayashi, H. (2010). Integrated materials for biomedical technology. Linkoping, Sweden: VBRI Press. Sasaki, K., Suzuki, O., & Takahashi, N. (2017). Interface oral health science 2016: Innovative research on biosis–abiosis interface. Berlin, Heidelberg, Germany: Springer.

BIOMATERIALS: IN VITRO AND IN VIVO STUDIES OF BIOMATERIALS Anatomy and Physiology for Biomaterials Research and Development

Inn Chuan Ng and Pornteera Pawijit, National University of Singapore, Singapore Jordon Tan, Temasek Polytechnic, Singapore Hanry Yu, National University of Singapore, Singapore; Agency for Science, Technology and Research (A*STAR), Singapore; BioSyM, Singapore-MIT Alliance for Research and Technology, Singapore; and Nanfang Hospital, Southern Medical University, Guangzhou, China © 2019 Elsevier Inc. All rights reserved.

Biomaterials Interact With Biological Systems Principles of Anatomy and Physiology Anatomy Levels of organization Physiology Processes of life Homeostasis Structure to Function Organ Systems and Organs The Circulatory System Heart and blood vessels Blood Skeletal System Bone Cartilage and ligaments Integumentary System Epidermis Dermis Tissues Cells Structure of a Cell Cell membrane Organelles Cytosol Cytoskeleton Cells Interact With Extracellular Environment (Input) Cell–ECM interaction Cell-soluble signals interaction Cell–cell interaction Signaling Cascade Cell Response (Output) Gene expression Cell differentiation Cell growth and division Cell spreading and migration Mechanistic Insights Into the Biomaterials Interaction With Biological Systems Further Reading Relevant Websites

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Glossary Adaptive immunity Specific resistance in which the body recognizes and memorizes antigens that it encounters. Response to the particular antigens improves each time the body encounters the antigens again. Anatomy A discipline of biology that studies the structure and organization of organisms.

Encyclopedia of Biomedical Engineering, Volume 1

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Cell The most basic functional unit of life. Cytokines Small proteins released by cells. They act as chemical mediators. Gene expression A process in which genetic information is used to synthesize functional products. This process is explained by the central dogma of molecular biology, which describes the flow of genetic information from DNA to RNA to protein. Gene expression involves gene regulation, transcription of DNA into RNA, RNA processing, translation of RNA into protein, and posttranslational modifications of protein. Epigenetic regulation of the genetic information is usually involved in regulating gene expression. Homeostasis The tendency of biological systems to maintain a relatively constant equilibrium in their internal environment. Hormones Chemical mediators produced by specialized glands. They can be proteins, steroids, amino acid derivatives, or fatty acid derivatives. Immune response A reaction of the body to protect itself against substances that appear foreign. Innate immunity General resistance to certain antigens. Response to the antigens is the same each time the body is exposed. Organ A collection of tissues that form a distinct structural unit and carry out a specific function together. Organ system A group of organs and tissues that perform similar functions. Physiology A discipline of biology that studies the functions and processes of organisms. Tissue A group of similar cells that carry out a specific function together.

Biomaterials Interact With Biological Systems Biomaterials are materials that have been designed to interface with biological systems, for the treatment, augmentation, or replacement of biological functions. Biomaterials and biological systems interact both ways. Biomaterials and any compounds released from them may induce positive or negative responses from biological systems, from increase in wound healing and improvement in biological functions to toxicity and activation of the immune response. Conversely, biological systems may modify the surfaces of biomaterials through the biomaterial corrosion, degradation, and deposition, which will affect the integrity and performance of biomaterials and in turn lead to subsequent biological responses. Since the interaction between biomaterials and biological systems is dynamic and occurs gradually over time, it may lead to different observed responses between short and long term. To understand the interaction of biomaterials with the relevant biological systems, knowledge of the principles of biology, especially in anatomy and physiology, is required. The knowledge is important during biomaterials research and development for improving biomaterials design and for evaluating the effects of biomaterials on biological systems.

Principles of Anatomy and Physiology Anatomy Anatomy is the study of the structure and organization of organisms, at both microscopic (histology) and macroscopic (gross) levels. Anatomy is typically studied through dissections, which is often accompanied by additional staining, optical, or imaging methods. The anatomy of an organism is understood as a hierarchical organization.

Levels of organization Through a reductionist approach, the anatomy of an organism can be simplified to a manageable level. An organism can be divided into levels of organization with descending complexity: from organism to organ systems, to organs, to tissues, and to cells (Fig. 1). An organ system is a group of organs and tissues that perform similar functions; an organ is a structurally distinct group of tissues that work together to perform a specific function, and a tissue is a group of cells. The following are the key characteristics of this hierarchical organization:

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Each component in a level is assembled from the more basic components of a lower level. Thus, each higher level in the hierarchy represents an increase in organizational complexity. For example, the digestive system constitutes organs such as the stomach, liver, pancreas, small intestines, and large intestines. Emergent functions appear at higher levels due to cooperation of different basic components of lower levels. These functions are not present at the lower levels. For example, digestion of food requires the cooperation of multiple organs and cannot be performed by a single organ. The mouth mechanically breaks down food; acids and enzymes in the stomach degrade partially digested food into chime; the liver produces bile, while the pancreas produces other enzymes to further digest food into its basic constituents. Each component may communicate with and regulate the function of another component from a similar or different hierarchical level. For example, the nervous and endocrine system can regulate the production of enzymes by the different organs in the digestive system.

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Fig. 1 Levels of organization. An organism is made up of different organ systems. Each organ system consists of different organs. Each organ is formed by different types of tissues. Each type of tissue is assembled from cells. Examples of components at each level are shown. Illustrations by Wong Chun Xi, MBI Science Communications Unit.

Physiology Physiology is the study of functions and processes of organisms and their components. The study includes normal and diseased states and maintenance of homeostasis under different environmental conditions. Physiology can be studied using physical, chemical, mechanical, and biological assays. Assays are investigative procedures for qualitative and quantitative measurement of biological functions. These assays can be performed in vitro in laboratory vessels or in vivo on animals. The interaction of biomaterials with biological systems can be evaluated using the mentioned assays for studying physiology. The interaction of a scaffold with cells can be studied in vitro in cell culture plates or in vivo by implantation of the scaffold into the animal models.

Processes of life An organism has seven key processes of life. These physiological processes are important to keep an organism alive. A biomaterial can be used to support or suppress some of these processes:

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Respirationda process of extracting energy out of food Movementdto search for food and shelter and to avoid danger Sensitivitydawareness of and the degree of response to changes in environmental conditions Growthddevelopment and repair of organism and their parts Reproductiondproduction of offspring Nutritiondacquisition of nutrients or its synthesis Excretiondremoval of harmful waste products

Homeostasis To ensure survival, a biological system needs to maintain a relatively stable internal environment, despite the continuous changes in the signals from within and outside the system. The maintenance of this dynamic equilibrium is called homeostasis. Homeostasis requires three components for constant monitoring and adjustments of the internal environment: 1. The receptor, which senses the changes 2. The control center, which receives and processes the information sent from the receptor 3. The effector, which increases or reduces its function in response to the commands from the control center Homeostasis is maintained through negative feedback: a response by the system to reverse the direction of change. Examples of homeostasis are the maintenance of body temperature and blood glucose level. Biomaterials are usually used to restore but seldom allowed to dysregulate homeostasis.

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Structure to Function Anatomy and physiology are a pair of related disciplines. They are often studied in combination to relate organism structure to function. Biomaterials interact with anatomy and impact physiology at different levels of organization. Positive effects of a biomaterial on one component at a particular level might not necessarily translate to similar effects on another component at the same level or other units at higher or lower levels.

Organ Systems and Organs An organism is made up of different organ systems. Each organ system performs a specific function, and all organ systems work in an interconnected manner to maintain the internal conditions essential to the survival of the organism. There are 11 key organ systems in the human body (Table 1). We will use three examples of the human organ systems of particular interest to the field of biomaterials, the circulatory, skeletal, and integumentary systems, to explain in detail how these organ systems work.

The Circulatory System The components of the circulatory system include the heart, blood vessels, and blood. The circulatory system delivers nutrients and hormones throughout the body and removes waste products and excess heat generated by the metabolic activities of the body. Biomaterials used in devices for the circulatory system often include catheters, pacemakers, heart valves, vascular grafts, and stents. Table 1

Function of human organ systems and their components

System

Functions

Key organs and tissues

Integumentary system

Forms a permeability barrier, regulates body temperature, reduces water loss, and helps produce vitamin D

• The skin (provides protection against abrasion and

Skeletal system

Provides structural protection and support, allows body movement, stores minerals and fat, and produces blood cells

Muscular system

Produces body movements, maintains posture, and generates body heat Removes foreign substances from the blood and lymph, combats disease, maintains tissue fluid balance, and absorbs fats from the digestive tract

Lymphatic system

Respiratory system

Exchanges oxygen and carbon dioxide between the blood and air, and regulates blood pH

Digestive system

Performs mechanical and chemical processes of digestion, absorption of nutrients, and elimination of wastes

ultraviolet light)

• Hair (regulates body temperature) • Nails (protect soft tissues from injuries) • Sweat glands (regulates body temperature) • Bone (provides mechanical support and body structure) • Cartilage (provides firm but flexible support) • Ligaments (prevent movements that may damage joints) • Joints (for movement and stability of limbs) • Muscle (generates contractile forces) • Tendons (connect skeletal muscles to bones) • Lymphatic vessels (drain tissue fluids back into the bloodstream)

• Lymph nodes (filtration of lymph to identify antigens) • Thymus (maturation site of T lymphocytes) • Spleen (filters blood) • Nose (primary organ of smell) • Larynx (sound generation) • Pharynx (provides airflow to and from lungs) • Trachea (provides airflow to and from lungs) • Bronchi (provide airflow to and from lungs) • Lungs (principal organ of respiration) • Mouth (chews/breaks down food) • Esophagus (channels food into the stomach) • Stomach (breaks down protein-rich food) • Liver (detoxifies and metabolizes chemicals and drugs) • Pancreas (secretes enzymes to digest food) • Gallbladder (stores and concentrates bile) • Small intestines (absorb nutrients and minerals) • Large intestines (absorb water) • Mesentery (attaches intestines to the walls of the abdomen)

Nervous system

Detects sensations and controls intellectual functions, movement, and physiological processes

• Rectum (temporary storage for feces) • Anus (digestive tract opening) • Brain (provides control over the actions of the body) • Spinal cord (connects the peripheral nervous system to the brain)

• Nerves (transmit nervous signals between different parts of the body)

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Function of human organ systems and their componentsdcont'd

System

Functions

Key organs and tissues

Endocrine system

Regulates metabolism, growth, reproduction, and many other functions

• Hypothalamus (links the nervous system to the endocrine system)

• Pituitary gland (secretes a variety of hormones) • Pineal gland (produces melatonin) • Thyroid gland (secretes hormones that regulate the growth and metabolism of the body)

• Parathyroid gland (secretes hormone that regulates calcium and phosphate levels)

• Adrenal glands (secrete hormones that regulate blood sugar levels and blood pressure)

• Pancreas (secretes hormones that regulate blood glucose levels)

• Ovaries (produce oocytes for reproduction and secrete estrogen and progesterone)

Circulatory system

Transports nutrients, waste products, gases, and hormones throughout the body, mediates immune response and the regulation of body temperature

Urinary system

Removes waste products from the blood and regulates blood pH, ion balance, and water balance

Reproductive system

Production of sex cells and hormones that influence sexual function and behaviors

• Testes (produce sperm and testosterone) • Heart (pumps blood for circulation) • Blood vessels (transport blood throughout the body) • Blood (carries oxygen, hormones, and waste products and is involved in the immune response)

• Kidneys (extract waste from blood and body fluids to form urine)

• Ureters (carry urine from the kidney to the bladder) • Urethra (carries urine from the bladder to outside of the body) Male • Testes (produce sperm and testosterone) • Prostate gland (secretes prostate fluid) • Penis (male sexual organ) Female • Ovaries (produce oocytes for reproduction and secrete estrogen and progesterone) • Uterus (nurtures fertilized ovum) • Vagina (birth canal)

Heart and blood vessels The heart is a muscular organ that contains four chambers. Valves at the entries and exits of the chambers maintain the correct unidirectional flow of blood. Sequential contraction of the heart chambers allows the heart to receive blood and pump it out to the lungs or to other parts of the body through closed loops of blood vessels. There are three types of blood vessels: the arteries, veins, and capillaries. Arteries take blood away from the heart, while veins bring blood toward the heart. Arteries have comparatively thicker walls because they experience higher blood pressure due to proximity to the heart. Veins contain valves that prevent the backflow of blood. Arteries decrease in diameter as they branch out from the heart throughout the body and end in capillaries. Meanwhile, veins begin from capillaries and increase in diameter as they converge toward the heart. Capillaries are tiny web of vessels of 5–10 mm in diameter that connect arteries to veins at the other end of the heart. The thin walls of capillaries allow the exchange of gas and nutrients between the blood and surrounding tissues. Blood flow and pressure are regulated by the heart rate and the contraction and relaxation of arteries and veins. When artificial heart valves or stent are implanted into the circulatory system, they interact with the flowing blood and the neighboring tissues that determine plaque deposition and blood flow dynamics in the local structures.

Blood Blood is a type of connective tissue. It is composed of plasma, a liquid matrix in which blood cells are suspended. The plasma is an aqueous solution consisting of water, blood plasma proteins and dissolved electrolytes, nutrients, gases, waste products, and other substances, while blood cells are grouped into red blood cells (erythrocytes), white blood cells (leukocytes), and platelets (thrombocytes). Red blood cells are biconcave-shaped cells with no nucleus and contain hemoglobin that helps in the transport of oxygen and carbon dioxide. White blood cells mediate immune response to protect the body against foreign substances and further categorized into two groups based on their morphology, namely, granulocytes and agranulocytes. Granulocytes consist of neutrophils, which phagocytose microorganisms and other substances; basophils, which promote inflammation by releasing histamine and prevent clot formation by releasing heparin; and eosinophils, which reduce inflammation by releasing cytokines and attack certain worm parasites. Agranulocytes consist of monocytes and lymphocytes. Monocytes leave the blood to become macrophages, whose function is to phagocytose bacteria, dead cells, cell fragments, and other debris within tissues. Lymphocytes can be divided into

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natural killer (NK) cells and B and T lymphocytes. Lymphocytes destroy microorganisms by producing antibodies and other cytokines; contribute to allergic reactions, tumor control, and graft rejection; and regulate the immune system. Platelets are cell fragments that take part in blood clotting. Immune response is a reaction of the body to protect itself against substances that appear foreign, such as microorganisms, parasites, harmful compounds, cancer cells, and some biomaterials. Activation of the immune response may lead to hypersensitivity of the body and rejection of a biomaterial. The immune response is mediated by certain plasma proteins in the blood and white blood cells in the circulatory and lymphatic systems. It is regulated by cooperation of two different mechanisms: the innate and adaptive immunities. The innate immunity is a general resistance that is activated immediately or within hours after the body encounters certain antigens, substances that activate an immune response, but the reaction to them is the same each time the body is exposed. It is activated by the chemical properties of the antigens. Innate immune response can manifest as local or systemic inflammation and is mediated by neutrophils, macrophages, NK cells, and complement cascade. Phagocytosis of antigens induces neutrophils and macrophages to release chemical mediators called cytokines that regulate the inflammatory response. These cytokines may either encourage inflammation by increasing vascular permeability and recruiting additional white blood cells to the site of inflammation or reduce inflammation and promote tissue repair. The former cytokines are released by M1 macrophages, while the latter cytokines are released by M2 macrophages. NK cells induce apoptosis of virus-infected and cancer cells. Complements bind to antigens and enhance their phagocytosis by neutrophils and macrophages, promote inflammation, and assemble to form membrane pores that induce lysis of pathogens. The adaptive immunity is a specific resistance in which the body recognizes and memorizes antigens, and the reaction to previously encountered antigens improves each time they meet again. It is mediated by B and T lymphocytes. B lymphocytes mature in the bone marrow and produce antibodies that bind to antigens, whereas T lymphocytes, which mature in the thymus, release cytokines and induce cytotoxicity. Antigen recognition by adaptive immunity involves cell surface proteins called the major histocompatibility complex (MHC). MHC receptors present processed fragments of antigens called epitopes at cell surfaces for recognition by T lymphocytes. There are three classes of MHC receptors: MHC class I, II, and III. MHC class I receptor is expressed by all nucleated cells and platelets. It presents epitopes to a subtype of T lymphocytes called cytotoxic T lymphocytes (CTLs). CTLs express CD8 receptor and T-cell receptor (TCR). CD8 receptor interacts with MHC class I receptor while TCR will only bind to a specific epitope. An antigenpresenting cell will be induced by CTL to undergo programmed cell death if the TCR could bind to an epitope presented by its MHC class I receptor while the MHC class I receptor itself is interacting with a CD8 receptor. MHC class II receptor is normally expressed by professional antigen-presenting cells such as macrophages, dendritic cells and B lymphocytes. It presents epitopes to a subtype of T lymphocytes called helper T lymphocytes (Th). Th lymphocytes express CD4 receptor and TCR. CD4 receptor interacts with MHC class II receptor. Naïve Th lymphocyte will undergo differentiation if its TCR could bind to an epitope presented by an MHC class II receptor while the MHC class II receptor itself is interacting with its CD4 receptor. The naïve Th lymphocyte will differentiate into either effector or memory Th lymphocyte that mediate immunization, or suppressor T lymphocyte that mediate immune tolerance. Biomaterials interaction with immune system plays a central role in any implantable materials or devices or even the extracorporeal devices that connect to the body via circulatory system.

Skeletal System The skeletal system consists of bone, cartilage, and ligaments. It protects and supports the body, helps with movement, produces blood cells, and stores important minerals. Examples of orthopedic applications of biomaterials include hydroxyapatite ceramics for bone grafts, biodegradable polymers for bone and cartilage substitution, and scaffolds for ligament reconstruction.

Bone Bone is a composite material made up of calcium hydroxyapatite, a mineral that provides strength and rigidity, and collagen, an elastic protein that provides flexibility and improves fracture resistance. Since bones are hard, they form the major structural support components of the body and protect organs such as the brain, heart, and lungs. There are two types of bones based on their structural arrangement: the compact and spongy bones. The compact bone is a dense, solid matrix that forms most of the shaft of long bones and provides mechanical strength. On the other hand, the spongy bone, which consists of a lacy network of bone with many small, marrow-filled spaces, provides compressional strength. Bone is a living organ that is constantly being remodeled. Bone remodeling occurs during bone growth and regulates calcium availability. Old bone is removed by cells called osteoclasts, and new bone is deposited by osteoblasts. Extensive work in biomaterial R&D has been conducted to stimulate bone regeneration in vivo or prosthesis to replace lost bone function with inert biomaterials.

Cartilage and ligaments The skeletal system also consists of connective tissues such as the cartilage and ligaments. Cartilage is a smooth, firm, resilient, nonvascular tissue, which provides structural support and reduces friction between bones at the joints. Ligaments are fibrous tissue that connect bones together at the joints, providing joint stability and prevent movement that might damage the joint. Their characteristics are mainly determined by their extracellular matrix (ECM) composition. The matrix always contains collagens and

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proteoglycans, but the types and quantities of these substances differ. Collagen is a tough, fiber-like protein that provides rigidity and flexibility. Proteoglycans are large molecules that consist of polysaccharides attached to core proteins that attract and retain large amounts of water. In the cartilage, the ECM consists of chondrocytes, polymeric components (type II collagen and proteoglycans), and water. The collagen makes the cartilage tough, whereas the proteoglycans make it smooth and resilient. Thus, the cartilage is relatively rigid but springs back to its original shape after being bent or slightly compressed, enabling it to absorb shock. On the other hand, the ligaments contain large amounts of collagen fibers, making these structures very tough. A key focus of cartilage and ligament tissue engineering for rehabilitation is to engineer constructs that mimic the mechanical properties and loadbearing characteristics of the cartilage and ligaments in vivo.

Integumentary System The integumentary system is an organ system that consists of the skin, its derivatives (sweat and sebaceous glands), nails, and hair. It forms a physical barrier between the body and the environment; it protects the body against invasion by infectious organisms and against ultraviolet light, reduces water loss, and provides thermal insulation. Other functions include excreting waste products through perspiration; sensing stimuli such as touch, pressure, pain, heat, and cold; generating vitamin D under ultraviolet light; and serving as a store for water and fat. Examples of biomaterial applications for integumentary system include skin grafts and skin substitutes such as epidermal covers and dermal replacements. There are commercial skin grafts of different levels of sophistication for implant or for drug testing in vitro. The current challenges are developing more sophisticated biomaterial support to improve the aesthetics with micro-/nanosized features, mechanical contractility, and interaction with immune or nervous systems.

Epidermis The human skin is composed of two major layers of tissue: the epidermis and dermis. The epidermis forms the outermost layer, providing the initial barrier to the external environment. Cells in this layer are increasingly filled with keratin, and they eventually die to form an outer layer of dead, hard cells that resist abrasion and also act as a permeability barrier. Another unique characteristic of this layer is its pigmentation. Melanin, the pigment that gives the skin its color and protects it against ultraviolet rays from the sun, is produced by melanocytes. Skin color has implications on the cosmetic value of biomaterials. Earlier generations of skin grafts involve only the epidermis substitutes.

Dermis Beneath the epidermis is the dermis that comprises of two layers: the papillary and reticular layers. The dermis contains connective tissues, vessels, glands, follicles, hair roots, sensory nerve endings, and muscular tissue. Beneath the dermis is the hypodermis that is primarily made up of adipose tissue. Substantial collagen bundles anchor the dermis to the hypodermis in a way that permits most areas of the skin to move freely over this tissue layer. Modern skin grafts involve both dermis and epidermis layers for mimicking the skin structure and functions.

Tissues Tissue refers to a group of cells and their microenvironment, which perform their functions together. There are four main types of tissues: muscle, epithelial, connective, and nerve tissues. Muscle tissue is involved in generating contractile forces, epithelial tissue contains tightly packed cells, connective tissue contains cells enveloped by the ECM, and nerve tissue contains nerve cells that conduct nerve impulses. Organs, for example, the heart, are made up of different types of tissues. Cardiac muscle tissue, a subtype of muscle tissue that is composed of aligned cardiomyocytes, mediates the contraction of the heart. The inner linings of the heart are covered by the endothelium, a subtype of epithelial tissue. The endothelium forms a barrier separating the cardiac muscle tissue from the blood, a subtype of connective tissue. Nerve tissue connects the heart to the nervous system, allowing heart rate to be regulated by the brain. In most applications, a biomaterial will interact with biological systems at the tissue level; rarely does it interact with the organ as a whole or only with a single cell. Tissue response is a collective response of individual cells to the environment. Guiding principles behind response of individual cells can be applied to the tissue level as well.

Cells Cells are the most basic component of life; they assemble and work together to form components at higher levels of organization. Millions of cells make up an organism. Although all called cells, there are a lot of different types of cells, each have different shape, distribution, and function. Cells respond to signals from the environment that may come in the form of ECM surrounding the cells, soluble signals, or physical interaction with neighboring cells (Fig. 2). Cells are the functioning unit to sense and process the signals from the environments and respond to the environmental cues. Cells are made up of many compartments that work together to make the cell a responsive functional unit. Cells respond locally to local stimulus outside the cells, in the form of physical or chemical stimulus, and convert this local signal into a cascade of intracellular signals amplified and transported to decision centers such as

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Fig. 2 Tissue and cell. Tissue is composed of cells and their environment. Cells are the most basic unit of life. Cells are made up of many compartments (blue): cell membrane, cytoskeleton, nucleus, mitochondria, endoplasmic reticulum, Golgi, and lysosome. It interacts with the environment through molecules highlighted in yellow: extracellular matrix, integrins, soluble signals, receptors, and cell–cell adhesion. Components involved in gene expression are labeled in black: DNA, RNA polymerase, RNA, ribosomes, mRNA, and polypeptide chain. Illustrations by Wong Chun Xi, MBI Science Communications Unit.

nucleus. Decision centers would integrate all the signals from the different stimulus from different parts of the cells over a specific period to decide a course of actions such as to move a cell, turn on a gene, and start cell differentiation or proliferation.

Structure of a Cell The cell is like a small machine; it consists of several compartments that work together to carry out proper cell functions. Although different cells have different shape, distribution, and function, they all have these compartments. Compartmentalization is a central concept in biology with lipid membranes separating aqueous compositions of living organisms into smaller compartments. Cellular compartments help to concentrate molecules on a two-dimensional space to facilitate molecular interactions, which speed up chemical reactions.

Cell membrane Cell membrane is a lipid bilayer that separates the inside of the cell from the outside environment. The lipid bilayer is made up of phospholipids that consist of a hydrophilic head and hydrophobic tails, which are assembled into two-layered sheet with the hydrophobic tails directed toward the center of the sheet. The cell membrane forms an impermeable barrier to most molecules, allowing only small nonpolar and lipid-soluble molecules such as oxygen, carbon dioxide, and steroids, to diffuse freely through it. Interspersed throughout the cell membrane are molecules such as certain proteins, sphingolipids, and cholesterol. The function of the cell membrane is to serve as a barrier between intracellular compartments and extracellular environment and prevent molecules from coming in or leaving the cell freely. This allows the internal cell environment to be stabilized and not fluctuate with the changes in the external environment. This stabilization is important as disruption to homeostasis in the internal environment can lead to cell death. Biomaterials design usually considers the interaction with the cell membranes as a critical issue. Other membrane compartments within a cell are in the form of organelles. Like the cell membrane, these internal membrane compartments separate their contents from other parts of the cell, allowing different environments to be maintained within the cell. Hence, substances that need to be delivered within the cell might have to cross multiple membranes to reach a particular location. Nanomaterials or biomaterials for drug and gene delivery can penetrate the cell membranes and concentrate on the membranes by taking advantage of the charge properties or hydrophilicity of the membrane to accelerate their intracellular delivery.

Organelles Organelles are membrane-enclosed compartments in the cells that perform different functions.

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Nucleus The nucleus is enveloped in a double membrane, which is contiguous to the endoplasmic reticulum (ER). On the membrane are pores that regulate the constant flow of entry and exit of molecules into and from the nucleus. Molecules like transcription factors translocate in and the transcribed mRNAs leave the nucleus. The nucleus contains genetic information of the cell and proteins that maintain and transcribe the deoxyribonucleic acid (DNA). The genetic information in the nucleus is kept in long strands of DNA molecules. The DNA is composed of four nucleotide bases, guanine, adenine, thymine, and cytosine, short formed GATC. The bases are complementary pairs, G with C and A with T. The bases are linked together to form a strand. The strand exists in pairs in a structure called the double helix. One strand is called the coding strand, and the other strand is called noncoding strand.

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Nuclear translocation of drug or gene delivery carriers can take advantage of the dynamic nature of the bidirectional transport of molecules across the nuclear pore complex and importin-/exportin-mediated trafficking machinery to improve the efficiency of the gene and drug delivery into the nucleus. Transfection of foreign genes can be much more effective during cell division when the nuclear envelop dissolves. Mitochondria The mitochondria consist of an outer and inner membrane, the outer membrane encasing the mitochondria and a curvy inner membrane which allows for more surface area for reactions to take place. The inner membrane is where glucose is converted to adenosine triphosphate (ATP). Mitochondria are the powerhouse of the cell where ATP, the energy currency the cell uses, is generated. Mitochondria also play a role in regulating cell death and survival, thus a target for biomaterial applications that involve these two processes. Endoplasmic reticulum ER is organized into rough and smooth ER based on local morphology. Rough ER has attached ribosomes on its outer surface and is involved in protein translation and folding; smooth ER plays a role in lipid metabolism. ER also acts as a calcium store, regulating calcium levels in the cytosol. ER is the largest organelle compartment extending to every corner of a typical cell. Besides the canonical role in protein synthesis, folding and modification, ER regulates local functions such as adhesion complex maturation, local protein synthesis in long-term potentiation of neurons, and transient local response to stress. Recent findings suggest that ER is a master regulator of cell functions by providing signaling highways connecting local responses to decision centers such as the nucleus. Golgi apparatus Golgi apparatus exists as many stacked vesicles in the cytoplasm. It is involved in intracellular vesicle trafficking, protein sorting, and modification. It is important for membrane receptor recycling that innovative biomaterials R&D can explore to tune the cell response to extracellular environmental cues. Lysosomes Lysosomes are membrane organelles that have a very low pH; this environment facilitates the breakdown process that occurs within lysosomes. pH-sensitive biomaterials have been developed to activate gene carriers to escape from organelles prior to reaching or escape from lysosomes.

Cytosol Cytosol consists of everything enclosed by the cell membrane, excluding the nucleus. The aqueous component in which organelles are floating in is the cytosol. Cytosol is a relatively dense media with high viscosity such that the molecules in cytosol moves by diffusion in short distance of a few nanometers and by active transport on cytoskeleton tracks over longer distances of micrometers.

Cytoskeleton Cytoskeleton provides structural support to maintain the shape of the cell and serve as tracks for the directional transport of cargoes within the cell. The cytoskeleton can be divided into three types based on the composition of their protein subunits: microfilaments, which are made up of actins; microtubules, which are made up of tubulins; and intermediate filaments, which are made of a variety of subunits. These subunits can reversibly polymerize into filaments or depolymerize into individual subunits, allowing the cytoskeleton to be dynamic and rearrange itself in response to different conditions, such as for cell migration, attachment, division, and polarization. Polymeric gene carriers can take advantage of the transport machineries along cytoskeleton to move foreign genes into different organelle compartments including nucleus.

Cells Interact With Extracellular Environment (Input) Cells respond to many environmental cues to survive and function, namely, cell–ECM, cell-soluble signals, and cell–cell interactions. Each signal is generated locally. For chemical signals, each ligand binding to a receptor initiates a local event and generation of one or more secondary messengers to diffuse to the nucleus. For physical signal such as substrate stiffness, a local pinching of the substrate by a sarcomere unit leads to the local activation of Src that transmits the amplified signal to the nucleus. The consequence of either the physical or chemical signals interacting with the cells locally is the generation of the amplified chemical messengers to be sent to the decision center such as nucleus to decide on whether and how the cell should respond.

Cell–ECM interaction Cells are usually in contact with the ECM, molecules that cells secrete into the extracellular space. ECM is made up of interlocking fibrous proteins and glycosaminoglycans, forming a mesh that provides structural and biochemical support to cells. Examples of ECM types are fibronectin, collagen, proteoglycan, elastin, and laminin. Structurally, ECM may affect cell function depending on its composition, stiffness, porosity, density, and alignment. Cells bind to ECM through cell membrane receptor proteins such as integrins. Integrins are linked to the cell cytoskeleton, allowing the cells to sense stiffness and other physical properties of the surrounding ECM, triggering different downstream pathways within the cell. ECM is also involved in presenting biochemical cues that are crucial for cell survival. Biochemical cues usually come in the form of soluble signals. ECM can tether the soluble signals to cell surface, increasing their effective concentration by reducing their degrees of freedom to only two dimensional space.

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It is important to understand the native ECM properties around cells and in tissues, so that biomaterial scaffold design can recapitulate and mimic the in vivo environments. Decellularized scaffolds can support the cell growth ex vivo to mimic in vivo conditions. Scaffolds should contain the appropriate type of ECM so that cells can bind to it. We should also monitor the scaffold stiffness. Mammary epithelial cells cultured in stiff matrices express genes correlated to the breast cancer phenotype, in contrast to those cultured in matrices with physiological stiffness that exhibit normal gene expression profiles.

Cell-soluble signals interaction Cells can communicate to other cells that are not in immediate contact with them, by releasing soluble signals into their surrounding and bloodstream. These soluble signals may be cytokines, growth factors, hormones, neurotransmitters, salt ions, exosomes or microvesicles containing lipids, DNA, and microRNA (miRNA). The soluble signal might either bind to receptors on the cell membrane and initiate signaling or can move into the cells and interact with proteins involved in cell signaling. Cells may receive signals released from other cells, in cases of paracrine and endocrine signaling, or those released by itself, in the case of autocrine signaling. In autocrine signaling, cells secrete soluble signals into its surrounding; these signals can in turn bind to receptors on the cell, acting on itself as a stimulus. An example of endocrine signaling is insulin, which is released by islet’s beta cells in the pancreas in response to an increase in blood glucose level. Insulin is secreted into the bloodstream to signal cells in other parts of the body such as the brain and liver to uptake glucose. Insulin does this by binding to insulin receptor, a protein expressed on the cell membrane. Binding of insulin triggers translocation of glucose transporter, GLUT4, to the cell membrane, enabling the cell to uptake glucose, as glucose cannot pass through cell membrane directly. Soluble signals such as growth factors are often conjugated to biomaterials, to confer additional properties needed such as to stimulate tissue regeneration. For example, VEGF blended in biomaterials have been used in implants to promote vascularization to facilitate wound healing and regeneration.

Cell–cell interaction Cell–cell interactions can be relatively stable such as those mediated by tight junction, adherens junction, desmosomes and gap junctions, or transient such as those mediated by the cell adhesion molecules (CAMs) from the immunoglobulin and selectin families. Tight junctions, adherens junctions, and desmosomes act as structural linkers that hold two cells together, whereas gap junctions form pores that link the cytoplasm of one cell to another, allowing small molecules to diffuse through. CAMs from the immunoglobulin and selectin families are involved in contact-dependent cell–cell signaling. Cell–cell contact is important for maintaining structural integrity of tissues, polarization, and proliferation of the cells. Cell–cell interaction is important for synchronizing responses in groups of cells at the tissue level, forming a communication channel between cells that do not rely on outside signaling. Examples include calcium signal passing through gap junctions in cardiac muscle tissue to synchronize beating of cardiomyocytes and supracellular actin cables that synchronize cell movements in development. Most cell lining parts of the body that contact the outside environment requires strong cell–cell interaction to form a barrier. Examples include air tracks that contact the air breathed in; gastrointestinal track contacts the food consumed. The barrier formed by tight cell–cell junction prevents leakage of outside molecules into the body. For many cell types, cell–cell interaction is essential for cell survival and function. Therefore, biomaterials are designed to allow and encourage cell–cell interaction in macroporous scaffolds or flexible nanofibers or ribbons to support cell and tissue functions required in applications.

Signaling Cascade All environment stimuli, whether physical or chemical, will activate the signaling cascade, a series of intracellular biochemical reactions that allow signals to be amplified. An example is the G-protein-coupled receptor, a receptor that activates Gproteins. Once a ligand binds to the extracellular domain of the receptor, the receptor will undergo conformational changes and be activated, which will then lead to activation of G-proteins bound to the intracellular domain of the receptor. An activated receptor can activate multiple G-proteins, and each of the activated G-protein can activate many downstream secondary messengers. This results in an exponential amplification of the signal, even with only one ligand bound. The understanding of signaling cascade is causative. Each node of the cascade is connected to the next node downstream through conformational changes of the downstream node via binding or modifications such as phosphorylation, palmitoylation, or methylation. Cells make decisions in a digital manner; a response is either triggered or not triggered. A cell makes decision based on the accumulation of analog signals diffused from elsewhere; once the accumulation reaches a certain threshold at the decision center, this triggers a response from the cell. For example, rapid changes in membrane potential in the cell body or dendrites of a neuron can diffuse to the decision center (in this case the axon hillock) to generate an action potential. Action potential is only triggered when the membrane potential at the decision center reaches a certain threshold, after which specific ion channels are activated that lead to rapid and massive influx of ions to generate an action potential. This can apply to other reactions in the cell as well: only when signals exceed a defined threshold is the response triggered. Decision is based on the accumulated signals from different locations of the cell over a defined period.

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Cell Response (Output) Once a decision is made by the decision center such as the nucleus, cells execute or implement a decision causing a phenotype change (such as proliferation, growth, and differentiation). One pathway does not occur in isolation from the others, and there are overlaps and dependencies on other pathways.

Gene expression Gene expression is the process in which protein is synthesized from genetic material. All cells in the body contain the same DNA but differ from one another due to a different subset of genes actively expressing the corresponding subset of proteins. A gene is a specific sequence of DNA that codes for a protein. The central dogma of molecular biology is a gene is transcribed to RNA; RNA is then read by ribosomes, which translates the information on the RNA to proteins. Gene expression encompasses all regulation steps that control the level of expression of a particular gene into protein. Levels of the expressed protein can be regulated at the DNA level, RNA level, and protein level:

• • •

•

Regulation at DNA level: Genetic information is stored as long strands of DNA packed in the nucleus; the level of compaction of DNA where a particular gene is located can affect expression levels. Histone modification or methylation of base DNA can promote or inhibit gene expression. Regulation at transcription level: At the transcription level, from DNA to RNA, there are many molecules that need to come together for transcription to occur. Transcription starts when RNA polymerase is assembled at promoter site, a sequence usually located upstream of the gene to be transcribed. Transcription factors can promote or inhibit recruitment of RNA polymerase to the gene. Enhancer sequences are DNA sequences that if transcription factors bind to can affect transcription rate. Regulation at RNA level: Transcribed RNA requires processing, to mRNA, before it can leave the nucleus and be translated to proteins. During transcription, both regions that will be translated, exons, and regions that will not be translated, introns, are transcribed. Intron regions have to be spliced out, so the final mRNA contains only exons. Regions need to be added on to both ends, 50 cap at the 50 end and poly-A tail at the 30 end. The 50 cap allows translation machineries to recognize the mRNA and translate it. An mRNA can be translated multiple times, each time an A is hydrolyzed off the poly-A tail; the longer the poly-A tail, the more times the mRNA can be translated before it is degraded. mRNA levels are also controlled by miRNA. These short strands of noncoding RNA bind complementarily to mRNA to affect mRNA stability, readability, and integrity. Regulation at translation level: mRNA is translated to a polypeptide chain. The polypeptide is not always fully functional and needs further modification. For example, insulin is translated to preinsulin polypeptide chain that is cleaved prior to being functional. Others may require posttranslational modifications and chemical modifications to the amino acid bases, before being fully functional.

Cell differentiation Cell differentiation is known as a process in which cells become specialized. Pluripotency means the potential to become any cell type and is a potential observed in stem or progenitor cells at the beginning of the differentiation process. This potential is slowly lost as the cell differentiates and gains specialized functions. The differentiation process alters the cell dramatically, its shape, size, and energy requirements. This process is not a linear and irreversible process. Differentiation selects a subset of genetic information to be expressed at different stages of the differentiation process. Therefore, differentiated cells can be manipulated back to a more primitive or stem cell-like state, by reprogramming it to express a particular set of genes. Cell differentiation is a result of the integration of different stimulus in a spatiotemporal fashion. Cell differentiation is sensitive to both mechanical and chemical stimulus from the environment. Improved understanding of the differentiation process can allow us to engineer the process to achieve desired outcome. This is especially true when developing biomaterials as delivery carriers of the stem or progenitor cells in vivo.

Cell growth and division Cell division requires a lot of energy and resources and is a highly controlled process in the cell. There are many checkpoints in the cell cycle: there are requirements to control whether cells can transition to the next phase. An example of what happens when the checkpoints are bypassed is cancer. Cancer is a result of uncontrolled cell growth and division. Normally, cells rely on several survival signals and factors to suppress the default cell-death process. In cancer, these checkpoints are bypassed, and cells grow and divide despite no growth or dividing signals. Majority of the cells exist in the nondividing and quiescent G0 phase. If there are environmental cues to instruct a cell to divide, the cell enters the division cycle. Cell division is split into two phases, interphase and mitosis. Interphase is further split into four stages, G0, G1, S, and G2. Interphase is when the cells prepare themselves for division in mitosis. The cells have to meet size requirements and should be able to synthesize another copy of DNA. Cells have to proceed strictly in a specific order and have to meet strict requirements before the cell cycle can proceed. Failure to meet requirements will force the cells to exit the division cycle. This ensures that cell division occurs only when needed and the dividing cells are healthy. Dividing cells express unique proteins such as topoisomerase IIB, PCNA, and Ki-67 required during the division process. These proteins are useful markers in the study of the implanted scaffolds on how the tissue is responding to the scaffold or whether cells are migrating into the scaffold and dividing.

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Cell spreading and migration Cell spreading and migration are a few of the many processes that involve cytoskeletal rearrangement. Cell spreading is when cells in its rounded morphology in suspension flatten out on a substrate. This is a sign of adherence to substrate, as cells need anchors on the substrate to allow it to flatten itself out. Two ways were observed that cells can spread isotropically and anisotropically. Cells usually spread in an anisotropic manner, where the edges extend due to actin polymerization. In isotropic spreading, edges spread out in a circular manner and flatten rapidly. Isotropic spreading can be induced by removing serum; cells spread at a faster rate isotropically. Cell migration is needed during development when cells have to move to its required location for tissue rearrangement. It is important in wound healing and responses to injury where cells need to migrate to close up wounds or move to site of injury, respectively.

Mechanistic Insights Into the Biomaterials Interaction With Biological Systems To understand the mechanism of how biomaterials interact with a biological system, researchers have conducted systematic evaluation using in vitro and in vivo assays. In vitro assays are conducted in laboratory vessels with isolated molecules or cells, not on the living organism itself. In vitro assays can be divided into cell-free and cell-based assays. The former evaluates the interaction of a biomaterial with biological molecules, while the latter evaluates the interaction of a biomaterial with cultured cells or tissues. In contrast, in vivo assays are conducted directly on living organisms to evaluate the effects of biomaterials on an organism at both local and system levels. The results from in vivo assays are more physiologically relevant, but due to the complexity in interpreting the results, in most cases, in vitro assays are the first assays of choice. In vivo assays will be performed to confirm the mechanistic insights revealed by the in vitro assays. A mechanistic study is performed to understand causation, not just correlation. Causation demonstrates the direct connection between the input stimulation (e.g., nanosized features of a polymer substrate) and output response (e.g., cell differentiation). A biological system exhibits short-term to long-term responses when exposed to biomaterials. Short-term responses involve early events, which occur almost immediately, in timescales of seconds to a few minutes upon stimulation. Studying such short-term responses should occur at local or subcellular levels in the spatioscale of a few micrometers. Such output responses are the direct consequence of the input stimulation. In contrast, long-term responses occur after significant time has elapsed, in timescale of hours, days, or weeks. By then, the cell phenotype changes observed are the consequence of complex feedback mechanism leading to a steady state. Therefore, long-term responses often are studied with the aid of computational model and system biology methodologies to understand what and how the decision-making centers integrate local signals and multiple signaling networks over an extended period into a coherent understanding of the input-out causative relationship.

Further Reading Alberts, B., Johnson, A., Lewis, J., Morgan, D., Raff, M., Roberts, K., & Walter, P. (2014). Molecular biology of the cell (6th edn.). New York: Garland Science. Badylak, S. F. (2015). Host response to biomaterials: the impact of host response on biomaterial selection. Oxford: Academic Press. Patton, K. T. (2015). Anatomy and physiology (9th edn.). London: Elsevier Health Sciences. Pocock, G., Richards, C. D., & Richards, D. (2013). Human physiology (4th edn.). Oxford: Oxford University Press. Sherwood, L. (2015). Human physiology: from cells to systems (9th edn.). Boston, MA: Cengage Learning.

Relevant Websites https://www.khanacademy.org/science/health-and-medicine/. http://www.nature.com/scitable/ebooks/essentials-of-cell-biology-14749010/contents. https://www.ncbi.nlm.nih.gov/books/NBK10757/. https://www.ncbi.nlm.nih.gov/books/NBK21054/. https://www.ncbi.nlm.nih.gov/books/NBK21475/. https://en.wikipedia.org/wiki/Human_body. https://en.wikibooks.org/wiki/Human_Physiology.

Animal Models in Biomaterial Development James M Anderson and Sirui Jiang, Case Western Reserve University, Cleveland, OH, United States © 2019 Elsevier Inc. All rights reserved.

Introduction Biomaterial and Device Perspectives in In Vivo Testing Selection of Animal Models for In Vivo Tests Implantation Species Differences Animal Models for Host Response Mechanisms Further Reading

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Introduction Animal studies with appropriate animal models are a necessary component in the pathway to ultimately determine clinical predictability in humans. The results of animal studies most commonly are utilized to determine the biocompatibility (safety) of the biomaterial/medical device in the intended application. This process involves the design and execution of animal studies, from the proof-of-concept stage to the pivotal preclinical study of the medical device required for regulatory submission. Preclinical testing in animal models is an important part of the regulatory process, used to determine the safety and efficacy of devices prior to human clinical trials. The choice of the animal model and the selection of in vitro tests should be made according to the intended use of the respective medical device, prosthesis, or biomaterial. Numerous factors contribute to the choice of animal experimental model. Animal models are generally divided into small animal models (mouse, rat, and rabbits) and large-animal models (dog, goat, pig, sheep, etc.). Small animal models are usually used for ethical, economical, and statistical considerations, whereas large-animal models are primarily used if small animal models are not suitable for the replication of a clinical application or for proof-of-concept testing before clinical translation. Recent reviews in the literature have discussed the advantages and limitations of small- and large-animal models and procedural and experimental considerations vital to the implementation and evaluation of materials in these models. The use of any model for an intended clinical application requires that the strengths and weaknesses of any model be discussed in the development of the experimental strategy and plan.

Biomaterial and Device Perspectives in In Vivo Testing Two perspectives may be considered in the in vivo assessment of tissue compatibility of biomaterials and medical devices. The first perspective involves the utilization of in vivo tests to determine the general biocompatibility of newly developed biomaterials for which some knowledge of the tissue compatibility is necessary for further research and development. In this type of situation, manufacturing and other processes necessary to the development of a final product, that is, the medical device, have not been carried out. However, the in vivo assessment of tissue compatibility at this early stage of development can provide additional information relating to the proposed design criteria in the production of a medical device. While it is generally recommended that the identification and quantification of extractable chemical entities of a medical device should precede biological evaluation, it is quite common to carry out preliminary in vivo assessments to determine if there may be unknown or as yet identified chemical entities that produce adverse biological reactions. Utilized in this fashion, early in vivo assessment of the tissue compatibility of a biomaterial may provide insight into the biocompatibility and thereby may permit further development of the biomaterial utilized in a medical device. Obviously, problems observed at this stage of development would require further efforts to improve the biocompatibility of the biomaterial and to identify the agents and mechanisms responsible for the adverse reactions. As the in vivo assessment of tissue compatibility of a biomaterial or medical device is focused on the enduse application, it must be appreciated that a biomaterial considered compatible for one application may not be compatible for another. The second perspective regarding the in vivo assessment of tissue compatibility of medical devices focuses on the biocompatibility of the final product, that is, the fabricated medical device in the condition in which it is to be implanted. Thus, issues related to desired fabrications and interactions between biomaterials, etc., may come into play. Although medical devices in their final form and condition are commonly implanted in carefully selected animal models to determine function and biocompatibility, it may be inappropriate to carry out all of the recommended tests necessary for regulatory approval on the final device. In these situations, some tests may initially be carried out on biomaterial components of devices that have been prepared under manufacturing and sterilization conditions and other processes utilized in the development of the final product.

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Selection of Animal Models for In Vivo Tests Animal models are used to predict the clinical behavior, safety, and biocompatibility of medical devices in humans (Table 1). The selection of animal models for the in vivo assessment of tissue compatibility must consider the advantages and disadvantages of the animal model for human clinical application. Several examples follow, which exemplify the advantages and disadvantages of animal models in predicting clinical behavior in humans. A single test animal species may not assess all pertinent clinically important applications. For example, as described earlier, sheep are commonly used for the evaluation of heart valves. This is based on size considerations and also the propensity to calcify tissue components of bioprosthetic heart valves and thereby be a sensitive model for this complication. Thus, the choice of this animal model for bioprosthetic heart valve evaluation is made on the basis of possible accelerated calcification, the major clinical problem, assessed in rapidly growing animals, which has its clinical correlation in young and adolescent humans. Nevertheless, normal sheep may not provide a sensitive assessment of the propensity of a valve to thrombosis, which may be potentiated by the reduced flow seen in abnormal subjects, but diminished by the specific coagulation profile of sheep. The in vivo assessment of tissue responses to vascular graft materials is an example in which animal models present a particularly misleading picture of what generally occurs in humans. Virtually, all animal models, including nonhuman primates, heal rapidly and completely with an endothelial blood-contacting surface. Humans, on the other hand, do not show extensive endothelialization of vascular graft materials, and the resultant pseudointima from the healing response in humans has potential thrombogenicity. Consequently, despite favorable results in animals, small-diameter vascular grafts (< 4 mm in internal diameter) usually yield early thrombosis in humans, the major mechanism of failure that is secondary to the lack of endothelialization in the luminal surface healing response. Originally, the porcine coronary artery model was considered the model of choice for the evaluation of arterial stents. More recently, the rabbit iliac artery model for the evaluation of drug-eluting stents has been considered to be more realistic, as endothelialization is slower in the rabbit model than in the porcine model and inflammation is not as extensive in the rabbit. Thus, endothelialization, healing, and inflammation in the rabbit iliac artery model may be closer to these responses in humans than the porcine coronary artery model. Table 2 is a short, incomplete list of animal tissue/organ modifications that have been used in biomaterial development. This is considered to be incomplete but does offer a perspective on the variety of animals that have provided, in some cases, appropriate and adequate study responses.

Implantation Implantation tests assess the local pathological effects on the structure and function of living tissue induced by a sample of a material or final product at the site where it is surgically implanted or placed into an implant site or tissue appropriate to the intended application of the biomaterial or medical device. In some cases, the anatomic site of implantation used for biocompatibility evaluation is not the same as the site of ultimate use, but has representative mechanisms and consequences of tissue–biomaterial interaction

Table 1

Animal models for the in vivo assessment of medical devices

Device classification Cardiovascular Heart valves Vascular grafts Stents Ventricular assist devices Artificial hearts Ex vivo shunts Orthopedic/bone Bone regeneration/substitutes Total jointsdhips, knees Vertebral implants Craniofacial implants Cartilage Tendon and ligament substitutes Neurological Peripheral nerve regeneration Electric stimulation Ophthalmological Contact lens Intraocular lens

Animal Sheep Dog, pig Pig, dog, rabbit Calf Calf Baboon, dog Rabbit, dog, pig, mouse, rat Dog, goat, nonhuman primate Sheep, goat, baboon Rabbit, pig, dog, nonhuman primate Rabbit, dog Dog, sheep Rat, cat, nonhuman primate Rat, cat, nonhuman primate Rabbit Rabbit, monkey

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Animal tissue/organ modifications in biomaterial development

Table 2 Species

Tissue modification

Study response

Canine Mice Rat Mice Rabbit Canine Others

Female/male blood cross transfusion Muscle crush, ischemia, laceration, or toxin injection

Endothelium development on synthetic vascular grafts Muscle nerve and vascular regeneration

Critical defect in the bone

Bone regeneration and remodeling

(e.g., subcutaneous implantation in rodents of bioprosthetic heart valve materials to study calcification that occurs as a major clinical limitation in humans). The most basic evaluation of the local pathological effects is carried out at both the gross level and the microscopic level. Histological (microscopic) evaluation is used to characterize various biological response parameters. To address specific questions, more sophisticated studies may need to be done. Examples include immunohistochemical staining of histological sections to determine the types of cells present and studies of collagen formation and remodeling of the fibrous capsule. For short-term implantation evaluation out to 12 weeks, mice, rats, guinea pigs, or rabbits are the most common animals utilized in these studies. For longer-term testing in the subcutaneous tissue, muscle, or bone, animals such as dogs, sheep, goats, pigs, and other animals with relatively long life expectancy are suitable. If a complete medical device is to be evaluated, larger species may be utilized so that human-sized devices may be used in the site of intended application. For example, substitute heart valves are usually tested as heart valve replacements in sheep, whereas calves are usually the animal of choice for ventricular assist devices and total artificial hearts. In all aspects of biocompatibility testing, it is important to recognize that the effects of the material on the surrounding tissues are generally superimposed on the events occurring during physiological wound repair induced by the surgery of implantation. This is particularly important in shorter-term experiments. The use of canines in determining the blood compatibility of candidate materials in cardiovascular applications has been shown to be affected by the activity of the respective coagulation system in these animals. Early studies on the blood compatibility of vascular grafts reveal that some canines have a high level of coagulation activity, whereas other canines have a low activation potential. This broad range of blood coagulation activity in canines initially led to problems in the appropriate determination of blood compatibility of vascular graft materials. An issue such as this requires prior identification and characterization of the tissue activation potential of the respective animals utilized in any choice of animal model. Currently, the Food and Drug Administration (FDA) strongly suggests that the nonanticoagulated venous implants (NAVI) usually in the shape of a catheter be utilized in canines for periods up to 4 h and then gross evaluation of the amount of apparent thrombus on the surface be performed. Information derived from the NAVI model may be compromised by variation in the implant position, the implant technique, the extent of device–vessel wall contact and the time/incubation, the explant technique, the material/material surface, the nonthromboadherent materials that may be labeled nonthrombogenic, recipient/subject thrombotic potential, the statistical power, and the expertise of the evaluator utilizing this model. These caveats, although suggested for the NAVI model, hold for the choice of any model system utilized in the early biocompatibility evaluation of biomaterial and the later biocompatibility evaluation of medical devices containing component biomaterials.

Species Differences Species differences may play a significant role in the early determination of biomaterial biocompatibility and especially in the latter biological response evaluation with risk assessment of medical devices. While rats, mice, and rabbits are most commonly used for the early in vivo assessment of biomaterial biocompatibility, the in vivo assessment of tissue compatibility of the medical device usually requires larger animals. The limitations and the strengths and weaknesses of the choice of any animal model for biological response (biocompatibility) assessment must be considered in the selection of an appropriate animal model for medical device evaluation. In addition to the biological response evaluation of a medical device, the appropriate choice of an animal model may provide new insight into the biological interactions that may be discovered. For example, mini pigs have been most commonly used to evaluate the tissue response to the implantation of coronary stents. However, recent studies have demonstrated that the evaluation of drug-eluting stents in the atherosclerotic rabbit arteries is more predictive of neointimal progression and healing than normal rabbit iliac or porcine coronary arteries when compared with these responses in atherosclerotic human coronary arteries. While endothelial regrowth is significantly greater in the atherosclerotic rabbit model and occurs more rapidly than in humans, the time frame for these responses in the rabbit is much closer to that in humans than in mini pigs. These findings demonstrate two significant principles in the selection of the appropriate animal model. First, the healing response to coronary arteries in mini pigs is much faster than in rabbits that are much faster than in humans. Animal models

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may not adequately and appropriately predict the time-dependent nature of biological responses in humans. Secondly, the utilization of a diseased model may provide more significant information regarding the biological response assessment than when nondiseased animal models are utilized. However, rarely are diseased animal models used in biocompatibility, that is, biological response evaluation with risk assessment studies. Another example where the appropriate choice of animal models has revealed new information regarding the human response is studies from the early 1960s with the identification of blood-borne progenitor cells, presumably stem cells, in canines provided for the endothelialization of vascular grafts, and the endothelialization generally observed in animal studies with grafts and canines is not due to ingrowth either through the vascular graft or from the anastomosis of the vascular graft with the artery. Considering the early time frame, 1960s, these blood studies with canines were novel and innovative at that time. These canine studies utilize the identification of chromosomal markers in male recipients of vascular grafts after bone marrow radiation combined with exchange transfusions of female blood at the time of implantation of the vascular graft. The developed endothelial surfaces of these vascular grafts receiving female blood were identified to originate from the canine female blood. These cross transfusion studies in canines and the identification of the source of the endothelial cells may explain the lack of endothelialization in humans where blood-circulating progenitor cells, that is, stem cells, in humans are at a significantly lower concentration than that found in the various animal studies with vascular grafts that, in addition to canines, include rabbits, pigs, sheep, and significantly primates such as chimpanzees, baboons, and monkeys. These two examples clearly demonstrate that the appropriate choice of animal model may provide insight into biological responses and findings from such studies must be incorporated into the predictability of the animal model for human clinical application of the medical device. Infection is one of the most common complications associated with medical devices in clinical application. Some studies have been developed to address this common complication of infection and are considered to be significantly important in the development of wound dressings and tissue engineering. When selecting or designing an effective animal model, it is important to consider the host animal species and strain, the host animal effect, the pathogen species and strain, the inoculum concentration and vehicle, and many other factors specific to the disease state of interest. It is critical to fully explore previous animal models available in the literature and often necessary to conduct pilot studies to ensure an infection that has been created is self-sustaining but does not overwhelm the host. In this regard, specific attention to detail is required in the early development of the experimental plan and strategy for evaluation.

Animal Models for Host Response Mechanisms Table 3 provides an incomplete list of transgenic models in biomaterial development where the focus or target of the study has been the identification of mechanisms of interaction in the early inflammatory responses leading to final end-stage healing with fibrous capsule formation. These “knockout” animal models have been used to identify the participation of the various types of inflammatory cells that may contribute to the ultimate end-stage healing or the cytokines and other humoral factors that may play a significant role in the cellular and humoral responses leading to end-stage healing. Currently, this is an active area that is attempting to identify those important factors that may control or modulate the focal foreign-body reaction or the development of the final fibrous capsule that surrounds or encapsulates the implant. As such, the identification of these factors offers a target for intervention and inhibition of factors that may affect macrophage activation and adhesion to biomaterial surfaces. Macrophage fusion forms foreign-body giant cells that are adherent to the surface of the biomaterials or the inhibition of fibrous capsule formation. These are important considerations in the subsequent development of biosensors where the focal foreign-body reaction and the fibrous capsule have been inhibited in the development of appropriate sensors for clinical application.

Table 3

Transgenic models in biomaterial development

Species

Genetic modification

Biological response

Mice Mice Mice Mice Mice

Inhibits macrophage fusion (FBGC/osteoclast formation) No macrophage fusion (FBGC formation) Reduced accumulation and fusion reduced MMP-9 Nb IL-1b secretion, reduced scaffold cell, infiltration and phenotypes Deficient in cytotoxic T lymphocytes and IL12, induction, reduced IL6 and G-CSF

Mice

DAP12 deficient Syk deficient Monocyte chemoattractant protein-1 (MCP-1) NIrp3/ inflammasome Bat 3/ CD8 (þ) dendritic cell knockout TGF-b receptor II knockout

Mice Mice Mice Mice Mice

SCID muMT Reg2/IL2ry a-MF a-CXCL13

Photoactivation of latent transforming growth factor-b1 on dental stem cell differentiation Severe combined immunodeficiency B-cell deficient, decreased fibrosis T- and B-cell deficient, no fibrosis, macrophage dysfunction Macrophage depletion or inhibition, no fibrosis Decreased B cells, decreased fibrosis

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Knockout animals provide insight into molecular mechanisms and potential targets for the intervention of selected host responses. However, it must be understood that knockout and transgenically modified animal models may have other molecular and cellular host responses that are modified but not identified in the host response assessment. Therefore, caution is recommended in the use of these systems, and the experimental study should be as comprehensive as possible to identify all possible host responses. Particular attention should be given to the biocompatibility of bioactive biomaterials and tissue-engineered systems.

Further Reading An, Y. H., & Friedman, R. J. (1999). Animal models in orthopedic research. Boca Raton, FL: CRC Press. Anderson, J. M., & Schoen, F. J. (2013). In vivo assessment of tissue compatibility. In B. D. Ratner, A. S. Hoffman, F. J. Schoen, & J. E. Lemons (Eds.), Biomaterials science, an introduction to materials in medicine (3rd edn, pp. 609–617). Oxford, England: Elsevier Academic Press. Boutrand, J.-P. (Ed.). (2012). Biocompatibility and performance of medical devices. Oxford: Woodhead Publishing Limited. Nakazawa, G., Nakano, M., Otsuka, F., Wilcox, J. N., Melder, R., et al. (2011). Evaluation of polymer-based comparator drug-eluting stents using a rabbit model of iliac artery atherosclerosis. Circulation. Cardiovascular Interventions, 4, 38–46. Shah, S. R., Young, S., Goldman, J. L., Jansen, J. A., Wong, M. E., & Mikos, A. G. (2016). A composite critical-size rabbit mandibular defect for evaluation of craniofacial tissue regeneration. Nature Protocols, 11(10), 1989–2009. Tatara, A. M., Shah, S. R., Livingston, C. E., & Mikos, A. G. (2015). Infected animal models for tissue engineering. Methods, 84, 17–24. Veiseh, O., Doloff, J. C., Ma, M., Vegas, A. J., Tam, H. H., et al. (2015). Size- and shape-dependent foreign body immune response to materials implanted in rodents and nonhuman primates. Nature Materials, 14(6), 643–651. Wolf, M. F., & Anderson, J. M. (2012). Practical approach to blood compatibility assessments: general consideration and standards. In J.-P. Boutrand (Ed.), Biocompatibility and performance of medical devices (pp. 159–200). Oxford: Woodhead Publishing Limited. Yamamoto, S., Urano, K., Koizumi, H., Wakana, S., Hioki, K., et al. (1998). Validation of transgenic mice carrying the human prototype c-Ha-ras gene as a bioassay model for rapid carcinogenicity testing. Environmental Health Perspectives, 106(Suppl. 1), 57–69. Yang, Y., Jao, B., McNally, A. K., & Anderson, J. M. (2014). In vivo quantitative and qualitative assessment of foreign body giant cell formation on biomaterials in mice deficient in natural killer lymphocyte subsets, mast cells, or the interleukin-4 receptor f and in severe combined immunodeficient mice. Journal of Biomedical Materials Research. Part A, 102A, 2017–2023.

Blood–Biomaterial Interactions Nicholas P Rhodes, University of Liverpool, Liverpool, United Kingdom © 2019 Elsevier Inc. All rights reserved.

Introduction Biology of Blood In Vitro Observations Plasma Proteins Coagulation and Contact Phase Activation Platelets Leukocytes Complement Activation In Vivo Observations Rheological Factors Surface Topography Further Reading

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Glossary Biocompatibility The ability of a biomaterial to perform its desired function without eliciting undesirable local or systemic effects but generating the most appropriate response in that specific application. Complement system activation The activation of a protein cascade representing the noncellular immune system that ultimately results in the formation of a cell destruction complex. Contact-phase activation The surface-mediated autoactivation of factor XII (or Hageman factor) resulting in the initiation of the coagulation cascade which ultimately results in blood plasma clotting. Platelet activation Formation of an active form of platelets resulting from the attachment of surface-phase fibrinogen and collagen to glycoprotein receptors GPIIb/IIIa and Ib respectively, resulting in platelet spreading, the secretion of platelet granules and expression of procoagulant lipid on the plasma membrane of the platelets.

Introduction Researchers have, for many decades, expected that the performance of future blood-contacting biomaterials would match that of nonthrombogenic, natural endothelium, that is, with no blood coagulation and without blood cell adhesion or activation. Despite the wealth of knowledge accumulated regarding the interfacial phenomena of surfaces with protein mixtures and cells, it has not been possible to design such a surface, and the interactions of blood elements with biomaterials placed intravascularly can cause significant morbidity. Coronary arteries, for example, are routinely bypassed using saphenous vein autografts, as synthetic alternatives cannot be used. The failure of the medical engineering community over more than half a century to meet this challenge may, in part, be due to the lack of understanding that the phenomena observed in vitro, such as coagulation cascade protein activation or platelet adhesion, may be of relative unimportance in vivo at particular anatomical locations within the circulation and the fact that the response of vascular endothelium to intravascular biomaterial implantation is rarely considered. Described in this article are the in vitro phenomena usually considered when describing blood–biomaterial interactions and a selection of in vivo reactions that are often observed after intravascular placement of blood-contacting devices. Although not an exhaustive list, the range of possible effects due to poor blood compatibility of an intravascular device can be listed as:

• • • • • • • • • •

Thrombosis Embolism Thrombocytopenia and leukopenia Immune and allergenic reactions Infection Inflammation Fibrosis Material degradation Tumorigenesis Mutagenicity

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Thus, coagulation and thrombosis continue to be problematic in certain blood-contact applications, as no generalized theory currently exists that relates physical or chemical surface features to in vivo performance of a biomaterial. Introduction of a biomaterial into blood triggers a complex cascade of interrelated reactions that may, or may not, relate to the long-term performance of that biomaterial. The picture is made more complex by the recent discovery of surface interactions with biological molecules not previously identified, and the greater complexity of coagulation cascade protein cleavage than previously hypothesized.

Biology of Blood Blood is composed of red blood cells (RBCs), also known as erythrocytes, white blood cells (or leukocytes) and platelets (originally known as thrombocytes), proteins, ions, lipids and carbohydrates in an aqueous medium (Tables 1 and 2). RBCs are both physically large and great in number, comprising almost half of the volume of blood. Their role in blood–biomaterial interactions is largely limited to transport phenomena in flowing blood. Platelets are physically small, outnumbered by RBCs by a factor of 20:1 and so represent a small percentage of blood volume, but play a central role in the control and propagation of blood reactions, interacting with the coagulation system proteins. Leukocytes are similar in size to RBCs but represent only 0.1% of the total blood cell number. These interact with the host’s immune and inflammatory systems and play an important role in reactions with the endothelium. There are several hundred different plasma proteins within blood, but albumin, fibrinogen and immunoglobulins represent 98% of the total protein mass. There are a number of important protein systems: the coagulation cascade, complement cascade, protease inhibitors and fibrinolytic system being predominant in potentiating the reactions of blood with biomaterials.

In Vitro Observations Developers of intravascular biomaterials must initially test the biocompatibility of their designs rigorously in vitro. So while phenomena that can be observed in vitro may not be clinically relevant, material scientists have a requirement to perform a basic set of tests that may point to aspects of a biomaterial’s performance in vivo. It clearly follows that the surface chemistry, energetics and morphology of a biomaterial will have some bearing on its long-term clinical performance, but only when the significance of all the component parts is fully understood will a quantum leap in device performance be possible.

Plasma Proteins All cellular interactions with implanted biomaterials are controlled by a layer of proteins formed after the initial contact of the material with a biological fluid. When a biomaterial is placed into blood, dissolved plasma proteins adsorb to its surface within milliseconds, following a small quantity of water and inorganic ion adsorption. This is largely due to the polar nature of amino acids, the relative lack of their solubility and the nonhydrated nature of the majority of implantable biomaterials. Three-dimensional rearrangement of the plasma proteins allow the hydrophobic domains to retreat towards the material surface, hydrophilic domains facing the aqueous milieu. Typical blood-contacting polymers have a relatively narrow range of surface energies, that is they are moderately hydrophobic. Placing materials that are very much more hydrophilic, such as hydrogels, or extremely hydrophobic, can have a dramatic influence on the course of protein adsorption and subsequent cellular events. Researchers have over the years attempted to reach conclusions about the best, or an acceptable range of surface energy a biomaterial should possess, but unfortunately, their conclusions vary massively, and in some cases directly contradict each other. That is, some studies have concluded that long-term biocompatibility improves with increasing surface energy, while other studies conclude the opposite. Hydrophobic materials, however, tend to absorb more protein molecules than a surface which is correspondingly hydrophilic.

Table 1

Cellular composition of blood

Cell type

Approx. diameter (mm)

Approx. concentration in blood (number per mL)

Red blood cell Platelet Leukocyte Lymphocyte Monocyte Neutrophil Basophil Eosinophil

7.5 3 6–10 7 9 8 8 8

5000
106 250
106 7.5
106 2
106 0.5
106 5
106 0.2
106 0.1
106

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Noncellular composition of blood

Constituent

Approx. concentration in plasma (g per L)

Proteins Albumin Immunoglobulins Fibrinogen Ions Lipids Carbohydrate

720 56% of total protein 38% of total protein 4% of total protein 98 83 17

The composition of the protein layer is initially dependant on the concentration of proteins within blood, that is albumin, fibrinogen and immunoglobulins predominate. Where a surface has higher affinity for proteins that are less abundant, including those involved in the coagulation or complement cascades (e.g., C3 or factor XII), the concentration of proteins in that initial layer is gradually reduced in favor of those higher-affinity proteins due to the random arrivals and departure of individual molecules, a process known as the ‘Vroman Effect’. This change in the identity of proteins adsorbed to the biomaterial surface results in selected proteins to become susceptible to cleavage and others that are able to form bonds with cell surface receptors, thereby causing cellular activation. Fibrinogen, when adsorbed in the correct conformation, will cause the adhesion of platelets, as described in detail below, and subsequently lead to their activation, an interaction that is not generally observed, at least initially, on surfaces that preferentially adsorb albumin. Thus, the actual concentrations of specific proteins adsorbed on a biomaterial surface is of less importance than the precise conformational rearrangements they subsequently adopt, which develops dynamically over time. The specificity of the interactions required for cleavage and receptor activation, as described above, are difficult to predict, making the design of surfaces that have good blood compatibility challenging to achieve.

Coagulation and Contact Phase Activation Within the closed environment of a blood vessel in which a biomaterial has been placed, coagulation is initiated when factor XII (fXII) is cleaved to form fXIIa, a process known as contact phase activation (Fig. 1) of the intrinsic coagulation cascade. This reaction occurs when fXII, a serine protease in an inactive form, adopts a conformation that renders it susceptible to cleavage. The three principal routes of activation are autoactivation (i.e., activation by other molecules of fXIIa), activation of fXII by kallikrein that has been formed by the cleavage of prekallikrein by other fXIIa molecules, both of which are surface-mediated reactions the rate of which are dependent on the procoagulant nature of that surface, and by activated platelets, which could be either a surface or cell-mediated process. Other physiological activators of fXII include platelet-derived microparticles, acellular RNA molecules and collagen. The precise mechanisms of fXII cleavage have been intensely debated over the past decades. Most seminal literature concluded that negatively charged surfaces are most liable to activate fXII although recent studies have now shown this to not be the case. In fact, these studies have shown that both highly hydrophobic and highly hydrophilic surfaces are potent activators. Many textbooks indicate that fXII can be cleaved to form two separate activated forms, a small b-fXIIa cleavage product, and a larger activated fragment denoted a-fXIIa, both able to cleave factor XI. Current research, however, has shown that there at least 6, and possibly an even greater number of cleavage points on fXII, depending on the activator, and resulting in the production of multiple protein fragments of fXII due to autoactivation, some which initiate plasma coagulation, and some that do not. The major confusion in defining how susceptible a biomaterial is to initiating plasma clotting that has persisted for more than two decades has been due to the use of a popular method of quantifying contact activation. Some fragments of fXII that do not coagulate plasma have been found to activate a widely used chromogenic substrate assay, utilizing S-2302. Finally, activated fXII fragments also appear to undergo a form of autoinhibition and suppression of autoactivation, although the mechanisms of this remain unclear. The main feature of the coagulation pathway is the cascading effect of factors, whereby small concentrations of initiating stimulus generate large concentrations of thrombin through a series of reactions of ever increasing magnitude. Each factor circulates within blood in an inactive state but is then cleaved by its activator (the previous factor in the pathway), conveying proteolytic activity on it. Each active factor is a serine protease, whose active site is composed of the same amino acid sequence: Gly-AspSer-Gly-Gly-Pro which cleaves an Arg-X bond within the inactive factor. Despite the similarity of the active sites, the factors are highly specific, the three-dimensional architecture and chemistry of the molecules providing steric and intramolecular forces which are important in coagulation factor recognition. The cascade is controlled by the anticoagulant and fibrinolytic pathways, independent mechanisms which prevent excess generation of fibrin. There are also a variety of protein inhibitors, the most important being Antithrombin III, C1-esterase inhibitor, a2macroglobulin and a1-antiproteinase inhibitor. These react with active factors and inhibit them. The activity of antithrombin III is enhanced by its interaction with heparin.

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Fig. 1

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Intrinsic coagulation pathway.

Not shown in Fig. 1 is the extrinsic pathway, which has physiological relevance in the prevention of hemorrhage. Both pathways activate factor X and have a common terminal pathway, but with different initiating steps.

Platelets Platelets are small, anucleate cells that undergo a variety of interactions which, physiologically, are central to hemostasis and an element of the inflammatory response, broadly described as: adhesion, aggregation, activation and microparticle shedding. The ability of a biomaterial to influence the magnitude of any of these interactions is critically dependant on the rate of collision of the platelets with the surface, as discussed in “Rheological Factors” section. However, the following interactions have always served as a central part of any blood compatibility assessment of a biomaterial. Platelets become adhered to biomaterials predominantly via the reaction of adsorbed fibrinogen with the aIIbb3 integrin receptor, also known as GPIIb/IIIa. Platelets also possess surface receptor affinity for many other adhesive proteins, including collagen, fibronectin, vitronectin and von Willebrand’s factor (vWf). In its natural conformation, fibrinogen does not react with this receptor, but adsorption to a surface can expose a binding site for resting aIIbb3 on platelets. Fibrinogen has a total of 4 potential platelet-binding sites, two 12 amino acid sites (HHLGGAKQAGDV), the most commonly-observed binding sites, and two RGD sequences. Due to their small physical size, the adhesion of platelets onto a biomaterial rarely presents a problem in its own right, but subsequent reactions can initiate thrombosis, embolism, inflammation and immune responses. Platelet activation occurs when the level of stimulation of aIIbb3 is great enough and is characterized by the influx of extracellular calcium into the cytoplasm and release of the contents of alpha, dense and lysosomal granules via the open canalicular system to the extracellular medium. Alpha granules contain coagulation factors, chemotactic and growth factors which allow interaction with inflammatory cells and proteins. Actin filaments are mobilized within the platelet cytoskeleton causing spreading and expression of granule membrane proteins such as P-selectin (CD62P) and GP-53 (CD63) on the platelet surface, potential ligands for the adhesion of leukocytes. Prior to the full activation of platelets, however, a procoagulant stage is expressed when the membrane phospholipids convert to phosphatidylserine, which supports the complexation of coagulation factors on the surface, supporting the propagation of the clotting cascade. Strong activation of the coagulation cascade, even in flowing blood, can result in occlusion of the vessel.

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Platelet aggregation occurs when two or more platelets are joined together via fibrinogen or vWf. This type of reaction is observed in conditions of high shear stress when a platelet receptor GPIb-binding domain is exposed on vWf. The vWf-GPIb interaction causes a stimulation of the aIIbb3 receptor via internal signaling which allows it’s interaction with soluble and nonconformationally disturbed fibrinogen, and vWf. Large aggregates are potentially lethal by forming emboli within the narrow capillary structures of major organs and the brain. Platelets also produce microparticles, small vesicles of procoagulant lipid membrane which also express some platelet receptors, including aIIbb3. These have a volume many orders of magnitude smaller than a platelet, but have the potential for propagating reactions in the coagulation cascade, inflammatory and immune responses, in much the same way that do platelets. The release of microparticles has generally been associated with the activation of aIIbb3 by fibrinogen, or stimulation of the platelet by strong agonists. However, shear forces alone can also be responsible for the shedding of microparticles which express procoagulant activity. The significance of microparticles is that their behavior in flowing blood is different from that of platelets. Evidence suggests that the cyclo-oxygenase pathway is not utilized in the shedding of microparticles, but rather the phosphorylation of protein kinase C. Treatments such as aspirin therapy do not, therefore, affect the production of microparticles, nor is it a calcium-dependant mechanism.

Leukocytes The adhesion of white blood cells (leukocytes) onto an implanted biomaterial is normally considered to be controlled by the generation of activated fragments of complement protein at the biomaterial surface. Many in the field of blood compatibility do not consider the interaction of white blood cells with biomaterials of any great importance. This view has largely come about due to the observations of profound platelet activity after the contact of a biomaterial with blood in vitro whilst at the same time failing to observe any dramatic leukocyte activity. However, when an implant is deployed in vivo it is not possible for there to be solely blood contact without its interaction with blood-derived leukocytes in the surrounding tissue. Activation of resident macrophages and blood leukocytes recruited to the implantation site will ultimately lead to inflammation and the generation of a thrombosis via activation of the endothelium. Activation of leukocytes and complement in blood will often result in the same cycle of events occurring. Release of species such as cathepsin G from granulocytes directly activate platelets, whilst elastase secretion results in fibrinolytic activity. A thrombosis that is observed in vivo is generally attributed to the action of the coagulation cascade and platelets, whereas there is mounting evidence that the initial impetus for the reaction is often initiated outside of the circulation, a result of inflammation in the tissue surrounding the blood vessel.

Complement Activation The system of complement proteins represents the humoral arm of the innate immune system. In a process similar to coagulation, a number of factors interact without cellular involvement in such a manner as to generate a multimeric complex called MAC (membrane attack complex), comprising C5b, C6, C7, C8 and C9. This structure resembles a tube and is naturally energetically able to reside within the lipid structure of a cell membrane. This neatly causes egress of the cell cytoplasm contents and results in the death of the cell. Thus pathological activation of the complement cascade is not a desirable outcome of placing a biomaterial into blood. Activation of complement proteins on biomaterials results in the adhesion and activation of leukocytes, and the possible involvement of an immune response. The complement system is split into alternate and classical pathways, but the involvement of the alternate pathway is more important when considering the reaction of blood with biomaterials. The alternate pathway relies on the continuous hydrolysis of C3. The fragment C3b thus formed will adsorb to a nearby surface. The conformational rearrangements which take place if the surface is complement protein activating allow the attachment of factor B and the initiation of the pathway. The conformation that C3b adopts when adsorbed on a nonactivating surface allow the attachment of factor H, leading to pathway inactivation. The presence of C3b on the surface of a biomaterial is thought to be the main focal point for the adhesion of neutrophils, via the Mac-1 receptor. A number of strategies have been proposed to limit complement activation on blood-contacting biomaterials. The development of heparin coatings has been utilized extensively, which aim to capture factor H and so inactivate any pathway activation. However, heparin has been shown to have some undesirable consequences following its interaction with a number of plasma proteins, and ultimately, activated complement fragments are nevertheless produced by such surfaces. Creating nanostructures that sterically inhibit the formation of activated complement complexes on the surface have shown some success. In one example, reducing the diameter of surface pores from 200 to 20 nm successfully reduced complement activation, despite the surface area of the biomaterial being greater in the small-pore version. Each surface area both inside and around the pores was simply too small for the molecular assembly of the C3 convertase.

In Vivo Observations Biomaterials and vascular devices placed in the circulation that fail overwhelmingly suffer from thrombosis or embolism, both events resulting from the activation of the coagulation cascade and involving the activation of platelets. There are

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a number of obvious differences between in vitro and in vivo testing of a biomaterial, for example, fresh blood continuously arrives at the site of blood-material contact during in vivo contact, resulting in the dilution of activated factors or cells in their unactivated form. In addition, the blood-material interaction is combined with a response from endothelial cells and other extrinsic factors. The usual definition of biocompatibility refers to the ability of a biomaterial to successfully fulfill its function in a particular application at the anatomical site relevant to that device. For large-diameter vascular grafts, Dacron has successfully been utilized as a vascular graft for many years, despite experiments in vitro demonstrating that Dacron has rather unremarkable biocompatibility. However, the same material cannot be used for the replacement of a coronary artery due to rapid occlusion observed in experimental animals. Thus, important factors that must be considered when designing a material for use in a blood-contacting application in vivo are protein adsorption leading to coagulation cascade activation, platelet adhesion, activation and aggregation and activation of the endothelium, but taking into account the dynamics of blood flow and the rate of reactant turnover. Studies have also shown that thrombosis can be initiated without contact activation of the coagulation cascade. In addition to these, the importance of a number of additional factors in thrombosis has recently emerged. Polyphosphates having both short and long-chain lengths (100–1000 phosphates) are secreted from both platelet dense granules and microorganisms, respectively. Microbial polyphosphates, having a large charge density are potent autoactivators of fXII, leading to rapid contact-activation. Cell-free nucleic acids similarly activate fXII and are secreted from microorganisms and apoptotic or necrotic host cells. Activated neutrophils release fibers comprising a structure knows as Neutrophil extracellular traps, intended to entrap microbes but which are highly prothrombotic due to the presence of DNA-histones that simultaneously activate fXII and platelets, and tissue factor that is a potent extrinsic activator of the coagulation cascade. Any placement of a vascular biomaterial within the circulation necessarily results in a degree of damage to the endothelium. Such damaged endothelial cells release tissue factor leading to the activation of the extrinsic coagulation cascade. Inflammation within the tissue surrounding the site of implantation results in the expression of intercellular adhesion molecule 1 (ICAM-1) and vascular cell adhesion molecule 1 (VCAM-1) on the endothelial cells following the expression of TNFa and IL-6 within the soft tissue outside the circulation. Upregulation of these adhesion molecules results in the extraction of leukocytes through a process known as rolling and diapedesis, which perpetuates the inflammatory reaction within the external tissue. Activation of the endothelium also causes reduced secretion of prostacyclin, leading to reduced resistance to thrombosis.

Rheological Factors The type of reaction observed when a biomaterial is placed in blood is dependent on the level and type of blood flow that is experienced. Blood is a non-Newtonian fluid and under steady flow conditions will result in a greater concentration of red blood cells within the central portion of the flow profile. This has the effect of pushing platelets to the outer wall of the vessel, making contact with a biomaterial more likely. In steady state flow within a tube (the blood’s natural environment), platelet adhesion onto a surface will be reaction rate limited, that is the number of platelets adhering will be proportional to the affinity of the cells to the surface. At physiological wall shear rates the effect of the flow will not in itself be a source of cell activation. The distribution of velocities within the vessel has the effect of unfolding vWf, which allows platelets to aggregate via the interaction of vWf with platelet receptor GPIb/IX. At very high wall shear rates the shedding of platelet-derived microparticles is observed. When the flow is disturbed (turbulent), however, the degree of platelet activation, aggregation and microparticle shedding is dramatically increased. The presence of high blood flow rates has the effect of diluting activated cells and proteins, preventing thrombosis. However, these conditions increase the probability of embolism occurring, where small fragments of clots that begin to form, especially in combination with highly prothrombotic platelet-derived microparticles, break away and become coagulated at distant sites in low flow environments. This is particularly dangerous within the capillary beds of organs such as the lungs or brain.

Surface Topography Generally, surface features on biomaterials give rise to increased material surface area, resulting in protein and cellular reactions that are increased relative to a smooth surface of the same biomaterial. Roughness on the surface of a static device within flowing blood is likely to capture cells that would otherwise flow downstream away from the site of activation. Thus, in most cases, surface features are not helpful in improving the biocompatibility of blood-contacting devices. However, there are some notable exceptions. First of all, surfaces that are mobile, in which, for example, long side chains of polyethylene glycol, constantly move in relation to one another and so provide an unstable foundation for the complexation of procoagulant peptides, thus preventing coagulation and possibly complement activation occurring. Nanotopographical structures could be used to more precisely control the adsorption of specific proteins, in defined orientations and conformations. However, such investigations are at an early stage of research, but may in future provide better control of longterm performance of blood-contacting biomaterials.

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Further Reading Baier, R. E., & Kurusz, M. (2012). Understanding blood/material interactions: Contributions from the Columbia University biomaterials seminar. ASAIO Journal, 58, 450–454. Bauer, J., Xu, L.-C., Vogler, E. A., & Siedlecki, C. A. (2017). Surface dependent contact activation of factor XII and blood plasma coagulation induced by mixed thiol surfaces. Biointerphases, 12, 02D410-1-10. Ekdahl, K. N., Lambris, J. D., Elwing, H., Ricklin, D., Nilsson, P. H., et al. (2011). Innate immunity activation on biomaterial surfaces: A mechanistic model and coping strategies. Advanced Drug Delivery Reviews, 63, 1042–1050. Golas, A., Yeh, C.-H. J., Siedlecki, C. A., & Vogler, E. A. (2011). Amidolytic, procoagulant, and activation-suppressing proteins produced by contact activation of blood factor XII in buffer solution. Biomaterials, 32, 9747–9757. Kalathottukaren, M. T., & Kizhakkedathu, J. N. (2018). Mechanisms of blood coagulation in response to biomaterials: Extrinsic factors. In C. A. Siedlecki (Ed.), Hemocompatibility of biomaterials for clinical applications (pp. 29–49). Elsevier. Xu, L.-C., Bauer, J., & Siedlecki, C. A. (2014). Proteins, platelets, and blood coagulation at biomaterial interfaces. Colloids and Surfaces. B, Biointerfaces, 124, 49–68.

Interaction Between Mesenchymal Stem Cells and Immune Cells in Tissue Engineering Rong Huang, Yinghong Zhou, and Yin Xiao, Queensland University of Technology, Brisbane, QLD, Australia © 2019 Elsevier Inc. All rights reserved.

T Cell B Cell Macrophage Natural Killer (NK) Cell, DC Cell and Neutrophil References Relevant Websites

251 252 252 253 254 256

Glossary Antigen-presenting cells Cells which displays antigen complexed with major histocompatibility complexes on their surfaces and play an important role in adaptive immune response. Cognate immunity Also known as the acquired immunity or adaptive immunity, which creates immunological memory after an initial response to a specific pathogen, and leads to an enhanced response to subsequent encounters with that pathogen. Histocompatibility Compatibility between the tissues of different individuals, so that one accepts a graft from the other without giving an immune reaction. Major histocompatibility complex A set of cell surface proteins essential for the acquired immune system to recognize foreign molecules in vertebrates, which in turn determines histocompatibility. Non-immunogenic Not capable of inducing an immune response; not antigenic. Paracrine action Relating to or denoting a hormone which has effect only in the vicinity of the gland secreting it.

Abbreviations ALP Alkaline phosphatase APCs Antigen-presenting cells C3 Complement component-3 cAMP Cyclic adenosine monophosphate CCL C–C motif chemokine ligand CD Cluster of differentiation CD45RO Cluster of differentiation 45 isoform DAMPs Damage-associated molecular patterns ELISA Enzyme-linked immunosorbent assay EP Prostaglandin E receptor FGF Fibroblast growth factors GM-CSF Granulocyte/macrophage colony-stimulating factor GvHD Graft versus host disease HLA-DR Human leukocyte antigens HLA-G5 Human leucocyte antigen-G5 IDO Indoleamine 2,3-dioxygenase IFN-g Interferon gamma Ig Immunoglobulin IL Interleukin LPS Lipopolysaccharide M1 Classically activated macrophages M2 Alternatively activated macrophages M-CSF Macrophage colony-stimulating factor MHC Major histocompatibility complex MMP Matrix metallopeptidase

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MSCs Mesenchymal stem cells MyD88 Myeloid differentiation factor 88 NK cells Natural killer cells PAMPs Pathogen-associated molecular pattern molecules PBMCs Peripheral blood mononuclear cells PD-1 Programmed death 1 PGE2 Prostaglandin E2 RT-PCR Real-time reverse transcription-polymerase chain reaction TC cells/CTLs Cytotoxic T cells TCR T cell receptor Teff cells Effector T cells TGF-b Transforming growth factor-beta Th cells T helper cells TLR Toll-like receptor TNF-a Tumor necrosis factor-alpha Tol DC Tolerogenic DC TSG Tumor necrosis factor-inducible gene

The definition of tissue engineering evolved from biomaterials covers interdisciplinary fields involving basic principles of life science and engineering to replace or improve tissue structure and function (Langer and Vacanti, 1993). Currently, the treatment for cardiovascular failure (Surder et al., 2013; Yerebakan et al., 2011), liver diseases (Kharaziha et al., 2009), degeneration of the spinal cord (Frolov and Bryukhovetskiy, 2012; Cashman et al., 2008), and the use of substitutes to repair damaged tissues, such as cartilage (Wakitani et al., 2011) and bone (Nagata et al., 2012) have offered insights to the tissue engineering and regenerative medicine. Although tissue function can be restored to a certain level using conventional modalities, some drawbacks such as limited donor tissue, implant rejection, and procedure-related complications remain major challenges for the first-line clinicians (Herford and Dean, 2011; Paul et al., 2009). One of the highlights of tissue engineering is to repair and regenerate without donor site morbidity which is different from the traditional surgery that needs to sacrifice tissues from the donor site (Dimitriou et al., 2011). Extensive studies have confirmed the feasibility of generating bone to repair large defects (Viateau et al., 2007; Liu et al., 2008; Cao et al., 2012), and the outcomes of other ongoing clinical trials are quite promising (Ardjomandi et al., 2015; Yamada et al., 2008). To date, the use of autografts serves as a gold standard for tissue repair and regeneration. Combined with the cell-based therapy with autologous cells that can be manipulated to repair or replace damaged tissues, the field of tissue engineering is being explored with less ethical issues (Dimitriou et al., 2011). Recent advances in regenerative medicine, in particular the stem cell biology have enabled the development of stem cell-based therapies which are histocompatible and non-immunogenic. For example, bone regeneration has shown that more reliable and effective results have been found in the management of bone defects using stem cell delivery compared to the traditional methods (Lo et al., 2012; Dupont et al., 2010). Mesenchymal stem cells (MSCs) is one of the useful cell types which can be easily isolated from patients and transplanted back with less ethical issues. The generation of sufficient MSCs for transplantation can be simply achieved by in vitro expansion, suggesting the possibility of their broad application in cell-based therapy (Kuroda, 2016; Chanda et al., 2010). The anti-inflammatory and immunosuppressive properties (MacFarlane et al., 2013; Kuo et al., 2012), migratory ability to defect sites (Mokbel et al., 2011; Sordi, 2009) and capability to modulate the microenvironment by paracrine actions (Meyerrose et al., 2010) make them a promising candidate for tissue engineering. MSCs are fibroblast-like plastic cells with self-renewal and pluripotency. After their initial discovery in bone marrow, MSCs have been identified from a wide range of tissues including dental pulp, umbilical cord, adipose tissue, skin, periosteum, muscle, skeleton, liver, spleen, brain, tendon, and synovial. They are positive for expression of stem cell surface markers such as cluster of differentiation 73 (CD73), cluster of differentiation 90 (CD90) and cluster of differentiation 105 (CD105) and negative for expression of the hematopoietic cell surface markers such as cluster of differentiation 11 (CD11), cluster of differentiation 34 (CD34), cluster of differentiation 45 (CD45), cluster of differentiation 79a (CD79a) and major histocompatibility complex (MHC) class II cell surface receptor encoded by human leukocyte antigens (HLA-DR). They own the ability to differentiate into osteoblasts, chondrocytes, adipocytes and many other cell types (Kolf et al., 2007; Pittenger et al., 1999). MSCs have been well characterized with their effects on immune disorders, including graft versus host disease (GvHD), type I diabetes, rheumatoid arthritis, and inflammatory bowel diseases (Hinsenkamp et al., 2012; Polchert et al., 2008; Urban et al., 2008; Hayashi et al., 2008). It is reported that MSCs secret proinflammatory cytokines such as interleukin (IL)-1, interleukin (IL)-6, interleukin (IL)-8 and tumor necrosis factor-alpha (TNF-a), which are involved in inflammation-related diseases (Kim et al., 2005). Meanwhile, MSCs are able to inhibit T cell and NK cell activity and play an immunosuppressive role in dendritic cell differentiation and maturation (Rasmusson et al., 2003; Zhang et al., 2004). Most importantly, MSCs are non-targeted by MHC-mismatched immune cells, establishing an environment free from cellular attack.

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The mechanisms of MSC immunoregulation are based on two paradigms. First of all, MSCs can sense the microenvironment change and act accordingly. Secondly, they become polarized toward different phenotypes depending on the toll-like receptor (TLR) signals received (Auletta et al., 2012; Waterman et al., 2010). MSCs can regulate the survival and activation of immune cells on one hand; on the other hand, they can inhibit inflammation and promote tissue repair (Fig. 1). When tissue injury occurs, MSCs migrate to the injury site where they are directly activated through TLR stimulation including pathogen-associated molecular pattern molecules (PAMPs) or damage-associated molecular patterns (DAMPs) (Prockop and Youn Oh, 2012). MSCs are indirectly activated by several pro-inflammatory cytokines such as TNF-a and interferon gamma (IFN-g) which bond to MSC receptors. The activation of complement component-3 (C3) is triggered to target MSCs for complement lysis in normal circumstances. However, MSCs express a complement regulatory protein cluster of differentiation 59 (CD59) and release complement factor H to protect them from complement lysis. In response to the pathogen, MSCs either help clear pathogen or modulate immune responses depending on the stimuli received (Fig. 1). MSCs may promote pathogen clearance through secretion of IL-6 and IL-8, evoke the polarization of the pro-inflammatory phenotype of macrophages, increase bacterial clearance through the enhanced survival and function of neutrophils and induce the apoptosis of effector T (Teff) cells. In addition, MSCs are reported to regulate immune responses through the secretion of immunosuppressive soluble factors and evoke the polarization of the anti-inflammatory phenotype of macrophages (M2), tolerogenic DC (Tol DC) and Tregs (English, 2013). The following article will introduce the interaction between MSCs and immune cells to further understand the immunoregulatory properties of MSCs in tissue engineering.

T Cell Lymphocytes originate in the bone marrow and migrate to other parts of the lymphatic system such as the thymus, lymph nodes and spleen. In our immune system, T cells are the primary lymphocyte effectors and their functional properties are central to antigen specificity and memory associated with cognate immunity. T cells consist of T helper cells (Th cells), cytotoxic T cells (TC cells or CTLs), memory T cells, Tregs and a few other categories. T helper cells are also called cluster of differentiation 4 þ (CD4 þ) T cells because they express the CD4 glycoprotein on the cell surfaces. Antigenic peptides bound to self MHC class II molecules which are expressed on the surface of antigen-presenting cells (APCs) induce the activation of T helper cells. Once activated, T helper cells divide rapidly and secrete cytokines that assist other immune cells in immunologic processes, including the differentiation of B cells into plasma cells and memory B cells, as well as the activation of macrophages and cytotoxic T cells (Gutcher and Becher, 2007). Cytotoxic T cells are known as CD8 þ T cells because they express the CD8 glycoprotein on the cell surfaces. They are responsible for destroying virus-infected cells and tumor-related cells, and are also implicated in transplant rejection. Cytotoxic T cells bind to antigen associated with MHC class I molecules and secret cytokines for cell lysis. Memory T cells are a subset of antigen-specific T cells that persist long-term after an infection has resolved. They can be either CD4 þ or CD8 þ T cells and typically express cluster of differentiation 45 isoform (CD45RO) on the cell surfaces (Willinger et al., 2005; Akbar et al., 1988). Tregs are formerly known as suppressor T cells which are important for the maintenance of immunological tolerance. Tregs come in many forms and CD4 þ, CD25 þ, and FoxP3 þ are three major classes of Tregs. They play an important role in T cell-mediated immunity and suppression of auto-reactive T cells (Abbas et al., 2013; Singh et al., 2013). MSCs are reported to suppress T cell proliferation induced by alloantigens, nonspecific mitogens and anti-CD3 and anti-CD28 antibodies in vitro (Di Nicola et al., 2002). MSCs have a similar effect on naive T cells, memory T cells, as well as CD4 þ and CD8 þ T cells, which do not require major MHC restriction. MSCs can selectively suppress T lymphocytes by a cell-to-cell contact mechanism via the inhibition of programmed death 1 (PD-1) protein and its ligands PD-L1 (also known as B7 homolog 1 (B7-H1)) and PD-L2 (Freeman et al., 2000). PD-1 is a cell surface membrane protein of the immunoglobulin superfamily and negatively regulates T cell receptor (TCR) signals. PD-1 also inhibits T cell proliferation and IFN-g secretion via the inhibition of IL-2 secretion (Carter et al., 2002). Additionally, several reports suggest that cell-to-cell contact is not a compulsory requirement for MSCs to suppress the immune response of T cells (Groh et al., 2005). Cell inhibition is also due to cytokines secreted by MSCs such as IFNg, IL-1b and growth factors such as transforming growth factor (TGF)-b1 and hepatocyte growth factor (HGF). The immunomodulatory activity of MSCs is through indoleamine 2,3-dioxygenase (IDO) and prostaglandin E2 (PGE2). These growth factors and cytokines bind to their receptors on cell surfaces and trigger enzyme cascade of T cells (Glennie et al., 2005). The secretion of human leucocyte antigen-G5 (HLA-G5) by MSCs is also regarded essential for their effects of T cell suppression, shift of the

Fig. 1 Illustration of MSC immunoregulation. When tissue infection or injury occurs, MSCs migrate to the injury site where they are directly activated through TLR stimulation including pathogen-associated molecular pattern molecules (PAMPs) and damage-associated molecular patterns (DAMPs). MSCs also express other receptors such as Fas, TNFR and IFNGR depending on the signals received. MSCs are able to regulate the pathogen clearance, inhibit inflammation and promote tissue repair.

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activated T cell response to a T helper type 2 (Th2) cytokine profile and induction of CD4 þ, CD25 þ, and FoxP3 þ regulatory T cells (Selmani et al., 2008). MSCs also modulate immune responses through the induction of regulatory T cells. Regulatory T cells actively suppress activation of the immune system and prevent pathological self-reactivity in autoimmune diseases. MSCs are thought to induce the formation of CD8 þ regulatory T cells that are responsible for the inhibition of allogeneic lymphocyte proliferation. An increasing number of CD4 þ CD25 þ T cells, displaying a regulatory phenotype such as FoxP3 þ cells, have been demonstrated in mitogen stimulated peripheral blood mononuclear cell (PBMC) cultures with MSCs. On the contrary, depletion of CD4 þ CD25 þ regulatory T cells have no effect on the inhibition of T cell proliferation by MSCs (Aggarwal and Pittenger, 2005). The molecular mechanism by which regulatory T cells exert their regulatory activity has not been identified. The in vitro experiments have given mixed results regarding the requirement of cell-to-cell contact for effector cells to be suppressed. The immunosuppressive cytokines TGF-b and IL-10 have also been implicated in regulatory T cell function (Agarwal et al., 2014). Recent studies have shown that the MSC-induced formation of regulatory T cells is driven by MSC-derived TGF-b1 and by CeC motif chemokine ligand 18 (CCL18) which is produced by monocytes upon the interaction with MSCs (Melief et al., 2013). Further understanding of the interplay between MSCs and T cells is helpful in the development of promising approaches to improve cell-based regenerative medicine.

B Cell In mammals, immature B cells are derived from the bone marrow and distinguished from other lymphocytes by the presence of a B cell receptor (BCR) on the cell surfaces. This specialized receptor protein allows B cells to bind to the specific antigens. B cells have been considered mainly as positive regulators of immune responses and central contributors to the pathogenesis of immune-related diseases because they are able to produce antibodies such as immunoglobulin A (IgA), immunoglobulin G (IgG) and immunoglobulin M (IgM). B cells are characterized by the capacity of antigen presentation, the production of multiple cytokines, and a suppressive capacity by secretion of IL-10 (Mauri and Bosma, 2012). MSCs may also regulate the immune response through interaction with B cells. Co-culture of MSCs and B cells with stimuli can inhibit the proliferation of B cells and immunoglobulin expression. In addition, MSCs can suppress immunoglobulin production as a result of MSC-derived CeC motif chemokine ligand 2 (CCL2) and CeC motif chemokine ligand 7 (CCL7) which have influence on the chemotactic property of B cells. Inhibition of CCL2 in MSC impaired their suppressive capacity for B cells (Lee et al., 2017). Inhibition of matrix metallopeptidase (MMP) enzymatic activity may abolish the suppressive effect of MSCs because MMPprocessed CCL2 can suppress the signal transducer and activator of transcription3 (STAT3) activation in plasma cells (Rafei et al., 2008). However, depending on the level of stimulation, the secretion of IgG by induced B cells can be stimulated or inhibited after the addition of MSCs, which leads to different results (Corcione et al., 2006). Interestingly, MSCs possess the capacity to inhibit the proliferation of activated B cells in the presence of pro-inflammatory cytokines such as IFN-g. MSCs negatively regulate antibody production depending on the activation state of the B cells and the dose of MSCs themselves. Although the expression of migration related chemokines for B cell is inhibited by MSCs, MSCs have no influence on the expression of co-stimulatory molecules and cytokines (Krampera et al., 2006; Bernardo et al., 2009). Some in vivo and in vitro investigations have also illustrated that MSCs can modulate the expression of B cells and protect the bone healing process in the defect sites (Corcione et al., 2006).

Macrophage Macrophages are a type of circulating monocytes derived from hematopoietic progenitors or myeloblasts that engulf microbes and antigens in a process called phagocytosis. Macrophages perform an important role in tissue homeostasis and regeneration by acting as scavengers for potential pathogens. They participate in non-specific innate immunity and also initiate specific adaptive immunity by recruiting other immune cells (Ovchinnikov, 2008). In response to infection or injury, macrophages migrate to the local sites and contribute to different inflammation processes (Varin and Gordon, 2009; Lichtnekert et al., 2013). The differentiation, activation, and polarization of macrophages depend on specific cytokines such as granulocyte/macrophage colonystimulating factor (GM-CSF), macrophage colony-stimulating factor (M-CSF) and pro-inflammatory cytokines. Based on distinct functional properties, surface markers and inducers, macrophage phenotypes have been characterized broadly into M1 (classically activated) and M2 (alternatively activated), mirroring the Th1 and Th2 nomenclature described for T helper cells (Mills et al., 2000). Cytokines such as IFN-g and bacterial endotoxins such as lipopolysaccharide (LPS) can activate macrophages. Activated macrophages have increased metabolism, enhanced levels of lysosomal proteins and a greater potential to phagocytosis. They also undergo many changes which allow them to engulf invading bacteria and infected cellular debris. Proteases, neutrophil chemotactic factors, superoxide, cytokines such as TNF-a, IL-1 and IL-8, eicosanoids and some growth factors are released which result in tissue destruction. Another important role of macrophages in the immune system is antigen presentation. Macrophages are APCs that present antigens to effector T lymphocytes. Macrophages degrade protein exogenous antigens into peptides and attach them to MHC-II molecules. The MHC-II peptide epitopes are then placed on the surfaces of the macrophages where they can be recognized by complementary shaped TCRs and CD4 molecules on the surfaces of T cells to initiate the antigen presenting process (Hume, 2008).

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Increasing evidence demonstrates that MSC-mediated regulation of macrophages is critical for inflammatory response and tissue repair. MSCs interfere with the acquisition of M1 phenotype, and promote M2 polarization. By increasing IL-10 and decreasing IL-2 and TNF-a production, human MSCs can promote the generation of alternatively activated macrophages, characterized by a high expression of cluster of differentiation 206 (CD206) and a typical cytokine profile of M2 macrophages (Kim and Hematti, 2009). When inflammatory stimuli occur, bacterial toxins such as LPS act on the toll-like receptor 4 (TLR4) on the cell surface of MSCs and macrophages. MSCs then receive a second signal from macrophage-derived TNF-a. These events initiate a signaling cascade to activate myeloid differentiation factor 88 (MyD88) and nuclear factor-kappa B (NF-kB), upregulate cycloxygenase 2 (Cox2) expression and further increase the synthesis of PGE2. Thereafter, PGE2 secreted by MSCs binds to its receptors EP2 (prostaglandin E receptor 2) and EP4 (prostaglandin E receptor 4) on the surfaces of macrophages. Activation of EP2 and EP4 increases intracellular cAMP levels leading to direct suppression of multiple pro-inflammatory cytokines and promotion of anti-inflammatory cytokine IL-10 (Tyndall and Pistoia, 2009). Bone macrophages are a discrete population of resident macrophages involved in bone homeostasis. They are found to be intercalated throughout murine and human osteal tissues and distributed among other bone lining cells within both endosteum and periosteum. Researchers have demonstrated that bone macrophages are required for efficient osteoblast mineralization in response to the physiological remodeling stimulus. Removing them from calvarial cultures can down-regulate osteocalcin expression and inhibit osteoblast mineralization in vitro. Depletion of bone macrophages in vivo can cause complete loss of osteoblast boneforming surface at this modeling site. The previous study has also shown that it is bone macrophages and not osteoblasts that respond to pathophysiological concentrations of LPS (Chang et al., 2008). Our study has shown that MSCs regulate the activation of macrophage during osteogenesis (Zhou et al., 2017). These observations implicate that macrophages are closely associated with MSCs in bone tissues. In terms of acquiring anti-inflammatory properties, expanding lifespan, and increasing motility with proper engraftment of MSCs, the interaction could be explored in therapeutic applications.

Natural Killer (NK) Cell, DC Cell and Neutrophil NK cells are first identified as large granular lymphocytes with natural cytotoxicity against tumor cells and are later recognized as a separate lymphocyte lineage with cytokine-producing effector functions. Activating NK cell receptors detect the presence of ligands such as the stress-induced self-ligands, infectious nonself-ligands and TLR ligands. The in vitro exposure to TLR ligands promotes IFN-g production and enhances cytotoxicity of NK cells. However, this process is more efficient when accessory cells are present in the cultural environment of NK cells, suggesting that the TLR on cell surfaces of NK cells may play an indirect role in vivo (Vivier et al., 2008). MSCs have been shown to differentially affect the cytotoxicity of NK cells. MSCs inhibit IL-2-induced NK cell proliferation by secretion of the soluble factors such as TGF-b, HLA-G and PGE2 as well as direct cell-to-cell contact. MSCs also prevent effector functions of NK cells, such as cytotoxicity against virus-infected cells, by the inhibition of IFN-g secretion. This inhibitory effect is primarily mediated by IDO and PGE2 to reduce the expression of the activating NK cell receptors. Although MSCs do not directly inhibit the functions of NK cells, reduced cytolytic potential has been demonstrated, which is more pronounced against HLA class I-positive than HLA class I-negative cells (Sotiropoulou et al., 2006; Shi et al., 2011). DCs are APCs of the mammalian immune system. DCs play a role in the initiation, maintenance and regulation of immune responses by promoting antigen specific T cell activation and inducing cells of the innate immune system. DCs act through antigen presentation and activation of T cells. It is also suggested that a different class of DCs exists with the function of initiating and maintaining immune tolerance. The third category of DCs, known as follicular DCs, appears to maintain immune memory in tandem with B cells (Wu et al., 2012). In contrast to their effects on effector immune cells, MSCs are able to down-regulate the expression of the co-stimulatory molecules such as cluster of differentiation 40 (CD40), cluster of differentiation 80 (CD80) and cluster of differentiation 86 (CD86) on the surfaces of DCs and negatively affect DC differentiation in vitro. In fact, MSCs prevent DC maturation through PGE2 secretion. Recent studies show that MSCs inhibit DC maturation and function through the secretion of tumor necrosis factor-inducible gene-6 (TSG-6) (Liu et al., 2014). DCs cultured in the presence of MSCs release lower amounts of IL12 and TNF-a, but higher levels of IL-1b and IL-10, and further express low levels of MHC-II surface antigens (Banchereau and Steinman, 1998). Neutrophils are derived from stem cells in the bone marrow which can be subdivided into segmented neutrophils and banded neutrophils. They form part of the polymorphonuclear cell family which are short-lived and highly motile. In response to the inflammatory stimuli, neutrophils migrate from the circulating blood to the infected tissues, where they efficiently engulf and inactivate bacteria. Neutrophils also degranulate and secret antimicrobial factors into the extracellular medium (Witko-Sarsat et al., 2000). MSCs significantly inhibit apoptosis of resting and IL-8 activated neutrophils. As shown by transwell experiments, the anti-apoptotic property of MSCs does not require cell-to-cell contact. IL-6 expression is detected in MSCs by real-time reverse transcription-polymerase chain reaction (RT-PCR) and enzyme-linked immunosorbent assay (ELISA). Antibody neutralization experiments have demonstrated that IL-6 is responsible for neutrophil protection from apoptosis and it is signaled by activating STAT3 transcription factor. Researchers have also found that recombinant IL-6 can protect neutrophils from apoptosis in a dosedependent manner. However, MSCs perform no effect on neutrophil phagocytosis, expression of adhesion molecules and chemotaxis (Raffaghello et al., 2008; Fig. 2). In order to bypass the ethical controversy surrounding embryo destruction, stem cell obtained from differentiated somatic cells is an alternative way for tissue engineering. MSCs are ideal transplanted cells not only because of their own self-renewal capacity and

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Fig. 2 Illustration of the effect of MSCs on immune cells. Immune cells including T cells, B cells, neutrophils, NK cells, dendritic cells, monocyte and macrophages are presented in the frame on the left-hand side, and the main cytokines secreted by MSCs to regulate immune responses are presented on the right-hand side. MSCs inhibit T cell and B cell proliferation, induce neutrophil apoptosis, inhibit dendritic cellular maturation and limit division of NK cells. At the same time, MSCs induce differentiation of Treg cells. The green arrows indicate the cytokines produced by MSCs. The black line indicates the positive effect and the blunt-end line indicates the negative effect of MSCs.

pluripotency, but they are also a powerful tool in the regulation of immune cells for assisting transplantation (Fig. 2). For example, the transplantation of allogeneic skin grafts has been associated with a potent inflammatory immune response leading to the rapid elimination of donor cells and rejection of the graft (Benichou et al., 2011). When MSCs are applied in the surgery, MSCs can suppress lymphocyte proliferation and prolong skin graft survival (Bartholomew et al., 2002). Researchers have also found out that autologous bone grafts with MSCs and fibroblast growth factors-2 (FGF-2) can accelerate bone union in large bone defects (Murakami et al., 2016). To sum up, research on stem cell biology has demonstrated exciting opportunities and great potentials for cell-based tissue regeneration. Combined with new delivery strategies, desired properties for tissue grafts can be created along with less ethical concerns. As the field of tissue engineering continuous to evolve, more ideal cell sources will be developed for promising therapeutic applications.

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Relevant Websites http://www.bloodjournal.org/?sso-checked¼true. https://www.jci.org/. http://www.jimmunol.org/. http://www.liebertpub.com/overview/stem-cells-and-development/125/. https://www.nature.com/ni/. http://www.nature.com/nature/index.html. http://onlinelibrary.wiley.com/journal/10.1002/%28ISSN%291932-7005. http://onlinelibrary.wiley.com/journal/10.1002/(ISSN)1526-968X. http://www.sciencemag.org/journals. http://stemcellsjournals.onlinelibrary.wiley.com/hub/journal/10.1002/(ISSN)1549-4918/.

Osseointegration of Permanent and Temporary Orthopedic Implants JS Hayes, NUI Galway, Galway, Ireland RG Richards, AO Research Institute, Davos, Switzerland © 2019 Elsevier Inc. All rights reserved.

Historical Review of Internal Fixation Plate Design Metals Used in Orthopedics The Cause of Bone Loss under Plates – Stressing Shielding or Temporary Porosis? Implants – Remove or Not to Remove? Reasoning for Elective Implant Removal Reasoning against Elective Implant Removal Biological Performance – Empirically Determining the Cell/Tissue–Implant Interaction Surface Chemistry Surface Topography Substrate Stiffness Porosity/Pore Size Bioresorbable Implants – Temporary Implants to Provide a Permanent Solution? Treatment of Nonunions – A Brief Perspective into Scaffolds for Tissue Engineering Summary References

257 257 258 260 260 261 262 262 264 264 265 265 266 266 267 267

Glossary Anodization Electrolytic passivation process to increase the thickness of the natural oxide layer. Biocompatible Perform with an appropriate host response in a specific application. Ductile Plastically deform without fracture. Electropolishing Electrochemical process that removes material from a metal implant. Microtopography Micron scale variations in the height and roughness of an implant surface. Osseointegration Direct structural and functional connection between bone and the surface of a load-bearing implant. Paste polishing Mechanically abrasive process to remove material from a metal implant. Stiff Resist deformation by an applied force. Strong Withstand applied stresses without failure.

Historical Review of Internal Fixation Plate Design Despite the medical management of fractures being described for thousands of years, it is only in the last hundred years or so that the metal internal fixators that we rely on today were introduced. Since the first introduction of the metal plate by Lane in 1895, there have been several improvements on design, mechanical properties, and material choice that have allowed for great advances in fracture fixation. The initial plate developed by Lane was flat and because of issues with corrosion, subsequent improvements ensued. In the early 1900s, Lambotte introduced a curved plate in an effort to better fit the natural curvature of bone, as well as a threaded screw which eventually gave rise to the current threaded cortical screws found in clinics today. Shortly afterward in 1912, Sherman was credited with the development of self-tapping screws. While material and design advancements did demonstrate improved corrosion resistance over the Lane plate they too were ultimately abandoned due to inadequate strength and fixation instability. An important development in internal fracture fixation came in 1949 with the development of Danis’ ‘coapteur’ or compression plate (Figure 1). This design limited interfragmentary movement while improving the stability of the fixation. This new approach of absolute stability resulted in what Danis described as soudure autogène, a process we now recognize as primary bone healing (healing without the presence of fracture callus). While Danis’ compression plate was introduced less than a century ago, it is this design that influenced all subsequent plate designs. In fact, Danis’ description of diaphyseal fracture healing in the absence of callus formation spurred the interest of a young Swiss surgeon by the name of Maurice E Müller, who subsequently met with Robert Danis to discuss his observations in 1950. This meeting resulted in the establishment of the Arbeitsgemeinschaft für

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Figure 1 The evolution of internal fixation plate design. Lambotte’s plate (a) was introduced subsequent to Lane’s plate due to issues of corrosion. Insufficient strength and problems with stability eventually led to the naissance of Danis’ ‘coapteur’ plate (b). The latter allowed for interfragmentary compression, the result of which is primary bone healing (absence of callus formation). (c) It was this design that revolutionized all subsequent designs including the dynamic compression plate and later the limited contact dynamic compression plate which arose due to perceived issues relating to bone necrosis (see Section The Cause of Bone Loss under Plates – Stressing Shieding or Temporary Porosis? for discussion). (d) The point contact fixator (PC-fix) is another example of subsequent plate designs that focused on preserving the periosteal blood supply underneath the plate for appropriate healing to occur. This type of design limits the plates ‘footprint’ along the bone. (e) The development of the locking compression plate came from the continuous requests of surgeons for a synergistic system of a threaded and a conventional screw hole allowing for the greatest possible flexibility for the treatment of fractures. Images (a) and (b) reproduced with kind permission of the Journal of Orthopaedic Science (Uhthoff, H.K., Poitras, P., Backman, D.S., 2006. Internal plate fixation of fractures: short history and; recent developments. J. Orthop. Sci. 11, 118–126). Images ((c)–(e)) reproduced with kind permission of the AO Foundation from AO principles of fracture management, 2nd Edition.

Osteosynthesefragen/Association for the study of Internal Fixation (AO/ASIF) by Müller and others in 1958 and the subsequent Laboratory for Experimental Surgery in Davos, Switzerland which is responsible for much of what is known today relating to fracture healing and fixation. The AO contributed to the development of the dynamic compression plate (DCP) in 1969 (Figure 1). However, observations of interrupted cortical blood perfusion leading to cortical necrosis (Akeson et al., 1976; Perren et al., 1988) led to the development of the limited contact dynamic compression plate (LC-DCP) by Perren in 1990. With plating, most of the iatrogenic disruption to bone has been traced to the regions of direct bone–implant contact and is believed to result directly from periosteal blood supply disruption (Perren et al., 1988). This finding was the main motivation behind the large investment in the development of new generation internal fixation systems that effectively reduce the extent of contact of the plate with bone consequently avoiding disrupted periosteal vascularity. Here the AO again pioneered much of the developments within this area firstly with the development of point contact fixator or PC-fix system (Figure 1), which allowed fixation by shear forces between the screw and bone rather than frictional forces generated between the bone and plate as with conventional plating methods (Tepic and Perren, 1995). Subsequent designs include the less invasive stabilization system (LISS) which was developed for application specifically within metaphyseal and epiphyseal regions as the PC-fix was unsuitable in these areas. The locking compression plate (LCP; Figure 1) was developed in response to increasing requests from orthopedic surgeons for the availability of a system that could offer a synergy between a threaded and a conventional screw hole allowing for the greatest possible flexibility for the treatment of fractures (Frigg, 2003). The operating surgeon thus has the scope of deciding which variation of the system would be applicable for treatment depending on type of fracture and quality of the bone since the system can be applied in either a conventional technique (compression), bridging method (internal fixator), or combination of both (Wagner, 2003).

Metals Used in Orthopedics Desirable properties such as high stiffness, strength, biological toleration, and reliable function have meant that metal implants have gained unrivaled success in fracture fixation for many years. Before the advent of metals in orthopedics fracture management

Biomaterials: In Vitro and in Vivo Studies of Biomaterials j Osseointegration of Permanent and Temporary Orthopedic Table 1

Despite the excellent clinical success of stainless steel and titanium in fracture fixation, there are several key properties which differ between materials. As noted in Figure 2 stainless steel has a smooth mirrorlike finish compared to the microrough finish of titanium. In addition, differences in density, oxide chemistry, thickness, and regeneration time as well as modulus of elasticity all influence the materials clinical performance (Section Metals Used in Orthopaedics) Density (g cm3)

Young’s modulus (GPa)

Stainless steel

7.9

Titanium

4.5

Material

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Surface

Main components of oxide layer

Thickness of natural oxide later

Innate oxide regeneration time (min)

186

Smooth 0.1 ms

O, C, Cr, Fe, Mo, Si

2 nm

110

Microrough 1 mm

Ti, O, C

6 nm

Annealed

Cold-worked

Cold-drawn

35

UTS ¼ 490– 690 MPa Cerclage wire

UTS ¼ 860– 1100 MPa Bone screws, plates, IM nails

8

UTS ¼ 200– 550 MPa CMF plates

UTS ¼ 680

UTS ¼ 1350– 1600 MPa Schanz screws, Krischner wire –

Bone screws, plates, IM nails

Content adapted from Hayes, J.S., Richards, R.G., 2010b. The use of titanium and stainless steel in fracture fixation. Expert Rev. Med. Devices 7 (6), 843–853.

mainly involved immobilization in plaster or by traction but this approach greatly inhibited function during healing. The introduction of stainless steel and titanium in the early twentieth century would see internal fixation transformed. The reader is directed to the review paper by Hayes and Richards (2010b) for a more comprehensive insight to the properties and clinical applications of stainless steel and titanium that will be summarized here. Despite their similar clinical success, stainless steel and titanium have many different properties that are important for fracture fixation. Ideally, an implant for internal fixation applications should be ductile, strong, and stiff, perform well under different mechanical demands and importantly be biocompatible. Both stainless steel and titanium fulfill these requirements to different degrees. Stainless steel has a density of 7.9 g cm3 which for small implants is not generally considered to be problematic, however, it is worth noting that it is more than twice as dense as titanium implants of the same design (Disegi and Eschbach, 2000). The modulus of elasticity of steel is also higher than titanium (186 GPa compared to 110 GPa). In terms of transferring load to the adjacent bone (approximately 20 GPa) this means stainless steel is less effective than titanium which may be significant in terms of stress shielding. In contrast, stainless steel is more ductile than titanium which permits more deformation of the implant for intraoperative contouring by the surgeon to compensate for anatomical requirements. To counter this limitation of titanium anatomically designed implants are produced to restrict the need for contouring. Three main tensile versions of stainless steel are fabricated for clinics namely (listed from lowest to highest strength), annealed (i.e., cerclage wire, reconstruction plates), cold-worked (i.e., bone screws, bone plates, intramedullary nails), and cold-drawn (i.e., Schanz screws, Kirschner wire). These conditions vary in strength allowing for use in several different implant applications depending on strength and ductility requirements. For instance, annealed implants are generally the ‘softer’ version of the material and thus allow for deformation and are normally used in low stress loading applications whereas cold-worked materials represent the increased strength version of the material where ductility and high strength are favored. Titanium is fabricated for clinics in two mechanical forms, annealed (i.e., craniofacial plates, mini-fragment plates) and cold-worked (i.e., bone plates, bone screws, intramedullary nails) (Table 1). In terms of their biological reaction, stainless steel and titanium differ in two major areas. Firstly, stainless steel is fabricated for clinical applications with a smooth electropolished surface (an approximate microroughness of 0.1 mm) whereas titanium and its alloys are produced with a microrough surface of approximately 1 mm (Figure 2). As outlined in Section Surface Topography, the surface microtopography can have a profound biological effect on tissue integration. Another important difference of these materials is the chemical makeup of their surface oxides. This is especially important as it is the presence of an oxide film that is responsible for the implants’ increased corrosion resistance and biocompatibility. The reader is directed to the review by Hayes et al., (2010b) for more detail. Briefly, the passive film of stainless steel consists mainly of iron, nickel, and chromium. The chromium in stainless steel reacts with oxygen to form the 2 to 3-nm-thick characteristic passive film that provides increased corrosion resistance for these devices. In contrast, titanium oxide layer mainly consists of titanium, oxygen, and carbon which produces an innate oxide layer of approximately 5–6 nm. However, to improve clinical performance, the commercial technique of anodizing is normally employed for both stainless steel and titanium devices which can result in oxide layers of approximately 200 nm. As mentioned, it is this delicate oxide layer that is responsible for maintenance of biocompatibility by preventing the release of toxic ions from the underlying highly reactive bulk material. If abraded both stainless steel and titanium maintain the ability of regenerating the layer. Stainless steel devices tested in saline solution take approximately 35 min to regenerate compared with approximately 8 min for titanium and titanium alloys (Hanawa, 2004). This can have consequences for stainless steel compatibility given that the ions within the bulk material are more highly reactive than titanium and its oxide layer regenerates at a slower rate, thereby potentially permitting increased release of toxic ions once abraded.

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Figure 2 Examples of the surface of electropolished stainless steel (a) and commercially pure titanium (b) used for orthopedic devices. Note the extremely smooth surface of stainless steel compared to the 3D microrough surface of titanium. Despite unrivaled clinical success in orthopedics, both materials differ in several key areas (Table 1). The difference in surface appearance is translated to differential tissue responses at the implant interface in vivo. Generally, stainless steel supports a fibro–osseous interface (soft and hard tissue interface) while titanium supports direct osseointegration.

The Cause of Bone Loss under Plates – Stressing Shielding or Temporary Porosis? The occurrence of temporary osteoporosis adjacent to the implant contact area has long been observed and recognized. Specifically, there appears to be an enlarging of the intramedullary cavity coupled with cortical bone loss via endosteal resorption (Akeson et al., 1976). During the nineteenth century, Julius Wolff developed the concept that bone will adapt to the load under which it is placed. According to Wolff’s law if the load applied to bone is decreased then its density and structure will diminish (and vice versa). Applying this theory to osteosynthesis initially appeared logical: the implant would inadvertently transfer load away from bone to the device, thereby, protecting or shielding the bone from load therefore adapting its structure accordingly. Hence, experimental evidence to support the theory of stress shielding, also known as stress protection has been published for over 40 years. However, the term ‘stress protection’ in terms of bone plate fixation has become debated since the theory of disturbed vascularity has come to light. Arguing against stress protection, Perren and colleagues (1988) suggest that early osteoporosis observed within the first 6 months of fixation was due to cortical necrosis, secondary to excessive plate–bone contact interfering with cortical perfusion. Their hypothesis was that the initial increase in bone porosity was a result of the host attempting to eliminate necrotic bone and ultimately remodel the site. Studies demonstrating the ability of fixation devices to hamper perfusion have been available for over 50 years (Rhinelander, 1965; Gunst et al., 1979). These findings were later supported in a study where railed and contact plates were used in a dog model (Uhthoff et al., 1994). As a result of Perren and colleagues’ findings, significant developments were made to plate design with subsequent designs incorporating a limited contact approach as demonstrated by PC-fix, the LC-DCP, and LISS systems (Figure 3). Based on Perren’s theory it would seem that porosis is transient and is reversed upon completion of remodeling of the necrotic area. As a major critic of the theory of temporary porosis Uhthoff and colleagues (1994) addressed this issue and found no evidence to support the notion (Uhthoff et al., 1994). They did, nevertheless, show in other studies that the immobilization of canine limbs with plaster casts led to cortical thinning and widening of the medullary canal (Uhthoff and Jaworski, 1978; Jaworski et al., 1980). Uhthoff also maintains that if the initial hypothesis put forth by Perren and colleagues was true then the increased porosity would be found at the area of necrosis. Recently, Tepic (2008) showed that for PC-fix plates the area of porosity was increased in the vicinity of necrotic tissue. However, it has also been demonstrated that bone porosity is increased more in the endosteal part of the cortex rather than the periosteal part where necrosis is most evident (Akeson et al., 1976; Uhthoff et al., 1994). Despite the ambiguity surrounding this issue, these observations and clinical outcomes have resulted in the routine removal of many of these internal fracture repair devices.

Implants – Remove or Not to Remove? The ubiquitous use and success of biomaterials in medical applications is undeniable. Nevertheless, when referred to in the context of medical applications it is generally accepted as referring to long-term implantation. Almost all hardware removal procedures in adults are a result of a complication. In contrast, for pediatric patients, the accepted protocol was removal of the device once its purpose had been spent (Schmalzried et al., 1991). In this section we examine the main clinical considerations for and against elective implant removal.

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Figure 3 Decreased cortical perfusion is evident with increased bone–plate contact with contact plates (left) compared to railed plates (right) as demonstrated with intravital disulfine administration. Two main arguments exist to debate the bone changes noted beneath plates (Section The Cause of Bone Loss under Plates – Stressing Shieding or Temporary Porosis?). Perren and colleagues suggest that early porosis is due to cortical necrosis, secondary to excessive plate–bone contact interfering with cortical perfusion. Others advocate the theory of ‘stress protection’ which outlines that load is transferred away from bone as it is shielded by the plate and as a result undergoes a weakening in the absence of appropriate mechanical stimulus. Images from Uhthoff, H.K., Poitras, P., Backman, D.S. 2006. Internal plate fixation of fractures: short history and recent developments. J. Orthop. Sci. 11, 118–126.

Reasoning for Elective Implant Removal Many complications surround the use of metal fixation devices but are normally alleviated upon extraction and are therefore used as grounds for electively removing devices. Common ailments experienced include pain upon implantation, irritation of adjacent soft tissues, and acute aseptic inflammation. The formation of restrictive adhesions and subsequent limited motion is one of the most challenging areas of hand surgery. Specifically the tendon sheath if irritated produces a severe inflammatory response (Khan et al., 2000). Here, stainless steel devices are favored (Sinicropi et al., 2005) although recent developments in the modification of titanium alloy plates have proved successful (Hayes et al., 2012a). Moreover, the device may encroach into the joint space resulting in pain, loss of motion, and fracture instability. In severe cases an implant may migrate. Devices implanted originally in the humerus, mandible, and shoulder have been reported to migrate to the mediastinum, spinal cord, and jugular foramen, respectively (Seipel et al., 2001). Infection is practically impossible to treat unless the cause, most often the implant, is removed. This complication can arise at any time and is not restricted to the early events that occur upon implantation. Highland and LaMont (1985) report the occurrence of deep late infection from 7 to 24 months after implantation. In a case study cited by Peterson (2005), an eighteen-year-old male developed osteomyelitis some 50 years after initial treatment. It is possible that this was a result of a recent hematogenous infection rather than the organism remaining quiescent for those years, nevertheless, the persistence of bacteria intracellularly has been reported (Vaudaux and Lew, 2006; Broekhuizen et al., 2008). Orthopedic device-related infections are notoriously difficult to treat and several studies have reported the varying effect of antibiotics in treating pyogenic (pus producing) infections (Lew and Waldvogel, 2004). In many instances antibiotics are given prophylactically, however, the cost of continued treatment for a patient with an infected device can be staggering. Some patients demonstrate hypersensitivity to metal devices. However, it is not believed that metal hypersensitivity contributes to endoprosthesis loosening (Milavec-Puretic et al., 1998). A retrospective study reported elevated levels of metal ion (titanium, vanadium, and aluminum) concentrations in the serum (35% of patients (70–90 ppb while 20 ppb was considered ‘normal’) and hair (24% of patients (20–30 mg g1) compared to 3 and 19 mg g1) considered as ‘normal’ for titanium and aluminum, respectively) of patients with titanium alloy fixation devices (Kasai et al., 2003). Nevertheless, stainless steel devices tend to demonstrate more problems than titanium counterparts due to the presence of chromium, cobalt, and nickel. Current stainless steel implants comprise approximately 13–16 wt% nickel despite this being the most common skin-contact allergen. Approximately 1–2% of patients present with allergic reactions to Ni-containing stainless steel devices after internal fixation. Furthermore, nickel, chromium, and cobalt have been shown to produce reactive oxygen species (Valko et al., 2006) and are thought to contribute to the development and potential progression of neurodegenerative disorders (Olivieri et al., 2002). Directives are in place to define acceptable concentrations of metal debris within patients (Morgan and Shaller, 1999); however, in reality these are often exceeded (Schaffer et al., 1999; Pilger et al., 2002; Lhotka et al., 2005; Witzleb et al., 2006). Moreover, the concentration of metal ions for adverse responses to occur varies considerably. For instance, Hallab et al. (2008) found that metals such as cobalt and vanadium were toxic at concentrations as low as 0.5 mM, while other metals (aluminum, chromium, iron, molybdenum, and nickel) were toxic at much higher concentrations (>10 mM). While there is still much to learn regarding the tolerated concentrations of metal debris/ion release, evidence does exist to suggest that exposure can be influential in a variety of biological systems. Kanojia and colleagues (1998) has reported increased levels of cobalt and chromium in the cord blood of pregnant women. These patients all had metal fixation devices and had become pregnant subsequent to implantation. This is of particular concern given that these metal ions are known inducers of developmental toxicity (Domingo, 1994; Elbetieha and Al-Hamood, 1997; Keegan et al., 2007). The negative biological impacts of metal debris and ions have also been observed in cardiovascular, respiratory, urinary, and

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reproductive systems (Elbetieha and Al-Hamood, 1997; Oliveira et al., 2006; Keegan et al., 2007). Regardless, stainless steel maintains one of the leading materials used and is generally considered biocompatible. In pediatric patients the risk of premature physeal closure is increased when an implant transverses a physis. Thus removal of the device is strongly advocated in the case of patients who have significant growth potential unless the growth cessation is desired (Peterson, 2005). Unfortunately, the occurrence of premature unwanted physeal closure due to the presence of a fixation device is high. For instance, Chen and colleagues (2002) observed a 41% occurrence of premature closure of physis while Bagatur and Zorer (2002) report its presence in 65% of the patients included in their retrospective study. Other factors such as return to sports, carcinogenicity and peri-implant fracture are also concerns to be taken into account. In terms of carcinogenicity and peri-implant fracture, the literature indicates that their association with metal implants appears to be negligible and thus do not warrant routine removal. Nevertheless, some of those reported hold reason for revision. For instance, Keel and colleagues (2001) reported that sarcomas related to fixation devices tend to be high grade, to behave aggressively and to metastasize frequently. In the 12 cases they reviewed, 10 were found in bone (the other 2 were found in adjacent soft tissues) and follow up was permitted in only 8. Of these, 7 died between 2 and 30 months after diagnosis. In the case of patients wishing to return to sports, there seems to be a lack of consensus here too. Some believe that devices that restrict joint or muscle action should be removed prior to the patient returning to sport (Labosky et al., 1990). In contrast, 87% of professional rugby players returned to full training with retained fixation devices (Evans and Evans, 1997).

Reasoning against Elective Implant Removal Many surgeons, engineers, and researchers alike fail to find sufficient relevant clinical or experimental evidence to convince them to adopt an attitude that routine elective hardware removal is the best option in the end treatment of fixation. Although few reasons exist for this failure those that do are just as valid and compelling. From a purely economic view point elective removal can be costly and time consuming. Hospitalization, anesthesia, imaging, antibiotics therapy, and operating equipment are all points that should be considered. Cost to patient earning is rarely considered but is equally valid. Although the carcinogenicity risk associated with retained devices in animal models has been well documented there appear to be few cases of implant-associated tumors in humans. Thus, advocates of implant retention suggest that this risk is impossible to quantify accurately and does not warrant the application of routine removal based on its occurrence. A retrospective study of more than 116 000 patients reported no correlation between implant retention and increased cancer risk when compared to the general population (Signorello et al., 2001; Fryzek et al., 2002). They did, however, report an increase in prostate cancer and melanoma (standardized incidence ratio of 1.16 and 1.86, respectively). Of the patients that had implants retained no increase was noted for bone or other connective tissue-related cancers (Signorello et al., 2001). A later study also found that patients with knee prostheses did not present with an increase risk of cancer compared to the general population (Fryzek et al., 2002). Again, elevated risk was observed for prostate cancer (standardized incidence ratio of 1.20) and may warrant due attention. A risk in bone cancer was detected in this study (standardized incidence ratio 6.00), however, the authors believe these incidents (3 of the total 120 000 þ patients studied) to be unrelated to the implant itself (Fryzek et al., 2002). It is maintained that conclusions regarding the carcinogenic potential of implants in patients are particularly difficult to make since variables such as the quantity of wear debris produced and patient age may be possible confounding factors (Heath et al., 1971). The question remains whether malignancy is correlated to the presence of long-term implants or if it is in fact caused by them. Katzer and colleagues (2003) reported that neither chromium cobalt molybdenum nor titanium aluminum wear particles produced toxic or mutagenic effects, whereas Doran and co-workers (1998) document significant increases in cell transformation for soluble forms of the same materials and appeared to relate directly to toxicity. Furthermore, macrophages have been previously shown in vitro to phagocytose laboratory produced debris from steel and titanium and a titanium alloy in a dose dependent manner (Wilson, 1999). Steel particles are reported to be approximately 0.5 mm in size (ap Gwynn and Wilson, 2001), and thus, are easily dispersed throughout the body and have been noted in organs remote from the implantation site (Case et al., 1994). In contrast, particles generated from titanium are approximately 10 mm in size (ap Gwynn and Wilson, 2001) and have been reported to be found in tissue adjacent to the implanted site giving the harmless discoloration of the tissue. Nevertheless, the correlation to cancer remains unfounded (Signorello et al., 2001; Fryzek et al., 2002). Complications associated with the surgical aspect of device retrieval are a principal basis for hardware retention for many surgeons. Inherent risks such as iatrogenic impairment and complications with anesthesia are considerations for any surgical procedure. In addition, the actual procedure of implant retrieval can add to the already serious preexisting concerns. For instance, refracture, nerve damage, excessive blood loss, infection, and difficulty in removing the device e.g., stripping of the screw head or even breakage at the screw neck are all potential obstacles that must be reviewed carefully for one to carry out a successful retrieval. There have been steps in recent years to address this issue (Pearce et al., 2008; Hayes et al., 2009; Hayes et al., 2010a) which is of particular benefit for pediatric fixation.

Biological Performance – Empirically Determining the Cell/Tissue–Implant Interaction There are several definitions available to define a biomaterial in terms of tissue regeneration. One such definition by the National Institute of Health defines a biomaterial as “Any substance (other than a drug) or combination of substances, synthetic or natural in

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origin, which can be used for any period of time, as a whole or as a part of a system which treats, augments, or replaces any tissue, organ, or function of the body.” This encompasses several variations of biomaterial for several different clinical applications. Regardless of application, one commonality is that the surface properties play a determining role in the tissue response upon implantation. Ultimately, therefore, it is this delicate interaction between material and host tissue that has resulted in a billion dollar industry emerging to empirically design implants in an effort to regulate the ensuing biological response. In terms of a substrate dependent response, it is all down to what a cell ‘sees’ once it comes into contact with a material, be it gel, polymer, demineralized matrix, or specifically in this case, metal. The primary interaction upon implantation of a metal device is the formation of a water biolayer (Figure 4). This process occurs within nanoseconds of the device impacting the surrounding biological niche and is followed instantaneously by contact with blood. As the implantation of a device is essentially recognized as an injury response, blood coagulates at the implant interface forming a hematoma. This hypoxic environment, similar to fracture healing itself, provides a protein and cell rich milieu at the surface of the device resulting in an inflammatory response eventually leading to the recruitment of repair cells. A more in-depth explanation can be found in several additional resources (Hayes et al., 2010a, 2012). In terms of tissue repair at the implant interface, there are several key factors that knowingly elicit specific tissue responses. Here we will deal briefly with topography, chemistry, stiffness, and porosity/pore size. It is also important to note that while all factors individually will bring about a specific cell response, in reality it is a combination of several surface properties that elicit the overall biological response.

Figure 4 Summary of the cell surface interaction. (a) A standard microrough titanium surface. (b) The same titanium surface with a cell to demonstrate what a cell ‘sees’ in terms of its underlying surface. As outlined in Section Surface Topography, the cell will be sensitive to substrate microdiscontinuities which exert effect via the cytoskeleton resulting in differential transduction of signals to the nucleus to determine genotypic and phenotypic responses. (c) Depicts the initial interaction of a device when implanted. When implanted the device comes into instant contact with biofluid i.e., plasma, blood, water, etc. The primary interaction upon implantation is the formation of a water biolayer which is followed instantaneously by contact with blood. As the implantation of a device is essentially recognized as an injury response, blood coagulates at the implant interface forming a hematoma. This hypoxic environment, similar to fracture healing itself, provides a protein and cell rich milieu at the surface of the device resulting in an inflammatory response eventually leading to the recruitment of repair cells. (d) Owing to lack of anchorage a smooth surface does not permit fibrin to remain attached during wound healing contraction. In contrast, the microrough implant provides increased anchorage for fibrin via its 3D morphology thereby permitting improved resistance to the forces generated during contraction. This allows for continuous cell migration to the implant interface and consequently, direct osseointegration. Content of D adapted from Hayes, J.S., Richards, R.G., 2010a. Surfaces to control tissue adhesion for osteosynthesis with metal implants: in vitro and in vivo studies to bring solutions to the patient. Expert Rev. Med. Devices 7 (1), 131–142.

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Surface Chemistry Focusing on metal implants, the biocompatibility of these devices relies primarily on their oxide layer. This layer forms innately, measures a few nanometers, can spontaneously regenerate given time, and determines the primary protein interaction upon implantation. Ultimately, the responses elicited by the oxide layer mean that cells never truly interact with a ‘bare’ metal surface and consequently, they are the principle players responsible for the accepted biocompatibility of metal implants. To empirically determine a cell–implant interaction chemical (i.e., calcium phosphate coatings) and/or biological modifications (i.e., peptide-based strategies) can be made to the surface chemistry of an implant via biological or chemical processes. The majority of biological modifications involve the immobilization of extracellular proteins or peptide sequences to improve cell attachment and subsequent tissue response. A more comprehensive review is covered in Hayes et al. (2012). Briefly, coating implant surfaces with the Arg-Gly-Asp (RGD) sequence is currently the most common peptide-based strategy. The RGD sequence has been identified as a cell attachment motif present on several plasma and extracellular matrix (ECM) proteins, including collagen type I, fibronectin and vitronectin, among others. In turn these proteins interact with integrins, cell membrane proteins that bind a myriad of cell surface and extracellular matrix-associated ligands. In vitro efficacy of peptide immobilization strategies have long been recognized with several studies demonstrating improved attachment and increased osteospecific gene expression in vitro (Zreiqat et al., 2003; Rivera-Chacon et al., 2013; Mendes et al., 2013). This effect is further advocated in vivo where peptide immobilization was shown to improve bone healing and direct bone-implant contact while also expediting remodeling around the implant (Rammelt et al., 2004). In terms of chemical modification, the incorporation of calcium phosphates has dominated (Schlegel et al., 2009; O’Hare et al., 2010; Costa et al., 2013; Sun et al., 2013). This approach of incorporating hydroxyapatite to improve osteoconductivity has demonstrated efficacy for improving osseointegration. A major drawback of this approach is that during long-term implantation the surface coating can delaminate from the underlying implant thereby compromising stability. Some efforts have been made to address this issue. For instance, chemical modifications of the surface to incorporate the hydroxyapatite into the oxide layer of the implant have been made (Schlegel et al., 2009; O’Hare et al., 2010). In essence, the hydroxyapatite alters the chemical composition of the surface thereby eliminating issues of delamination associated with conventional coating technologies.

Surface Topography When referring to surface topography in the context of metal implants both nano- and microtopography are key. In fact, this interplay is so delicate that quantitative and qualitative responses vary from cell type to cell type for any one particular microor nanofeature. Mesenchymal stem cells appear more reactive to nanofeatures than microtopography and fibroblast growth is hindered on a titanium alloy while primary osteoblasts are not. In our own experience we have shown that surface topography influences fibroblast growth, spreading and cytoskeletal organization (Meredith, 2006; Meredith et al., 2007). Furthermore, liquid filled capsule formation (specifically important for free gliding of tendons over hand implants) can be directly controlled via surface microtopography manipulation (Hayes et al., 2012). For bone, we and others have demonstrated in vitro that microrough surfaces increase ‘osteospecific’ factors, local factor production and upregulate key factors in osteoblast differentiation (Olivares-Navarrete et al., 2011; Gittens et al., 2011; Hayes et al., 2010; Boyan et al., 2003) and consequently, bony integration (osseointegration) of a device can be directly influenced by the microtopography of a material (Pearce et al., 2008; Hayes et al., 2009; Hayes et al., 2010; Davies et al., 2013). Several of the aforementioned studies demonstrate that the same material can elicit differing tissue responses depending on whether it is smooth (approximately 0.2 mm) or microrough (approximately 1 mm). This is poignant in terms of implant removal. Titanium is generally favored for internal fracture fixation but its microrough surface is known to innately induce bone on-growth. It is this affinity to bone that makes titanium and its alloys particularly difficult to remove. In a series of preclinical studies using bone screws, compression plates, and intramedullary nails we have shown that surface polishing (removal of surface microroughness via mechanical (paste) or electropolishing) can significantly reduce the mechanical force required for implant removal compared to clinical microrough counterparts (Pearce et al., 2008; Hayes et al., 2009; Hayes et al., 2010, 2012). In clinical terms, this approach has seen great efficacy in hand surgery (Hayes et al., 2012) with commercial adaptation under way for other sites (Figure 5). Again the reader is directed to more in-depth focused reviews (Schwartz et al., 2005; Hayes and Richards, 2010, Hayes et al., 2012). More recently, nanotopography has been implicated as a means of controlling an osteogenic response. One such study investigated 15, 55, and 100 nm titanium nanopillars in terms of adhesion, spreading, cytoskeletal organization, and differentiation of mesenchymal stem cells (Sjöström et al., 2009). Sjöström and colleagues claim that the 15 nm pillars proved the most reproducible for osteogenesis specifically in terms of nodule formation after 21 days in vitro. While several other studies have emerged in recent years advocating the role of nanotopography for eliciting an osteogenic response in vitro, data supporting this role in vivo are yet to prove conclusive. A possible explanation for this is that the nanotopographical dependent adhesive, cytoskeletal and phenotypic changes observed in vitro may be masked in vivo by the protein layer that adsorbs onto the surface upon implantation (Hayes and Richards, 2010). It is more likely that the benefit of cell differentiation dependent on nanotopography will be found within the remit of tissue engineering. Recently, for example, McMurray and colleagues (2011) have demonstrated that ordered nanotopographical surface features were effective in maintaining mesenchymal stem cell (MSC) phenotype for up to 8 weeks in culture. Interestingly, the authors showed that subsequent to prolonged culture on the ordered surfaces, cells could be removed and replated on culture

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Figure 5 Bone has a high affinity for titanium and its alloys. (a) Here we show a titanium bone plate and screws that has been implanted on a sheep tibia. Note how the bone has overgrown the device. This excessive bony overgrowth is a significant clinical problem for device removal especially in pediatric patients were device removal is favored. The reason for this direct and excessive osseointegration is the microrough surface of clinical titanium implants (inset (b)). As shown in (b), which is an image of a standard microrough titanium alloy intramedullary nail implanted in sheep tibia for 12 months, this results in direct and strong bone contact. (c) By reducing the surface topography via mechanical polishing (inset (c)) we have shown that this excessive bone overgrowth can be markedly reduced causing a significant reduction in the force required to remove the implant which is partly a consequence to the fibro–osseous interface produced by the smoothened surface. This effect has been noted for screws, bone plates, and intramedullary nails.

plastic and retained multipotency until the appropriate media supplementation. Furthermore, the authors cultured MSCs on the ordered surfaces in the presence of osteogenic supplemented media and observed minimal differentiation, thus supporting the idea that specifically ordered surfaces could maintain cell phenotype. It is an important finding in the context of cell propagation for regenerative medicine where appropriate cell number and the cost associated with media additives can be crippling.

Substrate Stiffness Substrate stiffness elicits its effect on cell response via the delicate interplay of cell adhesion, force-sensing, and signal transduction. The reader is referred to Dalby (2005) and Hayes et al. (2012) for a more comprehensive explanation. In brief, integrins bind extracellular matrix ligands and link to the cytoskeleton. Signaling proteins are then differentially regulated to ultimately influence cell adhesion. One analogy to help is to visualize a spider’s web. The points attaching the web to its surroundings would be the cell adhesion proteins, namely integrins. The force transmitted through the individual silk fibers relates to the transduction of mechanical signals through the cytoskeleton, specifically the actin filaments. Similar to a spider sensing these disruptions within the web, the forces exerted on the cell are ultimately relayed to the nucleus which signals the appropriate phenotypic response. Several studies eloquently demonstrated how manipulation of substrate stiffness, or elasticity, can be used to control cell differentiation. One such study by Engler and colleagues (2006) reveal that neurogenic (nerve), myogenic (muscle), and osteogenic (bone) lineage specification pertaining to cell morphology, RNA profiles, cytoskeletal markers, and transcription factors are dependent on matrix stiffness. In a more sophisticated recent model, Young and Engler (2011) use time dependent increases in substrate stiffness to mimic the developmental stiffening of mesoderm into adult myocardium (heart muscle). In brief, they show that cells cultured on dynamic substrates exhibit a threefold increase in mature cardiac specific markers and up to 60% more maturing muscle fibers than cells grown on compliant but static polyacrylamide hydrogels.

Porosity/Pore Size While not directly implicated in traditional metal implants, surface porosity (percentage of void space in a solid) and pore size have been implicated in regulating the cell–surface interaction. It is worth briefly covering here as many second generation metal implants now consist of coatings such as hydroxyapatite that include porous surfaces. Furthermore, advancements in biomaterial design have seen the introduction of porous metal devices. Recent studies suggest comparable healing of segmental defects to hydroxyapatite scaffolds after 24 weeks implantation in rabbit (Jung et al., 2013; Zhang et al., 2013). This is particularly promising given that titanium porous implants have superior biomechanical properties compared to hydroxyapatite scaffolds. More recent developments have also seen successful incorporation of bioactive factors to further enhance osseointegration (Shim et al., 2013). Porous surfaces provide an interconnected network that distributes oxygen and nutrients as well as eliminating cytotoxic waste material. Currently, definitive pore sizes for the regeneration of different tissues are not reported. Consequently, porosity and pore sizes of varying degrees are published for tissues such as bone, cartilage, and muscle. Nevertheless, several key guidelines have emerged and are highly dependent on the tissue under study. For instance, limited nutrient exchange and cell migration is achieved if pores are too small. On the contrary, if pores are too large this results in a significant decrease in surface area thereby limiting cell adhesion and increasing porosity affects the load-bearing capacity of the material (Khoda et al., 2011). Importantly, the selection of the biomaterial with certain pore size depends on the location to where it is going to be transplanted.

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For example, a minimum pore size of 100 mm was initially suggested to be required to generate mineralized bone, however, pores above 300 mm result in vascularized tissue formation (Karageorgiou and Kaplan, 2005). The latter is imperative for appropriate bone formation. Even within a tissue such as bone, location will determine the porosity required (Cortical bone has porosity of only 5–10% whereas trabecular bone ranges between 78% and 92%). Finally, the biomaterial itself may influence effect of porosity/pore size relating to tissue response (Karageorgiou and Kaplan, 2005).

Bioresorbable Implants – Temporary Implants to Provide a Permanent Solution? Metal remains the material of choice in loaded regions. The main reason for this monopoly is simply due to the fact that no alternative material has yet been developed to rival the strength and established clinical success of metal fixators. One promising material to emerge for use in long bone fracture fixation is the nondegradable polymer poly-(ether–ether)-ketone, or PEEK. Several characteristics of PEEK have proven desirable for osteosynthesis namely a similar Young’s modulus to bone (thereby limiting the potential occurrence of stress protection) and its radiolucency (allowing real time imaging of fracture sites without the occurrence of artifacts commonly associated with the presence of metal implants). A major clinical limitation of its application, however, is the fact that PEEK is not bioactive. In essence, therefore, it does not interact with bone to produce an integrated surface which is necessary for stability in loaded regions. As the market leader of PEEK for medical device applications, Invibio (2013) have explored several approaches of surface modification for ‘biologically activating’ the surface of PEEK to improve osseointegration including the incorporation of osteoconductive hydroxyapatite. More recently, it has also been shown that a PEEK biomimetic surface can be achieved using plasma treatment, a method of chemically altering the outermost surface layers. This method successfully incorporates biomolecules that have been shown to enhance biological functionality (Hayes et al., 2013). In terms of temporary devices, in that the device is implanted with no intention of removal and in the hope that it will resorb subsequent to its function being fulfilled, success is mainly limited to nonloaded regions such as craniomaxillofacial bones. Again, this is a direct result of the current lack of bioresorbable materials that have the strength to withstand high recurring loads. Some recent studies have shown promise in preclinical in vivo models but have noted that full-weight bearing may not be indicated initially to allow healing to occur (Neumann et al., 2013). Nevertheless, experimental data within craniomaxillofacial regions have been promising (Turvey et al., 2011) and is advocated for pediatric fractures particularly (Ahmed et al., 2013; Degala et al., 2013). However, there remains some concern relating to inadequate fracture stability and implant breakage (Singh et al., 2013). The most popular polymers used in these applications are variations of polylactic and polyglycolic acid given that their products are considered innocuous although inflammation and osteolysis have been reported (Simon et al., 1997; Turvey et al., 2011). The developmental leap to long bone fracture fixation remains a major advance that has yet to be achieved and as such metal remains the material of choice for the foreseeable future, with the exception of nondegradable polymer composites such as biologically modified PEEK.

Treatment of Nonunions – A Brief Perspective into Scaffolds for Tissue Engineering While the evolution of fracture fixation devices has occurred for over a century it is only since the 1970s that cells were combined with scaffolds in an effort to ‘engineer’ tissue. Since then the field of tissue engineering, or regenerative medicine as it is more commonly referred to nowadays, has expanded at a staggering rate. In terms of research alone, this has been represented by a mere 12 manuscripts being published in 1990, increasing to over 400 in 2000 and more recently over 11 000 in 2013 alone. This trend is also reflected in the billion dollar investments within this field and the equal revenue generated. In keeping with bone, approximately 5–13% of fractures result in delayed/nonunion which results in the need for secondary intervention such as bone grafting. Unfortunately, current reconstructive strategies have yet to produce a satisfactory level of clinical predictability as well as encountering several other limitations such as immune rejection and severe tissue donor shortages. Therefore, research focus has shifted largely to degradable and nondegradable synthetic and natural matrices as a cell therapy approach for addressing this issue. As will be noted in previous sections, scaffolds for bone tissue engineering can be crudely divided into degradable (temporary) or nondegradable (permanent) materials. Materials include ceramics such as hydroxyapatite and or tricalcium phosphate (in various forms and chemical ratios), natural (collagen, platelet rich plasma) or synthetic polymers (polylactic acid, polyglycolic acid) and more recently composite or hybrid materials. Several studies exist that have used combinations of the approaches listed above to produce bone (Chen et al., 2010). A limitation of this strategy however, is use of osteoconductive materials (a material that will support cell attachment but will not drive differentiation). The inclusion of bone ‘inducing’ growth factors such as bone morphogenetic protein (BMP)-2 and BMP-7 has provided an osteoinductive (ability to induced osteoblast differentiation/function) dimension to tissue engineering. While limited success has been achieved the new generation of bone tissue constructs tend to favor multiphase materials that improve overall mechanical strength and allow different rates of degradation and thus differing rates of growth factor release to improve tissue integration. Again, the data to emerge from these studies have attracted interest, nonetheless, experimental inconsistencies across laboratories have meant the ‘correct’ combination of factors, their concentration, time of release, and carrier type are all widely debated (Chen et al., 2010). Moreover, preclinical data show promise for cell free (Lee et al., 2011; Lyons et al., 2010) and cell seeded scaffolds (Janicki et al., 2010) for de novo bone formation, the major clinical limitation for widespread

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application of these therapeutic strategies is lack of vascularization. Bone cannot survive without the presence of local vasculature. Therefore, the recently revised approach for bone tissue engineering incorporates the need for angiogenesis, the formation of blood vessels. It is fair to state that research in this area is relatively exploratory with some noteworthy data emerging (Wernike et al., 2010; Amini et al., 2012; Nguyen et al., 2012). Nevertheless, reproduction at a clinical scale is yet to be achieved and remains the major restraint for routine clinical application of engineered bone constructs (Nguyen et al., 2012, for review).

Summary Many noteworthy developments have been made for improving and controlling tissue integration upon implantation of a device. Despite advances, the widespread clinical application and success of these developments are relatively limited. Currently metal still remains the material of choice for fracture fixation given that its desirable mechanical properties are yet to be equaled by its alternatives. Nonetheless, more eloquent surface modifications have seen a new wave of materials such as PEEK becoming more serious contenders. There remains much debate regarding the reasoning behind elective removal of devices. However, microtopographical polishing of titanium and its alloys have proved to ease the intraoperative issues of excessive bony on-growth thus allowing the debate to fundamentally focus on secondary health issues. Bioresorbable implants have proven effective for low stress loaded regions such as craniomaxilliofacial; however, their adaption for long bone fracture remains exploratory. In the field of bone tissue engineering, the overwhelming differences in experimental ideals have meant that progress to the clinics has been slow. It is predicated that this will remain to be the case until more definitive requirement guidelines that incorporate the need for vascularization are agreed.

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Tissue Response to Biomaterials Jiao Jiao Li, University of Sydney, Sydney, NSW, Australia; Kolling Institute, Northern Sydney Local Health District, St Leonards, NSW, Australia; and Sydney Medical School Northern, University of Sydney, St Leonards, NSW, Australia Hala Zreiqat, University of Sydney, Sydney, NSW, Australia © 2019 Elsevier Inc. All rights reserved.

Tissue Response to Biomaterials Sequence of Tissue Responses to Biomaterials Initial Response to Injury Provisional matrix formation Inflammation Acute inflammation Chronic inflammation Granulation tissue formation Foreign Body Reaction Fibrous Capsule Formation Influence of Biomaterial Properties on Directing Tissue Responses Bioinert, Bioactive, and Biodegradable Materials Metals Polymers Ceramics Summary Further Reading

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Glossary Apoptosis Process of programmed cell death. Autograft A tissue graft transplanted from one part of the body to another in the same person. Cytokines Small proteins secreted by cells with specific effects on the interactions and communications between cells. Extracellular matrix A collection of extracellular molecules secreted by cells that provide structural and biochemical support to the surrounding cells. Exudation Escape of fluid, proteins, and blood cells from the vascular system into tissues in response to injury. Growth factors Proteins that can stimulate or regulate many aspects of cell function, such as survival, proliferation, migration, and differentiation. Neovascularization Formation of new blood vessels or capillaries through the proliferation, maturation, and organization of endothelial cells. Osseointegration Direct structural and functional connection between living bone and the surface of a load-bearing artificial implant. Phagocytosis Process by which living cells ingest or engulf other cells, particles, or material fragments. Scaffold Biomaterial(s) engineered into a three-dimensional construct to support or induce cellular activities necessary for tissue repair or regeneration. Stress shielding Bone resorption and reduction in bone density due to the removal of typical levels of stress from the bone by an orthopedic implant.

Tissue Response to Biomaterials Biomaterials are natural or synthetic materials that are intended for use in a biological environment, often in the form of implants or medical devices. They are the central element of biomedical engineering solutions to replace, restore, or improve the function of diseased or damaged tissues or organs. Bioactive biomaterials can even be used to induce natural tissue healing. All biomaterials placed inside the body will elicit tissue responses, the nature and extent of which depends on the biocompatibility and bioactivity of the biomaterial. This article will introduce the sequence of tissue responses to biomaterials, including injury, provisional matrix formation, acute inflammation, chronic inflammation, granulation tissue formation, foreign body reaction, and fibrous capsule development. This will be followed by a summary of characteristics of the common classes of biomaterials for use in medical

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applications, including metals, polymers, and ceramics, and the influence of these characteristics on directing the tissue response to different classes of biomaterials.

Sequence of Tissue Responses to Biomaterials The implantation of a biomaterial into the body triggers a cascade of tissue responses involving inflammatory and wound healing responses to the injury, as well as the body’s responses to the foreign material (Fig. 1). The inflammatory and wound healing responses to injury are common across different classes of biomaterials, but the nature and extent of the body’s responses to foreign material are determined by the properties of the biomaterial, which include but are not limited to its composition, degradation rate, morphology, shape and size, porosity, roughness, and surface chemistry.

Initial Response to Injury Immediately following the injury elicited by biomaterials implantation, changes in blood flow occur that cause the exudation of fluid, proteins, and cells from the vascular system into the injured tissue. The response to injury in vascularized connective tissue is similar across different tissues and organs. This response triggers the cellular events that later occur as part of the inflammatory response.

Provisional matrix formation The initial blood–biomaterial interactions result in blood protein deposition on the biomaterial surface, and are the first step in provisional matrix formation. The provisional matrix, or blood clot, forms within minutes to hours of biomaterial implantation and consists of several components, including fibrin, inflammatory products released by the complement system, activated blood platelets, inflammatory cells, and endothelial cells. Fibrin is the blood-coagulating protein produced from fibrinogen by activation of the coagulation and thrombosis systems. It is the key component of the provisional matrix and has an important role in facilitating neovascularization to support the wound healing process. Platelets are activated during the formation of the fibrin network, and contribute to fibroblast recruitment by releasing a number of growth factors such as platelet-derived growth factor and

Fig. 1

Sequence of tissue responses following biomaterials implantation.

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transforming growth factor b. The fibrin network is also decorated with adhesive proteins such as fibronectin and thrombospondin, providing substrates for cell adhesion and migration. The provisional matrix hence acts as a sustained release system, supplying a rich mixture of substances that initiate tissue repair, recruit inflammatory cells and fibroblasts, and facilitate ongoing wound healing processes.

Inflammation Inflammation is defined as the reaction of vascularized living tissue to local injury. It serves to contain, neutralize, dilute, or isolate the agent or process causing the injury. In the context of biomaterials implantation, inflammation triggers a series of events that may heal and reconstitute the implant site through the formation of fibroblastic scar tissue, or if possible, regeneration of the original tissue. The intensity and duration of the inflammatory response depends on the severity or invasiveness of the implantation procedure, and the biocompatibility of the implant. The inflammatory response comprises an initial acute phase and a subsequent chronic phase, which are followed by granulation tissue formation (Fig. 2).

Acute inflammation Acute inflammation is usually of short duration, lasting from minutes to days depending on the severity of the injury. It is marked by the release of fluid and blood plasma proteins, and the arrival of leukocytes which initially comprise neutrophils and later macrophages. Excess blood flow into the site of injury due to vessel dilation results in the influx of numerous blood and tissue proteins such as cytokines and growth factors. These attract the infiltration and accumulation of leukocytes from the blood vessels into the injury site, which constitute the most important feature of the inflammatory reaction. Neutrophils are the primary cell population during the first few days following injury. They arrive in large numbers and are primarily involved in the early prevention of infection, by phagocytosing microorganisms and foreign materials, and cleaning up injury-related debris. Neutrophils are relatively short lived and typically disintegrate and disappear after 24–48 h. In the following days to weeks, the neutrophils are replaced by monocytes which differentiate into macrophages. These cells are very long lived and can persist for up to months. The macrophages phagocytose microorganisms, while also cleaning up the remains of dead tissue cells and neutrophils. Acute inflammation usually resolves within a week, a prolonged course of which is probably indicative of infection.

Chronic inflammation Chronic inflammation is caused by persistent inflammatory stimuli, such as the continuous presence of a biomaterial implant. Chronic inflammation can persist for weeks to months or even years, and the extent of the response depends on the degree of injury. It is generally characterized by the presence of monocytes, macrophages, and lymphocytes. This is accompanied by the proliferation of blood vessels, which is essential for supplying the necessary nutrients to support wound healing, and proliferation of connective tissue to reconstruct the injury site. Of the cell types present in chronic inflammation, macrophages are the most important as they secrete a large collection of biologically active products. These products have a key role in promoting the growth of fibroblasts and blood vessels, as well as tissue regeneration and remodeling. The chronic inflammatory response to biomaterials may be caused by the properties or size of the biomaterial, or its motion within the implant site, but is usually of short duration and confined to the implant site. Due to their size disparity, the biomaterial can rarely be engulfed by the attached inflammatory cells and can lead to frustrated phagocytosis, where the continuous release of

Fig. 2 The temporal variation in the acute inflammatory response, chronic inflammatory response, granulation tissue development, and foreign body reaction to implanted biomaterials. The intensity and time variables depend on the extent of injury created by the implantation, as well as the size, shape, topography, and chemical and physical properties of the biomaterial. Reproduced from Anderson JM and Shive MS (2012) Biodegradation and biocompatibility of PLA and PLGA microspheres. Advanced Drug Delivery Reviews 64: 72–82, with permission from Elsevier.

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leukocyte products occurs in an attempt to degrade the biomaterial. Chronic inflammation with the presence of collections of lymphocytes and monocytes at extended implant times, from weeks to months to years, may be an indication of long-term infection.

Granulation tissue formation Following biomaterial implantation and initiation of the healing process by the action of monocytes and macrophages, the formation of granulation tissue commences through the proliferation of fibroblasts and vascular endothelial cells at the implant site. Granulation tissue may appear as early as 3–5 days following biomaterial implantation, with the characteristic features of proliferating fibroblasts and small blood vessels. It is found on the surface of healing wounds and has a soft and pink granular appearance. The small blood vessels are formed through neovascularization, which involves the proliferation, maturation, and organization of endothelial cells into blood capillaries. The connective tissue in the granulation tissue is developed by fibroblasts, which actively synthesize collagen and proteoglycans. The proteoglycans predominate in the early stages, and this later changes into collagen (especially collagen type I) which forms the fibrous capsule. Macrophages are almost always present in granulation tissue, but fibroblasts typically dominate. The mode of wound healing depends on the extent or severity of the injury or defect associated with the biomaterial implantation. In a clean wound with neat wound edges such as that created by surgical incision, healing proceeds with minimal infection and tissue loss. However, in a large defect associated with extensive tissue loss, the original architecture cannot be completely reconstituted by the regenerating tissue, resulting in significant fibrosis and the formation of large amounts of granulation or scar tissue that will often remain permanently.

Foreign Body Reaction While inflammation and granulation tissue formation are common to normal wound healing and biomaterial implantation, the foreign body reaction stage is unique to biomaterials. The foreign body reaction to biomaterials consists of foreign body giant cells (multinucleated macrophages) and the components of granulation tissue, including fibroblasts, capillaries, and macrophages. The nature of the foreign body reaction is determined by the form and topography of the implanted biomaterial. For example, implants with smooth and flat surfaces such as silicone breast prostheses have a foreign body reaction composed of a layer of macrophages one to two cells in thickness. In contrast, implants with rough surfaces such as expanded polytetrafluoroethylene vascular prostheses have a foreign body reaction consisting of individual macrophages and foreign body giant cells on the surface, with varying degrees of granulation tissue adjacent to the macrophage layer. A response similar to that for rough surfaces is seen for particles of implant debris or small fragments of wear debris generated by the implant. The foreign body reaction may be present at the tissue–implant interface for the lifetime of the implant. Multinucleated foreign body giant cells, which are coalesced tissue macrophages derived from circulating blood monocytes, typically persist on implanted biomaterial surfaces for decades and may contribute to implant biodegradation. There is evidence for the involvement of antiinflammatory Th2 helper lymphocytes in the development of the foreign body reaction to implanted biomaterials, which is mediated by interleukin (IL)-4 and IL-13. Downstream of this, the macrophage mannose receptor has a critical role in the fusion of macrophages to form foreign body giant cells. Macrophages and foreign body giant cells which remain adherent on biomaterial surfaces may continue to release cytokines or become quiescent, and their apoptosis is in part controlled by the surface chemistry of the biomaterial. The foreign body reaction is a precursor stage to fibrous tissue encapsulation of the biomaterial.

Fibrous Capsule Formation The end stage of the foreign body reaction and healing response to a biomaterial is often fibrous encapsulation. This involves fibroblasts creating a vascular and collagenous fibrous capsule that confines the implant and prevents it from interacting with the surrounding tissue. Fibrous capsule formation is common for large impervious implants such as breast prostheses, and typically occurs for bulk bioinert and biocompatible materials. A number of exceptions result from the wound healing cascade for biomaterials with specific properties. Porous biocompatible materials are not surrounded by a fibrous capsule because they allow the ingrowth of connective tissue within the pores. Porous or nonporous bioactive materials are also not surrounded by a fibrous capsule because they interact favorably with native tissues and encourage direct tissue apposition or bonding to the material surface. Biodegradable materials do not have a persisting foreign body reaction or fibrous capsule because they are eventually metabolized by the body and removed through physiological processes. Materials with toxic properties, which should not be used for biomedical applications, cause prolonged complications and ongoing cell death and often do not proceed to the end stages of the healing process.

Influence of Biomaterial Properties on Directing Tissue Responses All biomaterials implanted into the body will elicit tissue responses which in turn influence the function of the biomaterial. A critical component in the development of biomedical engineering devices is to select the appropriate biomaterials and processing methods that will result in favorable tissue responses, to enhance or at least not interfere with the function of the device. Bioinert,

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bioactive, and biodegradable materials each produce a different set of tissue responses, due to changes in material behavior that influence biological interactions. The classes of biomaterials used in biomedical engineering, which can be broadly divided into metals, polymers, and ceramics, also each exhibit specific properties that generate different tissue responses. Regardless of their classification, it is important to note that all biomaterials used for biomedical engineering applications should be biocompatible, meaning that they should not be toxic or harmful to living tissue and should not cause immunological rejection after implantation.

Bioinert, Bioactive, and Biodegradable Materials Bioinert materials are nontoxic, biologically inactive materials that cause the formation of a fibrous capsule with variable thickness around the implant. Highly bioinert materials, such as alumina, zirconia, silicon nitride, and gold, can have a very thin fibrous capsule if the implant has a tight mechanical fit, but the fibrous capsule can be hundreds of microns thick in the case of a loosely fitted implant that enables movement. Weakly bioinert materials, such as stainless steel and nitinol, have a proportionally much thicker fibrous capsule. Bioactive materials are nontoxic, biologically active materials that induce the formation of a direct chemical bond between the implant and host tissue by eliciting a biological response at the interface. These materials are often also biodegradable and are extensively used in tissue engineering and regenerative medicine. They are favored by the body and can play an active role in promoting tissue repair and regeneration at the implant site. Examples include calcium phosphates which can form a direct bond with hard tissues such as bone, and bioactive glasses which can form a direct bond with both hard tissues and soft tissues such as muscle and ligament. Biodegradable materials are nontoxic, biologically degradable materials that become gradually replaced by host tissue within a certain period of time after implantation. The material is resorbed through processes such as erosion and cell-mediated enzymatic degradation, and the products are removed through normal physiological processes. The formation of a thin fibrous capsule around the material is possible depending on its bioactivity. Fast-degrading materials such as collagen and b-tricalcium phosphate (b-TCP) generally have favorable biological interactions, but may be degraded within weeks to months and provide insufficient support for complete tissue repair. Slow-degrading materials such as polycaprolactone (PCL) and hydroxyapatite are generally more bioinert, often persisting for years within the implant site and may cause ongoing tissue remodeling and reorganization around the implant.

Metals Metals have a range of characteristics that are advantageous for biomedical engineering applications, including high tensile and yield strengths, and high resistance to fatigue, creep, and corrosion. Metals are commonly used in the form of alloys for biomedical engineering applications because most pure metals evoke very strong immune responses in the body. After implantation, the metal undergoes corrosion within the body and forms metal oxides on the surface. The formation of a dense oxide layer may passivate the metal and reinforce its corrosion resistance, such as in the case of titanium. Metal alloys that contain potentially toxic elements are safe for use in medical applications provided that they do not corrode or generate wear particles. For example, stainless steels (Cr and Ni) and cobalt–chrome alloys (Co and Cr) contain potentially toxic elements, but are strongly corrosion resistant and have a long history of use in clinical practice. Nevertheless, metal ions that leach from metallic implants into the blood can cause toxicity problems, which constitute an important safety consideration in implant design. Metals are commonly used for load-bearing orthopedic implants and internal fixation devices, such as hip and knee replacements, dental implants, cardiovascular prostheses, and surgical instruments. Stainless steels, titanium and its alloys, and cobaltbased alloys are the most common metallic materials used in biomedical engineering. Titanium alloys in particular have been the materials of choice for load-bearing orthopedic implants due to their superior mechanical properties, chemical stability, and in vivo biocompatibility. High corrosion resistance is a key selection criterion for the use of metals in medical applications, as seen in 316L stainless steel (corrosion resistance due to low carbon content), cobalt–chrome alloys (corrosion resistance due to presence of chromium), and Ti–6Al–4V (corrosion resistance due to passivating properties). Tissue responses to medical-grade metal implants depend largely on their surface properties. Metal implants with a smooth or polished surface tend to result in fibrous encapsulation (Fig. 3A), while implants with a rough or functionalized surface can promote osseointegration or the direct apposition of surrounding bone (Fig. 3B). More recently, a number of other metals have gained increasing popularity and research attention as candidate implant materials that provide unique properties for certain applications. For example, porous tantalum implants have been fabricated with structure and elastic modulus similar to that of natural trabecular bone, avoiding the common problems of stress shielding and insufficient bone ingrowth experienced by conventional metal implants. Biodegradable magnesium alloys are emerging as potential candidates for medical devices that benefit from complete resorption within a certain time period after implantation, such as screws and coronary stents.

Polymers Polymers have a long history of use in medical applications, featuring many different types with versatile properties and a wide range of processing options. All polymers are long-chain molecules composed of small repeating monomers, and can be derived

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Fig. 3 Light microscope images of titanium implants in the canine mandible at 4 weeks after implantation, showing the tissue interface for implants having (A) a smooth surface, with fibrous capsule formation around the implant, and (B) a surface functionalized using bone morphogenetic protein (BMP), with direct bone apposition to the implant. Hematoxylin–eosin stain, original magnification 33 . Reproduced from Xiang W, Baolin L, Yan J and Yang X (1993) The effect of bone morphogenetic protein on osseointegration of titanium implants. Journal of Oral and Maxillofacial Surgery 51(6): 647–651, with permission from Elsevier.

from natural or synthetic sources. Most polymers are relatively bioinert in the body, unless they are decorated with biologically active agents such as adhesion molecules or growth factors. Some polymers are biodegradable and undergo surface erosion or bulk erosion after implantation, which ideally should create space for regenerating tissue to bridge the implant site. Their degradation rate can usually be adjusted from weeks to years by modifying the molecular weight, crystallinity, and copolymer ratio. Polymeric materials have been extensively used in biomedical engineering applications in the form of surgical tools, implantable devices and coatings, catheters, vascular grafts, injectable materials, and tissue engineering scaffolds. Natural polymers, such as collagen (from biological tissues), chitosan (from crustacean shells), and alginate (from brown algae), are derived from natural sources. They have the advantages of possessing structures that mimic native extracellular matrix, and may already contain biochemical cues that enhance cell attachment or migration. They have been used in biomaterial systems involving cell encapsulation, injectable materials, and tissue engineering constructs and grafts, and are naturally accepted by the body. However, a common disadvantage of natural polymers is their weak mechanical properties and rapid degradation after implantation, which is partly due to their ability to be degraded by cell-mediated enzymatic digestion. Due to their extraction from natural sources, natural polymers may also exhibit large batch-to-batch variation in properties which affect their downstream functions. Synthetic polymers are an extensive class of materials that are artificially synthesized. They may be nondegradable, such as ultrahigh molecular weight polyethylene (used in hip and knee replacements), silicone (used in tube-shaped replacements such as trachea or gastrointestinal segments), polytetrafluoroethylene (used in blood vessel grafts), polyurethane (used in heart pacemakers), and poly(methyl methacrylate) (used as bone cements), or biodegradable, such as polyglycolic acid, polylactic acid and their copolymers (used as sutures and tissue engineering scaffolds), and PCL (used as long-term tissue engineering implants and drug delivery devices). Synthetic polymers are generally bioinert and have little interaction with host tissues unless they are functionalized with biological agents. They have a number of significant advantages for use in biomedical engineering applications, such as easy fabrication, flexibility in processing, and tuneable chemistry to produce desired properties. They also have a controlled chemical composition and can give highly consistent properties between batches. However, most synthetic polymers require organic solvents for processing, which may cause a toxic reaction in the body if not completely removed. Prolonged implantation of nondegradable synthetic polymers in the body can also result in the generation of wear particles (particularly for polyethylene) and the release of harmful monomers (particularly for poly(methyl methacrylate)), which may trigger chronic inflammatory and foreign body responses.

Ceramics Ceramics are inorganic, nonmetallic materials with high compressive strength. In biomedical engineering, several classes of ceramics have been the materials of choice for musculoskeletal tissue repair and regeneration, with applications as orthopedic implants and implant coatings, bone cements, tissue engineering scaffolds, prostheses for dental and maxillofacial restoration, and drug delivery systems. Commonly referred to as “bioceramics,” their main advantages for these applications are similarity in chemical composition and mechanical properties to natural hard tissues. Bioceramics may be biologically inert (alumina and zirconia), or biologically

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active and often biodegradable (calcium phosphates and bioactive glasses), each with a unique set of characteristics that make them useful for specific medical applications. Alumina and zirconia are both highly bioinert bioceramics, with high hardness, modulus of elasticity and compressive strength accompanied by excellent friction and wear properties. Their ability to avoid the generation of wear particles and sensitization of tissues make them good candidates for the articulating surfaces of artificial joint replacements, as well as dental implants. However, these materials exhibit the characteristic brittleness of ceramics and may fail under impact loading, which has been the subject of research into toughening mechanisms. Calcium phosphates and bioactive glasses are bioactive and often biodegradable classes of materials which have been extensively used as bone graft substitutes and tissue engineering scaffolds for hard tissue regeneration. These materials are inherently bioactive as they are similar in chemical composition and surface structure as the mineral phase of bone, which consists of plate-like hydroxyapatite crystals. After implantation, a series of surface reactions occur which result in the formation of a direct bond between the material and native bone. These surface reactions also mediate the adsorption of extracellular matrix proteins that promote the attachment, proliferation, and differentiation of bone-related cells, ultimately resulting in enhanced bone formation. For calcium phosphates, their bioactive properties originate from the ability to form a carbonate apatite (CHA) layer at the bone–material interface through a cell-mediated dissolution and precipitation process, which releases calcium and phosphate ions from the material into the surrounding microenvironment and encourages the precipitation of CHA microcrystals. The tissues surrounding the material hence become richly mineralized, creating an environment that is conducive to bone formation. For bioactive glasses, surface reactions involving ion substitution, glass dissolution, and mineral precipitation occur within 24 h of implantation, which accelerate CHA layer formation on the glass surface (Fig. 4). The reactivity of the glass surface is determined by the chemical composition of the glass. The CHA layer, which has a similar composition to bone mineral, initiates the adsorption of growth factors and a synchronized sequence of cellular events that result in the rapid formation of new bone. Bioactive glasses are the only bioactive materials with the ability to bond to soft tissues as well as hard tissues. The most commonly used calcium phosphate materials for clinical bone repair are hydroxyapatite, b-TCP, and biphasic calcium phosphate (BCP). Stoichiometric hydroxyapatite is relatively bioinert and slow-degrading, and can persist in the implant site with little biodegradation for 5–10 years following bone reconstruction procedures. In contrast, b-TCP is relatively fast-degrading and highly bioactive, with good ability to induce clinical bone incorporation and remodeling but may not persist within the defect site for a sufficiently long period of time for complete bone repair to occur. BCP is a two-phase ceramic composed of hydroxyapatite and b-TCP phases, which allows an optimal balance of bioactivity and biodegradability to be achieved between the more stable hydroxyapatite phase and the more soluble b-TCP phase. Bioactive glasses have been used clinically as particulates for filling contained bone defects, and have demonstrated excellent bone bonding and osteogenic ability that can exceed calcium phosphates, with comparable performance to bone autografts which are the gold standard for bone reconstruction. However, the progression from particulates to porous scaffold systems has been difficult due to the glasses losing their bioactivity after the high-temperature

Fig. 4 Scanning electron microscope image of carbonate apatite layer formed on the surface of a bioactive glass scaffold after soaking in simulated body fluid (SBF) for 28 days. The inset reveals the formation of rod-shaped crystals of biologically similar apatite. Reproduced from Rezwan K, Chen QZ, Blaker JJ and Boccaccini AR (2006) Biodegradable and bioactive porous polymer/inorganic composite scaffolds for bone tissue engineering. Biomaterials 27(18): 3413–3431, with permission from Elsevier.

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sintering process required to produce porous scaffold structures. This is due to crystallization of the glass which removes its surface reactivity and therefore ability to directly bond with adjacent tissues. In preclinical animal models, calcium phosphates with surface microporosity and bioactive glasses have demonstrated evidence of osteoinductivity, which is the ability to induce the homing of stem cells from the circulation to participate in bone formation at the implant site without the incorporation of growth factors.

Summary The implantation of a biomaterial or its prolonged contact with the body elicits a series of tissue responses that are broadly defined by the inflammatory and wound healing responses resulting from the body’s natural reaction to a foreign material. However, specific biomaterial properties such as composition, degradation rate, morphology, shape and size, porosity, roughness, and surface chemistry can also influence these tissue responses and lead to significantly different healing outcomes. When designing a new medical device that involves one or several biomaterials, the biomedical engineer must consider the tissue responses that are likely to occur for each biomaterial to ensure the long-term safety and efficacy of the device in the human body.

Further Reading Anderson, J. M. (2001). Biological responses to materials. Annual Review of Materials Research, 31, 81–110. Li, J. J., Kaplan, D. L., & Zreiqat, H. (2014). Scaffold-based regeneration of skeletal tissues to meet clinical challenges. Journal of Materials Chemistry B, 2(42), 7272–7306. Morais, J.M., Papadimitrakopoulos, F. and Burgess, D.J. Biomaterials/tissue interactions: Possible solutions to overcome foreign body response. AAPS Journal 12(2), 188–196. Rezaie, H.R., Bakhtiari L. and Öchsner A. Biomaterials and their applications. Switzerland: Springer International Publishing. Rezwan, K., Chen, Q. Z., Blaker, J. J., & Boccaccini, A. R. (2006). Biodegradable and bioactive porous polymer/inorganic composite scaffolds for bone tissue engineering. Biomaterials, 27(18), 3413–3431.

BIOMATERIALS: BIOMATERIAL APPLICATIONS AND ADVANCED MEDICAL TECHNOLOGIES Biomaterials in Dentistry

Li Wu Zheng, Prince Philip Dental Hospital, The University of Hong Kong, Hong Kong Jing Yi Wang and Ru Qing Yu, The University of Hong Kong, Hong Kong © 2019 Elsevier Inc. All rights reserved.

Restorative Dentistry Direct restorative Materials Amalgam Composites Glass ionomer cements Indirect Restorative Materials Metals Ceramics Cements Endodontic materials Denture bases and lining materials Impression materials Waxes Adhesive system Orthodontic Brackets Adhesives Archwires Elastomeric Modules and Chains Periodontics Barrier Membranes Bone Graft Materials 3D-Printed Scaffolds Implants Dentistry Implant Materials Metals Ceramics Polymers and carbon compound Oral and Maxillofacial Surgery Autogenous Grafts Bone Substitutes Titanium Tissue Engineering Further Reading

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The earliest dental materials science, dating back to the early 19th century at Northwestern University, began from the investigation of dental amalgam. For approximately a hundred years, synthetic restorative materials were the major focus in the field of dental materials until the end of the last century, when the real potential for biological engineering of tissues and organ systems was revealed. The comprehension of development and advances in dental biomaterials will benefit both dental practitioners and patients in selecting appropriate materials for clinical cases and improving treatment outcomes.

Restorative Dentistry Restorative dentistry can be traced back to ancient times. Materials used for restoration back then include cork, ivory, human teeth, and metal foils (lead and tin), etc. Nowadays, amalgam, composites, ceramics, metals, and cements are common restorative materials.

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Direct restorative Materials Amalgam Amalgam has been using to restore tooth structure since the 19th century. Dental amalgam mainly contains mercury, tin, copper, and zinc. With proper cavity preparation and manipulation of dental amalgam, this can provide adequate strength to resist masticatory forces with quite good longevity. Although amalgam has been a great success as a restorative material, its mercury content has raised concerns regarding the potential adverse effects such as neurotoxicity in recent decades. The exposure to mercury for dental personnel and environment also adds to the argument of reducing the use of dental amalgam. Moreover, esthetics has become an increasingly important aspect in the choice of materials, causing patients to be reluctant to choose amalgam’s gray color material over more esthetic materials, which are becoming more durable. Currently, with the development of other restorative materials (i.e., composites), the use of dental amalgam as a direct restorative material is phasing down globally, despite its low cost and good performance.

Composites Dental composites, or resin-based composites, are synthetic materials that combine polymeric matrix with a dispersion of glass, mineral, or resin filler particles and/or short fibers by coupling agents. Just like dental amalgam, they are used to restore tooth structure lost through trauma, caries, or other diseases. Composites can also be used as cements to cement crowns and veneers, etc. While the amalgam is phasing out in dentistry, composites have become one of the most widely used esthetic restorative materials. Traditional composites contain relatively large particles of ground amorphous silica and quartz, which gives them good mechanical properties but makes the surface of the restoration more likely to become rough from daily abrasion. In addition, many failures of composite restoration are seen at the interface between tooth and composite due to shrinkage or adhesive failure. To overcome this, microfilled composites, nanofilled composites, and other hybrid composites were developed, using much smaller particles (at the same time with a large variety in size) to fill in the matrix. With these developments, smoother surfaces are achieved, wear resistance is increased, and shrinkage is decreased without compromising the mechanical and physical properties. Composites can be classified as chemically activated (self-cure) resins and photochemically activated (light-cure) resins. The selfcure resins are supplied as two pastes; polymerization is activated when those two pastes are mixed together. Disadvantages are that the air may be incorporated into the mix during mixing, thus weakening the material, and the operator cannot control the working time after mixing. The light-cure resin is supplied as a single paste using the photosensitive initiator system and a light source for activation. It does not need mixing, which makes it stronger and less staining, and has totally controllable working time. However, it exhibits higher marginal stress during curing and only cures within limited depth (2–3 mm). Although this incremental curing demonstrates some advantages, the demand for bulk cure has never stopped. Some new products have claimed that the cure depth can be up to 4 mm (the cure depth of dual cure resins is unlimited since it is a combination of chemical and light-cure technology), but their clinical performance has not been fully assessed.

Glass ionomer cements Glass ionomer cements conventionally are acid-base materials that have been used to esthetically restore tooth structure since the late 20th century. The powder of a number of glass components mixes with the liquid of polyalkenoic acid to form a paste, then the acid-base reaction starts and stiffens the paste. The mechanical property does not suit the clinical requirements enough in the beginning, but slowly improves with time. One of the advantages of glass ionomers is the true bonding between materials and dentin/ enamel; thus they have been widely used for Class V restorations which have high requirements in adhesion, for Class II and Class III restorations in deciduous teeth, for luting of crowns, and they also can be used as bases or liners. Fluoride release is another merit of glass ionomers. The disadvantages are that they are moisture-sensitive and have relatively low strength. The newer glass ionomers are the resin-modified glass ionomers, which were introduced in the 1980s. These are a combination of conventional glass ionomer cement and light-cure resins to improve some characteristics of conventional glass ionomers such as increased strength, lower solubility, and less sensitivity to moisture. However, fluoride release of resin-modified glass ionomer is lower and the biocompatibility is not as good as that of conventional glass ionomers. Recently, glass ionomers were combined with bioactive glass (BAG) to improve their bioactivity and regenerative capacity. These materials might be a better choice in tooth restoration compared to conventional glass ionomers or resin-modified ones, due to their remineralization ability.

Indirect Restorative Materials Metals Crowns, inlays, cast posts and cores, and partial dentures are all examples of indirect metallic restorations. They are produced in the dental laboratory instead of being carried out in the dental chair. There is a variety of alloys used in dentistry; the main alloys include noble and precious metal alloys and various base-metal alloys, i.e., CoeCr alloys and titanium. Noble and precious metal alloys Gold, platinum, iridium, ruthenium, rhodium, and osmium are considered to make up the noble metals, which are very resistant to corrosion. Silver and palladium are usually referred to as the precious metals, which are expensive. High’gold alloys must have a gold content > 60% and precious metal content no < 75%. They can be classified according to their gold content into four types (I–IV): soft, medium, hard, and extra-hard. In the 1970s, alloys with reduced gold content were rising in the market because of the

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rapidly increasing price of noble metals. These alloys can be further classified into medium’gold (40%–60% gold content) and low’gold alloys (10%–20% gold content). Their mechanical properties are similar to the type III and IV gold alloys, but exhibit lower ductility. The white color makes them less popular with patients, but they can be used for posts and cores since they will be covered with other materials. Silver–palladium alloys are another kind of precious alloy that presents various properties due to their composition and are less popular than the above-mentioned ones. Base metal alloys There are two main kinds of base metal alloys: cobalt–chromium and titanium. Cobalt–chromium alloys have high hardness and modulus of elasticity, but low ductility. In addition, they are relatively low cost and have very good biocompatibility. Generally, they are a very popular choice of material in restorative dentistry. Titanium alloys exhibit high strength, low density, and have great biocompatibility. However, casting is a significant challenge for these alloys.

Ceramics Ceramics is a material that is opaque and porous, thus relatively weak. Dental porcelain mainly differs from traditional ceramics in terms of firing techniques, which make it more suitable for dental restoration. Dental porcelain has very stable chemical properties and outstanding esthetics which are unlikely to be influenced by time. It has similar thermal conductivity and coefficient of thermal expansion to enamel and dentine, and exhibits high compressive strength. However, the tensile strength of dental porcelain is very low (20–60 MPa). In addition, the surface microcracks caused by various reasons during manufacturing are the starting sites of catastrophic fracture. These drawbacks have limited their use to low-bearing areas, which are anterior regions of both mandible and maxilla, and made them unsuitable for multi-unit bridges. To overcome the disadvantages of dental porcelain, three types of dental ceramics have been developed:

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Metal-ceramics (porcelain fused to metal, or PFM), combine the positive mechanical properties of cast dental alloys and excellent esthetic property of porcelain. The choice of metals is a key element in PFM. Reinforced ceramic core systems, which are similar to PFM in that instead of using alloys to support the porcelain, they use another ceramic material with high strength and toughness yet do not offer esthetic qualities. Resin-bonded ceramics involve the ceramic being bonded to enamel and dentine directly, and thus the support comes from its own tooth structure by resins. Not only the increasing strength and toughness are necessary, but also the adhesive bond, which eliminates the surface flaws of dental ceramics and thereby reduces the probability of fractures.

Cements Cements are used in dentistry for various purposes. Some are for cavity lining or bases; others are used as luting agents to lute an indirect restoration to a prepared tooth. Basically, a liner is a thin layer of material (0.5 mm) placed on a prepared tooth to protect it, whereas a base acts as the dentin to withstand the forces applying on it. It is thicker than the liner and also protects the pulp from thermal and chemical stimulates and galvanic shock. Thus a liner or a base should have good thermally and electrically insulating properties, and not contain irritants. It should set rapidly, exhibit enough strength to resist fracture, and not move or flow easily while the filling material is being placed. Ideally, linings should be radiopaque so that any caries around the filling material can be seen. The liner or base must not interfere with the setting of the filling material or affect the properties of it. Luting cements share similar properties with linings except for the setting time. There should be enough setting time before the final seating of the restoration. In addition, it should be strong enough to assist retention and have low solubility. Commonly used cements are as follows:

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Zinc oxide/eugenol cements are mixtures of zinc oxide (powder) and eugenol (liquid). They are mainly used as a lining or base under amalgam restorations and as temporary luting cements or filling materials. Zinc–phosphate cements have zinc oxide as the major component of the powder and phosphoric acid solution as the liquid. They are widely used as luting cements and can also be used as linings with adjustment of the powder/liquid ratio to change the consistency. However, they may have an irritant effect as a liner in deep cavities. Calcium hydroxide cements have a low strength and high solubility, and therefore are usually used as linings beneath the base of zinc phosphate cements or other base materials, and are not suitable for luting. Nevertheless, this material does have other properties that make it crucial to dentistry such as the fact that it can be used for pulp capping and root canal sealing.

Endodontic materials Endodontic materials are used in endodontic treatment, which is the procedure to save the tooth when the pulp and/or periradicular tissues are injured. These materials can be generally classified into two groups: materials used to maintain the vitality of the pulp (pulp capping materials), and materials used to disinfect (irrigants and intracanal medicaments) and fill the pulp in root canal therapy. Pulp capping materials should be able to induce hard-tissue formation in a superficial way, protect the pulp from further invasion of bacteria, and not have side effects so that the pulp can be alive. Hard-setting calcium hydroxide cements are the most commonly used pulp capping materials. This material causes the formation of a 1–1.5 mm thick necrosis layer in the superficial pulp. The layer will undergo calcification eventually and it is called a dentine-bridge.

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Commonly used irrigants and intracanal medicaments include 0.5% sodium hypochlorite solution (as a disinfectant or antimicrobial), EDTA (as a chelating agent and lubricant), chlorhexidine (antimicrobial), non-setting calcium hydroxide (as an antibacterial agent), and some antibiotics and anti-inflammation drugs (non-setting Ledermix paste). The most widely used canal-filling material is gutta-percha. Gutta-percha point is one of the application of gutta-percha, warm lateral or vertical condensation can be applied to it to soften and compact it. The other use is the hot gutta-percha application system, which injects the heat-softened gutta-percha into the root canal system. In addition to gutta-percha, sealer cements are also needed in root canal filling procedures. These cements include zinc oxideeugenol cements, calcium hydroxide-containing sealers, glass-ionomer cements, polymers, and mineral trioxide aggregate, etc. Due to space limitations, details of these materials will not be discussed here.

Denture bases and lining materials Nowadays, most denture bases are fabricated using acrylic resins. The materials are usually supplied as a powder and liquid. The major component of the powder is beads of polymethylmethacrylate (PMMA), which can be added to the liquid (mainly methylmethacrylate monomer) to form a mixture. Shortly after the mixing, the material first has a sandy consistency, then becomes a sticky mass, and finally comes the dough stage. The dough at this point has lost most of its tackiness and should be packed into the mold to prevent it from becoming tough and rubbery. Acrylic denture base materials can be classified into different groups according to their method of curing. Generally, there are heat curing materials, autopolymerizing materials, thermoplastic, light-activated, and microwave-cured materials. With sufficient thickness, the material showed acceptable mechanical properties; however, the impact strength and fatigue strength are relatively poor. Dimensional stability is also a problem in acrylic resins in addition to its thermal insulated nature, which is not an advantage for denture base because it prevents the protective reflexes to stimuli in oral mucosa. Moreover, the unreacted monomer that remained in the denture base may cause mucosal irritation and sensitization of tissues. Denture lining materials can be roughly classified into soft acrylics and silicone rubbers. The hard reline material is to be used when the denture base has gone through some dimensional change and become less fit so that it needs relining of the fitting surface. The soft lining materials are to provide a cushion when the denture is under load, and a tissue conditioner is very similar except it is softer and can only remain soft for 1–2 days, while the softness of the temporary soft lining materials can last for 1–2 months. There also exist permanent soft lining materials used for patients who cannot bear a hard base. However, no soft lining materials are truly permanent in practice.

Impression materials Impression materials are used in dentistry to record the details of intraoral structures to fabricate a reproduction of teeth and soft tissues for the construction of dental prostheses. These materials should be able to produce an accurate replica of the intraoral structure, to prevent deformation and be atraumatic when removing from undercuts; they should also have proper setting time and biocompatibility. They can be categorized as rigid and elastic impression materials. Rigid impression materials include plaster and compo/zinc oxide-eugenol; however, since they cannot engage the undercuts, their application is limited nowadays. Elastic impression materials can be further divided into hydrocolloid and elastomeric impression. Hydrocolloid materials include agar, which is reversible, and alginate, which is irreversible. Elastomeric materials include polysulfide, polyether, condensation-cured silicone, and addition-cured silicone. The choice of which impression material to use in each case will depend not only on the specific needs of each case, but also on the impression technique and tray to be used. Alginate is currently one of the most popular impression materials. It is supplied as dust-free powders. After mixing with proper amount of water in a rubber bowl with a spatula, it is ready for impression taking. Two to three minutes after the surface tackiness has been lost, it can be removed from the oral cavity. However, it does not produce very accurate surface detail, and has poor dimensional stability. A snap-removal technique is required to minimize permanent deformation. It is thus not recommended for the fabrication of crowns and bridges. The need for more satisfying impression materials promotes the development of elastomeric impression materials. These materials have a much lower chance of permanent deformation when removed and are able to reproduce the surface detail very accurately, but they are hydrophobic and therefore contamination of saliva will result in loss of surface detail to some degree. Another problem of these materials is that they all undergo setting shrinkage due to polymerization, but in general, the shrinkage is very low (with polyether and addition-cured silicones being the lowest and condensation-cured silicones being highest).

Waxes Waxes have a number of applications in dentistry. A major use is as pattern waxdin other words, as modeling wax and/or inlay wax. In the process of fabrication of dentures or restorations, there is a stage which is to produce the wax pattern of the dental appliances on the model (indirect technique) or in the mouth (direct technique). This wax pattern determines the size and shape of the appliance needed. A lost-wax technique is then used to replace the waxes using alloys or polymers. The pattern wax must be able to shape the appliance accurately, and once it is formed there should be no dimensional change. Also, the wax should be removable either by boiling or burning, and does not leave a residue.

Adhesive system Facilitating adhesive materials in restorative dentistry instead of mainly relying on the retention form of restorative materials has many advantages. One of the most important advantages is the conservation of tooth structure.

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Enamel bonding The acid-etch technique is most widely used for bonding resins to enamel. In 1955, Buonocore found that by modifying the enamel surface with the phosphoric acid solution, he could make the surface more receptive to adhesion, which led to the development of the acid technique. The etching process increases the bonding area by creating etch pits, into which the resin is absorbed by capillary attraction. The adhesion is mainly based on this micro-mechanical interlocking. The acid-etch technique presents excellent bonding between the etched enamel and the resin. Dentine bonding Dentine is a heterogeneous tissue composed of about 70% inorganics (hydroxyapatite, HA), 20% organics (collagen), and 10% water by weight. This makes it much more difficult for bonding compared to enamel. Another problem is that it is a vital tissue and unlikely to create a dry dentine surface, and will damage the pulp if doing so. Furthermore, the dentine is covered by a smear layer. However, the strong demand for dentine adhesive systems has promoted the development of dentine adhesive systems. There are three principal components in dentine-bonding agents: conditioner, primer, and sealer. A conditioner is an etchant, which is used to remove the smear layer on the dentine. The smear layer is formed because of abrasion or burr cutting. Various acids were used as conditioners: phosphoric acid, oxalic acid, EDTA, etc. After the conditioner is applied to the dentine, it dissolves the HA and opens the dentinal tubules, producing a demineralized layer on the surface. The primer acts as an adhesive to bond the hydrophobic resin to the hydrophilic dentine. The primer should be able to saturate fully into the demineralized layer, otherwise the remaining demineralized dentine will become a weak region in the restoration. The coupling agent in the primer is carried in a volatile solvent which is for the water removal in the dentine so that a hydrophobic resin can be bonded tightly. A sealer is the resin used to fill the cavity, which bonds to the dentine through primer. Total-etch technique This technique involves etching the enamel and dentine at the same time, usually with 35% phosphoric acid for 20 s. Dentine-bonding agents can be categorized according to the number of steps used clinically as three-step, two-step, and singlestep systems. Those consisting of a dentine conditioner, primer, and the sealer are three-step systems. Two-step systems and singlestep systems are developed to make the process more efficient and easier to use. The two-step systems combine either the conditioner and primer together (self-etching primers) or the primer and the sealer together (one-bottle bond systems). The single-step systems are supplied in two bottles; the practitioner only needs to mix the two components and applies them to the enamel and dentine surface for the drying and light-curing. However, the bond strength is lower compared with the aforementioned two systems. Apart from the application in restorative dentistry, adhesive systems can also be used for the adhesion of orthodontic brackets using the acid-etch technique.

Orthodontic Materials used in orthodontic field mainly include brackets, adhesives used to bond the brackets onto teeth, archwires, and elastomeric modules and chains.

Brackets Orthodontic brackets are small orthodontic attachments (metal or ceramic) secured to a tooth for fastening an archwire. Each attachment is either soldered or welded to a previously placed band enclosing the tooth, or is bonded directly onto the tooth. Metallic brackets can be further categorized into stainless steel, non’nickel or low’nickel stainless steel, and titanium brackets. The nickel component in traditional stainless steel has revealed genotoxic effects and may cause some allergic reaction in patients, therefore non’nickel or low’nickel stainless steels are substitutes for the traditional ones. This type of steel exhibits similar or higher hardness but may show lower corrosion resistance. Recently, titanium has been used as an alternative to produce brackets because of its superior biocompatibility and higher corrosion resistance. However, titanium brackets have lower hardness compared to the stainless steel brackets. Thus, wear presents a problem in titanium brackets. Ceramic brackets are developed for esthetic purposes. Other esthetic materials have also been useddfor example, plastics and polycarbonate-based materials. Plastic brackets undergo extensive creep deformation and discoloration during use, and a low hardness limited their use in orthodontic area. As an alternative, ceramic brackets exhibit higher hardness and stiffness; however, they showed a higher incidence of fracture due to brittleness and aging in the oral environment. Newly developed ceramic brackets showed better mechanical properties and even more esthetics due to improved transparency, and are more compatible with intraoral environment.

Adhesives Adhesives used in the orthodontic field are similar to those in restorative dentistry. The acid-etch technique is commonly used for bonding. Details can be found in the relevant section.

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Archwires Archwires should move the teeth with light and continuous forces which increase patients’ comfort and optimize the treatment process. The material used for the archwire should have good elasticity and be elastic for weeks at least when a force is applied on it. To meet with these criteria and objectives, several alloys have been used as orthodontic archwires. Each has its own advantage in a particular stage of treatment, but no single alloy suits all the stages involved in a treatment.

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Stainless steel alloys

Stainless steel alloys are very strong and have high corrosion resistance; however, they have lost their popularity due to the new materials being introduced to the archwire market, but they are still in the market because the advantage of low cost.

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Cobalt–chromium alloys

This material contains mainly cobalt and chromium, but also some iron and nickel. Not only does it show similar stiffness to stainless steel, but also its formability can be changed by heat treatment. In other words, it can be formed to a specific shape and be heated to increase the resilience and strength of the material.

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Nickel–titanium alloys

This alloy has a shape memory effect and was first developed by W.F. Buehler in the 1960s. The shape memory effect means that the nickel–titanium wires remember its original shape before deformation and have the ability to return to that shape when not loaded. Compared with stainless steel wires, nickel–titanium wires exhibit higher strength and resilience, and a lower modulus of elasticity. Together with the ability to recover fully from 8% strain deformation (other alloys can recover from only 1% strain deformation), nickel–titanium wires show some superiorities as orthodontic wires: lighter forces can be applied on teeth, reducing the number of wire changes required during the treatment. Although they also have some limitations (e.g., restricted formability), they still are a great improvement in the history of orthodontic wires.

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Beta-titanium alloys

Compared with stainless steels, these alloys applied gentler forces on teeth during deactivation and showed higher springback. They exhibited better formability over nickel–titanium alloys. However, their stiffness is lower, and they are not resilient enough to withstand the friction created during the movement of teeth.

Elastomeric Modules and Chains The use of elastomeric modules and chains is being phased out due to the application of self-ligating brackets, but these modules are still used to close small gaps in the anterior teeth. They should be able to be stretched easily without great loss of energy, have high tensile strength and stiffness, and have high resilience to recover their original form fully and rapidly. Further reading is recommended for the detail of the materials.

Periodontics In recent years, the management of periodontal diseases has evolved from simply the debridement of periodontal pockets to the regeneration of periodontal tissues. The use of biomaterials has become crucial in the treatment of patients.

Barrier Membranes A barrier membrane is a device originally used to prevent epithelial migration into a specific area in the guided tissue regeneration (GTR) procedure. Barrier membranes are divided into two major types: resorbable and nonresorbable (Table 1). In order to achieve ideal function in a surgical site, the features of a barrier membrane need to meet certain criteria including biocompatibility, tissue integration, cell-occlusiveness, space-making, and clinical manageability. Cellulose and ePTFE were the first non-resorbable membranes developed and used in early GTR procedures. The titanium mesh was developed to reinforce the membrane and shaped for more space. The disadvantages of non-resorbable membranes are that they have to be removed, which inspired the development of resorbable membranes. Nowadays most practitioners prefer resorbable membranes, especially collagen membranes because they are biocompatible and user friendly. They are gradually degraded without additional surgery once positioned into the defect site. In recent years, acellular dermal matrix, which is human skin processed in such a way as to remove epidermis and all dermal cells, has been used in root coverage and socket preservation. The removal of cells of this graft warrant minimal risk of rejection and inflammation after being grafted. This grafting material acts as a bioactive scaffold for migration of fibroblasts, epithelial cells, and endothelial cells. It is also a decent material in periodontal plastic surgeries for increasing the zone of keratinized tissue. Compared to non-resorbable membranes, resorbable membranes are weaker in space maintaining, thus the appropriate selection of a membrane is crucial for a good clinical outcome.

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Types of membranes for periodontal regeneration

Resorbable

Non-resorbable

Synthetic membrane • Polylactic • Polylactic/polyglycolic • PL, PG and trimethylcarbonate • PG and TMC • Polyethylene glycol Natural membrane • Collagen • Acellular dermal matrix • Oxidized cellulose mesh

• Cellulose acetate filter • PTFE • ePTFE • Titanium mesh • Ethylene cellulose • Rubber dam

Bone Graft Materials Graft materials are nowadays widely used in bone regeneration and reconstruction. Table 2 lists the main categories of graft materials apart from autograft which is beyond the scope of this chapter. Autologous grafts perform well in osteogenic, osteoconductive, and osteoinductive cases, but require sophisticated operating skills and more surgical time. While autografts are harvested from the patients themselves, allografts are harvested from an individual other than the patient. Allograft bone is typically sourced from a bone bank. Allografts are basically osteoconductive and osteoinductive; however, they do not possess the ability of osteogenesis. The advantage of allograft over autograft is that the former eliminates donor site morbidity. However, even though the preparation of these freeze-dried bone allografts would reduce the immunogenicity of the grafts, they still carry a risk of infection and immunoreaction. In addition, the fresh frozen bone is more prone to cause disease transmission and rejection. Xenografts are derived from other species such as cows, horses, coral, etc. Deproteinized bovine bone minerals are the most generally used bone grafts to date, and are frequently used in site augmentation, sinus augmentation, ridge preservation, and materials for peri-implant defects. They are also often used in conjunction with autografts and resorbable membranes. Alloplastic grafts are completely synthetic and they usually contain calcium and phosphate. The most used alloplast is HA, which has a good quality of osteoconduction, hardness, and acceptability by bone. Alloplastic grafts such as calcium phosphate are often used as an osteoconductive matrix and should be tightly packed in the adjacent host bone to maximize ingrowth. Compared to bone grafts mentioned above, alloplastic grafts have less of an issue with limited supply and carry minimal risk of disease transmission. In addition to being used in oral and maxillofacial surgery, the most common application of bone grafting is in periodontics and implant dentistry. They are either in block or particulated to adapt better to a defect. In the development of tissue engineering, barrier membranes and graft materials have been combined with growth factors, for example, enamel matrix derivative (EMD), to generate better outcomes. EMD is a purified acid extract from the enamel matrix of porcine fetal tooth. It has been used in treatment of peri-implant and periodontal diseases due to its profound ability to stimulate the soft and hard tissues surrounding the teeth, and restore cementum, periodontal ligament, and alveolar bone. There are existing products which mix EMD and Straumann bone ceramic (Straumann Emdogain PLUS, Straumann Holding AG, Basel, Switzerland), which proved useful in regenerative periodontal surgery such as treatment of vertical bone defects.

3D-Printed Scaffolds Bone grafting materials are now the treatment of choice to obtain adequate bone volume in periodontal diseases and implant sites. However, as mentioned above, these grafting materials are not flawless, and all have their pros and cons. With the development of 3D-printing technology, 3D-printed scaffolds have become an attractive alternative to bone grafts in bone augmentation, socket preservation, and sinus augmentation. The manufactural technique of the “printing” includes inkjet printing, extrusion printing, and laser-assisted printing with different printing methods using specific biomaterials and generating different resolutions. Using

Table 2

Types of bone graft materials

Allograft

Xenograft

Alloplast

Fresh frozen bone (FFB) Mineralized freeze-dried bone allografts (FDBA) Demineralized freeze-dried bone allografts (DFDBA)

Bovine Porcine Equine Coralline hydroxyapatite Algae hydroxyapatite

Hydroxyapatite Tricalcic phosphate b-TCP and HA Bioactive glass Polylactic acid and polyglycolic acid

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computer-aided design and computer-assisted manufacturing technologies based on a CT scan of the specific defect to create a personalized grafting scaffold and well fitted at the defect site, this technique could be a solution worth considering. There is also great potential for the application of scaffolding matrices in periodontal tissue regeneration as a membrane and grafting material. Although the efficacy of 3D-printed biomaterials has been demonstrated preclinically, they still have a long way to go before being widely used in dental clinics.

Implants Dentistry The application of dental implants can be traced back to AD 600, when humans used bamboo stakes to replace missing teeth. Since then, dental practitioners struggled with materials until 1952, when Professor Branemark developed a threaded implant design made of pure titanium and found that titanium apparently bonded irreversibly to living bone tissue. Subsequently, the materials of dental implants have developed further to meet varied clinical requirements. To produce a high-performance dental implant to replace a missing tooth, the properties of the biomaterials, including the bulk properties, surface properties, and biocompatibility, have to be taken into consideration (Table 3).

Implant Materials Dental implants are usually classified according to their placement situation with the tissue. There are essentially three categories: endosseous implants, subperiosteal implants, and transosseous implants. Four classes of materialsdmetals, ceramics, polymers, and carbon compounddare used in modern dental implants (Table 4).

Metals Metals are currently the most popular material in dental implants and they are almost exclusively titanium based. The characteristic properties of titanium meet many requirements in a dental implant, such as primarily high strength and high corrosion resistance, Table 3

Properties of an implant biomaterial

Properties

Function

Bulk properties • Modulus of elasticity of 18 GPa

• High creep deformability • High tensile, compressive, and shear strength • High yield strength, high fatigue strength • A minimum ductility of 8% • Hardness and toughness Surface properties • Surface tension and surface energy • Surface roughness Biocompatibility • Good corrosion resistance

Table 4

Uniform distribution of stress Minimizes the relative movement at implant bone interface Materials with high creep values could better tolerate high masticatory forces Prevent fractures Improve functional stability Improved stress transfer Prevent brittle fracture under cyclic loading Crucial for contouring and shaping of an implant Decrease the incidence of wear Prevents fracture of the implants Determines the wettability of implant Alteration in surface roughness improve cell attachment to the implant Implant bio-material should be corrosion resistant as corrosion could result in roughening of the surface, weakening of the restoration, release of elements from the metal or alloy, and toxic reactions

Classification of implant materials

Types

Materials

Metallic

Pure titanium (CPTi) Titanium alloys (Ti–6AL–4U) Co–Cr alloys Stainless steel Precious metals Bio-inert ceramics Bioactive and biodegradable ceramics PMMA, PTFE, PE, PSF

Ceramic and ceramic coated Polymers Carbon compound

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which ensure good biocompatibility. Thus titanium is viewed as the material of choice in dentistry. Both commercial pure titanium (CPTi) and Ti–6AL–4U alloy possess excellent corrosion resistance; Ti–6AL–4U alloy has a slightly higher modulus of elasticity than pure titanium, but it is more expensive. The disadvantage of titanium material is that there might be esthetic issue due to its gray color; in addition to that, there is an unesthetic display of metal when soft tissue recession occurs. For high load-bearing zones such as in posterior areas with lower esthetic requirements, implant materials with high strength such as CPTi or titanium alloys should be considered. Stainless steel has also been used for many years; however, it is contraindicated in patients who are allergic to nickel, as it contains 7%–8% nickel to stabilize the austenitic structure. Meanwhile, the toxicity of nickel is still a concern. Apart from that, the low cost and good mechanical properties are its significant advantages.

Ceramics Because of their better biologic inertness compared with metals, and their good strength, ceramics have been used as dental implant materials. Ceramic materials are classified into two types: inert ceramics, which are non-reactive materials such as aluminum oxide (AL2O3), carbon, and zirconia (ZrO2); and bioactive and biodegradable ceramics such as glass ceramics, BAG, and calcium phosphate ceramics (CPCs). Inert ceramics from aluminum, zirconium oxides, and titanium have been used for endosteal plate form, root form, and pin type of dental implants. This type of material is high in compressive, tensile, and bending strength, with better physical characteristics such as similar color as natural tooth, and minimal thermal electrical conductivity. However, its inherent brittleness has limited its application. Modern dentistry has placed more emphasis on bioactive and biodegradable ceramics as these materials possess similar constituents to those of normal tissue, with excellent biocompatibility, and a similar modulus of elasticity to bone. They could be gradually resorbed and even replaced by tissue over time. As a matter of fact, they are primarily used as scaffolds or to coat metallic implants because of their good tissue bonding property. Bioactive and biodegradable ceramics, such as calcium phosphate ceramics (CPCs), have attracted attention in the development of implant materials. Calcium phosphate materials have been applied as bone augmentation and replacement materials as well as in implant materials. They could be formed to serve as structural support under relatively high-magnitude loading conditions for implant applications, or as alloplastic grafts and mixed with drugs, collagen, and growth factors such as bone morphogenic proteins. In addition, CPC coating on metallic surfaces has been used in endosteal and subperiosteal dental implant designs to improve the longevity and surface biocompatibility of an implant. Mixtures of HA, tricalcium phosphate (TCP), and teracalcium phosphate can be either plasma sprayed or coated to produce a bioactive surface on an implant. In fresh extraction site or a newly grafted site, HAcoated implants seem a better choice for its greater implant bone interphase and higher shear bond strength.

Polymers and carbon compound Polymeric implants were used in the past few decades. Due to their low mechanical strength and susceptibility to fracture, they are nowadays only used as adjuncts to enable stress distribution along with implants. Carbon compounds have an excellent quality of biocompatibility as well as a modulus of elasticity similar to that of bone. However, their low compressive strength results in their application only as coatings on metallic and ceramic materials.

Oral and Maxillofacial Surgery In terms of maxillofacial diseases, the most crucial concern, apart from treatment of the actual disease, is the esthetic requirements and functional demands of the oral and maxillofacial region. The need to reconstruct the facial region for the reason of tumor resection, trauma congenital disorders, or even esthetic need is one of the main challenges of maxillofacial surgeons. Autologous bone or soft tissue grafts are routinely used in reconstruction design nowadays. However, the donor site morbidity, the medical condition of the patient to tolerate a major surgery, the special training and sophisticated technical requirements of a surgeon, and the relatively high rate of infection and graft resorption are always concerns. Apart from the widely used autologous grafts, many other biomaterials have been used alone or in combination with bone grafts. As mentioned above, allografts, xenografts, and alloplastic materials have been used as bone substitutes in periodontics and implant surgeries to restore small defects. However, will these biomaterials be an appropriate choice of treatment for major head and neck reconstruction?

Autogenous Grafts To date, autogenous bone is the gold standard for maxillofacial reconstruction. An autogenous graft is prepared from a healthy part of the patient’s own bone and grafted onto the defect area. A graft of cortical bone ensures a good structural support and a reduced resorption, while a graft of cancellous bone ensures early revascularization. The types of autogenous grafts include free bone graft, particulate cancellous bone marrow graft, pedicle composite flaps, microvascular free-flap transfer, bone marrow aspirates, and plasma-rich protein. Among these, microvascular free-flap transfer performs noticeably well in major defect reconstruction since it is vascularized with adequate blood supply and excellent in functional and esthetic reconstruction. However, the morbidity associated with the harvesting procedures of these autogenous grafts, the limited availability when the volume of the defect is huge, and the prolonged operation time have all inspired the continuous seeking of alternatives.

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Bone Substitutes Bone substitutes include allografts, xenografts, and alloplastic grafts. Allografts are costly and proved to have a high rate of resorption with poor mechanical properties due to chemical and radiation pre-treatment. Therefore, allografts are limited to small and medium defects, and are not suitable for reconstruction of critical-size defects in the maxillofacial region. Xenografts are less costly than allografts. They gained popularity in reconstruction of small bony defects because of their slow resorption property. A common example of a xenograft is a kind of natural porous bone mineral called Bio-OssÔ (Geistlich, Wolhusen Switzerland), which is derived from bovine bone and is commonly used in periodontics. However, as with allografts, xenografts are not suitable for the reconstruction of critical-size defects due to their poor mechanical properties and lack of vascularization. Alloplastic grafts contain a wide range of materials (Table 2), some of which have been used for bone regeneration. Ceramics such as HA and calcium sulfate have been used in the restoration of periodontal defects and small correction of maxillofacial regions such as sinus augmentation, but they still possess some limitations when it comes to repairing major defects in the maxillofacial region. While ceramics slowly find their way to repair small defects in maxillofacial surgery, BAG has also proved its usefulness and bright future in the reconstruction of small cranial-facial bone defects. BAG is a silicate glass-based material which undergoes a specific surface reaction. It forms a bond with hard and soft tissues when implanted. BAG demonstrates controlled resorption in optimal time, efficient bioactivity, and the ability to modulate cell migration, with a cortical bone-like modulus of elasticity. However, similar to ceramics, it is not suited for reconstruction of continuity defects of jaw bones due to the lack of the required mechanical properties. It has been applied for contour refinements as inlay for frontal bone, nasal dorsum, mandibular angle, and alveolar ridge augmentation. Polymeric materials used in maxillofacial surgery are mainly synthetic polymers. Although natural polymers such as collagen (used in periodontal regeneration in combination with other grafting materials), alginate, hyaluronic acid, peptide hydrogen, and chitosan are biocompatible, they could only be used in combination with other materials because of their disadvantage in being water soluble. Synthetic polymers could be prefabricated as well as contoured with a burr in a surgery. Prefabricated polymers that have played a vital role in maxillofacial surgery include hard-tissue replacement polymers, PMMA, and porous polyethylene (MEDPOR; Stryker, Kalamazoo, Mich., USA). PMMA, as the first substitute used for adult cranial reconstruction, presented a 9.5% infection rate in cranial reconstruction. However, in recent years it has also been used as a framework in combination with bioglass fabricated into a customized porous implant to repair calvarial and midface bony defects. This promising material proved good functional and esthetic outcomes with no observation of long-term complications. Porous polyethylene implants are also commercially available and have been used for nasal and malar augmentation, orbital floor reconstruction, and calvarial defect reconstruction. They are dense implants with a pore size of 100–250 mm. However, these materials still have limitations such as the risk of infection, exposure, and extrusion. Resorbable poly(L-lactide) (PLLA) and polyglycolide polymers (PGA) have also been widely used in maxillofacial surgeries. Their combination product copolymers, which take in the most positive characteristics of them both, are poly(lactide-co-glycolide) (PLGA) and PLLA-PGA; these are widely used as bioresorbable plates and screws for the fixation of bones. These resorbable plates have gained popularity in treatment of pediatric mandibular complex fractures, since they function well in realignment and stable positioning of rapidly healing fracture segments with no secondary removal operations needed. They have also been used in internal fixation of adult maxillofacial fracture in certain anatomy sites when strength is not much of a concern. The limitations of these resorbable materials are their limited ability to withstand masticatory forces and the chances of inflammation when the materials begin to degrade. In addition, PGA and PLGA have been combined with HA or b-TCP to form a composite scaffold in an attempt to increase the degradation time and improve the material’s mechanical properties.

Titanium Titanium not only has outstanding performance in implant dentistry, it is also a common choice for maxillofacial surgeon for its desirable mechanical properties and good biocompatibility. It is used as a bone tray for mandibular reconstruction, for cranioplasties, orbital floor implants, condylar reconstruction, and titanium plates and screws for internal fixation of fractures. Its high mechanical properties and lightness render it a good choice when used as fixation plates/screws for major force-bearing fracture sites such as mandibular angle fractures. This metallic material has also been used in combination with polymeric materials such as polyethylene; the combined material possesses not only the mechanical strength offered by titanium but also the porous biocompatible surface offered by polyethylene. Titanium has the advantage of visibility on post-op with minimal distortion in MRI images, yet the other side of the coin is that it produces artifacts and interfere with the interpretation of MRI images just like other metallic objects. Other limitations of titanium are that the thermal conductivity of the metal would bring discomfort to patients in cold weather, and the gray color would be an esthetic issue when soft tissue is thin and pervious to light. The functional and esthetic reconstruction in the maxillofacial region has always been a vital task for surgeon and researchers. The application of biomaterials for maxillofacial repair and reconstruction are summarized in Table 5.

Tissue Engineering Biomaterials have been widely used in almost all fields of dentistry; however, none of them is able to completely restore or replace the structure and function of missing tissues. The burning needs in restoration and reconstruction in clinic inspire and promote the

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Maxillofacial region Extra-oral Maxillofacial fracture Pediatric complex fracture Maxillofacial bone defect Facial augmentation Intra-oral Ridge augmentation Maxillary sinus lift Temporomandibular joint replacement

Repair/reconstruction requirement

Materials commonly used

Alternatives

Strong mechanical strength, stable positioning Stable positioning, biocompatibility Reconstruction, protection

Titanium plates/screws

–

Resorbable plates/screws Autogenous grafts

Restoration of bone volume and contour, esthetic

Autogenous grafts, titanium, bioactive glass, polymers

Titanium plates/screws Titanium mesh, polymers, injectable cement, tissue engineering Resorbable polymer, polymer face filler, 3D scaffolds

Restoration of function and esthetic, implant placement Implant placement Function, esthetic

Titanium, deproteinized bovine bone minerals Mineral composite Cast cobalt–chromium–molybdenum alloy, titanium with polyethylene

Bioactive glass, injectable cement ceramic, tissue engineering Tissue engineering –

Modified from Deb, S. (2015). Biomaterials for oral and craniomaxillofacial applications. S. Karger AG.

development of tissue engineering, a brand new technique using a combination of scaffolds, cells, and biologically active molecules to assemble functional constructs that restore, maintain, or improve damaged tissues for medical or dental purposes. Tissue engineering evolved from the field of biomaterials, but has been growing in scope and importance and is now an independent field. The great potential of tissue engineering in dentistry has advanced the study and clinical trials in periodontal regeneration, dental pulp regeneration, and maxillofacial reconstruction; however, the imperfect technique, high cost, risk of biological contamination, and ethical issues are concerns that need be solved before the engineered tissues can be widely used in clinics.

Further Reading Agarwal, S., Gupta, A., Grevious, M., & Reid, R. R. (2009). Use of resorbable implants for mandibular fixation: A systematic review. Journal of Craniofacial Surgery, 20(2), 331–339. Anusavice, K. J., Shen, C., & Rawls, H. R. (2013). Phillips’ science of dental materials. London, UK: Elsevier Health Sciences. Asa’ad, F., Pagni, G., Pilipchuk, S. P., et al. (2016). 3D-printed scaffolds and biomaterials: Review of alveolar bone augmentation and periodontal regeneration applications. International Journal of Dentistry, 2016, 1239842. Carpena Lopes, G., Narciso Baratieri, L., de Andrada, C., Mauro, A., & Vieira, L. C. C. (2002). Dental adhesion: Present state of the art and future perspectives. Quintessence International, 33(3). Craig, R. G., & Powers, J. (2002). Restorative dental materials, 11th edn. St. Louis: Mosby. Darby, I. (2011). Periodontal materials. Australian Dental Journal, 56(Suppl. 1), 107–118. Deb, S. (2015). Biomaterials for oral and craniomaxillofacial applications. S. Karger AG, Basel. Dhuru, V. B. (2004). Contemporary dental materials. Oxford: Oxford University Press. Eliades, T. (2007). Orthodontic materials research and applications: Part 2. Current status and projected future developments in materials and biocompatibility. American Journal of Orthodontics and Dentofacial Orthopedics, 131(2), 253–262. Jones, J. R. (2015). Reprint of: Review of bioactive glass: From Hench to hybrids. Acta Biomaterialia, 23(Suppl), S53–82. Lyngstadaas, S. P., Wohlfahrt, J. C., Brookes, S. J., et al. (2009). Enamel matrix proteins; old molecules for new applications. Orthodontics & Craniofacial Research, 12(3), 243–253. McCabe, J. F., & Walls, A. W. (2013). Applied dental materials. Chichester: John Wiley & Sons. Misch, C. E. (2008). Contemporary implant dentistry. St. Louis: Mosby Incorporated. Payne, K. F., Balasundaram, I., Deb, S., et al. (2014). Tissue engineering technology and its possible applications in oral and maxillofacial surgery. British Journal of Oral & Maxillofacial Surgery, 52(1), 7–15. Profeta, A. C., & Huppa, C. (2016). Bioactive-glass in oral and maxillofacial surgery. Craniomaxillofacial Trauma and Reconstruction, 9(1), 1–14. Sam, G., & Pillai, B. R. (2014). Evolution of barrier membranes in periodontal regenerationd"Are the third generation membranes really here?”. Journal of Clinical and Diagnostic Research, 8(12). Ze14–17. Van Noort, R., & Barbour, M. E. (2013). Introduction to dental materials4: Introduction to dental materials. Elsevier Health Sciences. von Recum, A. F. (1998). Handbook of biomaterials evaluation: Scientific, technical and clinical testing of implant materials. London, UK: Taylor & Francis, 2nd edn.

Biomaterials in Ophthalmology Rachel L Williams, Hannah J Levis, Rebecca Lace, Kyle G Doherty, Stephnie M Kennedy, and Victoria R Kearns, University of Liverpool, Liverpool, United Kingdom © 2019 Elsevier Inc. All rights reserved.

Introduction Refraction Contact Lenses Intraocular Lenses Space Filling Vitreous Substitutes Scleral Buckles Orbital Implants Flow Control Electric Stimulation Tissue Regeneration Natural Scaffolds Biological Polymers Synthetic Polymers Conclusion References

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Introduction Innovations in biomaterials science and engineering have the potential to make a significant contribution to the development of treatments for ophthalmic diseases and thus to reduce the burden of vision loss on the global community. The loss of vision has major social and economic consequences not only to the individual but also to society at large, and with the increasing age of the population, this will have greater consequences. There are functional requirements of the eye and visual system that allow vision to occur, and various diseases and trauma can disrupt these processes. Biomaterials have a role in addressing several of these functional problems such as refraction, space filling, flow control, electric stimulation, and tissue regeneration. This article will discuss the biomaterials used in each of these categories.

Refraction The cornea is the transparent window that allows light to enter the eye. It also has a major role in focusing light, providing about 80% of the eye’s refractive power, with the rest being achieved by the lens. To overcome problems with the ability of the eye to refract the light and focus it onto the retina, biomaterials have been used in the form of contact lenses and intraocular lenses.

Contact Lenses Contact lenses not only are traditionally used to correct refractive error but also can be used to assist wound healing as a bandage contact lens or, alternatively, as a drug reservoir. The material properties that must be considered for contact lens selection are high oxygen permeability so oxygen can reach the cornea and prevent vascularization, high water content for tear film wettability and comfort, and resistance of protein/lipid/mucus deposits on the contact lens (Refojo, 1996). There are four main groups of materials used for contact lenses; these can be classified as either “hard” or “soft.” Soft flexible contact lenses include hydrogels (based on poly(hydroxyethylmethacrylate) (pHEMA)), silicones, and silicone hydrogels. Gas-permeable materials (usually made from fluorosilicone acrylates) are rigid and classed as hard contact lenses. The largest proportion of contact lenses on the market today are silicone hydrogel at 64%. These lenses combine different ratios of each material to achieve both the benefits of the high oxygen permeability from the silicone and increased wear comfort from the hydrogel (Caló and Khutoryanskiy, 2015; Kirchhof et al., 2015). Soft contact lenses have also been used for bandage contact lenses; these aim to prevent necrosis after surgery and facilitate wound healing by protecting the eye from external assault and sources of infection; they also keep the ocular surface hydrated and relieve pain by isolating friction during blinking. An example of when these are routinely used is after keratoprosthesis (Carreira et al., 2014; Thomas et al., 2015). A bandage contact lens that offered a therapeutic antimicrobial effect could be beneficial as a replacement for conventional antimicrobial eye drops in treating infectious keratitis. Gallagher et al. have demonstrated this with a hydrogel synthesized from poly-ε-lysine with additionally bound biomolecules to achieve a hydrogel with optimized

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mechanical and antimicrobial properties (Gallagher et al., 2016). Another potential application for contact lenses is their use in drug delivery to the front of the eye; this would be beneficial as there is a low bioavailability of ophthalmic eye drop (about 5%) due to the high drainage via the tear duct or down the cheek. One of the challenges is to achieve a sustained therapeutic release of the drug over a period of time. Simple methods involve soaking commercial hydrogel contact lenses in solutions containing drugs; however, the release profile of the drug is uncontrolled, leading to an initial high overdosing period over a few hours followed by a long underdosing period, making them unsuitable for long-term release. An alternative method is to immobilize drugs onto the surface of hydrogel contact lenses that may require modification using polyethylene glycol (Kirchhof et al., 2015; Xinming et al., 2008). Other methods to trap drugs and release them from contact lenses include molecular imprinting, colloid encapsulation, and polymeric nanoparticles (Dixon et al., 2015; Maulvi et al., 2016). Ciolino et al. examined the properties of a copolymer of poly(hydroxyethylmethacrylate) and methacrylic acid (pHEMA/MAA) contact lens that encapsulated a poly(lactic-co-glycolic) acid (PLGA) film that incorporated the glaucoma drug latanoprost. They demonstrated that as the PLGA drug film degraded, the contact lens delivered a therapeutic amount of the drug over a 4-week period in vivo without any signs of cytotoxicity (Ciolino et al., 2014). Drug delivery in the eye, and the biomaterials used in general, is a large topic and will not be covered further in this article.

Intraocular Lenses Cataracts are the leading cause of preventable blindness worldwide. During surgery, the cloudy lens is removed from its capsular bag and replaced by a polymeric intraocular lens (IOL). Over the years various polymers and designs have been used for IOLs, in an attempt to prevent postoperative complications associated with scarring. The polymers used can be grouped into three main groups: hydrophobic acrylic (e.g., phenylethyl methacrylate (PEMA) and phenylethyl acrylate (PEA)); hydrophilic acrylic (e.g., poly hydroxyethylmethacrylate (pHEMA)); and silicone (e.g., poly dimethylsiloxane (PDMS)). Due to the nature of the surgery, the chosen material should be foldable, so it can be inserted into an incision of 3–4 mm diameter or less to minimize the damage caused by the surgery and postoperative scarring. How the IOL responds to the native tissue plays a key role in scarring. Postoperative complications occur when residual lens epithelial cells migrate onto the previously cell-free posterior capsule in which the lens is housed. Once here, these cells dedifferentiate into fibroblast-like cells causing the posterior capsule to wrinkle and disrupt the path of light to the back of the eye (Apple et al., 1992) (Fig. 1). This is known as posterior capsule opacification (PCO). The incidence of PCO varies between studies, both the IOL material (hydrophobic vs hydrophilic) and design (sharp- or round-edged IOLs) can play a part in this (Auffarth et al., 2004; Yuen et al., 2006). Although IOL material and design choice may slow down the development of PCO, it does not prevent it. Rønbeck et al. demonstrated in a 12-year postoperative review of three IOLs including a round-edged heparin surface-modified poly(methyl methacrylate) IOL, a round-edged silicone IOL, and a sharp-edged hydrophobic acrylic IOL that there was no significant difference in the incidence of PCO after 12 years between the acrylic and silicone IOL regardless of edge shape. There was, however, a significant difference short term (5-year postsurgery), in which patients with acrylic IOLs had significantly less PCO than those with silicone IOLs (Rønbeck and Kugelberg, 2014). With an increasing ageing population, other methods need to be investigated to prevent the long-term effects of PCO; one possible strategy could be the use of IOLs as a drug delivery system (see Liu et al., 2013). Injectable gel-like polymers, based on polysiloxane, have been investigated as an alternative to conventional IOLs. The gel structure can be injected inside the capsule bag, leaving it intact. The gel conforms to the shape of the capsular bag, while the capsular bag supports the gel. These gels have the added benefit of accommodation, similar to the natural lens, which is not fully possible with

Fig. 1 Photograph of a donor eye with intraocular lens implant and early posterior capsule opacification formation. Early fibrosis (white arrows) was observed as scar tissue at the periphery and the development of Soemmering’s ring (red arrow).

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the conventional IOLs (de Groot et al., 2001). By varying Young’s modulus of polysiloxane gels to 0.8 kPa, similar to the native human lens (1 kPa), lenses can undergo changes in refractive power during equatorial stretching, which simulates accommodation and have a higher lens power than the age-matched natural lenses (Koopmans et al., 2003). However, the incidence of PCO remains a problem and in addition, Nd:YAG laser capsulotomy, which is the removal of the opaque posterior capsule via a laser, cannot be performed as this would affect the accommodation capabilities of the soft lens. Alternative methods to eliminate PCO have been investigated, for example, aggressive cytotoxic chemicals to kill remaining LECs prior to injecting the gel-like lens (van Kooten et al., 2006); however, due to the level of aggression needed to kill all remaining LECs, further complications of cornea opacification have been observed (Koopmans et al., 2011).

Space Filling Vitreous Substitutes Replacement of the vitreous humor is required if the native vitreous body needs to be removed, either to facilitate surgical treatment for retinal detachment or if the vitreous body itself is compromised due to other conditions including infection, tumor, or trauma (Kleinberg et al., 2011). Vitreous substitutes also have the potential to be used as drug delivery devices. The vitreous humor is predominantly composed of interpenetrating networks of collagen fibrils and hyaluronan molecules (with various other components, including cells) forming a clear hydrogel. It provides support to the surrounding structures, absorbs mechanical trauma, and is involved in the circulation and regulation of oxygen, metabolites and nutrients. The composition and structure of the vitreous humor and vitreoretinal interface has been described in detail by Sebag (1998, 1992). Silicone oils (Giordano and Refojo, 1998) were first used to treat retinal detachment in the 1960s and have been more commonly used since the availability of vitrectomy. They are currently the only class of vitreous substitute available for longterm use (it is usually removed within 3–6 months, although there are patients with permanent silicone oil (Morphis et al., 2012) but are used much more in Europe than in the United States (D’Amico, 2016). The conventional belief is that their primary mode of action is to block the flow of fluid through the retinal breaks, excluding the inflammatory factors that can lead to the formation of scar tissue that can cause the retina to distort and detach (Wong and Williams, 2005). The use of silicone oil has, however, been associated with the development of cataract (Heimann et al., 2008) and corneal endothelial graft failure (Friberg and Guibord, 1999). Early formulations caused problems relating to the presence of low-molecular-weight components (Pastor et al., 1998); all currently available products are highly purified to remove these components. Silicone oils have a propensity to emulsify. This is likely to be due to a combination of high shear forces, insufficient interfacial tension between the oil and aqueous phases in the eye, and the presence of proteins that both lower the interfacial tension and stabilize emulsified droplets. The formation of emulsions affects the passage of light but, more significantly, is associated with the development of complications such as glaucoma (Ichhpujani et al., 2009) and migration into the anterior chamber (Light, 2006). It has been suggested that emulsification is underreported, as droplets that can be observed clinically are relatively large compared with those that can be measured from samples retrieved from patients but studied in vitro (Chan et al., 2015). A number of strategies have been developed to attempt to reduce emulsification, including the use of oils with higher shear viscosity (Scott et al., 2005; Chan et al., 2011); there are no randomized clinical trials, however, showing superior emulsification resistance of 5000 mPa$S oil over 1000 mPa$S oil. A more recent strategy has been the development of silicone oils with increased extensional viscosity, achieved by adding a small percentage of a high-molecular-weight component to standard 1000 mPa$S oil (Williams et al., 2010). These oils, which are in clinical use, also experience shear thinning, making them relatively easy to inject (Williams et al., 2011). Silicone oils have a lower density than the aqueous solutions that fill the remainder of the cavity; thus, they are unsuitable for treatment of pathology in the lower part of the vitreous cavity. A range of “heavy” silicone oils have been developed, and a few are in clinical use (Heimann et al., 2008). They are based on standard silicone oil with the addition of perfluorohexyloctane or a partially fluorinated olefin. The resulting oils, which have densities of 1.06 and 1.02 g/cm3, respectively, are reported to result in anatomical success, but complications such as emulsification exist (Ozdek et al., 2011; Wickham et al., 2010). A more recent development is the combination of the addition of perfluorohexyloctane and the high-molecular-weight additive, resulting in a “heavy oil” that should have increased emulsification resistance (Caramoy et al., 2015), although a large body of clinical evidence is not yet available for this tamponade. An alternative strategy, which aims to take advantage of the light and oxygen diffusion properties of silicone oil while mimicking the natural structure of vitreous and vitreoretinal interface, is to contain the silicone oil (or other fluid) within a silicone rubber capsule within the eye (Lin et al., 2012; Yang et al., 2014). This technique has been tested in clinical trials, although has not been widely adopted. Hydrogels (Swindle-Reilly et al., 2016; Chirila et al., 1998; Su et al., 2015) are the most widely researched alternative to silicone oils, because of their potential to more closely mimic the structure of the vitreous humor. Hydrogels should be designed to have appropriate physical, chemical, and biological properties, including the ability to gel in situ; however, to date, none has made it into clinical practice. Unlike silicone oils, they are not able to exert a significant tamponade effect (Liang et al., 1998), so are limited to conditions that require space filling and as drug delivery devices. Many of them are reported to degrade once implanted into the vitreous cavity. Hyaluronan- and collagen-based gels have been widely investigated, but even following the addition of other components and cross-linking, these materials do not have the required physical properties and have a lower density than water, limiting their use for tears in the upper area of the vitreous cavity or are still susceptible to degradation in vivo (Liang et al., 1998;

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Schramm et al., 2012; Pruett et al., 1979; Barth et al., 2016). Work on various other natural and synthetic polymers, including polyacrylamide (Santhanam et al., 2016), polyethylene glycol (Annaka et al., 2011), and chitosan (Yang et al., 2008), has been reported.

Scleral Buckles Scleral buckles are devices that are placed within or on the sclera and are used to treat retinal breaks that are associated with retinal detachment. The principle is that by displacing the sclera, vitreoretinal traction is reduced, and fluid, that includes inflammatory factors, is moved away from the retinal breaks. Although it remains a successful treatment, with clinical outcomes comparable with other treatments (Khan et al., 2015), it is being used less frequently. Devices based on silicone are the only materials that are currently used clinically. They may be solid structures, often bands between 2 and 5 mm wide (Schepens and Acosta, 1991), or sponges. One of the advantages of silicone is that it does not encourage the attachment of tissue, which can cause complications if the implant needs to be removed. Furthermore, they are sufficiently tough and pliable to allow the surgeon to manipulate them during implantation. The tendency of other materials to allow tissue attachment is one of the reasons why they are not clinically successful. Hydrogels, biologically derived polymers and degradable polymers have insufficient mechanical strength for the required duration (Baino, 2010).

Orbital Implants It is sometimes necessary to remove the eye of the patient, for example, following cancer, severe infection, or trauma. An orbital implant (Baino and Potestio, 2016; Sami et al., 2007) can be used to restore the resulting cavity, which is important for aesthetic and psychological reasons. Orbital implants can be classified as integrated or nonintegrated. From a materials perspective, this can be linked to whether the implant surface is porous or not, although some are designed to allow tissue ingrowth on the posterior surface while usually having a smooth anterior surface. A recent Cochrane review (Schellini et al., 2016) has been unable to determine whether either strategy is better. Polymethyl methacrylate is in clinical use to produce orbital implants of various designs (Baino and Potestio, 2016) as it has a good track record and is cheap. Polyethylene, particularly in porous form with a smooth anterior surface or coating to reduce irritation to surrounding tissue such as the conjunctiva and to allow connection of extraocular muscles, is also used (Karesh et al., 1994; Jung et al., 2012). Porous silicone (Son et al., 2012) and expanded polytetrafluoroethylene have been investigated (Dei Cas et al., 1998), but are not used clinically. Ceramic implants (Baino and Vitale-Brovarone, 2015), both hydroxyapatite and alumina, are used in various clinically available implants (Suter et al., 2002; De Potter et al., 1994; Jordan et al., 2003), both allowing desirable tissue ingrowth. Promising early results have also been reported for patients receiving a bioactive glass–ceramic implant (Crovace et al., 2016). Optimizing microstructure and topography may be important in determining and enhancing the clinical response (Mawn et al., 1998; Patel et al., 2010), although this is not yet widely studied. Composites of the earlier materials have been investigated, although many have not been successful in patients. In particular, bioactive ceramics have been used as coatings to encourage tissue ingrowth. Those that have made it into clinical use include one based on hydroxyapatite and silicone, and another based on porous polyethylene and 45S5 bioglass (Ma et al., 2011).

Flow Control Normal eye pressure ranges from 12 to 22 mmHg. The intraocular pressure is maintained in the normal range by the flow of aqueous humor, which is produced by the ciliary body, from the back of the eye into the anterior chamber through the pupil. It then flows out of the eye through the trabecular meshwork into Schlemm’s canal and is absorbed into the bloodstream. The production and flow of the aqueous humor is an active, continuous process that is needed for the health of the eye. If the flow is restricted, pressure builds up in the eye, which can cause damage to the optic nerve, leading to vision loss. Glaucoma is usually characterized by a high intraocular pressure that may be due to the reduction of fluid flow through the trabecular meshwork or in the collector channels or venous plexus further downstream. In many cases, glaucoma can be treated with eye drops that cause a lowering of the intraocular pressure either by reducing the amount of aqueous humor produced or by increasing its outflow. In some situations, however, this is not sufficient or eyedrops are no longer functional, and alternative strategies are needed to ensure the pressure remains in the normal range. In general, this involves surgery to produce a channel through the obstructed tissue to allow outflow of the fluid. Minimally invasive glaucoma surgery has been designed to make glaucoma surgery simpler to perform with less trauma to the eye and a more easily reproducible technique so that the procedure is available to all ophthalmologists rather than just glaucoma specialists (Manasses and Au, 2016). This has resulted in several different designs of glaucoma microstents that can be classified by the targeted outflow destination: into Schlemm’s canal, the suprachoroidal space, or the subconjunctival space. The iStent, which is 1 mm long and has a lumen of 120 mm, is designed to bypass the trabecular meshwork and allow outflow directly into Schlemm’s canal. It is made from a heparin-coated titanium alloy (Ti6AI4V), and it is possible to insert two or three at the time of cataract surgery. Placement of the istent in conjunction with cataract surgery in mild to moderate glaucomatous eyes has been shown to cause a moderate reduction in IOP and reduce the dependency of the patient on medication (Manasses and Au, 2016). The Hydrus Microstent is similarly a trabecular meshwork bypass device. It is much longer at 8 mm and is designed to follow the curve of Schlemm’s canal preventing its compression. It is manufactured

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from NiTi alloy and is fenestrated along its length to aid outflow of aqueous into Schlemm’s canal. One year clinical results show that this can maintain a reduced IOP, rarely achieving less than 15 mmHg, and a reduction in use of medication over at least 12 months. The Cypass stent is designed to bypass the trabecular meshwork and provides a direct outflow channel into the suprachoroidal space. It is made from polyimide and is 6.35 mm in length with a 300 mm lumen with fenestrations along its length to facilitate outflow. Like the Schlemm’s canal devices, IOP is rarely observed below 15 mmHg and generally resides above 16 mmHg. The iStent Supra uses a similar approach and is a 4 mm tube made from fenestrated polyethersulfone with a Ti sleeve. The more traditional way to release fluid from the anterior chamber is into the subconjunctival space. There are two glaucoma microstents that are designed to take advantage of this route and bypass all potential outflow obstructions. The XEN 45 is manufactured from porcine-derived gelatin cross-linked with glutaraldehyde. It is a 6 mm long tubular structure with a lumen of 45 mm that has been designed specifically to control the rate of fluid flow to maintain the correct IOP. The InnFocus MicroShunt (Fig. 2) similarly drains into the subconjunctival space. It has been through various design iterations to ensure as little trauma to the tissues and a lumen size that minimizes hypotony (Pinchuk et al., 2017). The final design resulted in a stent that is 8.5 mm long with a 70 mm diameter lumen. The material that it is manufactured from was a key part of the design process. The objective was to design a synthetic thermoplastic elastomer that is biostable and caused less of an inflammatory response than conventional materials. The material is synthesized at InnFocus and is a triblock copolymer of a soft polyisobutylene central block with glassy polystyrene end blocks called poly(styrene-block-isobutylene-block-styrene) or SIBS. It has enhanced biocompatibility and long-term stability resulting in less inflammation than other materials. A lowering of the IOP to  14 mmHg was recorded after 3 years in 95% of the patients (Manasses and Au, 2016).

Electric Stimulation Damage to retinal tissue via various disease mechanisms has led to research to attempt to augment or replace the function of damaged tissues using retinal implants. Patients in many of the clinical trials that have been run to date suffer from retinitis pigmentosa (RP), a group of hereditary diseases that damage the photoreceptor cells of the outer retina but leave the nerve cells of the inner retina intact. Retinal implant devices provide electric stimulation to these remaining nerve cells that transmit the signals along their axons, which form the optic nerve, to the visual cortex in the brain. The devices aim to capture light and convert this to an electric pulse that is delivered to nerve cells, particularly the ganglion cells. These electric impulses cause patients to perceive flashes of light called phosphenes. Many groups are investigating various different approaches to light or image capture, processing and conversion to electric impulses, and electrode design and location (Maghami et al., 2014; Ha et al., 2016). The Alpha IMS, developed by Retinal Implant AG based in Reutlingen, Germany, uses a CMOS active pixel sensor array that is implanted subretinally, beneath the transparent retina, replacing the defunct photoreceptor cells. It received CE marking for sale in Europe in 2013, and the latest version, the Alpha AMS, obtained CE marking in 2016. The devices contain a photodiode-amplifying microelectronics electrode set within each pixel of the array. Incident light on the photodiode is converted and amplified into an electronic pulse and is delivered via the electrodes to nearby nerve cells. The circuitry is created using a CMOS process. This is encapsulated by an insulating material, such as a polymer, that will prevent fluids in the eye from shorting the circuitry and thus causing device failure. Electrodes are connected to the circuitry and made of conducting noncorrosive materials, such as TiN or IrO (Graf et al., 2009). Work by this group in 1999 investigated the biocompatibility of various semiconductor and electrode materials including SiO2, Si3N4, TiN, and Ir (Guenther et al., 1999). Rat retinal cells were seeded onto the materials, and biocompatibility was determined by examining cell attachment and growth. Cells grew on all materials, and although there were fewer cells on TiN compared with Ir, it appears that this material is still used for electrodes. The chip is mounted on a metallized

Fig. 2

Diagram of the InnFocus MicroShunt in situ in the anterior chamber angle. Provided by InnFocus Inc.

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polyimide ribbon cable that exits the eye and extends toward an induction coil located behind the ear, similar to a cochlear implant. The coil must be hermetically sealed, usually in silicone or ceramic, similar to other implants. Power is wirelessly transferred from an external supply to this coil and via the ribbon cable to the chip. Polyimide has previously been demonstrated as a biocompatible, flexible substrate for electrode arrays (Klinge et al., 2001; Seo et al., 2004; Richardson et al., 1993). The Argus II (Fig. 3), developed by Second Sight based in Sylar, CA, the United States, is an epiretinal device that has been granted permission to be marketed in the United States and Europe. It consists of a pair of glasses with a mounted camera that is connected to a small processing unit. Images are captured and processed outside the eye, and then, information and power are transmitted by transcranial induction to a paired coil and electronics unit that is held in place outside of the eye by a scleral band. The scleral band and antenna are encased in silicone (Second Sight Medical Products, 2013), which functions similarly to scleral buckles mentioned previously. The electronics unit is hermetically sealed and sends information and electric pulses, via a ribbon cable that passes through the eye, to an electrode array attached to the inner surface of the retina. The conducting wires of the electrode cable are encased in polyimide, but the end of cable and the electrode array are coated with silicone (Second Sight Medical Products, 2013). Electric impulses are delivered to the retina via platinum electrodes. The electrode array is held in position with a spring-loaded titanium alloy retinal tack (de Juan et al., 2013). One of the drawbacks of the systems outlined earlier and those similar is the fact that information is captured outside the eye and needs to be transferred, along with power, from an external source to the inside of the eye. A device that was contained completely within the eye at the site of operation would be beneficial as image capture would move with the patients’ eye, and not their head as with external mounted image capture devices and would not necessarily require materials to pass through different layers of tissue. Optoelectronic polymers are materials that could produce these benefits and are currently being investigated for their use in electric stimulation of the retina. These polymers create an electric impulse when struck by photons of light but are much more flexible compared to stiff silicon-based photovoltaic materials. Work using regioregular poly(3-hexylthiophene) blended with or without phenyl-C61-butyric acid methyl ester as a semiconducting layer and poly(3,4-ethylenedioxythiophene)-poly(styrenesulfonate) as a conductive layer has shown promising results. They have been demonstrated to restore light sensitivity to diseased explanted rat retinas (Ghezzi et al., 2013) and are stable for at least 5 months when implanted into rat retinas (Antognazza et al., 2016). These materials show promise, but much more work is required.

Tissue Regeneration The regeneration of damaged tissue using tissue engineering strategies is being investigated for tissues throughout the eye. The two major areas are the ocular surface, involving the cornea and conjunctiva, and the retinal pigment epithelial layer under the retina. The precise functional requirements are different in either case, but there are also many similarities. In cases of corneal damage, the current treatment is replacement of the dysfunctional tissue with a corneal transplant; however, due to the high demand for tissue, there is a shortage of donors, and so tissue engineering approaches are needed to provide alternative options. The cornea is composed of five layers (Fig. 4A): the outermost epithelial layer with its underlying Bowman’s membrane; the stroma whose regular arrangement of collagen fibers is responsible for the transparency of the cornea; and finally, the most posterior layer, the monolayer of corneal endothelial cells sitting on the Descemet’s membrane. The cornea is continuous with the tough, opaque sclera; therefore, it plays a key role in maintenance of the protective outer layer of the eye but must be transparent to allow light to reach the retina. The conjunctiva (Fig. 4B) is a transparent membrane that covers the inner surface of the eyelids and extends over the sclera to meet the limbus at the periphery of the cornea. It is a stratified, nonkeratinized epithelium that serves as a barrier to protect underlying tissue but also as a mucous membrane. Goblet cells within the epithelium secrete mucins that maintain the tear film protecting the cornea from infection and desiccation. One of the striking features of the conjunctiva is the flexibility of the continuous tissue over the large and cavernous regions of the eyelids.

Fig. 3 A cartoon of an implanted Argus II epiretinal implant, developed by Second Sight. Information and power are transmitted to the electronics case via the antenna. From the electronic case, electric impulses are sent to various electrodes on the array, depending upon the information received. The electrode array is pinned to the internal surface of the retina.

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Fig. 4 (A) Schematic showing the five layers of the cornea; outermost epithelium, Bowman’s layer, stroma with interspersed keratocytes, Descemet’s membrane and most posterior endothelial layer that contacts the aqueous humor of the anterior chamber (B) conjunctiva comprises a nonkeratinized, stratified squamous epithelium interspersed with goblet cells.

Disease and disorder can affect each of the components of the ocular surface, and the latest surgical techniques often target the specific layer affected. Therefore, recent tissue engineering strategies have followed the same path with most attempting to combine an ex vivo expanded cell population with a suitable material for transplant that meets the particular requirements of the damaged tissue. Retinal diseases such as retinitis pigmentosa (RP) and age-related macular degeneration (AMD) cause damage to the retinal pigment epithelium. The retinal pigment epithelium is a monolayer of cells that sits beneath this neural retina on a thin basement membrane known as Bruch’s membrane that separates it from the blood vessels in the choroid. The retinal pigment epithelial (RPE) monolayer of cells is the crucial component of the support tissue of the retina and is known to play a key role in maintaining its normal functions. In particular, the RPE cells phagocytose the spent outer segments of the photoreceptors. Damage to, or loss of RPE cells can have serious consequences, particularly as its ability to support the overlying photoreceptors will be compromised leading to irreversible vision loss. Cell transplantation of RPE cells offers a potential therapy to prevent photoreceptor loss. Successful cell transplantation requires the precise delivery, appropriate cell organization, alignment, integration, and differentiation of cells to form a functional epithelial monolayer.

Natural Scaffolds The complex structure of the corneal layers and cellular basement membranes pose a significant challenge to replicate. Some regenerative medicine strategies aim to circumvent this problem by simply reusing these perfected structures as starting scaffolds. Corneal stroma comprises predominantly type I collagen fibers arranged in a precise orthogonal fashion interspersed with keratocyte cells. Decellularization of stromal tissue retains the precise structure of the scaffold allowing it to be transplanted in cases of stromal scarring with reduced risk of rejection (Choi et al., 2011). A commonly trialed natural material in ocular surface tissue engineering is amniotic membrane, the innermost layer of the fetal membrane. This is because it has some very useful characteristics such as anti-angiogenic and antiinflammatory properties, its relative elasticity and flexibility, and its favorable basement membrane features that are conducive to epithelial cell growth (Rahman et al., 2009). For these reasons, it has been used in conjunctival reconstruction, as a carrier/substrate for corneal epithelial layer transplantation and also as a substrate for corneal endothelial cell expansion. As is the case with any biological tissue, there is significant donor variation in terms of thickness and related degradation rate and the presence of anti-inflammatory cytokines. In addition, the tissue preparation protocols are not standardized between centers, which make clinical data relating to its use in tissue engineering difficult to interpret as the membrane composition may differ widely. An alternative to tissue engineering using a biomaterial substrate is injected cell therapy, which is currently being explored as a potential therapy for damaged corneal endothelium (Okumura et al., 2012). This approach involves simply injecting a cell suspension into the anterior chamber and allowing the monolayer to form in situ on the posterior corneal surface. There are some concerns about this approach regarding the ultimate destination of the injected cells at aberrant ocular sites. Therefore, due to the relative simplicity of the corneal endothelium, attempts have been made to engineer a scaffold-free endothelial cell sheet ex vivo using a thermoresponsive poly (N-isopropylacrylamide) (pNIPAAm) substrate for cell growth. At 37 C, the surface is hydrophobic, and cells can adhere, but when the temperature is lowered to 20 C, the polymer chains of pNIPAAm hydrate to form an expanded structure detaching the cells as an intact sheet without the need to use enzymatic digestion (Tang et al., 2012). The major advantage of a scaffold-free sheet is that they have been associated with fewer inflammatory responses that are typically observed in the biodegradation of natural scaffolds; however, delivery of the intact sheet to the site of transplant still remains a significant challenge. In the subretinal position, injection of RPE cell suspensions and tissue grafts have resulted in poor cell survival and low integration of cells, with disorganized and incorrectly localized grafts and difficulty in retaining the cells in the targeted site (Tomita et al., 2005; MacLaren et al., 2006; Klassen et al., 2004). However, this continues to be a strategy employed in current clinical trials (Kimbrel and Lanza, 2015). Transplantation of RPE cells onto alternative natural material including decellularized Bruch’s membrane,

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amniotic membrane, Descemet’s membrane, and lens capsule (Hartmann et al., 1999; Lee et al., 2007; Turowski et al., 2004) that mimic the mechanical properties of natural tissue has been investigated with varying levels of success. Limitations with these approaches arise due to the limited expansion of cells and availability of donor tissue. Scaffold-free approaches using pNIPAAm, similar to that described earlier for corneal endothelium, have been also reported (Yaji et al., 2009).

Biological Polymers Collagen is widely used in corneal tissue engineering since the basement membranes of both the corneal epithelial and endothelial layers, as well as the stroma, are abundant in collagens including types IeV. Many research groups have developed stromal tissue equivalents using animal-derived collagen (Levis et al., 2010, 2013), but an alternative strategy has been reported that uses recombinant type III human collagen cross-linked with carbodiimide. This material was designed to mimic the collagenous extracellular matrix of the stroma to stimulate in situ repopulation of the graft with keratocytes and corneal nerves. The material has sufficient tensile strength and elasticity and allows keratocytes to repopulate the graft over time in addition to maintaining a tear film on the surface (Liu et al., 2008). These hydrogels have since been trialed in clinical studies with four patients all displaying stable corneal regeneration 4 years after implantation (Fagerholm et al., 2014). These materials have also been investigated for conjunctival epithelial and goblet cell growth (Table 1) (He et al., 2016). Other natural materials that have been used for ocular surface repair are chitosan, silk fibroin, and fibrin. The latter forms the basis of Holoclar, the first stem cell-based medicinal product to be approved for use in Europe. Human limbal stem cells are cultivated on a fibrin (thrombin and fibrinogen) membrane and transplanted onto the ocular surface to repair damage in cases where the native limbal stem cell population has been destroyed by chemical or thermal burns. The fibrin substrate supports the expansion and growth of a healthy epithelial layer and, after transplantation, slowly degrades leaving the stem cells to seed the regrowth of a transparent cornea. One additional advantage of the fibrin gels is that its constituents can be produced from autologous plasma (Pellegrini et al., 2016). In vitro studies have shown that collagen scaffolds are capable of supporting RPE cell adherence and proliferation, with phenotypic characteristics of differentiated RPE cells (Malafaya et al., 2007; Karwatowski et al., 1995). In vivo studies in rabbits have shown biocompatibility with no immune rejection or inflammation and the ability to phagocytose photoreceptors (Thumann et al., 1997).

Synthetic Polymers Synthetic materials have an advantage over natural materials because they can be produced with a fully defined composition and designed features and structures. Both poly-L-lactic acid (PLLA) and poly-DL-lactic-co-glycolic acid (PLGA) have been tested as Table 1

Conjunctival cells grown on various biopolymer and synthetic polymer substrates (He et al., 2016)

Material

Production method Cell interaction

Mechanical properties

Recombinant human collagen

Hydrogel

Good viability

Variable degradation rates

Recombinant human collagen 2methacryloylxyethyl phosphorylcholine

Hydrogel

Majority of cells CK7þ (goblet) >90% Good viability

Variable degradation rates

Arginine-glycine-aspartic acid (RGD)-modified fibroin (Bombyx mori)

Film

Majority of cells CK7þ (goblet) >90% Good viability

Variable degradation rates

Poly-D-lysine (PDL)-coated fibroin (B. mori)

Film

Majority of cells CK7þ (goblet) >90% Good viability

Variable degradation rates

Electrospun Electrospun

Majority of cells CK7þ (goblet) >80% N/A Reasonable viability

Poly (caprolactone)(PCL)

Electrospun

Majority of cells CK7þ (goblet) >90% Very low cell viability

Poly (vinyl alcohol) (PVA)

Electrospun

No cell attachment

Collagen Poly (acrylic acid)(PAA)

Too fragile Control over structure and chemical composition Difficult to handle Control over structure and chemical composition Control over structure and chemical composition

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potential substrates for corneal endothelial cell monolayer transplantation, as well as polyvinylidene fluoride coated with type IV collagen. Confluent monolayers formed on PLLA and PLGA substrates and degradation could be controlled by varying the monomer ratios to allow the carrier to remain intact for an interval sufficient for cells to lay down their own basement membrane (Hadlock et al., 1999). However, current surgical treatments commonly graft an endothelial layer with a thin (< 150 mm) piece of stroma attached to the posterior corneal surface. This stromal layer persists long term, and vision is still restored, so it is debatable as to whether a degradable carrier is required. The requirements for conjunctiva reconstruction differ from that of the cornea. The substitute must consist of a stable elastic matrix, but does not necessarily need to be transparent, which allows for a wider variety of materials to be investigated. Maintaining a mixed population of goblet and epithelial cells in a conjunctival tissue equivalent is an important consideration. It is reported that while biopolymers such as recombinant human collagen and coated silk have variable degradation rates, their cell compatibility and handleability was superior to that of the synthetic polymers tested (Table 1) (He et al., 2016). PLGA/PLLA scaffolds have been evaluated for their ability to support RPE cell delivery and survival and shown to promote integration and differentiation of the cells (Tomita et al., 2005; Lavik et al., 2005). Electrospinning of PLGA-generated 3-D nanofibrous scaffolds with mechanical properties similar to Bruch’s membrane and when coated with collagen I were able to support human RPE cells, with a correctly orientated monolayer of cells (Warnke et al., 2013). Fibrous PLA scaffold improved retinal ganglion cell survival and guided retinal axon orientation onto rat retinal explants in vitro (Kador et al., 2013). However, when the PLA/ PLGA scaffolds were thick (150–250 mm), they resulted in retinal detachments in rodent models and were an unsuitable scaffold for subretinal transplantation. In addition, in some studies, degradation of PLGA/PLLA (Tomita et al., 2005; Lavik et al., 2005; Warfvinge et al., 2005) caused a buildup of acidic degradation products within the subretinal space leading to inflammation, fibrosis, and cell death (Warfvinge et al., 2005; Sundback et al., 2005). PCL, which degrades more slowly than PLGA/PLLA, showed no increase in acidity within the subretinal space (Tao et al., 2007; Grayson et al., 2004). In vitro studies showed an increase in cell attachment and organization with a decrease in markers of early progenitors and an increase in photoreceptor markers (Steedman et al., 2010). Thin (5–6 mm) PCL scaffolds that were highly permeable resulted in minimal physical distortion and good permeability (Tao et al., 2007) and were shown to be highly compatible with the subretinal space in mice and pigs. They supported RPE cells (Steedman et al., 2010), and porous scaffolds with 1 mm continuous pores supported a functional monolayer of fetal human RPE cells that expressed mature markers, increased pigmentation, cell density, barrier function, polarized growth factor secretion, and metabolite transport (McHugh et al., 2014). Improving the hydrophilicity and surface morphology of PCL scaffolds using plasma treatments or alkaline hydrolysis increased the biocompatibility, wettability, cell adhesion, and viability of the cells and led to a functional RPE cell layer (Shahmoradi et al., 2017). Hybrid electrospun scaffolds fabricated with PCL, silk fibroin, and gelatin generated thin porous scaffolds that sustained a RPE morphological phenotype without signs of inflammation (Xiang et al., 2014). Some researchers have suggested that biostable scaffolds may be advantageous in the long term by providing support for the transplanted cells. This is because biodegradable substrates may eventually leave the cells in direct contact with compromised Bruch’s membrane and because the local response to degradation products may have negative effects. Biostable polymers investigated include parylene, polyethylene terephthalate, and polyimide (Ilmarinen et al., 2015; Pennington and Clegg, 2016; Stanzel et al., 2014). Surface-modified polytetrafluoroethylene (ePTFE), which is biocompatible, biostable, porous, and flexible and has good mechanical properties for surgical handling, has also demonstrated the ability to support a functional layer of RPE cells (Kearns et al., 2012; Krishna et al., 2011).

Conclusion A wide range of biomaterials from decellularized tissue, biological polymers, synthetic polymers, and metals to ceramics have been used in ophthalmic applications. The functional requirement of the biomaterial must be considered in each application and in the eye the array of requirements is also broad, including the ability to refract the light coming into the eye, passive or active space filling, controlling fluid flow out of the eye, electric stimulation of the neural retina or tissue regeneration. For some applications, the use of biomaterials has been well established for many years, for example, as contact or intraocular lenses or silicone oil tamponade agents to treat retinal detachments. In other applications, however, the optimal biomaterial is not yet defined, such as in relation to tissue engineering scaffolds for subretinal RPE or corneal endothelial cell transplantation, and in these areas, further studies are required.

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Biomaterials in Orthopaedics Emmanuel Gibon and Stuart B Goodman, Stanford University, Stanford, CA, United States © 2019 Elsevier Inc. All rights reserved.

Introduction Polyethylene Ceramics Metals Conclusion References

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Introduction Biomaterials in orthopedic surgery cover a broad range of different devices used in different procedures and for different goals. Among these procedures, total joint replacements (TJRs) of the hip and knee are probably the most successful operations. TJRs are projected to dramatically increase in number and cost for society (Kurtz et al., 2007). Studies have shown that for orthopedic surgery alone, the underlying industry is expected to grow to $41.1 billion by 2016 (Richards et al., 2012). Currently, the majority of the studies analyzing the survivorship of total hip arthroplasty (THA) show data ranging above 90% after 15 years (El Masri et al., 2010; McLaughlin and Lee, 2006). Similarly, the long-term outcome of total knee arthroplasty (TKA) has shown survivorship after 15 years ranging from 81.7% to 98.14% (Attar et al., 2008; Epinette and Manley, 2007; Melton et al., 2012). However, this outcome has not come without a great deal of research and development and both successes and failures. Indeed, different implants demonstrate a variable degree of durability (Keurentjes et al., 2014). Considering TJRs, biomaterials currently used include metals and their alloys, polymers, and ceramics. In TJRs, surgeons aim to replace the joint itself, that is, articular surfaces and the adjacent bone. To reach these goals, metal alloys are used to replace the bone that is located underneath the articular cartilage, whereas the articular surfaces are usually replaced either by polyethylene, ceramics, or in some cases metals. Implants are fixed to the bone with polymethyl methacrylate (PMMA) also called “bone cement” or without cement using “press-fit cementless” implants. The aim of this article is to discuss the biological response to bearing materials used in orthopedic surgery.

Polyethylene Metal on polyethylene (MoP) is the most commonly reported bearing used in the United States with up to 51% according to Bozic et al. (2009). Since the original failures when using the polymer polytetrafluorethylene by Sir John Charnley, the biomaterial of choice for one of the bearing surfaces has included polyethylene (PE). However, periprosthetic bone loss and implant loosening due to the chronic inflammatory and foreign body reaction to polyethylene wear particles limited implant durability (Gallo et al., 2013a,b). Therefore, after the first generation of “conventional” polyethylene also called “ultra-high molecular weight polyethylene (UHMWPE),” a second generation of PE was developed to improve wear and mitigate the production of wear particles. Hip stimulator studies showed that UHMWPE subjected to 10 Mrad of g-irradiation significantly produced fewer particles than 5 Mrad, g-inert, or conventional UHMWPE (Ries et al., 2001). Second-generation highly cross-linked polyethylene (HXLPE) (Goodman et al., 2009) includes the doping of the antioxidant vitamin E within the PE or repeated treatment with heating and annealing of the polymer (Essner et al., 2005). Wear particles are always generated during regular usage of joint replacements. Each individual takes around 500,000–2000,000 steps each year, and hundreds of thousands to millions of wear debris are generated with each step (Goodman and Ma, 2010). Generation of wear debris is dramatically increased with increased femoral head roughness (Wang et al., 1998). Roughening of the femoral head was shown to increase wear by an order of magnitude resulting in a two- to threefold increase in the wear rate that could be predicted by the following equation: k ¼ 7.21  106 (Ra)0.42 (mm3/Nm), with K being the specific wear rate and Ra the centerline average roughness of the femoral head. The consequences of such an excessive amount of generated particles might be harmful and led to periprosthetic osteolysis and potential aseptic loosening of the implants. Campbell et al. (1995) have shown that PE particles are needlelike in shape and are generally < 1 mm in size. PE wear particles migrate within the entire periprosthetic bed (Schmalzried et al., 1992), known as “the effective joint space.” Production of PE particles leads to a nonspecific macrophage-mediated foreign body reaction (Goodman, 2007). Macrophages become activated either by phagocytosis (Xing et al., 2002) of PE particles or simply by cell membrane contact without phagocytosis. Activation occurs through receptors present in the outer cell membrane (CD11b, CD14, Toll-like receptors, etc.). The Toll-like receptors (TLRs) are known to function in the innate immune response (Tuan et al., 2008). TLR2 and TLR4 have been found to be critical for particle-induced osteolysis (Valladares et al., 2014). TLRs are activated by different types of stimuli and act through an adapter protein called myeloid differentiation primary response gene 88 (MyD88) to induce activation of nuclear factors such as nuclear factor kappa-B (NFkB). Recently, Lin et al. (2016) using a well-established small-animal model of continuous PE particle delivery (Ma et al., 2008, 2009;

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Ortiz et al., 2008) showed that NFkB decoy oligodeoxynucleotide (ODN) mitigates wear-particle-associated bone loss. Moreover, NFkB decoy ODN reversed the loss of bone mineral density in the distal femur exposed to particles. Activation of NFkB in macrophages then triggers an inflammatory cascade leading to the release of various pro-inflammatory cytokines (Wang et al., 2010; Shanbhag et al., 2007) (IL-1, IL-6, and TNF-a), growth factors (macrophage colony-stimulating factor-1), and chemokines (MIP1a and MCP-1) that would ultimately lead to systemic recruitment and local infiltration of more macrophages to the area of inflammation (Ren et al., 2010, 2011). Retrieval studies have found high levels of pro-inflammatory factors released in the surrounding tissues of failed arthroplasty groups (Wang et al., 2010; Goodman et al., 1998). Locally activated macrophages undergo polarization toward the M1 (pro-inflammatory) phenotype (Mantovani et al., 2013). M1 macrophages produce primarily pro-inflammatory mediators including TNF-a, IL-1, and IL-6 and express inducible nitric oxide synthase (iNOS) (Murray et al., 2014). Subsequently, locally and systemically recruited activated macrophages differentiate into multinucleated giant cells and osteoclasts leading to bone resorption around implants within a foreign body reaction (Goodman et al., 1989) (Fig. 1). Particle generation remains an issue; however, newer highly cross-linked PE is now significantly better than conventional PE (Scemama et al., 2017; Langlois et al., 2015).

Ceramics Ceramics are considered “hard bearing surfaces.” Current ceramics used for joint replacement are actually composites of two of ceramicsdalumina (AL2O3) and zirconia (ZrO2) in which alumina is the primary or continuous phase (70%–95% composition) and zirconia is the secondary phase (30%–5% composition) (Kurtz et al., 2014)dand are called either “alumina-toughened zirconia” (ATZ) or “zirconia-toughened alumina” (ZTA). Zirconia or alumina alone is no longer used. Ceramic-on-ceramic (CoC) bearings currently represent 14% of bearing surface usage in the United States (Bozic et al., 2009). However, a ceramic head can also articulate with a PE liner. The reason for ceramic-on-polyethylene (CoP) bearing surfaces use is twofold. First, with this bearing combination, surgeons can avoid squeaking from CoC bearing surfaces. Second, with a ceramic head (instead

Joint replacement with polyethylene “MoP”

Polyethylene (PE) particles size: < 1 μm Macrophage = key cell

MΦ activation

Cell contact: TLRs, CD11b, CD14

Phagocytosis

MΦ activation Systemic recruitment Cytokines/chemokines /ROS release

MCP-1 MIP-1α TNF-α IL-1,6

MΦ polarization

M1 = inflammatory response MSCs chemotaxis

Foreign-body reaction, granuloma Differentiation into osteoclast

Bone defect

Fig. 1 Biological response to PE particles. Production of PE particles leads to a nonspecific macrophage-mediated foreign body reaction. IL, interleukin; MF, macrophage; MCP-1, monocyte-chemoattractant molecule 1; MGC, multinucleated giant cell; MIP-1, macrophage inflammatory protein 1; MSCs, mesenchymal stem cells; TLRs, Toll-like receptors; ROS, reactive oxygen species; TNF-a, tumor necrosis factor-alpha.

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of a metal femoral head) on a metal taper, surgeons avoid so-called trunnionosis caused by corrosion at the head–neck junction. Moreover, in vivo studies have shown polyethylene wear rates in CoP bearings to be lower than rates in MoP bearings (Wang et al., 2013; Meftah et al., 2013) albeit it remains controversial (Kawate et al., 2009). A drawback of CoP bearings compared with MoP bearings is the price, with CoP being more expensive than MoP. Recent work by Carnes et al. (2016) showed that CoP bearings might not be worth the cost given the economic burden for society. Hatton et al. (2002) characterized alumina wear debris by analyzing tissues retrieved at revision surgery. Interestingly, their work compared tissues from CoP implants and MoP implants. The authors found alumina wear debris to have a bimodal size range distribution. The particles were either nanometric (5– 90 nm) or micrometric (0.05–3.2 mm). Furthermore, histomorphometric analysis showed that tissues retrieved from CoC implants had significantly less macrophages and giant cells than MoP implants but significantly more neutrophils. Ding et al. (2012) challenged RW 264.7 macrophages with ceramic and titanium particles. They used three distinct sizes of ceramic particles (0.2– 1.2 mm, 1.2–10 mm, and > 10 mm). Both ceramic and titanium particles increased the expression of TNF-a in a time-dependent manner. However, ceramic particles provoked a significantly lower production of the pro-inflammatory cytokine TNF-a. The authors also investigated the effect of the sizes of the particles. Smaller ceramic particles (0.84 mm) are more bioactive and more proinflammatory as they induced significantly higher mRNA and protein expression of TNF-a and RANKL than larger particles. Similarly, Zhang et al. (2011) challenged osteoblast-like cells (MG-63) and murine RAW 264.7 macrophages with ceramic particles (AL2O3 and ZrO2) of a nanometric size (40–50 nm). The cytotoxic assay performed with the MG-63 cells showed no adverse effects at any time points (24, 48, and 72 h). Moreover, ZrO2 particles significantly increased the intrinsic alkaline phosphatase (ALP) activity of MG-63 cells at higher concentration (500 ppm), whereas AL2O3 particles had the same effect but at lower concentrations (115 ppm). Interestingly, the smallest AL2O3 particles were more potent to increase ALP activity. Tsaousi et al. (2010) used primary human fibroblasts to test the cytotoxicity and genotoxicity of AL2O3 particles. Particles used were either 0.2 nm mean size or 2 mm mean size or fiber sized (0.9 mm in diameter and 12.03 mm in length). The authors found no significant differences in cell viability between control and ceramic-treated cells, at all doses and time points studied. Moreover, AL2O3 particles were only weakly genotoxic (for high doses, involving chromosome loss and tetraploidy). Further studies have shown limited biological reaction to alumina particles and a weak local inflammatory reaction (Germain et al., 2003; Gutwein and Webster, 2004; Catelas et al., 1998, 1999a,b; Lucarelli et al., 2004). Ceramic particles appear to be low in number and well tolerated and elicit a minor cellular response (Fig. 2).

Metals Metal-on-metal (MoM) bearings are also “hard-on-hard” bearings. With the idea of decreasing the production of wear debris and periprosthetic osteolysis, MoM bearings became popular among the orthopedic community about one decade ago. This resurgence was supported by studies showing that the volumetric wear released by MoM articulations can be up to 100-fold lower than with MoP bearings (Greenwald and Garino, 2001; St John et al., 2004). Bozic et al. showed that between October 2005 and December 2006, MoM bearings were the second most used articulations in the United States comprising up to 35% of primary THA (Bozic et al., 2009). MoM bearings reached a peak in 2008 making up to 40% of all primary THA procedures in the United States. However, unexpected issues started to be reported with MoM bearings, and the Food and Drug Administration recalled many implants (Ng et al., 2011; Hug et al., 2013). MoM bearings are made of an alloy containing cobalt (Co) (62%), followed by chromium (Cr) (27%–30%) and a small amount of molybdenum (Amanatullah et al., 2016). Studies have shown that metal particles or

Joint replacement with ceramic bearings “CoC”

Ceramic (Al2O3 and ZrO2) particles

size: 5–90 nm or 0.05–3.2 μm (bimodal distribution)

Activation

CoC bearings: mild inflammatory response

TLRs

Phagocytosis

ALP activity RANKL TNF-α

Fig. 2 Biological response to ceramic particles. Production of ceramic particles leads to a more mild biological response. MF, macrophage; NFkB, nuclear factor-kappa-B; TLR, Toll-like receptor; TNF-a, tumor necrosis factor-alpha; ALP, alkaline phosphatase; RANKL, receptor activator of nuclear factor-kappa-B ligand.

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particulates generated are very small ranging from 20 to 90 nm (Doorn et al., 1998; Ingham and Fisher, 2005). Furthermore, ions of Co and Cr released from MoM bearing surfaces can travel via lymph and blood to various sites throughout the body including bone marrow, lymph nodes, the spleen, the liver, and the heart (Case et al., 1994; Langkamer et al., 1992; Urban et al., 2000). The effect of metal particles and ions has been widely studied. DNA can be damaged by Co and Cr through oxidative stress (Cannizzo et al., 2011). Similarly, free radicals released by Cr can cause breakage in DNA strands, and both Cr and Co ions can induce an inhibition of DNA repair pathways leading to defective gene expression (Daley et al., 2004; Polyzois et al., 2012). Cells exposed to Co and Cr had a decrease in DNA synthesis (Hodges and Chipman, 2002). Chromosomal aberrations and translocations have also been reported (Dunstan et al., 2008; Doherty et al., 2001). With these different studies reporting genetic aberrations induced by Co and Cr released by MoM joint, one may ask if there is an increased risk of developing cancer. In 40 tumor bioassays in which Co and Cr ion concentrations were significantly higher than those in hip implants, none reported a significant increase in systemic cancer (Christian et al., 2014). A Finnish study also looked at 2000 patients roughly 15 years after a first-generation MOM articulation and found that there was no increase in the incidence of cancer. The authors concluded that factors other than the TJR surgery are responsible for the development of cancer in hip implant patients (Visuri et al., 1996). Similarly, studies of patients with more modern metalon-metal articulations failed to suggest an increased risk of cancer (Smith et al., 2012). Another adverse reaction induced by MoM bearing is metal hypersensitivity. Metal hypersensitivity (or metal allergy) to jewelry is found in roughly 10%–15% of the general population (Thyssen et al., 2007; Thyssen and Menné, 2010). However, metal ions released by MoM joints by themselves are too small to mount a hypersensitivity reaction. The immune response is elicited by denaturized protein (e.g., albumin) bound to the ions and forming haptens (Adala et al., 2011). The key cell in the biological reaction to metallic particles and ions is the lymphocyte (T lymphocyte) (Hallab et al., 2001). More specifically, the immune response is a cellmediated type IV delayed hypersensitivity response (Hensten-Pettersen, 1993), as opposed to PE particles where it is a nonspecific nonantigenic immune response (foreign body reaction). The local tissue reaction associated with metallic by-products is now well established and is referred to as adverse local tissue reaction (ALTR), which is a contributing factor to THA failure (Prieto et al., 2014). ALTR is often associated with pain in the groin, hip, thigh, and buttock (Pivec et al., 2014). ALTR has been categorized in three major types: aseptic lymphocyte-dominated vasculitis-associated lesions (ALVaL), pseudotumors, and osteolysis (Fig. 3):

•

ALVaL describes a periprosthetic histological inflammatory reaction that is similar to a type IV hypersensitivity response in the soft tissue surrounding metal-on-metal implants (Kolatat et al., 2015; Lawrence et al., 2014).

Joint replacement with metal bearings “MoM”

Metal particulates size: 40–50 nm + Cr and Co ions

Fibroblast ROS MCP-1 Metalloproteinase

cytokines, PGE2 GM-CSF

MSCs MΦ

Denaturized proteins (haptens)

Lipid peroxidation

Inhibition of osteoblastic differentiation Lymphocyte proliferation

Osteoblast toxicity Osteoclast chemotaxis

DNA flaws Osteolysis

Hypersensitivity reaction: type IV delayed

Pseudotumor Fig. 3 Biological response to metal particulates and ions. Production of metal particulate and ions leads to a nonspecific immune response and in some cases to a type IV (cell-mediated) hypersensitivity reaction. GM-CSF, granulocyte-macrophage colony-stimulating factor; MCP-1, monocytechemoattractant molecule 1; MSC, mesenchymal stem cell; PGE2, prostaglandin E2; ROS, reactive oxygen species.

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Pseudotumors are large, cystic, or solid masses that are often seen in patients with ALTR. The rate of development of a pseudotumor is controversial, and the literature reports rates ranging from 1% to 60% (Pandit et al., 2008; Hart et al., 2012; van der Weegen et al., 2013). Osteolysis is the destruction of the bone, resulting from osteoclast resorption (Ries and Link, 2012). Swedish studies showed that osteolysis was the cause of revision surgery for 75% of patients with metal-on-metal articulation (Malchau et al., 2002; Dattani, 2007).

Conclusion Biomaterials in orthopedics are numerous, and their effects on the human body have been extensively studied. Newer generations of highly cross-linked PE have seen great success, whereas some hard-on-hard bearings such as metal on metal and alumina on alumina have been withdrawn. Future directions and current research aim to develop bioactive, “smart” materials that are well tolerated and cause less local inflammation, facilitate bone integration, and limit infection. With younger patients now receiving joint replacements and the overall aging of our population in general, developing safe, durable, and cost-effective bearing surfaces is of paramount importance. Surgeons, engineers, manufacturers, and regulatory agencies must work together to achieve these goals.

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The in vitro genotoxicity of orthopaedic ceramic (Al2O3) and metal (CoCr alloy) particles. Mutation Research, 697(1–2), 1–9. Tuan, R. S., Lee, F. Y., TK, Y., Wilkinson, J. M., & Smith, R. L. (2008). What are the local and systemic biologic reactions and mediators to wear debris, and what host factors determine or modulate the biologic response to wear particles? Journal of the American Academy of Orthopaedic Surgeons, 16(Suppl. 1), S42–8. Urban, R. M., Jacobs, J. J., Tomlinson, M. J., Gavrilovic, J., Black, J., & Peoc’h, M. (2000). Dissemination of wear particles to the liver, spleen, and abdominal lymph nodes of patients with hip or knee replacement. Journal of Bone and Joint Surgery. American Volume, 82(4), 457–476. Valladares, R. D., Nich, C., Zwingenberger, S., Li, C., Swank, K. R., Gibon, E., Rao, A. J., Yao, Z., & Goodman, S. B. (2014). Toll-like receptors-2 and 4 are overexpressed in an experimental model of particle-induced osteolysis. Journal of Biomedical Materials Research. Part A, 102(9), 3004–3011. van der Weegen, W., Brakel, K., Horn, R. J., Hoekstra, H. J., Sijbesma, T., Pilot, P., & Nelissen, R. G. H. H. (2013). Asymptomatic pseudotumours after metal-on-metal hip resurfacing show little change within one year. The Bone & Joint Journal, 95-B(12), 1626–1631. Visuri, T., Pukkala, E., Paavolainen, P., Pulkkinen, P., & Riska, E. B. (1996). Cancer risk after metal on metal and polyethylene on metal total hip arthroplasty. Clinical Orthopaedics and Related Research, 329(Suppl. S2), 80–89. Wang, A., Polineni, V. K., Stark, C., & Dumbleton, J. H. (1998). Effect of femoral head surface roughness on the wear of ultrahigh molecular weight polyethylene acetabular cups. Journal of Arthroplasty, 13(6), 615–620. Wang, C.-T., Lin, Y.-T., Chiang, B.-L., Lee, S.-S., & Hou, S.-M. (2010). Over-expression of receptor activator of nuclear factor-kappaB ligand (RANKL), inflammatory cytokines, and chemokines in periprosthetic osteolysis of loosened total hip arthroplasty. 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Cell Encapsulation and Delivery Stefani Mazzitelli and Claudio Nastruzzi, University of Ferrara, Ferrara, Italy © 2019 Elsevier Inc. All rights reserved.

Cell Encapsulation for Immunoisolation Biomaterials for Cell Encapsulation Bioencapsulation Procedures Cell Encapsulation by Gas Driven Mono-jet, Vibrating Nozzle and Jet Cutter Cell Encapsulation by Microfluidics Cross Junction Microfluidic Chips T-Junction Microfluidic Chips Microcapillary Coaxial Device Method Preparation of Fibrous Scaffolds Further Reading

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Glossary Alginate Alginate is a natural polysaccharide that comprises from 30% to 60% of brown algae (on dry weight basis). It is an anionic polymer which is present in cell walls of algae, where through binding with minerals from seawater it forms a viscous gum structure (named as “jelly bodies”). Alginate and its derivatives have been utilized as hydrocolloids in a variety of applications such as food additives, pharmaceuticals, cosmetics and textile manufacturing. Biopolymer Macromolecules are formed by living organisms. Biopolymers can be polynucleotides (such as the nucleic acids DNA and RNA), polypeptides (that is, proteins) or polysaccharides (that is, polymeric carbohydrates). These consist of long chains of repeating, covalently bonded units, such as nucleotides, amino acids or monosaccharides. Cell encapsulation A process in which cells or cell aggregates are surrounded by a coating materials to give small devices (typically in the dimensional range of hundreds of microns) with diverse useful properties. The process is used to incorporate food ingredients, enzymes, cells or other materials on a micro metric scale. A simple form of cell encapsulation device is microcapsule, a sphere with an uniform wall around it. The material inside the microcapsule is referred to as the core, internal phase, or fill, whereas the wall is generally called shell, coating, or membrane. Microfluidics It is a multidisciplinary science dealing with the manipulation and control of fluids, usually in the range of microliters to picoliters, in a network of channels with dimensions typically from tens to hundreds of micrometers. Microfluidics is a science at the intersection of engineering, physics, chemistry, biochemistry, nanotechnology, and biotechnology, with practical applications in the design of systems in which low volumes of fluids are processed to achieve multiplexing, automation, and high-throughput screening.

Cell Encapsulation for Immunoisolation Cell encapsulation involves the design and application of specific procedures for immunoisolating cells into appropriate scaffolds. It represents a major advance in cell based therapy as it avoids constraints associated with cell sources, making both allogenic and xenogenic cells putative alternatives to the scarcely available autologous cells. Cell based therapies require therefore the combined presence of a biomaterial based scaffold and selected cells to create a biocompatible device that, once transplanted, isolates cells from the host’s immune system, eliminating or reducing the requirement for immunosuppressant systemic drug administration. In this respect, the biomaterial plays a fundamental role in cell encapsulation for an efficient implantation of cells, providing functional groups to anchor cells and offering an environment in which the transplanted cells can survive, migrate to target sites and effectively function. Cells are combined with biomaterials following two main strategies: (a) “conventional” seeding of cells in preformed scaffolds or (b) embedding cells into scaffold during the scaffold production (i.e. cell encapsulation procedure). If compared to conventional seeding strategies, the encapsulation of cells in 3D scaffolds provides a better control on proliferation, cell morphology, and precise geometric constrains. The seeding of cells often leads to an inhomogeneous distribution of cells along the scaffold surface/structure, due to cell aggregation in clusters, thus leading to heterogeneous properties of final construct. The success of different encapsulation approaches is mostly related to properties of biomaterial used for scaffold production. Biomaterials provide indeed the support to cells to adhere, differentiate and/or grow both in vitro culture or in vivo after implantation. Materials employed to produce the scaffolds must be compatible with both the cellular components of transplanting tissue and cells of hosts and they should possess a porous structure for the effective transport of nutrients and metabolites.

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Intuitively, the best encapsulation material should be similar to the extracellular matrix (ECM) of target tissue in its native state. Unfortunately, the enormously complex molecular architecture and function of ECM make it extremely difficult to be mimicked. Therefore, a variety of biomaterials with different physical and chemical properties have been proposed and utilized. For instance, materials able to form hydrogels are particularly attractive as scaffolding materials. Hydrogels possess indeed great potential for cell encapsulation thanks to their distinct physical properties. They provide an appropriate environment that is mechanically and structurally similar to the topology of native ECM, supplying adhesion sites to the cells. Moreover, hydrogels can be generally obtained by mild reaction conditions and can be administered by minimally invasive protocols.

Biomaterials for Cell Encapsulation The essential function of a biomaterial for cell encapsulation is to provide a temporary 3D structure for cells. Criteria for scaffold selection include controlled biodegradability, suitable mechanical strength and appropriate surface chemistry. Another important feature of scaffold is related to its porosity which should be designed to enable nutrition, proliferation and integration into host tissue/organ. At the present, different scaffold dimensions and shapes have been proposed for cell encapsulation, generally defined as macro and microdevices, the latter generally in shapes as particles, fibers, flat sheets and disks (see Fig. 1). Among microdevice class, microparticles represent one of the most widely known and studied systems; over the years, therefore, various alternative cell encapsulation technology for the spherical immunoisolation have been described. Microcapsules with a spherical shape are considered beneficial as sphericity provides an optimal surface-to-volume ratio for efficient protein and nutrient diffusion, improving cell viability and functions compared to other scaffold geometries. Furthermore, small microcapsule dimensions, usually with a diameter comprised between 250 and 750 mm, facilitate the implantation at different body sites through small diameter catheters. The microcapsule forming materials are indeed expected to permit diffusion of nutrients and molecules including oxygen and growth factors essential for cellular metabolism, proliferation, differentiation and morphogenesis, while excluding entrance of all high molecular weight molecules such as immunoglobulins and immune system cells. (A)

medium core cells shell

micro-particles

increasing device size

(B)

micro-fibre (C)

macrodevice (milli-cylinder) Fig. 1 Scheme of hydrogel biomaterial for cell encapsulation protocols. The scaffolds can be tailored to possess different geometrical features (i.e., dimensions), namely: microparticles (A), microfibers (B) or macrodevices (C).

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Microcapsules are generally constituted of hydrogel matrix forming a highly hydrated microenvironments, similar to those of their native tissue; in addition matrix permeability provides a high degree of diffusion for low-molecular mass biochemical and physical stimuli for cellular processes such as differentiation, proliferation, and migration. Hydrogels are prepared by mild procedures, via thermo-, ion-, or photo-induced processes, leading to an uniform distribution of cells into the gel matrix. These features result in a high biocompatibility associate to minimal host-cell adhesion and protein adsorption phenomena. The success of therapeutic approaches based on cell encapsulation requires a detailed characterization of the biomaterials employed and of the cell–material and material-host tissue interactions, with special regard to the biocompatibility and immunogenicity of the cell-microcapsule assembly. Alginate or other polysaccharides with similar properties, such as pectins, carrageenans, xanthans, chitosans and natural and semisynthetic celluloses, occupy a prominent position for producing hydrogels intended as biomaterials for cell encapsulation. Microcapsular shape greatly influences the in vivo performances, therefore irregular geometries, such as fused or partially fused capsules or tear shaped capsules result, after implantation, in the formation of capsular fibrotic overgrowth. Other microcapsular defects such as presence of cracks or surface fissures, together with an irregular (rough or waved) surface, can also cause an obvious immunological in vivo response. Dimension and morphology represent also critical characteristics that influence the microcapsule performances, including: cell loading capacity, diffusive properties, method of implantation and biocompatibility. Porosity and pore size and structure are other important factors as they impact nutrient supply to transplanted cells. To regenerate highly vascularized organs such as liver, porous scaffolds with large void volumes and large surface-area-to-volume ratios are desirable for maximal cell seeding, growth, and vascularization. Small-diameter pores are preferred to yield high surface area per unit volume so long as the pore size is greater than the diameter of a cell in suspension (typically 10–15 mm). Finally, the mechanical behavior is a critical feature for microcapsules since they are usually exposed to several mechanical stresses during the processing, post processing (i.e., washing step), implantation and in vivo permanence. In this respect, the mechanical and elastic properties can be tailored to tune the microcapsule properties by adjusting the polymer source and concentration, processing methods, formulation and gelling conditions. Alginates, which are by far the most frequently used biomaterial in the field of cell microencapsulation, are a family of unbranched binary copolymers of 1 / 4 linked b-D-mannuronic acid (M) and a-L-guluronic acid (G), of widely varying compositions and sequential structures. The composition has a great effect on alginate properties including gelling behavior, permeability, mechanical resistance, in vivo stability and biocompatibility. When dispersed in water, alginates form a colloidal dispersion, which gels ionotropically by the addition of divalent cations (i.e., Ca2 þ, Sr2 þ, or Ba2 þ). The most important asset of alginate is indeed the ability to form gels under extremely mild conditions without chemicals or at specific pH. The polymer cross-linking occurs following the exchange of sodium ions from the guluronic acids with the divalent cations resulting in a chain-chain association that constitutes the junction zones of so-called “egg box model.” The buckled chain of guluronic acid units is shown as a two-dimensional analogue of a corrugated egg-box with interstices in which divalent ions may pack and be coordinated. Since hydrogel formation happens following selectively linkage between the carboxylic moieties on the G blocks of alginate and cations, high ratio of G:M outcomes in stiff gels. Alginate forms mechanical and chemical stabile gel with a defined pore size and a narrow pore-size distribution. Unfortunately, the above features do not pertain at all the commercially available alginates; for cell encapsulation procedure, only ultrapure alginates should be considered and an efficient purification process should be accomplished to obtain starting materials containing low number of contaminants including toxic, pyrogenic and immunogenic substances. Alginate microcapsules are generally prepared by dripping chosen cell/alginate suspension into a divalent crosslinking solution. Gelling bath with Ca2 þ or Ba2 þ are generally preferred for cell encapsulation procedures in reason of the higher biocompatibility. After preparation, the gelified microcapsules are usually coated by a cationic polyelectrolyte. The coating is performed to slow down the swelling and in vivo degradation, however it may cause immunological reactions and fibrotic growth; this latter can subsequently decrease the therapeutic efficacy of the entrapped cells sharply reducing the diffusive properties of alginate beads. In this respect, the use of the positively charged polyelectrolytes for the capsular coating (i.e., poly-L-lysine) can lead to a more intense overgrowth compared to negative ones. This behavior was attributed to activated macrophages that preferentially adhere to positive charged surfaces. Ba2 þ ions have also been proposed, as alternative to calcium, for the alginate cross-linking. It has been proved that Ba2 þ ions provide stronger gels, allowing the transplantation of microcapsules without the need of the coating procedure. As drawback, barium is sometimes believed to be toxic but recent studies have shown that when used at low concentration, with short time of gelling and intensive rinsing of the beads, no barium leakage and cytotoxic effect were observed.

Bioencapsulation Procedures As previously mentioned, numerous natural and synthetic polymers have been explored for microencapsulation design, but alginates have been the most used polymers for bioencapsulation due to easy gelling properties and biocompatibility.

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Hydrogel based microparticles, for cell encapsulation, can be prepared by different procedures, as schematized in Fig. 2, including emulsion method (Fig. 2A), vibrating nozzle procedure (Fig. 2B), gas driven mono-jet (Fig. 2C) or jet cutter (Fig. 2D). Irrespectively of the preparation strategy, the microparticle formation consists of three main steps, as below summarized. Step 1. Preparation of a suspension of viable cells in an aqueous colloidal alginate dispersion (the polymer is usually employed at a concentration ranging from 1% to 3%, w/v). Step 2. Generation of alginate droplets which represent microparticle precursors; this step is preferably achieved by a controlled procedure resulting in dimensionally homogenous microdroplets. Step 3. The consolidation of the droplets by a gelation process with final formation of soft hydrogel based microparticles.

Cell Encapsulation by Gas Driven Mono-jet, Vibrating Nozzle and Jet Cutter The encapsulation procedure based on the use of a gas driven mono-jet device (also known as coaxial bead generator), represents also one widely applied approach for the production of microcapsules for cell encapsulation. Different encapsulation hardware is commercially available for the production of alginate microbeads in a controllable manner. The general principle of the device is based on a coaxial air stream that blows droplets from a needle tip into a gelling bath. A typical instrument is composed of a gas driven mono-jet device connected to a precision pump for the alginate feeding (syringe or peristaltic pumps) and to a gas flask (usually nitrogen) equipped with a flow meter, providing gas for the atomization of alginate stream. Generated microdroplets are then consolidated to give the microparticles by a gelation procedure normally based on calcium or barium ion solutions. The feeding cell suspension is continuously stirred to prevent cell clumping, which could lead to inhomogeneous distribution of cells in the microparticles. Typically step 2 represents the most critical phase of the processes. During preparation process a stream of liquid alginate (containing the cells) is forced through a nozzle, generating the individual droplets by gravity or with the aid of various mechanisms. The alginate droplet break-off, at the nozzle tip is achieved by an air-jet directed to the forming droplet while in the case of vibrating nozzle procedure by vibrations or a mechanical cutting in jet cutter encapsulation (emulsion methods do not pertain to this discussion). Varying the experimental settings such as the nozzle diameter, the alginate pumping rate and the applied air-flow or vibration frequency, the droplet diameter can be adjusted. Unfortunately, all the aforementioned fabrication methods can lack of reproducibility in macro-scale environments and they are prone to the production of large and polydispersed microparticles, often characterized by an irregular surface. In this respect, aside the well described microparticle fabrication procedures in recent years, new microfluidic based techniques have been described in response to the need of preparing extremely homogenous microparticles.

Cell Encapsulation by Microfluidics Microfluidics is a new developing technological field dealing with the handling of fluids in micro/nano environments (i.e., the microfluidic chips). This technology has recently found various applications including chemical synthesis, biochemical assays, drug screening. In the case of cell encapsulation, the use of microfluidic chips can overcome some of the drawbacks associated to conventional encapsulation methods, resulting in the continuous formation of microparticles with an elevate control of their size and morphology. Moreover, microchip channels represent a controllable environment that does not need particular sterilization procedures, making the fabrication process suitable for cell encapsulation. Therefore, microfluidics provides a route to encapsulate cells in hydrogel microdevices with a superior control of the process of microencapsulation. Microfluidic strategies for cell encapsulation typically involve the initial formation of an emulsion constituted of aqueous polymer droplets dispersed in a nonmiscible phase (i.e., an oil phase), followed by the consolidations of the droplets. Depending on the chemical structure of the polymer, the liquid droplets are hardened by different procedures including the previously mentioned ionic or thermal gelation (Fig. 3). Homogeneity of microparticle dimension is strictly related to the emulsification step (i.e., the homogeneity of the microdroplets). Optimal microfluidic devices should be indeed able to produce a regular and stable stream of droplets with precise volume and rates. To this end, microfluidic chips have been designed based on different channel geometries such as: (a) cross junction (Fig. 4A), (b) T-junction (Fig. 4B) and (c) microcapillary coaxial devices (Fig. 4C). These different geometries share a common mechanism: the phase to be dispersed (water phase) is injected into a microchannel, where it encounters the immiscible oil phase coming from another inlet(s). The junction where the two phases encounter is designed to optimize the reproducibility of droplet production. Furthermore, the geometry of the junction, the relative flow rates and the physical properties of the fluids, such as viscosity and interfacial tension, lead to droplet pinch off. Notably, size of the formed droplets (into the microchannels) can be controlled by adjusting the flow parameters, therefore microparticles of any size can be virtually obtained. Microparticles or different geometries (e.g., disk-shape particles) can be

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cells in polymer

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gelling bath BaCl2

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alginate microbeads

Fig. 2 Schemes showing some procedures typically employed for production of microparticles for cell encapsulation: emulsion method (panel A), vibrating nozzle procedure (panel B), gas driven monojet (panel C) and jet cutter (panel D).

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liquid droplets

multiphasic flow (emulsion)

consolidation solid particles thermal gelation

ionic cross-linking suspension

agarose microparticles

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Fig. 3 Alternative gelling strategies (i.e., thermal gelation and ionic cross-linking) for hardening/consolidation of liquid droplets generated in microfluidic channels.

generated customing the droplet consolidation inside the channel. Moreover, properly varying hydrodynamic parameters droplets with interesting multiphase morphologies such as Janus microparticles can be also obtained. In the case of alginate, the consolidation of microdroplets into microparticles is accomplished in different ways; when the alginate droplets reside in the microchannels, the consolidation process is defined as “internal gelation.” This procedure involves the dispersion of an insoluble (or slightly soluble) calcium salt (e.g., calcium carbonate) into alginate solution. At the pH reduction by the surrounding continuous phase, Ca2 þ ions are released from the insoluble calcium salt, allowing the crosslinking of the alginate in microparticles. However, the starting of gelation process inside the channel can lead to the formation of clogs inside the chip, by the aggregation of the particle of insoluble calcium salt, or to alginate droplets not completely gelified due to the short resident time into the channel. On the contrary, the external gelation is based on the gelation outside the chip by transferring the obtained alginate droplets into a crosslinking solution where a uniform gelation occurs.

Cross Junction Microfluidic Chips The principle of a cross junction microchip is to create a stream of dispersed phase, later periodically broken up by a flow focusing with a second, immiscible carrier fluid, acting as continuous phase. When these two phases come in contact, droplets in a drop-bydrop fashion are generated by the shear force imposed by the stream of the continuous phase. The formation of droplets occurs, indeed, by a balance between the interfacial tension and the shear of the continuous phase, depending on fluid flow rate, viscosity and microchannel dimensions. In cross junction chip many different geometrical variants are available but in all channel geometries the dispersed phase is pumped through the central inlet and the second immiscible fluid through the two lateral inlets. In order to obtain hydrogel microparticles, the uniform droplets formed at the cross junction are subsequently hardened by different gelling mechanisms.

T-Junction Microfluidic Chips Although very simple in design, the T-junction represent a largely employed geometry for microfluidic devices used for microparticle production. In T-junction chip, the dispersed phase meets at 90 degrees, in a T-shaped junction, the continuous phase. The forming droplet of the dispersed phase obstructs the channel as it grows, resulting in a restriction of the flow of the continuous phase around it. This reduction in the gap through which the continuous phase can flow causes to a dramatic increase in the dynamic pressure which squeezes the dispersed phase leading to the droplet formation. Manipulating relative flow rates of the two fluids, the precise control of droplet size can be reached together with the synchronization of alternately formed droplets by two T-junctions. T-junction, for its simplicity, represent the most effective geometry for

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cross-junction device

external phase

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external phase

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internal phase

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external phase internal phase

Fig. 4 Alternative strategies for the production of microdroplets in microfluidic devices. Geometries currently used are: cross-junction (A) T-junction (B), and microcapillary coaxial (C).

dispersing droplets in a continuous phase. Moreover, coalescence probability of two consecutive droplets is low, depending greatly on flow rate, as droplets tend to flow out to the main channel one by one. Regarding material fabrication, microfluidic devices with T-junction geometry are generally constituted of glass or polydimethylsiloxane (PDMS).

Microcapillary Coaxial Device Method Capillary microfluidic devices are microfluidic systems assembled generally from the alignment of glass tubes: a cylindrical internal tube with a rectangular or cylindrical outer channels. The key advantages of glass microfluidic devices reside in the highly resistance to chemicals and the possibility to create three dimensional flows. Moreover, the glass surface can be hydrophobic or hydrophilic after specific treatment with surface modifiers. Generally, to build a microcapillary device a cylindrical glass capillary with an outer diameter of about 1 mm is used. The capillary is later heated and pulled to obtain a gradually narrowing shape with a fine orifice. The latter is then inserted into another glass capillary with square cross-section to assemble the final microfluidic device. To achieve the coaxial alignment of the two glass capillaries (the rounded and the squared ones) the outer diameter of the rounded capillary has the same dimensions of inner diameter of the squared one. By pumping one fluid through the circular capillary and another immiscible fluid through the square capillary, flowing in the same direction a coaxial flow is obtained. On the contrary, when the two immiscible fluids are introduced from the two ends of the same square capillary but in opposite directions a flow-focusing mechanism is achieved. The outer fluid hydrodynamically focuses the liquid stream of the inner fluid through the small orifice of the rounded capillary, provoking the generation of extremely regular droplets. The droplet size can be tuned by adjusting the fluid flow rates.

Preparation of Fibrous Scaffolds Scaffolds in a fibrous form could be an appealing substitute to spherical microparticles, representing an example of scaffold with controlled physical and architectural features for tissue engineering applications or cell delivery. This statement derives from the fact that fibrous scaffold could enable the guided growth, alignment and migration of cells and they could find many potential applications as small vascular grafts, nerve conduits, artificial kidney tubules as well as cell release vehicles.

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Over the years, different technologies have been developed to create fibers with well-defined physical and architectural features including melt spinning wet spinning and electrospinning. Many of these approaches, however, have limitations on the control of the morphology, dimension, or direction of alignment of obtained fibers. Moreover, when hydrogels and cells pass through conventional nozzles they are subjected to high shear forces causing cytotoxic effects. Conversely, the shear stress developed in microfluidic channel is significantly reduced during fiber production due to the envelop of a laminar core by a sheath flow. Furthermore, microfluidics, due to small channel dimensions and efficient mixing, allows a precise control of the dimensional and morphological characteristics of the produced microfibers, providing new routes for in situ fabrication of fibrous scaffold.

Further Reading Nicodemus, G. D., & Bryant, S. J. (2008). Cell encapsulation in biodegradable hydrogels for tissue engineering applications. Tissue Engineering Part B Review, 14, 149–165. Orive, G., Santos, E., Poncelet, D., Hernández, R. M., Pedraz, J. L., Wahlberg, L. U., De Vos, P., & Emerich, D. (2015). Cell encapsulation: Technical and clinical advances. Trends in Pharmacological Sciences, 36, 537–546. Krishnan, R., Alexander, M., Robles, L., Foster, C. E., 3rd, & Lakey, J. R. (2014). Islet and stem cell encapsulation for clinical transplantation. Review of diabetic Studies, 11, 84–101. Silva, R., Fabry, B., & Boccaccini, A. R. (2014). Fibrous protein-based hydrogels for cell encapsulation. Biomaterials, 35, 6727–6738. Kang, A., Park, J., Ju, J., Jeong, G. S., & Lee, S. H. (2014). Cell encapsulation via microtechnologies. Biomaterials, 35, 2651–2663. Mazzitelli, S., Capretto, L., Quinci, F., Piva, R., & Nastruzzi, C. (2013). Preparation of cell-encapsulation devices in confined microenvironment. Advanced Drug Delivery Review, 65, 1533–1555. Rokstad, A. M., Lacík, I., de Vos, P., & Strand, B. L. (2014). Advances in biocompatibility and physico-chemical characterization of microspheres for cell encapsulation. Advanced Drug Delivery Review, 68, 111–130. Hunt, N. C., & Grover, L. M. (2010). Cell encapsulation using biopolymer gels for regenerative medicine. Biotechnology Letters, 32, 733–742.

Drug Delivery Systems and Controlled Release Nicholas J Kohrs, Thilanga Liyanage, Nandakumar Venkatesan, Amir Najarzadeh, and David A Puleo, University of Kentucky, Lexington, KY, United States © 2019 Elsevier Inc. All rights reserved.

Introduction History and Progression of Drug Delivery Systems Topical Delivery Transdermal Delivery GI Delivery Nasal Delivery Mucoadhesive Delivery Parenteral Delivery Implantable Carriers Solid Dosage Forms Tablets and Capsules Microparticles and Nanoparticles Films Mechanisms of Controlled Drug Delivery Diffusion-Controlled Release Erosion-Controlled Release Targeted Release Active targeting Passive targeting Stimuli-Responsive Release pH Temperature Drug Delivery Materials Synthetic Polymers Polyesters Poly(ortho esters) Polyanhydrides Polyamides Polymeric prodrugs Biopolymers Lipids Surfactants Ceramics Hydroxyapatite systems Biphasic calcium phosphate systems Bioglasses Metallic Nanoparticles Future Materials and Systems Carbon Nanotubes Microchips 3-D Scaffolds 3D Printing Conclusion Further Reading

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Glossary Condensation polymerization A process in which two low-molecular-weight subunits combine to form a molecule with a higher molecular weight, typically generating by-products such as water or alcohol. Copolymer A higher-molecular-weight substance formed by polymerizing two or more monomers together. Cross-linking Act of bonding two or more polymers together.

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Crystallinity Referring to the degree of structural order of a material, crystalline (highly ordered) or amorphous (highly disordered). Emulsification The process of mixing together two or more liquids in which one liquid forms microscopic droplets in the other. Hydrolysis Process in which water breaks down chemical bonds. Ring-opening polymerization Process of opening cyclic monomers and covalently joining them to produce a highermolecular-weight chain. Solution polymerization Polymerization process in which the monomer, solvent, and initiator are mixed together prior to polymerization.

Abbreviations 3DP Three-dimensional printing BCNU Bis-chloroethylnitrosourea BCP Biphasic calcium phosphate CNT Carbon nanotubes EPR Enhanced permeability and retention FDA Food and Drug Administration GI Gastrointestinal HA Hydroxyapatite LCST Lower critical solution temperature PCL Poly(ε-caprolactone) PEG Poly(ethylene glycol) PGA Poly(glycolic acid) PLA Poly(D,L-lactic acid) PLGA Poly(D,L-lactic-co-glycolic acid) PLLA Poly(L-lactic acid) RGD Arg-Gly-Asp SiRNA Small interfering RNA TCP Tricalcium phosphate Tg Glass transition temperature

Introduction Drugs have been used throughout history to aid in the treatment of illnesses that plagued the populace of the period. During this time frame, modifications were made to drugs and their associated delivery systems, improving their bioavailability and efficacy. Fast-forwarding to modern times, drug delivery systems now also play a major role in wound healing and tissue regeneration, allowing for antibiotics, growth factors, hormones, and other therapeutic agents to be released in a controlled manner to the affected areas. In addition to treating systemic conditions, the need for controlled and localized delivery of drugs has never been more pertinent for site-specific applications. Fig. 1 shows how uncontrolled release permits variability in drug concentration, which can lead to either toxic or ineffective drug levels. Controlled release aims to prevent this variability and maintain an appropriate therapeutic dose for the desired duration. Remarkable material innovations, system formulations, and strategies have been developed for a myriad of local and systemic applications, such as heart disease, diabetes, wound healing, and cancer treatments. This article summarizes the history and progression of drug delivery systems, solid dosage forms, mechanisms of controlled release, and materials used in controlled drug delivery systems.

History and Progression of Drug Delivery Systems The history of controlled release originates from ancient times when drugs were either applied directly to the injury as a paste or poultice or administered orally, such as a medicinal wine. As history progressed, so too did the methods for drug delivery. Current

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DRUG CONCENTRATION

TOXIC LEVEL

UNCONTROLLED RELEASE

CONTROLLED RELEASE

MINIMUM EFFECTIVE CONCENTRATION

INEFFECTIVE LEVEL

TIME Fig. 1 Uncontrolled release versus controlled release. Controlled release reduces undesirable fluctuations of drug concentration in systemic circulation, thus diminishing side effects and improving treatment outcomes.

systems and methods of drug administration include gastrointestinal (GI), transdermal, topical, nasal, mucoadhesive, parenteral, and implantable carrier systems. Table 1 summarizes advantages and disadvantages of the different approaches.

Topical Delivery Topical systems focus on localized delivery to the top layers of the integumentary system, that is, the epidermis and dermis. Delivery of antiseptic, antifungal, or anti-inflammatory agents, along with emollients, provides immediate treatment and protection to the different layers of the skin. Topical systems attempt to locally accumulate drugs, via chemical permeation enhancers, deep within the strata of the skin. Systemic effects are still possible but are hindered due to the slow diffusion of the drug into the surrounding vasculature.

Table 1

Advantages and disadvantages of common drug delivery approaches

Approach

Advantages

Disadvantages

Transdermal/topical

1.Sustained delivery 2.Avoids first-pass metabolism 3.Reduced dosing frequency 4.Steady absorption 5.Noninvasive 1.Convenient 2.Simple design 3.Safest administration route 4.Low cost

1.Possible skin reaction 2.Delayed drug action 3.Variability in drug absorbance rates due to application location 4.Adhesion failure 1.Gastric retention time 2.Requires high levels of gastric liquid and food 3.Drugs with solubility and stability issues in highly acidic environment cannot be used 4.First-pass metabolism 1.Possible nasal irritation 2.Potential toxicity of absorption enhancers 3.Difficult to halt after administered 1.Potential for damage to mucosa 2.Patient discomfort

Gastrointestinal

Nasal Mucoadhesive

Parenteral

Implantable

1.Rapid absorption 2.Quick drug action 3.Avoids first-pass metabolism 1.Prolonged residence times 2.Localization of treatment 3.High drug flux at absorbing tissues 4.Avoids first-pass metabolism 1.Highest bioavailability 2.Immediate physiological response 3.Avoids first-pass metabolism 4.Dosing control 1.Eliminates need for patient compliance 2.Avoids first-pass metabolism 3.Potential for zero-order release

1.Pain 2.Potential injection site infection 3.Patient aversion to needles 4.Difficult to halt after administered 1.Invasive 2.Limited loading capacity 3.Biocompatibility considerations 4.Surgical removal at therapy termination 5.Possible device failure and drug dumping

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Transdermal Delivery The ease of starting, stopping, and restarting treatment makes transdermal and topical delivery convenient pathways. Transdermal delivery aims to rapidly introduce therapeutic agents into systemic circulation. This is a difficult task for certain drugs due to the protective integumentary system. Drug transport, across the skin, can be accomplished by chemical permeation enhancers, allowing < 400 kDa-sized drugs to pass, or electroporation, opening pores to allow passage of  4000 kDa-sized drugs. Transdermal treatment aims to increase drug bioavailability by avoiding first-pass metabolism.

GI Delivery Delivery of dosage forms, that is, pills, tablets, and capsules, through the GI pathway is the most convenient and commonly employed pathway for administering drugs. Depending on patient compliance, appropriate drug concentrations can be delivered daily, allowing for an extended time period of drug–tissue contact. This pathway is limited, however, by first-pass metabolism, wherein the drug is degraded or metabolized in the GI tract or liver prior to reaching systemic circulation, leading to decreased therapeutic blood concentrations and longer time to achieve said concentrations.

Nasal Delivery The nasal pathway is a logical choice for treating issues affecting the nose or paranasal sinus, such as sinusitis and rhinitis, and the ease of administration and quick transport into systemic blood circulation makes nasal delivery convenient for systemic conditions, such as drug overdose, diabetes insipidus, and prolonged labor, where therapeutic agents are needed rapidly. The large surface area, high permeability, and close proximity to blood vessels allow for a bypass of first-pass metabolism and rapid attainment of systemic therapeutic drug levels. This route of administration is limited by solubility and drug size, making it unattainable for most large or hydrophobic drugs.

Mucoadhesive Delivery Mucoadhesive systems allow for intimate contact between dosage form and the mucosa, resulting in localized release, increased residence time, rapid adsorption, and high drug flux at the site of adherence to provide high bioavailability. These systems can take the form of tablets, films, microparticles, and gels for application to sites such as the oral cavity, eye conjunctiva, nasal cavity, vagina, and GI tract.

Parenteral Delivery Parenteral delivery is defined by the US Food and Drug Administration (FDA) as drug administration by injection, infusion, and implantation or by some other route other than the alimentary canal. Parenteral injections and infusions allow for systemic drug delivery and can be divided into several categories, including intravenous, intramuscular, intraarticular, epidural, intrathecal, intraosseous, and subcutaneous injections. Because this approach avoids physical barriers such as the skin or mucosa and first-pass metabolism, parenteral delivery has high levels of drug bioavailability and quick drug action.

Implantable Carriers Implantable carrier systems, such as implantable rods, osmotic pumps, ceramics and glasses, and synthetic polymers, allow for both site-specific administration and systemic distribution of drugs. Advantages of implantable devices include elimination of need for patient compliance, lower dosage required due to avoidance of first-pass metabolism, and ability to tune drug release kinetics. Drug release can be actively or passively controlled, depending on the type of implantable system. For instance, implantable drug pumps, such as an insulin pump, can be programmed to administer an appropriate dose at a specific time or measureable level. Passive releasing systems, such as silicone-based birth control, allow for diffusion from the implant to provide long-term sustained release. Starting from simple solutions, pastes, and pills to deliver therapeutic agents, drug delivery and controlled release have developed into a complex field combining aspects of materials science and engineering, pharmacology, anatomy, and physiology. Many different dosage forms have been developed to treat specific tissues and locations.

Solid Dosage Forms The development of dosage forms has had a profound impact on the efficacy of controlled release systems. Insight to the physicochemical properties of a drug enables selection of certain materials and mechanism to obtain the desired release kinetics. Without a proper delivery system, adverse consequences, for example, premature release, tissue toxicity, and device failure, can occur. A variety of dosage forms have been developed to account for the different physiological and mechanical stresses that these systems will experience. Current solid dosage forms include tablets and capsules, micro- and nanoparticles, and films.

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Tablets and Capsules Tablets and capsules are the dosage forms of choice when it comes to drug delivery via the GI tract because they can be selfadministered and deliver an accurate dose. Tablets and capsules are designed to readily degrade once within the proper physiological environment, whether the stomach or intestines. These dosage forms are created either by compressing a granulated drug into a desired shape (tablets) or by encapsulating the drug into a gelatin casing (capsules). Current formulations use a variety of processing steps to both improve delivery and sustain release of their loaded agents. For example, some formulations use enteric coatings to prevent degradation until the dosage form has entered the small intestine, while other formulations use laser-drilled holes to enhance drug transport from the dosage form. Yet, others employ different materials to alter the physicochemical properties, for example, mucoadhesivity, of the dosage form.

Microparticles and Nanoparticles Nano- and microparticles are increasingly used in controlled release systems, and the difference in size has numerous effects. Nanoscale carriers offer enhanced versatility when compared with particles of larger size owing to their small sizes and improved dispersibility. In addition, circulation time of nanoparticulate vehicles is prolonged, and their size enables easier uptake for intracellular delivery of drugs. Nanoscale carriers are increasingly modified for attachment targeting ligands or for “stealth” capabilities by conjugation with poly(ethylene glycol) (PEG). Microspheres and nanospheres can employ many different materials, for example, polymers, glasses and ceramics, and metals, each with different processing techniques. Polymers are most prevalently used in the formulation of microspheres for drug delivery. Ceramics, glasses, and metals are gaining interest because they can be bioactive and have better mechanical properties. Typically, ceramics, metals, and bioactive glasses are loaded via adsorption, whereas polymeric particles are formulated via emulsification. Emulsions rely on the hydrophobicity or hydrophilicity of drug molecules to partition into the respective phase. For hydrophobic drugs, an oil-in-water (O/W) single emulsion is used because the oil phase dissolves the therapeutic agent for emulsification. For hydrophilic drugs, either a water-in-oil (W/O) single emulsion or a water/oil/water (W1/O/W2) double emulsion is used. In either of these latter cases, the therapeutic agent is dissolved in an aqueous solution. Particles are collected for storage after solvent evaporation. Particle size is determined by multiple factors, including organic solvent, polymer concentration, surfactant, and emulsification energy. Large surface area-to-volume ratios expose a greater percentage of the loaded drug to the external phase, which can lead to adverse outcomes, such as small particles having a lower maximal drug loading and a sizable loss of payload.

Films Film delivery systems are typically flexible sheets of polymeric matrices incorporating the therapeutic agent. Interest in polymeric films as dosage forms has increased as an alternative approach to conventional dosage forms. Films can be used for systemic or local action via several approaches, for example, oral, topical, GI, and transdermal delivery. An example with significant clinical impact is found in drug-eluting stents. These devices encompass thin polymeric films deposited onto the metallic struts of a stent to release a drug, such as paclitaxel or sirolimus, to control cell proliferation and prevent restenosis of the blood vessel. Film formulations offer several advantages, including convenient administration through noninvasive routes, that is, transdermal or oral mucoadhesive patches, and ease of handling during manufacture and transportation. Commonly used techniques for the preparation of films include solvent casting, photopolymerization, hot-melt extrusion, and three-dimensional printing (3DP). In solvent casting, a solution of polymer and drug is dispensed into a suitable mold, and the solvent is allowed to evaporate, which leaves a drug-loaded polymeric film. The rheological properties of the polymeric mixture affect the drying rate, film thickness, morphology, and content uniformity of the films. Hot-melt extrusion is a substitute method that is especially useful when an organic solvent system must be avoided. A mixture of pharmaceutical ingredients is melted and charged through a die to obtain homogeneous matrices. 3DP is briefly described in the section “Future Materials and Systems.” Microparticles, nanoparticles, films, and other drug delivery systems each have advantages and disadvantages. Different properties can be achieved by combining two or more of these systems together to confer the advantages of one class of vehicle on another. This approach has been used, for example, with microparticles in films, nanoparticles in a cross-linked hydrogel matrix, microparticles in scaffolds, and many other combinations. All steps, including the synthesis method, encapsulation process, or final surface modification for targeted delivery, determine the characteristics of these systems and their main goal, that is, the controlled release of bioactive agents.

Mechanisms of Controlled Drug Delivery Drug delivery systems are used to enhance the therapeutic effect of the drug at the target site by protecting the drug under physiological conditions, controlling its release, increasing or reducing systemic circulation, and assisting in reaching the site of action. These functions are achieved by complex and diverse mechanisms that work individually, in combination, or sequentially.

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Polymers, ceramics, glass, metals, and metal alloys are often used as vehicles in the drug delivery systems. Control of drug release from dosage forms is multifactorial, involving diffusion of drug out of or water into the system, dissolution of the drug, partitioning, system/matrix swelling, degradation, or erosion.

Diffusion-Controlled Release Most polymeric drug delivery systems are diffusion-controlled, where the drug diffuses from the vehicle along a concentration gradient and the release is either Fickian or non-Fickian. Glass transition, also known as second-order transition, is a characteristic nonphase transition that occurs at a temperature specific to each polymer. Below the glass transition temperature (Tg), polymers turn rigid, hard, and dimensionally stable and are considered to be in a glassy state, but above Tg, they turn soft and are said to be in a rubbery state. Polymers in a rubbery state allow better infiltration of the surrounding medium, allowing dissolution and subsequent diffusion of the drug. When diffusivity of the drug is independent of concentration, the release is Fickian, but if diffusion depends on time and concentration, then the release is non-Fickian. Two kinds of diffusion-controlled systems are monolithic and reservoir-based (Fig. 2A). In monolithic systems, the drug is uniformly distributed within the vehicle. The rate of diffusion depends on the diffusion coefficient of the molecule, and it decreases as a function of time because the concentration of drug in the core decreases with time. In reservoir systems, the diffusion coefficient is controlled by the rate-limiting membrane, because the drug partitions from the core into the membrane and then diffuses from the membrane due to a concentration gradient. When saturation of the core drops, diffusion from the membrane decreases.

Erosion-Controlled Release Polymers can be degraded actively (by enzymes) or passively (by hydrolysis), resulting in surface or bulk erosion (Fig. 2B). Surface erosion is a heterogeneous process wherein degradation of the polymer happens at only the surface and the rate is proportional to the surface area. Drug release in surface-eroding systems is often correlated with a predictable erosion rate, which is considered ideal for many drug delivery applications. Erosion begins from the outside and progresses inward. In most cases, thicker systems have longer erosion times, and hydrophilic polymers degrade faster compared with hydrophobic materials. With bulk-eroding systems, degradation is homogenous throughout the material, and the size of the system remains constant in most cases. The drug is released in three stages: burst release from the surface, release from initial degradation of the polymer, and release of residual drug during complete degradation/homogeneous erosion of the polymer. Polyanhydrides and polyesters are typical examples for the surfaceand bulk-eroding systems, respectively. Hydrophobic or medium-insoluble drugs are generally released by erosion.

Targeted Release Targeted drug delivery, postulated as a “magic bullet” in the 1990s, is widely studied to deliver drug to a specific tissue or organ to achieve a desired therapeutic outcome without activating side effects. Efficiency of this approach depends on three important factors: specific ligand that is characteristic to the tissue of interest, effective drug for the condition to be treated, and appropriate mechanism for delivery of the drug. Targeted delivery systems are achieved by two different mechanisms, including ligand–receptor interaction (active) and physiological changes (passive).

Active targeting Active targeting involves conjugating the drug/carrier system to a specific ligand that guides the drug to the target site. The ligand acts as a “homing signal” that improves selective drug delivery to a specific organ. Antibodies are commonly conjugated to drugs

(A)

(B)

Monolithic

Surface Erosion

Bulk Erosion

Time

Reservoir

Time

Fig. 2 (A) Schematic representation of monolithic and reservoir drug delivery system. (B) Changes in the polymeric delivery system during surface and bulk erosion.

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because of their specificity. For example, some of the promising studies include a prostate-specific membrane antigen antibody J591 conjugated to a polymer-containing drug that is used to target prostate cancer cells. Folate receptors are also widely targeted because of their overexpression in most cancer tissues, and folate-conjugated doxorubicin polymer aggregates actively target the folate receptor in cancer cells and inhibit tumor progression. Similarly, Arg-Gly-Asp (RGD) peptide is used to target integrin avb3 receptors that are overexpressed in neovasculature and activated endothelial cells to deliver nucleic acids constructed with polymer aggregates (siRNA–polyethyleneimine–polyethylene glycol–RGD) to achieve tumor-specific and also gene-specific targeted therapy.

Passive targeting Passive targeting refers to drug accumulation at the target site via systemic circulation without any specific ligand. Accumulation is achieved through physiological changes, such as the dense vasculature with loose or defective architecture associated with most solid tumors. The enhanced permeability and retention (EPR) effect is a molecular-weight phenomenon in which highmolecular-weight molecules are retained in the dense vasculature due to poor lymphatic drainage. This characteristic leaky vascular architecture is widely used to deliver the drug selectively to the tumor. For example, increased retention of the low-molecular-weight antitumor protein neocarzinostatin was achieved by increasing its weight by conjugation with poly(styrene-co-maleic anhydride). This effect has also been studied with other antitumor drugs, such as doxorubicin encapsulated in PEG-coated liposomes, which increase circulation and localization within the tumor.

Stimuli-Responsive Release The release pattern of the drug delivery system depends not only on the stimulus but also on the material from which the drug is released. Responsive polymers undergo a conformational change responding to specific physiological stimuli. There are two types of responsive drug delivery systems, including open-loop (externally regulated) and closed-loop (self-regulated) systems. Open-loop systems are nonpolymeric drug release systems, where the delivery of drug is predetermined, for example, an insulin pump. In the case of closed-loop systems, drug delivery is determined by physiological conditions, such as pH and temperature.

pH The characteristic differences in pH at various sites, for example, GI tract, blood, and intracellular environment, along with the acidic conditions associated with most cancers and infected tissues, make pH an attractive stimulus for responsive polymers. pHresponsive polymers are polyelectrolytes with weak acidic or basic groups that undergo a change in their ionization state by accepting or donating protons. Changes in the ionization state result in conformational changes and swelling behavior of the polymer that consequently modulate drug delivery. For example, studies on chitosan/heparin nanoparticles tested against Helicobacter pylori protect the drug from gastric secretion (the positive surface charge is stable at pH 1.2–2.5), infiltrate through the mucus layer (pH 4.5–7), and disintegrate at the infection site (nanoparticles are deprotonated at pH  7) to release the drug. Similar studies have shown that hydrogels with 2,4,6-trimethoxybenzaldehyde groups exhibit a hydrophobic nature at pH 7.4 and transition to hydrophilic at pH 5, releasing a drug found to be effective against lung cancer.

Temperature Certain polymers exhibit phase transition at a specific temperature that alters solubility. Changes in the solvation state are attributed to the inter- and intramolecular hydrogen bonding of the polymer. This transition can be used for drug release from polymers that expand or collapse at body temperature. Poly(N-isopropylacrylamide) is a common thermoresponsive polymer that is widely tested for drug delivery applications. Below the lower critical solution temperature (LCST) of 32–33 C, water molecules bond to the polymer network to cause swelling, which enables loading of drugs within the polymer. Above the LCST, a hydrophilic to hydrophobic transition causes the polymer network to collapse and become insoluble, thereby releasing all the water molecules and entrapped drug during transition.

Drug Delivery Materials Materials used in controlled release systems can be classified into many categories, but this section will focus on both biodegradable materials, for example, biodegradable polymers, lipids, and ceramics, and nonbiodegradable materials such as magnetic nanoparticles. Nonbiodegradable materials can be used in therapeutic applications only if the material can be recovered or excreted after drug release, for example, intraocular implants, buccal patches, and chemotherapy nanoparticles. Most therapeutic applications use a biocompatible material that will degrade within the body and whose degradation products are not cytotoxic. Due to these constraints, the study of synthetic polymers has increased and has yielded several successful systems. Table 2 summarizes the advantages and disadvantages of different materials, while Table 3 lists the examples of commercially available products that employ these materials.

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Drug delivery material advantages and disadvantages

Material

Advantages

Disadvantages

Polyesters

1.Simple synthesis 2.Physical and chemical properties can be tuned 3.Easily fabricated into a variety of dosage forms 1.Zero-order release kinetics 2.Surface erosion is slow 1.Easy synthesis by low-cost resources by one-step synthesis. No need of any purification step 2.Degrade into nontoxic carboxylic acids at a predictable rate 3.Degradation can be tuned 1.The carboxyl and amino groups provide synthetic handles for further functionalization of the polymer backbone 1.High drug loading 2.Prolonged and localized release 1.Mucoadhesive property suitable for use in both oral and nasal formulations 2.Ability to condense DNA to form complexes to form nonviral gene delivery vehicles 1.The most abundant biomaterial 2.Can be processed into different forms 1.Pharmaceutical stability 2.Ability to deliver hydrophilic and lipophilic drugs 3.Long drug circulation times

1.Bulk erosion causes nonlinear degradation kinetics 2.Acidic degradation products can cause irritation

Poly(ortho esters) Polyanhydrides

Polyamides Polymeric prodrugs Chitosan

Collagen Lipids

Hydroxyapatite

1.Tunable morphology leading to higher drug loading 2.Localized delivery (implanted)

Biphasic calcium phosphate

1.Tunable dissolution rates and mechanical properties 2.Synthesis comparatively straightforward, easy, and inexpensive 3.Localized delivery (implanted) 1.Delivers hydrophilic drugs 2.Ability to carry high dose of drug in nanoparticles 3.Regenerative properties with therapeutic effects 1.Can penetrate biological barriers 2.Biocompatible and inert (depending on material) 3.High drug loading

Bioglass Metal nanoparticles

Table 3

1.Synthetically complex 2.Need for excipients to accelerate degradation 1.Hydrolytic instability. Requires storage under cold, moisture-free conditions 2.Low mechanical strength and film or fiber-forming properties 1.Immunogenic 2.Poor mechanical properties 3.Require enzymes to degrade 1.Synthetically complex 1.Slow enzymatic degradation

1.Mildly immunogenic 1.Low encapsulation efficiency 2.Premature membrane degradation 3.Low solubility 4.Difficult to scale up 1.Brittle 2.Quick clearance (of nanoparticles) 3.Loading limited by absorption and electrostatic interactions 1.Brittle 2.Low loading efficiency 1.Poor mechanical properties 1.Potential toxicity 2.Removal issues 3.Susceptible to physical instability

Examples of commercially available drug delivery systems using different materials

Material

Commercially available products

Polyesters

Lupron Depot®, Decapeptyl®, Risperdal Consta®, Eligard® Alzamer® Gliadel®, Septacin® TransCon Growth Hormone, PolyAspirin INFUSE® Lipoquin®, Neoral®, Kaletra®, Sustiva® KaliLight Ferumoxytol

Poly(ortho esters) Polyanhydrides Polymeric prodrugs Collagen Lipids Calcium phosphates Metallic nanoparticles

Synthetic Polymers Synthetic polymers consist of long chains of repeating molecules that are designed with a variety of chemical and physical properties in mind. Most of these materials possess ester, anhydride, or amide linkages that enable degradation by hydrolysis or enzymatic activity in vivo. Biodegradability makes it possible to administer the controlled release system without the need for a second

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procedure to remove the exhausted delivery vehicle. Because of the large variety available, synthetic polymers are the most widely used materials for developing controlled release systems.

Polyesters Polyesters are the most commonly studied biodegradable polymeric materials. They can be easily prepared by ring-opening polymerization of the corresponding cyclic lactone monomer or by condensation polymerization. They mainly degrade via hydrolysis of ester bonds. Polyester-based biodegradable polymers include poly(D,L-lactic acid) (PLA), poly(glycolic acid) (PGA), poly(D,L-lacticco-glycolic acid) (PLGA), and poly(ε-caprolactone) (PCL). PLA is prepared by polymerizing lactic acid monomers or by ring-opening polymerization of cyclic lactide monomer. Lactic acid has an asymmetrical a’carbon, which gives rise to R and S enantiomeric forms commonly referred to as the D and L forms of lactic acid. The extra methyl group in the side chain makes PLA more hydrophobic compared with its chemically similar analogue PGA. Poly(L-lactic acid) (PLLA), which is prepared using all L-lactic acid as the monomer unit, has a glass transition temperature of 60–65 C and a melting temperature of  175 C. Due to its good tensile strength, high modulus ( 4.8 GPa), and slow degradation ( years) due to high crystallinity of PLLA polymers, they are considered appropriate for load-bearing applications, such as orthopedic fixation (e.g., Phantom Suture AnchorÒ (DePuy)). The amorphous racemic PLA polymer that consists of both D and L form of lactic acid has a glass transition temperature of 55–60 C, lower tensile strength ( 1.9 GPa), and relatively fast degradation time ( months) compared with PLLA. Note that degradation is highly dependent on the molecular weight of the corresponding polymeric material, its composition, and the processing method, with higher-molecular-weight polymers exhibiting slower degradation. The amorphous form of PLA has applications in drug delivery and scaffolding materials for tissue regeneration because of its low strength, lower crystallinity, and comparatively faster degradation profile. PLA is used in dosage forms such as microparticles or nanoparticles for delivery of drug molecules, peptides/proteins, and nucleic acids. To alter crystallinity and degradation kinetics, lactic acid is often copolymerized with other monomers, oligomers, or polymers, for example, glycolide, caprolactone, and PEG, to prepare di- or triblock copolymers. PGA is also obtained by ring-opening polymerization of the cyclic diester of glycolic acid known as a glycolide. PGA shows low solubility in common organic solvents (e.g., dichloromethane, tetrahydrofuran, acetone, chloroform, and ethyl acetate) due to its hydrophilicity. With a melting temperature of 225 C, PGA is a hard, tough, crystalline polymer. Due to its fiber-forming properties, PGA was used as the first synthetic absorbable suture material (DEXONÒ). Poor solubility, high melting point, and high rate of degradation limit its uses in drug delivery. Copolymerization of lactic and glycolic acids gives rise to the polymeric material commonly known as PLGA. Properties of PLGA polymers can be easily tuned by controlling the ratio of lactide (L) to glycolide (G) (L is more hydrophobic due to the a-methyl group). For example, PLGA polymers with 50:50, 75:25, and 85:15 (L/G) have degradation timescales of 1–2, 4–5, and 5–6 months under in vitro conditions (isotonic saline, pH 7.4, and at 37 C). The degradation products produced may lead to an increase in acidity that can cause irritation at the implant site. This can be overcome by incorporating basic salts to control the pH. PLGA has been used to form various drug delivery vehicles, such as microspheres and nanospheres, nanofibers, and films, for controlled release of therapeutic agents. For example, Lupron DepotÒ is a PLGA microsphere-based drug delivery system that releases gonadotropin to treat prostate cancer. The nontoxic, water-soluble PEG can be introduced into PLGA to form PLG–PEG diblock (AB) and PLGA–PEG–PLGA (ABA) triblock, multiblock, or branched block copolymers exhibiting different physicochemical properties, including microphase separation, crystallinity, water solubility, and biodegradability. The presence of PEG in the backbone results in a significantly more hydrophilic and protein-resistant polymer, allowing longer circulation time in the body. The amphiphilic PLGA–PEG block copolymers can also form micelles used for delivering hydrophobic drugs. The size and morphology of micelles can be fine-tuned by adjusting the chemical composition, total molecular weight, and ratio of the PLGA and PEG block lengths. Various hydrophobic drugs, such as the chemotherapeutic paclitaxel, have been incorporated into these micelles. Another commonly used degradable aliphatic polyester is PCL. PCL is synthesized by ring-opening polymerization of the sixmembered ring ε-caprolactone using a catalyst such as stannous octoate. Due to extremely slow degradation rates, PCL is often used as a long-term scaffold or drug/vaccine delivery material, for example, the contraceptive device CapronorÒ. PCL has a glass transition temperature of approximately 60 C, making it semirigid at room temperature. Block copolymers are synthesized by copolymerizing PCL with other monomers or polymers, for example, PLGA and PLA, to decrease the degradation time. Hydrophilic block segments, such as PEG, can be conjugated to the hydrophobic PCL portion to prepare micelles. For example, PCL–PEG diblock copolymer micelles were used to formulate doxorubicin for targeted drug delivery.

Poly(ortho esters) This class of polymers is synthesized by transesterification or by addition of polyols to diketene acetals. The orthoester linkages that connect polymer segments can undergo rapid hydrolysis, but the bulk of the polymer is hydrophobic and minimizes the water penetration. Therefore, hydrolysis and the resulting drug release occur only at the surface of the material (surface erosion). The pH sensitivity of orthoester bonds allows adjustment of degradation rate; introducing acidic excipients, for example, suberic acid, to the polymer matrix accelerates degradation, while neutral and basic excipients stabilize the polymer matrix by neutralizing the acidic by-products formed through hydrolysis. The recent modification of the poly(ortho ester) backbone by incorporating glycolide or lactide avoids the need for excipients, because once hydrolyzed, the resulting lactic or glycolic acid monomers can catalyze

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orthoester bond breakage. As an example, a poly(ortho ester) synthesized from 1,10-decanediol and 1,10-decanediol dilactide was used to prepare a tetracycline delivery system for treating periodontal disease.

Polyanhydrides Polyanhydrides are mainly prepared by melt condensation of dicarboxylic acid with an acetic anhydride at a high temperature under vacuum. The most widely studied polyanhydrides are based on sebacic acid, adipic acid, and terephthalic acid. The anhydride bonds hydrolytically cleave and form water-soluble products. The hydrophobic polymer backbone prevents water penetration into the bulk of the polymer, which causes the polymer to degrade by surface erosion. Consequently, the polymer maintains nearly zeroorder degradation kinetics while retaining its structural integrity. Degradation can be tuned by introducing hydrophobic or hydrophilic monomers into the polymer backbone. Hydrophobic monomers stabilize the anhydride bond over hydrolysis. Relatively less hydrophobic aliphatic polyanhydrides based on sebacic acid degrade fast (within days), while polymers based on more hydrophobic aromatic (p-carboxyphenoxy)hexane take years to degrade. Poly(anhydrides) have many applications and can be fabricated into microparticles, nanoparticles, wafers, etc. For example, poly[(carboxy phenoxy propane)-(sebacic acid)] is used for controlled release of the chemotherapeutic agent BCNU to treat brain cancer (GliadelÒ). Salicylic acid-based polyanhydrides are able to stimulate new bone formation.

Polyamides The most common polyamides are poly(amino acids). Compared with polyesters and polyanhydrides, polyamides are more stable toward hydrolytic cleavage due to the amide linkage but instead rely on enzymes for degradation. The degradation rate can be manipulated by hydrophilicity of the amino acids. The most widely used polyamino acids are poly(g-glutamic acid), poly(3-Llysine), poly(aspartic acid), and their derivatives. These polymers consist of ionizable pendant groups, such as carboxyl and amino groups. The functional groups provide synthetic handles for further functionality of the polymer backbone, for example, glycosylated poly(g-glutamic acid) used for liver-specific targeted drug delivery.

Polymeric prodrugs In these polymers, the bioactive molecule, that is, drug, serves as a repeating monomer unit, or it can be attached as a pendant group. Incorporation of bioactive molecules into the polymer backbone allows for higher drug loading, controlled release, and easy processing. The most widely studied polymeric prodrugs, based on salicylate, are prepared through melt condensation or solution polymerization. These polymers degrade by surface erosion and have controlled and extended release. Ibuprofen, ketoprofen, and naproxen have been used as pendant groups on the polymer backbone. Rather than adding pendants, a polysimvastatin prodrug has been synthesized as a diblock copolymer of simvastatin monomers linked to different molecular-weight PEG moieties to control the hydrophilicity. The PEG moiety initiates the ring-opening polymerization through simvastatin lactone ring in the presence of a stannous octoate catalyst or other catalysts. Release of simvastatin is highly dependent on the PEG moiety incorporated into the polymer backbone.

Biopolymers Biopolymers are used in drug delivery applications due to their availability and biocompatibility. These materials include proteinbased polymers, such as collagen, albumin, and gelatin, and polysaccharide-based polymers, such as agarose, alginate, carrageenan, hyaluronan, dextran, chitosan, and cyclodextrins. Natural polymers have limitations in batch-to-batch reproducibility, uncontrolled rates of hydration, broad molecular-weight distributions, and susceptibility to microbial contamination. The two most widely studied protein- and polysaccharide-based polymers are collagen and chitosan, respectively. Collagen As the most abundant protein in the body, collagen is a major component of musculoskeletal tissues, the skin, and blood vessels. Collagen can be easily cross-linked using chemicals, such as aldehydes, carbodiimides, hexamethylene diisocyanate, and epoxy compounds; elevated temperatures; or high-energy irradiation. Cross-linked collagen is used as scaffolds for tissue engineering and drug delivery vehicles. The drug release profile from collagen matrices can be manipulated by varying the physical properties, such as density, porosity, and degradation kinetics. Cross-linked collagen has also been used as a vehicle for delivery of absorbed/ adsorbed bioactive proteins, such as recombinant human bone morphogenetic protein-2 (INFUSEÒ (Medtronic)). Chitosan Chitosan is a linear polysaccharide consisting of glucosamine and N-acetyl glucosamine units. In vitro degradation can occur by chitosanase, lysozyme, and papain, but in vivo degradation occurs mainly via lysozyme. Chitosan’s degradation kinetics mainly depend on the degree of acetylation and crystallinity; an increase in the degree of deacetylation causes an increase in crystallinity and a decreased degradation rate. Chemical modifications of chitosan can significantly affect its degradation rate. For example, inclusion of bulky side groups disrupts its strong hydrogen-bonding network and accelerates degradation. Due to the strong interaction with the negatively charged mucous membrane and low solubility in water, chitosan is widely used in mucoadhesive systems. Chitosan also shows both anti-inflammatory and antibacterial properties, which are beneficial in wound-healing applications. Chitosan is also used in the production of injectable thermosensitive carriers. Chitosan can be easily blended with other polymers or

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fabricated into nanoparticles, microspheres, gels, or fibers for drug delivery applications. Due to its ability to form electrostatic complexes with DNA, extensive research is ongoing to develop chitosan-based gene delivery systems.

Lipids Lipids are a large and diverse class of biological substances including surfactants, phospholipids, waxes, and oils. These materials do not readily solubilize in water and are typically composed of a long hydrocarbon tail attached to a polar head group. Due to molecular interactions in aqueous solutions, the formation of lipid bilayers is energetically favorable for them to sequester their hydrocarbon tails from water. This allows formulation of systems such as vesicles, multilamellar/unilamellar liposomes, micelles, and other membranes for use in physiological environments. Lipid systems have the ability to encapsulate hydrophobic and hydrophilic drugs, maintain longer drug circulation times, reduce systemic toxicity, and allow for targeting via ligand conjugation. Loading depends on the number of lipid layers and size of the carrier. Multiple layers present the opportunity to load multiple drugs, create different release profiles, or provide extra layers of protection. Most clinically used liposomes are 50–300 nm in size. In addition to allowing different sized drugs or proteins to be loaded into the system, the size range reduces uptake in the liver, which provides the liposomes a better chance of remaining in circulation. Delivered via parenteral injections, liposomal systems can actively target, via ligand conjugation, or passively target, via aggregation, their desired tissues, thus lowering cytotoxic or excess drug concentrations in healthy tissues. Release occurs by aggregation and destabilization of the lipid membrane. This method of release is energetically favorable compared with fusing with the cells. Environmental factors, such as pH and temperature, affect stability of the lipid membrane, thereby causing issues of premature release and short half-life. While encapsulation of hydrophilic drugs can be accomplished, it remains difficult and is often associated with poor encapsulation efficiency.

Surfactants Surfactants are a subgroup of lipid-like materials that have the ability to lower the surface tension at an interface. At low concentrations, surfactant molecules remained separate, but once past a critical concentration, they begin to aggregate into small particles called micelles. Depending upon the hydrophile to lipophile balance, different colloidal systems can be formed. Two such systems are O/W micelles and W/O reverse micelles. Micelles and reverse micelles are typically one of the smaller delivery systems, having particle sizes ranging from 10 to 100 nm and 1 to 10 nm, respectively. However, encapsulation is not the only use of surfactants in drug delivery. Facilitation of drug transport across the respiratory wall and enhanced skin penetration by target molecules depends on the formal charge of the surfactant head group. The charge on this head group can interact with the membrane and alters its structure to hydrate, expand, or electrically change the membrane, resulting in drug transport. In recent years, 90% of new drugs have been classified as poorly soluble in water. Surfactants enhance the permeation of drug delivery systems across physiological barriers, leading to rapid and efficient delivery of proteins and insoluble drugs to their targeted tissues. One adverse effect is the poor thermodynamic stability of low-molecular-weight surfactants in solution, which causes premature dissociation of the surfactant membrane prior to reaching the targeted tissues. While this system has flaws, delivery via surfactants allows for increased bioavailability, distribution, and permeabilization while also reducing adverse reactions.

Ceramics Ceramics are inorganic, nonmetallic materials formed by ionic and covalent bonding of metallic and nonmetallic elements. Their microstructure varies from crystalline (ceramics) to amorphous (glasses). In biomedical applications, the materials are often used to replace diseased, deteriorating, or damaged bone, with some also being used in soft tissues. Being nonimmunogenic and compatible with most incorporated therapeutic agents, these materials are suitable candidates for drug delivery systems.

Hydroxyapatite systems Hydroxyapatite (HA) is a pure, synthetic version of bone mineral. It is osteoconductive and biocompatible, and depending on processing, the material can possess similar compressive strength and nanostructure as bone. Processing to generate high surface areato-volume ratio allows for sorption of therapeutic agents onto and into the matrix. Manipulation of the porosity and pore structure allows for controlled loading and release kinetics. Commercial products such as nanoXIM release therapeutic agents from HA scaffolds and microspheres by diffusion. HA has an associated initial burst followed by a prolonged release. The initial burst comes from differences in concentration between the scaffold and interstitial fluid. With decreased drug concentration, the gradient between the scaffold and interstitial fluid drops, slowing release and maintaining a release until most of the drug has been delivered. One consideration in choosing this system is the rate of dissolution. HA systems are designed to remain in the body for an extended period of time. As such, HA would not be suitable for an application that needs relatively quick degradation.

Biphasic calcium phosphate systems Biphasic calcium phosphates (BCPs) are a mixture of HA and b-tricalcium phosphate (b-TCP). The ratio of HA/TCP plays a key role in the bioactivity and dissolution rate of the material. This combination of stable HA and resorbable b-TCP allows for controllable

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resorption and gives BCP its mechanical properties. Like most ceramics, BCP exhibits an inherent brittleness and inability to mechanically support load-bearing areas. Attempts have been made to create composite materials that increase the ability of BCPs to support load-bearing areas. Drug loading and release from BCPs depend upon the percent porosity, morphology, and purity of the material. The release kinetics of BCP are similar to those of HA, with an initial burst followed by a slowed rate of release. A tailorable pore structure and bioactive properties make BCP a versatile ceramic and a suitable option for antibiotics, hormones, insulin, cancer drugs, etc.

Bioglasses Bioglass is a family of bioactive glasses composed of silicon dioxide, sodium oxide, calcium oxide, and phosphorous pentoxide. This material yields a surface pore structure, 5–20 nm, with excellent surface area, pore volume, cytocompatibility, and ability to induce apatite formation. As with HA and BCP, clinical applications of bioglasses have focused on osteogenesis. From this came an increase in research on the loading of therapeutic agents to help enhance the regeneration of the surrounding tissues. Loading depends upon the crystallinity, porosity, and morphology of the implant, whereas release is dictated by the rate of degradation, typically ranging from a few hours to several weeks. This variability grants control over how quickly the loaded agents will release. Bioglass degrades via hydrolysis of the silica matrix to form Si(OH)4 and silanol. These products first cause surface pH levels to rise, leading to further hydrolytic degradation and nucleation of carbonated HA in the silica gel surface, which contributes to the development of new bone. Until recently, bioglass materials were not clinically used as drug delivery systems but as a material to aid in bone regeneration. However, as of recent, there has been success with therapeutically loaded bioglasses. In a clinical study, an enamel matrix protein derivative was loaded into bioglass and has resulted in the formation of new cementum with associated periodontal ligament and enhanced mineralization around the bioglass particles.

Metallic Nanoparticles Metallic delivery systems have potential for gene and drug delivery applications. These systems can vary from simple nanoparticles to complex nanoshells and nanocages. Nanoscale particles have the unique ability to more easily pass through biological barriers, for example, blood–brain barrier and cell membranes. Having release occur inside the cell promotes higher therapeutic efficacy than obtained via free drug absorption. Once the magnetic nanoparticles are injected near the targeted area, a high-powered magnetic field is turned on to magnetically direct the nanoparticles to the desired location. Nanoparticles release via diffusion, and direct conjugation of the drug and nanoparticle can be weak requiring the action of organic linkers such as amines, carboxylic acid, aldehyde, or thiol. These linkers allow for active targeting of cancer-specific markers that upon binding will allow release to occur. For passive release, aggregations of nanoparticles begin to attach or penetrate the targeted cells/tissues and release their payload via diffusion. In the early 2000s, clinical trials began to be initiated and have shown metallic nanoparticles to be successful carrier systems. Currently, a system of paclitaxel and gold nanoparticles (Aurmine (CytImmune Sciences, Inc.)) is being developed to increase the effectiveness of chemotherapy. Because the magnetic field is generated outside the body and its effective strength decreases with distance, targeting locations deep within the body is difficult. Another issue to consider is the toxic by-products left from the drug delivery system after release. Metal nanoparticles must be excreted from the body via alternative routes, that is, defecation and urination.

Future Materials and Systems Current drug delivery systems have made significant strides to improve issues such as targeting, loading, release, and circulation times. Future technologies will enable the creation of controlled release formulations that further improve upon these successes. However, there is a need not only to further study nanoparticles but also to look for new materials and new systems as means for drug delivery. Several technologies are being developed with ultimate goal of becoming commercialized.

Carbon Nanotubes Carbon nanotubes (CNTs) are structures comprising long graphene tubes with either single or multiple walls. CNTs can be up to several microns long and have diameters from 0.4–3 to 2–100 nm for single and multiwall structures, respectively. This system has drawn considerable attention due to its structural, electronic, and mechanical properties; thermal stability; high surface area; and size. One downfall of a larger surface area is the increased chance for protein opsonization. Drugs, proteins, and bioactive molecules are loaded onto or into CNTs via functionalization of the surface by methods of covalent and noncovalent bonding. Toxicity depends upon several factors, such as functionalization, purity, length, size, diameter, surface chemistry, and route of dispersion. Excess CNT accumulation, which can occur in different parts of the body such as the lung, heart, testes, and brain, leads to oxidative stress and cellular toxicity. Other researchers, however, suggest that once functionalized, such as with polycations, CNTs do not pose a significant threat to the body. Toxicity of CNTs remains a hotly debated topic.

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Microchips Implantable systems have been around for many years, but controlled release from these implants depends upon which material is used, how the system is processed, and the system mechanisms of release. Researchers have begun using implantable microchips to obtain precisely controlled release patterns rather than rely on standard mechanisms of release, that is, diffusion or degradation. The microchip contains drug-filled microreservoirs that are covered with a thin gold film. Once in place, release can be activated wirelessly by applying a voltage to rupture the reservoir membrane and release the loaded contents. Because this system is programmable, it allows control over the dose the patient receives and the time at which drug is released. Currently, this device is being clinically tested for delivering parathyroid hormone to osteoporosis patients.

3-D Scaffolds 3-D scaffolds are designed to be implanted, support the surrounding tissue while ingrowth and regeneration occur, and degrade over time. Different systems and materials have been developed for mechanical and physiological environments pertinent to specific applications, including the skin, cartilage, bone, and blood vessels. Ongoing development is attempting to provide controlled release of therapeutic agents from the scaffold to promote tissue regeneration or remodeling. In these systems, controlled release depends on modifying either the rate of degradation or the swelling behavior of the given scaffold to allow release to occur.

3D Printing Rapid prototyping (RP) techniques, including 3DP, were originally created to allow for quick fabrication of new devices during product development. In using computer-generated 3-D models, companies are able to quickly build, test, and analyze different models. A variety of RP methods have found several uses for biomedical applications. Germane to the present topic of drug delivery, three main techniques are being currently studied: laser-based systems, nozzlebased deposition systems, and printing-based inkjet systems. Selective laser sintering and stereolithography (SLA), two prominent strategies of laser-based systems, use a laser to either sinter or polymerize different materials. Fused deposition modeling, a nozzlebased system, extrudes hot thermoplastics layer by layer onto a loading platform in which the final product is formed. 3DP, an inkjet system, involves the use of particles and binding agent to form the product. Computer-aided design allows for a geometrically precise, layer-by-layer design of the delivery system. Both loading and release of therapeutic agents can be controlled by altering the porosity and structure of the dosage form. 3DP allows for specific drug loadings that were otherwise unobtainable by traditional methods. Each system has limitations, however. Laser-based systems are limited by material because it must be able to withstand the extreme temperatures and potential denaturation of the fabrication process. Another limitation, mainly with SLA, is issues of scarce biocompatible resin materials, cytotoxic photoinitiators, and entrapment of unreacted monomer and residual photoinitiator. Fused deposition modeling is limited by the type of thermoplastic, as the polymer must have good melt viscosity with low enough viscosity to be extruded from the nozzle but viscous enough to maintain the printed shape.

Conclusion Drug delivery systems are designed to achieve a desired therapeutic effect at the target tissues without systemic toxicity. These systems are engineered to load different types of hydrophilic or hydrophobic drugs and enhance transportation across physiological barriers, without any compromise in bioactivity and while maintaining controlled release of therapeutic agents. Controlled release improves patient compliance by reducing the invasive procedures and oral drug intake. Controlled release technologies have increased the quality and longevity of life for patients. Many ailments, such as cancer, diabetes, tobacco addiction, and tissue infections, have been ameliorated or cured through the use of drug delivery technologies. Currently, research and clinical trials are attempting to treat diseases once thought incurable with cutting-edge controlled release systems. Future formulation strategies and materials used in the development of new controlled release systems will be at the forefront of disease treatment.

Further Reading Agnihotri, J., Saraf, S., & Khale, A. (2011). Targeting: New potential carriers for targeted drug delivery system. International Journal of Pharmaceutical Sciences Review and Research, 8(2), 117–123. Bae, Y. H., & Park, K. (2011). Targeted drug delivery to tumors: Myths, reality and possibility. Journal of Controlled Release, 153(3), 198–205. De Jong, W. H., & Borm, P. J. A. (2008). Drug delivery and nanoparticles: Applications and hazards. International Journal of Nanomedicine, 3(2), 133–149. Eltorai, A. E. M., et al. (2016). Microchips in medicine: Current and future applications. BioMed Research International, 2016, 1743472. Fang, J., Nakamura, H., & Maeda, H. (2001). The EPR effect: Unique features of tumor blood vessels for drug delivery, factors involved, and limitations and augmentation of the effect. Advanced Drug Delivery Reviews, 63(3), 136–151.

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Gujral, S., & Khatri, S. (2013). A review on basic concept of drug targeting and drug carrier system. International Journal of Advances in Pharmacy, Biology and Chemistry, 2(1), 130–136. Heller, J. (1993). Polymers for controlled parenteral delivery of peptides and proteins. Advanced Drug Delivery Reviews, 10(2), 163–204. Huang, X., & Brazel, C. S. (2001). On the importance and mechanisms of burst release in matrix-controlled drug delivery systems. Journal of Controlled Release, 73(2–3), 121–136. Jain, S., Jain, V., & Mahajan, S. C. (2014). Lipid based vesicular drug delivery systems. Advances in Pharmaceutics, 2014, 574673. Lian, T., & Ho, R. J. (2001). Trends and developments in liposome drug delivery systems. Journal of Pharmaceutical Sciences, 90(6), 667–680. McBain, S. C., Yiu, H. H., & Dobson, J. (2008). Magnetic nanoparticles for gene and drug delivery. International Journal of Nanomedicine, 3(2), 169–180. Middleton, J. C., & Tipton, A. J. (2000). Synthetic biodegradable polymers as orthopedic devices. Biomaterials, 21(23), 2335–2346. Miller, R. A., Brady, J. M., & Cutright, D. E. (1977). Degradation rates of oral resorbable implants (polylactates and polyglycolates): Rate modification with changes in PLA/PGA copolymer ratios. Journal of Biomedical Materials Research, 11(5), 711–719. Rastogi, V., et al. (2014). Carbon nanotubes: An emerging drug carrier for targeting cancer cells. Journal of Drug Delivery, 2014, 670815. Schmaljohann, D. (2006). Thermo- and pH-responsive polymers in drug delivery. Advanced Drug Delivery Reviews, 58(15), 1655–1670. Tan, M. L., Choong, P. F. M., & Dass, C. R. (2010). Recent developments in liposomes, microparticles and nanoparticles for protein and peptide drug delivery. Peptides, 31(1), 184–193. Tan, M. L., Choong, P. F. M., & Dass, C. R. (2010). Recent developments in liposomes, microparticles and nanoparticles for protein and peptide drug delivery. Peptides (New York, NY, United States), 31(1), 184–193. Torchilin, V. P. (2001). Structure and design of polymeric surfactant-based drug delivery systems. Journal of Controlled Release, 73(2–3), 137–172. Torchilin, V. P. (2005). Recent advances with liposomes as pharmaceutical carriers. Nature Reviews Drug Discovery, 4(2), 145–160. Tsung, J., & Burgess, D. J. (2012). Biodegradable polymers in drug delivery systems. In J. Siepmann, R. A. Siegel, & M. J. Rathbone (Eds.), Fundamentals and applications of controlled release drug delivery (1st edn, pp. 107–123). New York: Springer US. Uhrich, K. E., Cannizzaro, S. M., Langer, R. S., et al. (1999). Polymeric systems for controlled drug release. Chemical Reviews (Washington, DC), 99(11), 3181–3198. Verron, E., et al. (2010). Calcium phosphate biomaterials as bone drug delivery systems: A review. Drug Discovery Today, 15(13), 547–552. Wu, C., & Chang, J. (2012). Mesoporous bioactive glasses: Structure characteristics, drug/growth factor delivery and bone regeneration application. Interface Focus, 2(3), 292–306.

Electrospinning and Electrospray for Biomedical Applications Min Wang, The University of Hong Kong, Pokfulam, Hong Kong Qilong Zhao, Shenzhen Institutes of Advanced Technology, Chinese Academy of Sciences, Shenzhen, China © 2019 Elsevier Inc. All rights reserved.

Introduction Biomedical Requirements for Fibrous or Particulate Devices Electrospinning History, Principle, and Apparatus Advanced Electrospinning Techniques Materials for Electrospinning Influencing Factors Tissue Engineering Scaffolds and Applications Fibrous Drug Delivery Systems and Applications Biosensors Other Biomedical Applications Electrospray Principle and Apparatus for Electrospray Electrosprayed Solid Microparticles and Applications Electrosprayed Core-Shell Structured Microparticles for Biomolecule Delivery Electrosprayed Microparticles for Cell Delivery Combining Electrospinning and Electrospray for Fabricating Novel Medical Devices Concluding Remarks Acknowledgments Further Reading

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Glossary Aptamer Oligonucleotide or peptide molecules binding to a specific target molecule. Antibiotic A type of drug used for killing or inhibiting the growth of bacteria. Collagen A main structural protein of extracellular matrix and connective tissues with a triple helix structure. Extracellular matrix Nanofibrous networks formed by extracellular molecules secreted by cells that provide structural and biochemical support to the surrounding cells. Gelatin An irreversibly hydrolyzed form of collagen. Glycosaminoglycan A series of long unbranched and highly polar polysaccharides consisting of a repeating disaccharide unit. Progenitor cell A type of cell enabling limited self-renewal and having the capability to differentiate into a specific type of cell. Stem cell A type of cell enabling self-renewal as also differentiation into specialized cell types.

Nomenclature AgNP Silver nanoparticle AuNP Gold nanoparticle bFGF Basic fibroblast growth factor BCP Biphasic calcium phosphate BMP Bone morphogenetic protein Ca–P Calcium phosphate DNA Deoxyribonucleic acid ECM Extracellular matrix EDTA Ethylene diamine tetraacetic acid EGF Epidermal growth factor FDA Food and drug administration GelMA Gelatin methacryloyl HA Hyaluronic acid

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IGF Insulin-like growth factor NGF Nerve growth factor PAN Polyacrylonitrile PCL Poly(ε-caprolactone) PDGF Platelet-derived growth factor PDMS Polydimethylsiloxane PEO Poly(ethylene oxide) PGA Poly(glycolic acid) PHBV Poly(hydroxybutyrate-co-hydroxyvalerate) PLA Poly(lactic acid) PLACL Poly(lactide-co-caprolactone) PLGA Poly(lactic-co-glycolic acid) PM2.5 Particulate matter
2.5 mm in size PS Polystyrene PU Polyurethane PVA Poly(vinyl alcohol) PVDF Poly(vinylidene fluoride) PVP Poly(vinyl pyrrolidone) RGD Arginine/glycine/aspartic acid SEM Scanning electron microscope SERS Surface-enhanced Raman scattering SF Silk fibrin TCP Tricalcium phosphate TEM Transmission electron microscope TGF Transforming growth factor VEGF Vascular endothelial growth factor

Introduction Electrospinning and electrospray are facile electrohydrodynamic (EHD) techniques governed by similar principles, which use electrostatic forces to overcome the surface tension of charged liquids. With the emergence of nanotechnology in the 1990s, electrospinning and electrospray have become popular research topics in the academic world as well as in the R&D of fibrous and particulate products for various applications. Electrospinning enables high-efficiency fabrication of ultrathin fibers (from nanofibers with diameters of tens of nanometers to microfibers with diameters of hundreds of micrometers), while electrospray provides continuous production of small-diameter particles (from nanoparticles with diameters of tens of nanometers to microparticles with diameters of hundreds of micrometers). In a typical electrospinning or electrospray process, an electrostatic force is applied to overcome the surface tension of a charged liquid which normally comes out of a syringe with a metallic needle connected to a high-voltage power supply. The electrostatic force stretches or breaks up the charged liquid to be viscoelastic filaments or jets, resulting finally in dry fibrous or particulate products after the evaporation of solvent in the liquid (if the liquid is a polymer solution) during the journey towards a grounded fiber or particle collector. A broad range of materials can be processed by electrospinning or electrospray. The morphology and structure (diameter, surface morphology, interior structure, etc.) and properties of electrospun or electrosprayed products can be effectively controlled, which makes electrospinning and electrospray very promising for a variety of applications, ranging from the energy field to biomedical and healthcare applications. Owing to the ease of conducting electrospinning or electrospray and also the unique characteristics of electrospun or electrosprayed products, electrospinning and electrospray techniques are widely investigated and used in the biomedical field. Electrospun or electrosprayed products can be used for tissue engineering, drug delivery, wound dressing, biosensor, filtration, mask, etc. (Fig. 1). Continuous investigations into electrospinning and electrospray will certainly assist the progress of biomedical research and generate improved and advanced biomedical devices for the rapidly growing healthcare market.

Biomedical Requirements for Fibrous or Particulate Devices Biomedical devices need to possess appropriate properties and provide required functions for targeted applications. Fibrous and particulate materials fabricated by electrospinning and electrospray have their respective advantages as biomedical devices. For

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Biomedical applications of fibrous and particulate products fabricated via electrospinning and electrospray, respectively.

tissue engineering applications, porous scaffolds resembling the structure and features of native ECM are advantageous for cell attachment and proliferation. In spite of variations in composition and structure, ECMs are normally composed of intermeshed nanofibers of proteins (collagen, elastin, etc.) and glycosaminoglycans with different fiber assemblies, offering structural and biochemical support for cells of the tissue. Consequently, producing mechanically stable artificial scaffolds with nanofibrous and porous structures is an important goal in the tissue engineering field. Although some other methods such as molecular selfassembly and phase separation are studied for making nanofibrous tissue engineering scaffolds, many drawbacks, particularly the low efficiency in scaffold manufacture, prevent them from being widely adopted in the tissue engineering field. As a simple, versatile and effective technique for nanofiber production, electrospinning has shown its superiority in making nanofibrous and porous tissue engineering scaffolds. Compared to the molecular self-assembly and phase separation techniques, electrospinning is capable of processing a broader range of materials to nanofibers and requires only simple control or device for fiber alignment. Therefore, a variety of nanofibrous scaffolds with different compositions, morphologies and architectures are fabricated via electrospinning for different tissue engineering applications. Through electrospinning and electrospray, solidified fibrous and particulate materials with diameters down to nanometer range can be directly formed from liquid precursors through a simple one-step procedure. Furthermore, electrospinning and electrospray are amiable to incorporating bioactive agents in fibers and particles. These distinctive features make these two techniques highly attractive for encapsulating bioagents (drug, growth factor, gene, etc.) for their controlled delivery. Good drug delivery vehicles should have excellent capability for drug loading, provide controlling drug release and protect the bioactivity of delicate bioagent (e.g., growth factor). Electrospinning and electrospray are capable of processing both water-phased solutions and oil-phased solutions, resulting in highly efficient encapsulation of hydrophilic bioagents and hydrophobic bioagents, respectively. And the ease to control the composition, morphology, and structure of electrospun and electrosprayed products through using different electrospinning and electrospray techniques makes it relatively easy to control the release of bioagents in specific spatiotemporal manners. With appropriate modification, electrospinning and electrospray can achieve efficient biomolecular and even live mammalian cell encapsulation in fibers and particles for their subsequent controlled release. Fibrous and particulate biomaterials can also find other biomedical applications such as biosensor, filtration, and personal healthcare products. Nanofibers and nanoparticles, due to their high surface-area-to-volume ratios and other useful properties, are excellent for high-efficiency adsorption and enrichment of various species such as biomolecules to be detected, thereby possessing high potential as biosensors with improved detection sensitivity and wide detection range. Electrospun nanofibrous meshes can also be excellent materials for removing unwanted biomolecules as well as pollutants with good air and water permeation rates, exhibiting their potential as high-efficiency, cost-effective filtration membranes and personal protective masks/surgical masks with light weight, high flexibility, and good mechanical stability.

Electrospinning History, Principle, and Apparatus An early investigation conducted by Lord Rayleigh in 1882 revealed the amount of charges required for overcoming the surface tension of a liquid drop. Inspired by this study, the devices which used electrostatic stretch to spray liquids into small droplets were invented by Cooley and Morton in 1902–1903. They were the early explorations of the electrospray technique. Formhals in 1934 applied for the first patent, demonstrating the process and apparatus of electrospinning for producing artificial threads.

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However, electrospinning remained as a neglected area for a long time until the early 1990s when research on electrospinning produced different polymer fibers with diameters down to nanometers, which revitalized this old technique. Since late 1990s, studies on the manufacture of nanofibrous biomaterials by electrospinning for potential biomedical applications increased rapidly. Over the past 10 years, the number of publications on electrospinning has been growing exponentially, reaching > 3000 in 2016, with the majority of these publications reporting different biomedical applications of electrospun materials. Electrospinning shows convenience, simplicity, and versatility for making nanofibers and nanofibrous products (non-woven, aligned, stacked, etc.) through the use of a variety of materials. The simple principle and apparatus for electrospinning make it easy to perform and control, offering the possibility to fabricate nanofibrous materials with high efficiency as well as versatility to meet customers’ requirements. In a typical electrospinning process (Fig. 2), a high-voltage power supply, a syringe filled with a precursor solution (normally a polymer solution) and equipped with a metallic needle, a syringe pump, and a ground collector are used, where the metallic syringe needle is connected to the high-voltage power supply for generating an electrostatic field. Droplets of the precursor solution flowing out the needle tip are firstly charged and then stabilized by the force between the applied electrostatic force and the surface tension of the solution. The excess of the electrostatic force over the surface tension stretches and deforms the charged droplets at the needle tip, forming a stable cone-shaped structure, namely, “Taylor cone”. A single and rapidly whipping viscoelastic jet with decreased diameter and elongated length is generated from the tip of Taylor cone, which then splits into many small jets in accordance with the charge repulsion effects after traveling a small distance. The small jets are further stretched in the electrostatic field, finally resulting in dry fibrous products after the evaporation of the solvent (which has been used to make the solution for electrospinning) in the trajectory of jets towards the collector. Non-woven mats can be formed through such a simple electrospinning process. The ease manufacture of nanofibers offers electrospinning a bright perspective for a broad range of biomedical applications such as tissue engineering, drug delivery, biosensor, wound dressing, etc. Electrospun aligned fibers and other structures can be obtained using different (and specially designed) fiber collection devices.

Advanced Electrospinning Techniques In addition to the ability and ease to form nanofibers, other significant advantages of electrospinning include high adaptability and ease for modification. After some specific modification(s), advanced electrospinning techniques have been developed, which are capable of making nanofibrous products to meet specific requirements of different biomedical applications. Conventional electrospinning using a polymer solution normally results in solid (i.e., non-porous) nanofibers with smooth surface or nanoporous surface (which is caused by the rapid evaporation of solvent in the polymer solution). While a stable water-in-oil emulsion is electrospun, nanofibers with core-shell structures are fabricated. Electrospinning using an emulsion is termed emulsion electrospinning. For biomedical applications, nanofibers made via emulsion electrospinning usually have aqueous cores, which are suitable reservoirs for the encapsulation and protection of bioactive molecules. Therefore, emulsion electrospinning has been intensively studied for making nanofibrous delivery vehicles for the controlled release of drugs, growth factors, genes, etc. Another electrospinning technique for producing core-shell structured nanofibers is coaxial electrospinning, in which a coaxial spinneret is employed with the inner capillary being fed with an aqueous solution (containing bioactive agent) and the outer capillary being fed with a polymer solution. During emulsion electrospinning, two immiscible solutions are spun out of the simple spinneret (commonly a metallic needle) and are both stretched by electrostatic force. The solution with a relatively high conductivity (usually the inner aqueous solution) is the driving liquid to transfer the tangential electrostatic force through

Fig. 2

A schematic diagram showing the experimental setup for electrospinning.

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viscosity on the liquid interface and to deform the merged droplets to become small jets with aqueous interior, finally resulting in core-shell structured nanofibers upon sufficient stretch. Compared to emulsion electrospinning, coaxial electrospinning minimizes the contact between the “water phase” and “oil phase”, providing enhanced protection for the bioactive molecules to be delivered, whereas emulsion electrospinning displays jetting instability. With appropriate modification, modified coaxial electrospinning even enables the encapsulation and delivery of live mammalian cells, which is termed cell electrospinning. In one study of cell electrospinning, the outer capillary and the inner capillary of a coaxial spinneret are fed, respectively, with a CaCl2 aqueous solution and a mixture of an aqueous alginate sodium solution and an aqueous cell suspension. Cell-encapsulated hydrogel microfibers are made after the crosslinking of alginate by Ca2 þ ions. As-spun cell-encapsulated hydrogel microfibers exhibit tunable fiber diameter and good cell viability. Fiber arrangement/pattern is a key topographical feature of electrospun scaffolds, which can be made/created using specially designed devices or following physics laws/principles. Scaffolds made by conventional electrospinning normally show randomly arranged nanofibers. Aligned nanofibers can be produced by using either a special fiber collector (e.g., a high-speed rotating drum collector or a parallel electrode collector) or a directional magnetic field (which requires adding magnetic nanoparticles into the electrospinning solution). Detailed studies on the effects of processing parameters of electrospinning on fiber alignment when using a high-speed rotating collector indicate the major effects of rotating speed and polymer solution conductivity on the alignment of nanofibers (Fig. 3). Electrospun scaffolds with aligned nanofibers resemble the ECM of some human tissues such as cardiac tissue and ligament/tendon, which provide contact guidance on the cells grown on the scaffolds and give topographical cues for directing the elongation and cytoskeleton development of cells. Electrospinning employing a point electrode and a ring electrode as the collector can produce scaffolds with radically aligned nanofibers, which are potentially useful for wound closure as the migration of skin cells can be promoted by these scaffolds. Electrospun nanofibrous scaffolds normally have small interconnected pores, which is a shortcoming with regard to cell infiltration and vascularization. It has been determined that a scaffold with an average pore size of at least 20 mm will enable cell infiltration and the diameters of fibers need to be larger than 4 mm. However, the increase of fiber diameter normally results in unsatisfactory cell attachment since the ECM-like nanofibrous architecture is changed. Electrospinning techniques for increasing the pore size and porosity of electrospun nanofibrous scaffolds are thus investigated. Methods such as particulate-leaching-associated electrospinning, wet electrospinning (i.e., collecting electrospun fibers in a liquid bath, followed by a post-electrospinning freeze-drying treatment), cryogenic electrospinning (i.e., using a freezing collector for fiber collection), and selective removal of sacrificial electrospun fibers or electrosprayed microparticles are developed. For example, using the removal of sacrificial fiber approach, removing water-soluble gelatin fibers from gelatin and PLGA bicomponent scaffolds fabricated by dual-source dualpower electrospinning results in scaffolds with significantly enlarged micropores and well-preserved nanofibrous architecture (Fig. 4). However, the increase of pore size and porosity by using these electrospinning techniques sometimes causes new problems such as decreased structural stability and low mechanical properties of electrospun scaffolds. Inspired by additive manufacturing, an advanced electrospinning technique that can produce fibrous scaffolds with accurately deposited fibers and designed porous architecture in a programed manner is established, which is termed melt electrospinning. It employs molten polymer for electrospinning and an automated stage as the collector. Melt electrospinning makes 3D scaffolds by the programed stacking of fibers, which addresses the problems in conventional electrospinning such as the difficulty to form thick scaffolds with large interconnected pores and well-defined fiber pattern. But the range of materials suitable for melt electrospinning is limited and bioactive species such as cells and bioactive molecules that are susceptible to high temperature damages cannot be incorporated into scaffolds during melt electrospinning.

Materials for Electrospinning The range of materials suitable for electrospinning is wide, including synthetic polymers, natural polymers, ceramics, and composites. Among them, polymers are the most frequently used materials for electrospinning. Water-insoluble synthetic polymers such as

Fig. 3

An SEM image showing the morphology of electrospun aligned PHBV nanofibers formed on a high-speed rotating cylinder.

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Fig. 4 Morphology of nanofibrous scaffolds fabricated by dual-source dual-power electrospinning: (A) bicomponent scaffold consisting of PLGA fibers and gelatin fibers (1:1) and (B) PLGA scaffold with enlarged pore size after post-electrospinning water-immersion treatment.

PS, PAN, PVDF, and PU can be electrospun into fibers for different applications. Water-soluble synthetic polymers such as PVA, PVP, and PEO can also be made into nanofibers via electrospinning, which can be nanofibrous carriers for bioactive agents or used as template or sacrificial fibers to form fibrous products with complex structures. Biodegradable synthetic polymers, such as PCL, PGA, PLA, and their copolymers (e.g., PLGA and PLACL), have different advantages in the biomedical field. Electrospun tissue engineering scaffolds using these biodegradable polymers resemble the nanofibrous feature of native ECM and degrade through hydrolysis into nontoxic products those can be subsequently cleared by metabolic pathways with no or very limited side-effect on the human body. However, the quick degradation of some of these polymers leads to local acidic microenvironment and provokes inflammation. Scaffolds fabricated through the electrospinning of natural polymers such as chitosan, collagen, gelatin, and hyaluronic acid (HA) show improved biocompatibility for tissue engineering applications whereas some problems such as poor mechanical properties exists. Unlike most of natural polymers, SF exhibits high mechanical strength, which can be directly used for some applications requiring high mechanical loading or stability such as bone tissue engineering. However, making solutions by dissolving of SF for electrospinning is a challenge. Some modified natural polymers such as norbornene- or methacrylate-modified hyaluronic acid and GelMA can be electrospun into nanofibrous hydrogel scaffolds with a post-electrospinning photo-crosslinking treatment. Unlike normal electrospun hyaluronic acid or gelatin scaffolds with chemical crosslinking where the nanofibrous architecture of scaffolds is affected by chemical crosslinking, photo-crosslinked electrospun hyaluronic acid or GelMA hydrogel scaffolds show well-preserved nanofibrous structure with good swelling behavior. Ceramic fibers with diameters ranging from hundreds of nanometers to tens of micrometers can be fabricated via electrospinning. For making ceramic fibers, either a ceramic particles-dispersed polymer solution with an appropriate ceramic particle concentration and solution rheology or a suitable precursor enabling in situ chemical synthesis of ceramic by sol–gel, coprecipitation or other methods is needed for electrospinning. When using a ceramic particles-dispersed polymer solution for electrospinning, a postelectrospinning annealing treatment is required for the removal of polymer component and for the sintering of ceramic, resulting finally in ceramic fibers. To date, a variety of fibrous ceramic products (BaTiO3, Fe3O4, hydroxyapatite, etc.) have been fabricated by electrospinning for their potential applications in the energy, sensor, and healthcare fields. Electrospinning can also produce composite fibers, including metal-polymer composites and ceramic–polymer composites, when a metallic particle-dispersed polymer solution or a ceramic particles-dispersed polymer solution is used. For example, fibrous silver nanoparticle-incorporated polymer scaffolds are electrospun, which have anti-bacterial property and are intended for tissue engineering and wound dressing applications. Biodegradable scaffolds consisting of ceramic-polymer composite fibers that contain bioceramic nanoparticles such as hydroxyapatite, calcium phosphate, etc. are made via blend electrospinning. They exhibit excellent bioactivity and good mechanical properties for bone tissue engineering.

Influencing Factors There are many factors influencing electrospinning and electrospun products, which can be categorized into three areas: polymer solution properties, processing parameters, and environmental factors. Polymer solution properties include polymer molecular weight, polymer solution concentration, viscosity of polymer solution, solution conductivity, solvent property, etc. The effects of different polymer solution properties are sometimes difficult to be isolated and studied because some of these parameters are highly related. For example, changing the conductivity of a polymer solution may also change its viscosity. Processing parameters in electrospinning include solution feeding rate, applied voltage and field strength, nozzle configuration, etc. Environmental factors refer to temperature, humidity, air velocity, etc. All of these factors have effects, small or large, on the electrospinning process and consequently determine the morphology and structure of electrospun fibers. For most polymers, a polymer with different average molecular weights that are within a reasonable range, as-spun fibers fabricated from the polymer with a higher average molecular weight generally tend to form less beads on fibers and have a more uniform fibrous structure. For example, the electrospun scaffolds made from chitosan with an average molecular weight of 30 kDa show many beads, while the use of chitosan with an average molecular weight of 398 kDa results in uniform fibers. But not all polymers

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follow this “law” on fiber morphology. For example, PEO fibers electrospun from PEO with different molecular weights exhibit little difference between them. Solution viscosity (also polymer solution concentration, which is highly related) is a dominant polymer solution property affecting the electrospinning process. In general, a very low viscosity of the polymer solution (which also means very low polymer solution concentration) leads to a very high surface tension of the polymer solution, which cannot be overcome by electrostatic force and hence no solidified fibrous products can be collected. With a gradual increase in viscosity of the polymer solution (i.e., increasing the polymer solution concentration and decreasing the surface tension of the polymer solution), polymer particles or beads start to form. When the viscosity of polymer solution further increases to be above a critical point, a stable electrospinning process is established, resulting in the production of fibrous products. For polymer solutions with a very high viscosity (i.e., at a high polymer solution concentration), droplets of the polymer solution at the tip of the syringe needle are easy to dry out before jets are initiated and hence electrospinning cannot be performed. A stable electrospinning process requires viscosity (and polymer solution concentration) to be in an appropriate range. Within this range, the increase in viscosity (and polymer solution concentration) usually results in the increase of diameters of as-spun fibers. Conductivity is another key polymer solution property influencing the electrospinning process. Usually, the increase of polymer solution conductivity or charge density leads to relatively high mobility of ions and increasing whipping instability of polymer jets during electrospinning. Strong and homogeneous electrostatic stretch tends to occur and thin jets tend to form, thereby resulting in uniform fibers with relatively small diameters. For example, the addition of a volatile salt, pyridium formiate, into a PLA solution can significant increase the conductivity of the polymer solution and subsequently minimize the formation of beads on fibers. Likewise, the use of solvents with different polarities, for example, alcohol and tetrachloromethane, can also modulate the conductivity of a polymer solution, resulting in fibrous product with a uniform structure and that with concomitant beads, respectively. Solvent properties such as the boiling point of the solvent determine the evaporation rate of solvents during electrospinning and can affect the morphology of electrospun fibers, resulting in either smooth surface fibers or nanoporous surface fibers. Applied voltage and electrical field strength are similar parameters, which indicate the strength of electrostatic force, are also major processing parameters affecting the electrospinning process. During electrospinning, the applied voltage (field strength) has significant effects on tuning the shape of Taylor cone. The increase in applied voltage (field strength) results in relatively small Taylor cone and relatively high surface tension, which may increase the difficulty to initialize jetting and the frequency of bead formation. However, the increase in applied voltage (field strength) enhances electrostatic stretch at the same time. Consequently, the influence of the field strength on the diameter of electrospun fiber is not predictable. For polymers such as PLA, PLGA, and PVA, high applied voltage (field strength) results in large diameter electrospun fibers, whereas the trend for other polymers such as cellulose acetate is the opposite. As for electrospray, some electrosprayed microparticles (e.g., alginate microparticles) with the smallest diameter can be formed in a medium range of applied voltage (field strength), while an increase and decrease in applied voltage (field strength) out of this range causes an increase in diameter of as-sprayed microparticles. The polymer solution feeding rate is another processing parameter that can affect electrospinning even though the influence may not be large. Usually, a relatively low feeding rate results in the ease of solvent evaporation, thereby leading to small diameter electrospun fibers. In electrospray, the diameter of as-sprayed microsphere and polymer solution feeding rate generally follows the same relationship. The influences of environmental factors, including temperature and humidity, on electrospinning should not be ignored. Since the increase in temperature may slightly decrease the viscosity of a polymer solution, relatively thin fibers are sometimes electrospun. For electrospinning performed at a sufficiently low temperature (e.g., on ice or into liquid nitrogen), which is termed cryogenic electrospinning, the evaporation of solvent can be significantly delayed, which induces solid-liquid phase separation and hence the generation of tunable porous surface morphologies of as-spun fibers. The increase in humidity has similar effects on the solvent evaporation and pore formation on the fiber surface. It also increases the difficulty to generate desirable uniform fibers.

Tissue Engineering Scaffolds and Applications As a simple, effective, and versatile method to form ECM-like nanofibrous scaffolds, electrospinning has been extensively investigated for making many types of tissue engineering scaffolds since late 1990s. These nanofibrous scaffolds have shown great potential and also usefulness for assisting the regeneration and reconstruction of different types of human body tissues and organs, ranging from tissues such as bone, skin, and blood vessels to organs such as liver. Early research revealed that electrospun nanofibrous scaffolds of biodegradable polymers could closely mimic the hierarchical architecture of native ECM and facilitate good attachment and proliferation of cells. Later, the potential of highly biocompatible electrospun scaffolds for engineering many types of human body tissues with relatively simple structures such as bone, skin, and blood vessels was shown and is now generally realized. With the assistance of electrospun nanofibrous scaffolds, the repair or regeneration of some avascular tissues such as cartilage and tendon/ligament which have limited self-healing or self-renewal capability becomes reality. In both in vitro and in vivo situations, electrospun tissue engineering scaffolds as artificial extracellular microenvironments can offer specific structural, mechanical, and biochemical cues for anchoring cells. For example, electrospun scaffolds with specific topography have significant influences on behaviors of attached cells, where cells response to the fiber orientation, cell morphogenesis, and cell migration are directed. Electrospun scaffolds of aligned nanofibers can induce the anisotropic elongation, spreading, and arrangement of cells along to the fiber direction, showing their benefits for nerve, tendon/ligament, and smooth muscle tissue engineering. Fiber diameter is also a key structural factor affecting cell behaviors on electrospun scaffolds. With decrease in fiber diameter, electrospun scaffolds normally exhibit increasing surface-area-to-volume ratio, showing significant advantages in 3D cell attachment, whereas cell attachment on microfibrous scaffolds is sometimes on single fibers that is more in a 2D manner, which

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is different from the native cell-ECM organization. For engineering some tissues requiring the load-bearing ability such as bone, cartilage, and tendon/ligament, suitable mechanical properties, particularly mechanical strength, are required for electrospun tissue engineering scaffolds. The composite approach, with the incorporation of reinforcing component such as ceramic particles having relatively high mechanical strength, is also used for fabricating mechanically reinforced tissue engineering scaffolds. The control over fiber alignment is also effective for making electrospun scaffolds with anisotropic mechanical properties. The crucial point for the mechanical properties of electrospun scaffolds is their compatibility with the target/host tissues. For some cells from soft tissues, electrospun scaffolds with relatively low mechanical stiffness result in favorable cell attachment and cell spreading, whereas scaffolds consisting of fibers with high rigidity lead to relatively poor cell response. Normally, topographical cues and mechanical properties of electrospun scaffolds work together to contribute to the major impact on cell cytoskeleton development, where cells can sense the physical and mechanical signals and respond by changing their adhesion and morphogenesis, subsequently exhibiting specific phenotype and differentiation potential. Biochemical cues, such as surface chemistry, incorporation and delivery of bioactive agents, etc., are also critical factors influencing the performance of electrospun tissue engineering scaffolds. Electrospun scaffolds with specific surface functionalization can enhance cell attachment by ways of hydrophobic interaction (e.g., having specific chemical group terminals that promote the interaction with cell membrane), avidin–biotin covalent coupling, affinity bonding (i.e., functionalized by specific ligands such as antibody or aptamer that can recognize their targets on cell membrane), etc. For example, a commonly employed method for enhancing cell attachment is to functionalize electrospun scaffolds with cyclic RGD sequences or to incorporate components containing RGD motifs (e.g., collagen and gelatin) since RGD can be recognized by almost one-half of cell integrins to readily form stable RGD-integrin ligament-receptor pairs. Biochemical properties of electrospun scaffolds also have influences on cell phenotype and differentiation, particularly for progenitor cells and stem cells. For example, the incorporation of bioactive bioceramics (e.g., Ca–P including hydroxyapatite, TCP, and BCP) in scaffold fibers is effective for enhancing the osteogenic properties of electrospun scaffolds in directing mesenchymal stem cell functions for bone tissue engineering. Another direct but highly effective way for improving the biological activity of electrospun scaffolds is to incorporate specific growth factor(s) in electrospun fibers via emulsion electrospinning or coaxial electrospinning, where growth factor molecules are encapsulated in the aqueous core of core-shell structured fibers with well-preserved bioactivity for their later in vitro or in vivo release in a specific spatiotemporal manner. Controlled local delivery of specific growth factor(s) (Table 1) has provided significant benefits for a range of tissue engineering applications (e.g., bone, skin, blood vessel, cartilage, nerve, and tendon/ligament). For mesenchymal stem cells, similar electrospun nanofibrous scaffolds encapsulated with different types of growth factors (e.g., BMPs and NGF) are able to stimulate cell differentiation towards different cell lineages (e.g., osteogenic cells by BMP and neuronal cells by NGF, respectively). In many cases, the delivery of multiple growth factors is required. For example, in blood vessel tissue engineering, the sequential delivery of VEGF and PDGF from bilayer electrospun scaffolds consisting of different types of biodegradable polymers with different degradation rates promotes the regeneration and remodeling of endothelial layer and smooth muscle layers of blood vessel, respectively. Currently, it is still a great challenge to construct complex electrospun scaffolds for the regeneration of complex human body tissues/organs such as liver, kidney, gastrointestinal tract, cardiac tissue, etc. Good vascularization in newly formed tissue by regenerative medicine is essential for the survival of cells and successful tissue remodeling of engineered complex tissues. A promising method of concurrent emulsion electrospinning and coaxial cell electrospray has been developed recently. Endothelial cells can be placed inside the nanofibrous matrix of scaffolds that encapsulate VEGF and provide its controlled and sustained release, which shows the potential to build 3D vascularized cell-laden constructs for engineering complex human body tissues.

Fibrous Drug Delivery Systems and Applications Electrospun multifunctional nanofibrous scaffolds providing controlled release of growth factors for tissue engineering applications have been presented in the last section. Electrospinning is indeed an excellent technique for producing fibrous delivery systems for a variety of cargos, ranging from small molecule drugs, biomacromolecules, genes, and even live mammalian cells. Different strategies, including adsorption and encapsulation, can be used to load different types of cargos into/onto electrospun fibers. Table 1

Growth factors for use in tissue engineering

Growth factor

Major functions

Major applications

bFGF BMP-2 BMP-7 EGF IGF NGF PDGF TGF-b

Promoting proliferation of many cell types Stimulating chondrogenic and osteogenic differentiation of MSCs Stimulating bone and cartilage maturation Promoting proliferation of mesenchymal, glial, and epithelial cells Promoting proliferation of many cell types Promoting neurite outgrowth and neural cell survival Promoting proliferation of connective tissue, glial, and smooth muscle cells Anti-inflammatory, promoting wound healing, and inhibiting macrophage and lymphocyte proliferation Regulating endothelial cell proliferation, angiogenesis, and vascular permeability

Skin, musculoskeletal system Bone, cartilage Bone, cartilage Skin Cartilage, tendon/ligament Nerve Musculoskeletal system Cartilage, skin

VEGF

Blood vessel

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Owing to their high surface-area-to-volume ratio, electrospun nanofibers are naturally suitable for loading bioagents via adsorption. Bioactive agents such as small molecule drugs, biomolecules, and inorganic nanoparticles can be loaded onto electrospun fibers through different adsorption strategies, including physical adsorption and chemical bonding. Chemical bonding usually causes permanent immobilization of bioagents onto electrospun fibers, resulting in difficulties to release bioagents and sometimes in the damages or dysfunctions of immobilized biomolecules arising from their conformational changes. Physical adsorption through hydrogen bonding, hydrophobic/hydrophilic interactions or electrostatic interactions are popular for loading bioagents, particularly biomolecules as their bioactivity is not obviously or significantly affected. However, using surface adsorption to load bioagents has inherent problems in the amount of loaded bioagents (as compared to the encapsulation approach) and release behavior of bioagents (again, as compared to the encapsulation approach), thereby limiting their scope of applications. Blend electrospinning is a straightforward way to encapsulate bioagents in electrospun fibers as bioagents are directly suspended in the polymer solution to form a blend solution for electrospinning and subsequently encapsulated in resultant fibers. Blend electrospinning has been popularly for fabricating fibrous delivery vehicles for small molecule drugs (antibiotics, anti-inflammation drugs, anti-cancer drugs, etc.). Inorganic nanoparticles can also be encapsulated in electrospun fibers through blend electrospinning. For example, as shown earlier in this chapter, bone tissue engineering scaffolds of composite fibers containing homogeneously dispersed bioceramic nanoparticles can be fabricated by blend electrospinning, which provide sustained release of Ca2 þ ions and possess desirable osteogenic properties for promoting bone regeneration. AgNP-encapsulated fibrous scaffolds can also be made via simple blend electrospinning, which possess anti-bacterial property with the release of silver ions and hence can be used for wound treatment. Blend electrospinning requires the use of a polymer solution in which bioagents are dispersed. This significantly restricts its application for delivering delicate biomolecules such as growth factors, enzymes, and genes. In order to avoid the denaturation/degradation of biomolecules to be delivered, polymers requiring toxic organic solvents for making polymer solutions cannot be used to form fibrous delivery vehicles. While water-soluble polymers such as gelatin, PVA, and PEO can be used in blend electrospinning for forming fibers to deliver these biomolecules, specific post-electrospinning crosslinking treatment is necessary for these polymer fibers for subsequent handing and applications, which may cause damages too to encapsulated biomolecules. Another drawback of blend electrospinning is that encapsulated molecules tend to migrate to near the surface of fibers during electrospinning due to the charge repulsion effect, leading to initial burst release of bioagents in the in vitro and in vivo situations. Emulsion electrospinning and coaxial electrospinning are two frequently techniques for creating fibers for biomolecule delivery, employing a simple spinneret, and a coaxial spinneret, respectively, as both techniques can effectively minimize the contact between biomolecules and toxic organic solvent in polymer solutions with the help of core-shell structured fibers. Despite the small differences in the manufacturing procedures and also the difference in core-shell structures of resultant fibers, emulsion electrospinning and coaxial electrospinning produce biomolecule-encapsulated fibrous scaffolds which have similar biomolecule release profiles that are governed by the same release mechanisms. These fibrous delivery vehicles normally exhibit two-stage release profiles, that is, initially rapid release and subsequently sustained release, which are mainly controlled by biomolecule diffusion in the initial stage and polymer degradation in the late stage. The initial rapid release is undesirable and increases the risk of cancer. Since biomolecules (e.g., growth factors) usually carry a certain type of electrical charges (positive or negative) in the physiological environment, strategies such as incorporating biodegradable polyelectrolytes bearing a specific type of electrical charges or emulsion electrospinning using power suppliers with different polarities (positive applied voltage or negative allied voltage) have therefore been investigated for modulating the release of biomolecules through electrostatic interaction. These strategies can work well, effectively modulating growth factor release to be steady and sustained under specific conditions and therefore providing appropriate stimuli on cell functions (Fig. 5). Coaxial electrospinning, as compared to emulsion electrospinning, involves less diffusion between two separated liquids (i.e., the water phase and oil phase) during electrospinning and therefore may cause less disturbance/damage to biomolecules. It has been shown that coaxial electrospinning is feasible for producing non-viral fibrous delivery vehicles, providing successful gene (e.g., plasmid DNA) delivery. With good protection on bioactive substances to be delivered, coaxial electrospinning can be even be used for the encapsulation and delivery of live mammalian cells. This is demonstrated in a study where a sodium alginate solution containing suspended cells and a CaCl2 aqueous solution are fed respectively to the inner capillary and the outer capillary of a coaxial spinneret for electrospinning, resulting in cell-encapsulated hydrogel microfibers.

Biosensors The simplicity and versatility of electrospinning for fabricating nanofibrous meshes of a variety of materials make it attractive for preparing biosensors. For example, electrospinning of a viscous solution containing PVA, glucose, prehydrolyzed tetramethyl orthosilicate and horseradish peroxide enzymes is conducted for fabricating nonwoven meshes consisting of enzyme-encapsulated silica nanofibers. Nanofibers in the meshes have nanoporous surface that is caused by the glucose template, enabling sufficient contact between the encapsulated enzymes and hydrogen peroxide. These nanofibrous meshes, as potential biosensors, provide effective and reliable determination of hydrogen peroxide and simultaneously exhibit good mechanical flexibility, thermal stability, and reusability. Electrospun nanofibrous meshes are also investigated as non-enzymatic amperometric biosensor for hydrogen peroxide determination. In this study, polyurethane nanofibers filled with carbon nanotubes and AgNPs are fabricated via blend electrospinning, resulting in hybrid nanofibrous meshes with enhanced electrocatalytic activity on hydrogen peroxide, which is a desirable electrochemical sensor for hydrogen peroxide detection with an excellent sensitivity and a wide linear range. In another study, through

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Fig. 5 Emulsion electrospun PLGA scaffolds fabricated at different applied voltages (positive or negative) for controlling VEGF release: (A) electrical charge retention by different scaffolds, (B) in vitro release behavior of VEGF for scaffolds made at different applied voltages, (C) confocal microscopy of vascular endothelial cells cultured on two types of scaffolds, showing the effect of polarity of applied voltage on VEGF release and consequently cell behavior.

electrospinning of a PVP/SnCl2/AgNO3 solution which is followed by sintering, Ag/SnO2 composite nanotubes are formed. Meshes consisting of these Ag/SnO2 composite nanotubes exhibit sensitive amperometric response to hydrogen peroxide with a wide linear range and a low detection limit, making them promising electrochemical sensors for hydrogen peroxide. Electrospun fibrous membranes are also investigated as biosensors for immunoassay. For example, electrospun PCL nanofibrous membranes are able to effectively adsorb antibody proteins via hydrophobic interactions and then sufficiently capture target antigens for immunoassay. With the assistance of repeated fold-press procedures, immunoassay based on these free-standing electrospun membranes can obtain significantly amplified fluorescence signals, leading to reliable detection of human serum albumin with a good linear range and a low detection limit. Another example is an electrospun membrane of a special polymer with the anti-fouling property, which shows clear advantages for immunoassay. This nanofibrous structure allows sufficient binding and immobilization of antibodies and its anti-fouling property remarkably reduce the non-specific adsorption of proteins, which together contribute to the fast and reliable detection of target proteins (25% shorter detection time as compared to conventional immunoassay methods, and a wide linear detection range from) together with optimized signal-to-noise ratio. Electrospun membranes with specific functionalization can be developed as free-standing and flexible SERS substrates for molecular detection with extremely high sensitivity. For example, blend electrospinning of a PVA/AgNPs solution produces nanofibers with particular assembly and arrangement of AgNPs, which can provide reliable detection of reporter molecules at a very low dosage by SERS.

Other Biomedical Applications Electrospun fibrous structures are also investigated and developed for other applications, such as wound dressing and air/water filtration. There are already some commercial products on the market. Electrospun meshes have been extensively investigated as wound dressings since they can be effective barriers for inhibiting bacteria but allowing air and water permeation, thereby offering a favorable environment for wound healing. In order to have the anti-infection property for electrospun wound dressings, various strategies are explored. Polymers naturally having the anti-bacterial property such as chitosan and PVP are electrospun into wound dressings which are capable of inhibiting bacteria/fungal adhesion to some extent. However, to achieve active antibacterial property

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for electrospun wound dressings, anti-bacterial agents such as antibiotics and AgNPs are incorporated into fibers for their controlled released later. Between the use of antibiotics and AgNPs, electrospun wound dressings with incorporated AgNPs can achieve more durable anti-infection effect. Through electrospinning of an emulsion consisting of a silver ion-containing polymer solution and an aqueous PVP solution, fibrous AgNP-incorporated polymer scaffolds with evenly distributed AgNPs in fibers are formed by the postelectrospinning in situ reduction of silver ions by PVP (Fig. 6). Other considerations for wound dressings include rapid hemostasis and accelerating wound closure, which can also be fulfilled by electrospun functional wound dressings. For example, electrospinning of a medical adhesive, cyanoacrylate, results in wound dressings consisting of ultrathin cyanoacrylate nanofibers, which have the ability to immediately stop aortic bleeding within dozens of seconds. In other studies, wound dressings fabricated by emulsion electrospinning with sustained releases of growth factors facilitating specific cell migration and growth are effective in treating some wounds that are difficult to heal by ordinary treatments as in the case of chronic wounds of diabetic people. Electrospun nanofibrous membranes having relatively small interconnected pores are not good for cell infiltration but they are suitable candidates as filtration membranes for water/air purification. Electrospun membranes can be improved by incorporating active agents or be modified by other methods for enhancing the filtration of toxins/contaminants. For example, electrospun membranes consisting of PS nanofibers coated by a mesoporous silica layer are made for water purification purpose. A frequently employed method is to functionalize or to modify the nanofibers of electrospun filtration membranes with reactive groups such as oximes, cyclodextrins, and chloramines that are able to bind and detoxify hazardous agents. In recent years, air pollution, particularly the problems arising from PM2.5 air pollutants, is critical public health issue in China and also in some other developing countries. Electrospun polymer membranes (e.g., polyamide-6) with nano-sized pores have been proven effective as highefficiency air filters for PM2.5. Electrospun superhydrophilic polyacrylonitrile/silica nanofibrous membranes have a high moisture vapor transmission rate and are excellent in PM2.5 capture even in a high humidity environment, exhibiting their potential for use in personal protective masks.

Electrospray Principle and Apparatus for Electrospray Electrospray is an electrohydrodynamic technique similar to electrospinning. It is governed by similar principle and usually uses identical apparatus, that is, a high-voltage power supply, a syringe filled with a precursor solution (normally a polymer solution) and equipped with a metallic needle, a syringe pump controlling the feeding rate of the solution and a ground collector. During electrospray, a stable Taylor cone is also formed, which is stabilized by the liquid surface tension, electrostatic force and gravity. Electrostatic force acts as the major force against the surface tension of emitting droplets and deforms them to spherically shaped jetting beads. Compared to electrospinning, the degree of electrostatic stretch over the surface tension is relatively low during electrospray, leading to the formation of particulate products (nanoparticles or microparticles) instead of fibrous products. Consequently, a polymer solution with a relatively low polymer concentration (and hence low viscosity) is normally used in electrospray. But a very low viscosity of the solution will result in the whipping mode and also still liquid droplets when they arrive in the collector, whereas a relatively high viscosity of the solution will lead to intermittent formation of fibers. Polymer

Fig. 6 A TEM image of the fiber internal structure of an AgNP-incorporated emulsion electrospun polymer scaffolds, showing well-dispersed AgNPs in the fiber.

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Fig. 7

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A schematic diagram showing the experimental setup for simple electrospray (left) and coaxial electrospray (right).

solution concentration (and viscosity) needs to be in a specific range for electrospray and for determining the geometry of electrosprayed products. Electrospray can be categorized into two modes, that is, simple electrospray and coaxial electrospray, with the use of different types of spinneret (either a simple spinneret or a coaxial spinneret) (Fig. 7). Coaxial electrospray is more complicated because two immiscible liquids are involved. In the coaxial electrospray process, two immiscible liquids (usually one “oil phase” and one “water phase”) are merged out of spinneret to form conically shaped cone-jet under the excess force between the tangential electrostatic force and solution surface tension. Owing to different charge densities of the two immiscible liquids, different tangential electrostatic forces are exerted on them, where the liquid with relatively short electrical relaxation time (usually the water phase) bears the high tangential electrostatic force and therefore acts as the major driving phase deforming the compound jets during coaxial electrospray. Coaxial electrospray with the inner driving liquid are favorable for attaining the stable cone-jet mode, resulting in uniform core-shell structured particles with aqueous cores. But coaxial electrospray is not the only method for fabricating core-shell structured particles. By using a water-in-oil emulsion in simple electrospray, core-shell structured microparticles with multiple cores can also be produced. Due to its simplicity and high versatility in fabricating microparticles with tunable structures, electrospray has great potential for drug delivery applications.

Electrosprayed Solid Microparticles and Applications Simple electrospray of a single-phase polymer solution usually results in non-porous (“solid” or “monolithic”) nanoparticles or microparticles. Due to the high surface-area-to-volume ratio of electrosprayed solid particles and the relatively low shear stress exerted during electrospray as compared to electrospinning, electrospray has been extensively investigated for fabricating carriers for drug delivery, as well as other bioactive agents. Bioagents can be loaded in/on electrosprayed solid particles through either encapsulation or adsorption. For achieving efficient and high loading of bioagents by adsorption, electrosprayed nanoparticles are preferred owing to their very high surface-area-to-volume ratio. Bioagents, usually small molecule drugs, can be loaded onto the surface of nanoparticles through weak interactions (hydrogen bonding, electrostatic interaction, etc.) or layer-by-layer assembly. Regardless of adsorption mechanisms for loaded drugs, relatively low loaded amount and low loading efficiency are inherent problems, and drugs adsorbed on nanoparticles are easily to desorb, which normally provides release profiles with an initial burst release. Using the encapsulation approach, bioagents are firstly dissolved or dispersed in a polymer solution for electrospray, which either directly make bioagent-incorporated solid nanoparticles/microparticles or initially produce bioagent-containing droplets that are subsequently gelled/solidified by crosslinking in a particle collection bath. Both routes can achieve high drug loading amount and high encapsulation efficiency, for example, nearly 100% encapsulation efficiency for doxorubicin by electrosprayed PLGA microparticles. Drug-incorporated electrosprayed nanoparticles/microparticles meet different therapeutic requirements, for example, antibiotic-releasing electrosprayed PLGA microparticles for anti-infection purpose and paclitaxel-releasing electrosprayed PLGA microparticles for cancer therapy. However, the simple electrospray method for incorporating drugs also has shortcomings. Difficulties occur for dispersing some drugs in particulate carriers, which arise from the difference in the hydrophilicity/hydrophobicity between the drug and the carrier polymer. Another issue is the direct contact between delicate bioagents (e.g., growth factors) and the organic solvent in the polymer solution, which causes damages to biomolecules.

Electrosprayed Core-Shell Structured Microparticles for Biomolecule Delivery Electrosprayed core-shell structured microparticles are very useful for the delivery of bioactive agents such as growth factors and genes. The fabrication of electrosprayed core-shell structured microparticles involves bi-phased liquids, one being the water phase and the other the oil phase. The bioagent is contained in the water phase (which is not necessarily water; it can be an aqueous solution) and the polymer solution acts as the oil phase. Such structural arrangement provides good protection for bioagents and a high retention level (approaching 70%) for their bioactivity. Another advantage of using core-shell structured microparticles as delivery vehicles is the release profiles of biomolecules can be effectively tailored. Compared to the normally diffusion-controlled release kinetics of bioagents from solid particles, the release kinetics of bioagents from core-shell structured microparticles are more complicated, which are determined by the diffusion of encapsulated bioagents, swelling, and degradation of the carrier polymer, and coreshell structure of the microparticles. For example, both emulsion electrospray and coaxial electrospray can produce core-shell

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Fig. 8 PLGA microparticles fabricated by emulsion electrospray and coaxial electrospray, respectively: (A) TEM images showing the internal structure of emulsion electrosprayed (left) and coaxial electrosprayed (right) microparticles, (B) in vitro VEGF release profiles for different microparticles.

structured PLGA microparticles, showing multi-compartmented cores and one single core, respectively (Fig. 8a). Consequently, the release characteristics of biomolecules (e.g., VEGF) from these two types of microparticles are different (Fig. 8b), meeting the requirement for different therapeutic purposes.

Electrosprayed Microparticles for Cell Delivery The research to deliver live cells using electrosprayed microparticles for therapeutic purposes emerges in recent years. Cell delivery methods must meet stringent requirements for biocompatible materials and cell-friendly fabrication processes. An early study indicates the suitability of electrospray for cell encapsulation and delivery. In this study, rabbit aorta smooth muscle cells are successfully encapsulated in electrosprayed PEO microdroplets, achieving a high cell survival rate (70%). PDMS solution and aqueous solutions of ECM protein that have suitable viscosity are also used to make particulate carriers for mammalian cells. Using cell electrospray techniques, the encapsulation and delivery of many types of cells, including matured somatic cell isolated from different tissues (heart, skin, blood vessels, etc.) and stem cells, are investigated. Results show excellent cell viability and no significant changes for stem cell functions. A recent investigation studies a new cell electrospray technique that enables cell encapsulation in solid hydrogel microparticles. Sodium alginate is used as the carrier polymer which is crosslinked post-electrospray by Ca2 þ ions for generating cell-encapsulated hydrogel microparticles (Fig. 9). Cells encapsulated in electrosprayed hydrogel microparticles exhibit high viability (> 90%), particularly for core-shell structured hydrogel microparticles fabricated by coaxial cell electrospray. Importantly, these electrosprayed cell-encapsulated microparticles have sufficient mechanical integrity, providing both the ease for handling and enhanced protection for the cells even under subsequent harsh device fabrication processes. Cells released from the electrosprayed hydrogel microparticles also exhibit high cell viability.

Combining Electrospinning and Electrospray for Fabricating Novel Medical Devices Electrospinning and electrospray have respective advantages in different biomedical areas. An integration strategy to combine electrospinning and electrospray can lead to formation of novel medical devices for different applications. Concurrent electrospinning

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Fig. 9 A microscopic image of cell-encapsulated alginate hydrogel microparticles made by cell electrospray: cells are labeled by using a Live/Dead cell staining kit (live cell: green; dead cell: red).

and electrospray is thus investigated for fabricating drug- or biomolecule-releasing scaffolds for tissue engineering. Although electrospun scaffolds alone with the capability to deliver specific drugs or biomolecules can be made and used, nanofibrous scaffolds embedded with homogeneously distributed electrosprayed microparticles can have high drug or biomolecule loading and hence provide enhanced therapeutic effects. To solve the cell infiltration problem in conventional electrospun tissue engineering scaffolds, concurrent electrospinning and electrospray can be used to fabricate nanofibrous scaffolds embedded with cell-encapsulated biodegradable microspheres, resulting in cell-laden scaffolds that resemble the native cell-ECM organization. Early studies on the concurrent fabrication method use direct electrospray of cell suspensions, which incurs several problems for cells. A new concurrent emulsion electrospinning and coaxial cell electrospray technique has now been then developed. It firstly place structurally stabled electrosprayed cell-encapsulated hydrogel microparticles into the nanofibrous polymer scaffolds. The cells are subsequently released in the scaffolds through selective disruption of the hydrogel microparticles, which forms 3D cell-scaffold constructs with well-preserved cell viability. The nanofibrous scaffolds in the meantime can be emulsion electronspun scaffolds containing chosen growth factor. Such advanced scaffolds effectively stimulate cell functions by providing both biomimetic structural and mechanical cues of the scaffolds and biochemical signals with the sustained release of specific growth factor from the scaffolds. Concurrent electrospinning and electrospray can be used to create novel medical devices. For example, novel tissue engineering scaffolds embedded with theranostics are investigated for post-surgery cancer patients. Theranostics are multifunctional nanodevices that simultaneously provide diagnostic and therapeutic functions for cancer. They are now intensively investigated as future medical technologies. Current treatment for cancer includes surgery, chemotherapy, and radiotherapy. After surgical removal of the tumor, cancer patients face tissue dysfunction and cancer recurrence. Advanced multifunctional medical devices for promotion tissue regeneration and for detecting and treating recurrent cancer are therefore developed. In a recent investigation, AuNP-based theranostics can be encapsulated in coaxial electrosprayed core-shell structured polymer microspheres for their controlled release. Through concurrent electrospinning and electrospray, the AuNP theranostic-encapsulated microspheres are embedded in nanofibrous PLLA scaffolds. Controlling the biodegradation of microspheres enables controlled release of theranostics in the scaffolds, making the scaffolds multifunctional for cancer patients. While the nanofibrous scaffolds assist local tissue regeneration, the released AuNP-based theranostics detect and kill cancer cells. Such a comprehensive treatment holds great promise for postoperative cancer patients.

Concluding Remarks The simplicity and versatility of electrospinning and electrospray for respectively producing fibrous products and particulate products for a variety of biomaterials make these techniques highly attractive in the biomedical field. With advances in electrospinning and electrospray techniques, fibrous and particulate products possessing well-defined composition, size, morphology, and structure can be made, leading to numerous promising medical devices with distinctive but controllable characteristics and properties for meeting the requirements of different biomedical applications. Electrospun scaffolds with different nanofiber assemblies

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mimicking the features of native ECM of different human body tissues and with high surface-area-to-volume ratios have been extensively investigated, demonstrating their great potential and effectiveness in tissue engineering, drug delivery, biosensor and many other applications. Electrosprayed nanoparticles and microparticles have also been investigated as delivery vehicles for the controlled release of drugs, biomolecules, genes, and even live cells. Through integrating electrospinning with electrospray, novel multifunctional medical devices can be made, such as novel tissue engineering scaffolds embedded with theranostics for postsurgery cancer patients. Even though great successes have been made by developing and using electrospinning and electrospray in the biomedical field, the potential of these techniques has not been fully realized and there are still many issues that need to be dealt with for currently electrospun or electrosprayed products. A future development of electrospinning and electrospray may go in the direction of combining them with other manufacturing technologies for fabricating better medical devices and novel medical devices for millions of patients.

Acknowledgments Min Wang thanks The University of Hong Kong, Hong Kong Research Grants Council and the National Natural Science Foundation of China and Qilong Zhao thanks China Postdoctoral Science Foundation, Guangdong Innovative and Entrepreneurial Research Team Program and the Fundamental Research Program of Shenzhen for funding their research in the biomaterials and tissue engineering field.

Further Reading Bosworth, L., & Downes, S. (2016). Electrospinning for tissue regeneration (1st ed.). Cambridge: Woodhead Publishing. Brown, T. D., Dalton, P. D., & Hutmacher, D. W. (2011). Direct writing by way of melt electrospinning. Advanced Materials, 23(47), 5651–5657. Chakraborty, S., Liao, I. C., Adler, A., & Leong, K. W. (2009). Electrohydrodynamics: A facile technique to fabricate drug delivery systems. Advanced Drug Delivery Reviews, 61(12), 1043–1054. Cui, W. G., Zhou, Y., & Chang, J. (2010). Electrospun nanofibrous materials for tissue engineering and drug delivery. Science and Technology of Advanced Materials, 11(1), 014108. Greiner, A., & Wendorff, J. H. (2007). Electrospinning: A fascinating method for the preparation of ultrathin fibers. Angewandte Chemie International Edition, 46(30), 5670–5703. Jiang, T., E. Carbone, J., Lo, K. W. H. and Laurencin, C. T. (2015). Electrospinning of polymer nanofibers for tissue regeneration. Progress in Polymer Science, 46: 1–24. Khorshidi, S., Solouk, A., Mirzadeh, H., Mazinani, S., Lagaron, J. M., Sharifi, S., & Ramakrishna, S. (2016). A review of key challenges of electrospun scaffolds for tissueengineering applications. Journal of Tissue Engineering and Regenerative Medicine, 10(9), 715–738. Liu, W., Thomopoulos, S., & Xia, Y. (2012). Electrospun nanofibers for regenerative medicine. Advanced Healthcare Materials, 1(1), 10–25. Macagnano, A., Zampetti, E., & Kny, E. (2015). Electrospinning for high performance sensors (1st ed.). Switzerland: Springer International Publishing. Mitchell, G. R. (2015). Electrospinning: Principles, practice and possibilities (1st ed.). Cambridge: Royal Society of Chemistry. Peng, S., Jin, G., Li, L., Li, K., Srinivasan, M., Ramakrishna, S., & Chen, J. (2016). Multi-functional electrospun nanofibres for advances in tissue regeneration, energy conversion & storage, and water treatment. Chemical Society Reviews, 45(5), 1225–1241. Pham, Q. P., Sharma, U., & Mikos, A. G. (2006). Electrospinning of polymeric nanofibers for tissue engineering applications: A review. Tissue Engineering, 12(5), 1197–1211. Pisignano, D. (2013). Polymer nanofibers: Building blocks for nanotechnology (1st ed.). Cambridge: Royal Society of Chemistry. Rnjak-Kovacina, J., & Weiss, A. S. (2011). Increasing the pore size of electrospun scaffolds. Tissue Engineering Part B Reviews, 17(5), 365–372. Sun, B., Long, Y. Z., Zhang, H. D., Li, M. M., Duvail, J. L., Jiang, X. Y., & Yin, H. L. (2014). Advances in three-dimensional nanofibrous macrostructures via electrospinning. Progress in Polymer Science, 39(5), 862–890. Szentivanyi, A., Chakradeo, T., Zernetsch, H., & Glasmacher, B. (2011). Electrospun cellular microenvironments: Understanding controlled release and scaffold structure. Advanced Drug Delivery Reviews, 63(4–5), 209–220. Uyar, T., & Kny, E. (2017). Electrospun materials for tissue engineering and biomedical applications (1st ed.). Cambridge: Woodhead Publishing. Wang, C., & Wang, M. (2014). Electrospun multifunctional tissue engineering scaffolds. Frontiers of Materials Science, 8(1), 3–19. Wendorff, J. H., Agarwal, S., and Greiner, A. (2012). Electrospinning: Materials, processing, and applications. (1st ed.). Chichester: Wiley. Zhao, Q. L., & Wang, M. (2017). Smart multifunctional tissue engineering scaffolds. In Q. Wang (Ed.), Smart materials for tissue engineering: Applications (pp. 558–595). Cambridge: Royal Society of Chemistry. Zhong, S., Zhang, Y., & Lim, C. T. (2012). Fabrication of large pores in electrospun nanofibrous scaffolds for cellular infiltration: A review. Tissue Engineering Part B Reviews, 18(2), 77–87.

Gene Delivery and Clinical Applications Mahboob Morshed, Independent University, Dhaka, Bangladesh Ezharul Hoque Chowdhury, Monash University, Clayton, VIC, Australia © 2019 Elsevier Inc. All rights reserved.

Introduction Transport Vehicles for Genetic Materials Adenoviral Vectors Adeno-Associated Viral Vectors Retroviral Vectors Nonviral Vectors Strategies for long-term and tissue-specific expression of nonviral DNA Lipid-based DNA Polymeric DNA vectors Determinants of polyplex stability Determinants of stability of lipid-based complexes Passive targeting and endothelial escape Active targeting and uptake by target cells Endosomal escape Nuclear targeting Gene Therapy Applications Gene Therapy of Hereditary Diseases X-linked severe combined immune deficiency Hemophilia B X-linked adrenoleukodystrophy b-thalassemia Gene Therapy for Cardiovascular Diseases Gene Therapy for Neurodegenerative Diseases Gene Therapy for Human Immunodeficiency Virus Gene Therapy for Viral Hepatitis Gene Therapy for Cancer p53 gene TK gene TNF-related apoptosis-inducing ligand gene p21 gene Proinflammatory cytokine genes Angiotensin gene References

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Introduction Gene therapy is reversing of pathological processes through introduction of protein-coding or noncoding nucleic acids (DNA or RNA) in order to provide the missing protein function(s), correct the abnormal function(s), or silence the overexpression of protein(s). A gene is transcribed within the nucleus of a cell into an mRNA, which is subsequently transported to cytoplasm and translated into a protein or polypeptide. Cell functions are predominantly carried out by proteins and as a result mutation in a single or multiple genes could lead to suppression or overexpression of protein(s) with consequential perturbation of cellular functions. Many human diseases are caused by germline mutation(s), giving rise to inherited (genetic) diseases, or somatic mutations, causing acquired diseases, such as cancer. Introducing a functional gene ex vivo which involves extraction and culture of a specific cell type from a patient, genetic modification of the cultured cells and finally, transfer of the genetically altered cells into the same patient, and in vivo which means administration of the gene into the patent via a suitable route, such as intravenous or intramuscular or intranasal route, is a powerful approach to treat various critical human diseases at the genetic level, with many preclinical and clinical studies currently undergoing. The ultimate purpose of gene delivery into the target cells is to provide a new function or restore the missing function or suppress an unwanted function, by expressing new protein(s), or silencing harmful protein(s). Like a plasmid carrying a protein-coding gene(s), a short hairpin RNA (shRNA), which provides long-term effects on gene-silencing compared to a small interfering RNAs (siRNAs), is transcribed as an RNA transcript in the nucleus from plasmid DNA and subsequently transported to the cytoplasm. The exogenous nucleic acid or functional gene is either inserted into the genome of viral

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particles (viral vectors) or in the plasmid DNA used as a naked form or in association with a synthetic or semisynthetic particles (nonviral vectors). Long-term gene expression enables to avoid the repeated administration of gene therapeutics, while shortterm expression is considered relatively safer.

Transport Vehicles for Genetic Materials Since nucleic acids are vulnerable to degradation and unable to be transported across the cell membrane to reach the desirable site of target cells, they often require nanocarriers for being safely carried, overcoming the extracellular and intracellular barriers. Despite possessing advantages and disadvantages, both viral and nonviral vectors have been extensively explored as nucleic acid carriers in many preclinical and clinical trials of gene therapy.

Adenoviral Vectors Employed in approximately 25% of all gene therapy trials, adenoviral vector is the most clinically used vector that can efficiently transduce both dividing and nondividing cells. Following cellular uptake usually promoted by interaction between the virus fiber “knob” and cellular coxsackievirus–adenovirus receptor, and subsequent nuclear translocation, the viral genome remains episomal, replicating in synchrony with the host genome. As a result, when the vector carrying a therapeutic protein-coding gene was transferred into relatively quiescent tissues, such as the brain, liver, or muscle, the therapeutic protein, such as coagulation factors, a1 antitrypsin, or erythropoietin, was stably produced throughout the lifetime of a mouse (Roth et al., 1996). Since the vector can induce severe toxicity owing to an immediate innate immune response and a secondary antigen-dependent response, the secondand third-generation adenoviral vectors have been produced with less toxicity than the first-generation vectors, by deleting additional viral genes. The majority of gene therapy trials (> 400) carried out or undergoing with human adenoviral vectors aims at treating various cancers. Among the large number of different transgenes incorporated into the adenoviral genome for clinical trials, cystic fibrosis transmembrane conductance regulator (CFTR), E. coli cytosine deaminase, granulocyte macrophage colony stimulating factor (GM-CSF), human immunodeficiency virus (HIV) proteins (gag, pol, nef), human IL-2, human IL-12, human interferon (IFN), human fibroblast growth factor 4 (FGF-4), heat shock protein 70 (HSP 70), hepatitis C virus (HCV) nonstructural proteins, ornithine transcarbamylase, p53, thymidine kinase (TK), tumor necrosis factor (TNF), prostate-specific antigen, malaria circumsporozoite protein, Her-2, and vascular endothelial growth factor (VEGF) are worth mentioning (Wold and Toth, 2013).

Adeno-Associated Viral Vectors Adeno-associated virus (AAV), one of the most attractive gene therapy vectors, usually interacts with heparan sulfate proteoglycans on cell membrane as the first step for cellular internalization. Like an adenovirus, it also induces strong immune responses, infects both dividing and nondividing cells, and remains episomal without being integrated into host chromosome. The short genome size (4 kb) of the virus poses a limitation in carrying a long foreign gene. In clinical trials for familial lipoprotein lipase (LPL) deficiency, intramuscular injection of an AAV1 vector encoding the gain-of-function LPLS447X variant led to persistent gene expression, with consequentially sustained decrease in the incidence of pancreatitis. Considering the efficacy as well as the safety profile, AAV1LPLS447X vector received marketing approval in 2012 in the European Union as the first approved gene therapy product in Western nations under the name “alipogene tiparvovec” (Glybera). In another trial, administration of an AAV1 vector with ATP2A2 gene encoding sarcoplasmic/endoplasmic reticulum calcium ATPase 2 was found to improve the key outcomes in the patients with advanced heart failure (Kotterman and Schaffer, 2014).

Retroviral Vectors Only lentiviruses among the retroviruses are capable of replicating in nondividing cells, making them attractive for transducing human cells comprising both dividing and nondividing cells. In addition, as a member of retroviruses, they can stably integrate their genomes into the host chromosome, which although enables sustained gene expression, simultaneously increases the risk of proto-oncogene activation and cancer development. Genetic engineering enables production of targetable lentiviral vectors through fusion of a ligand protein or antibody to viral glycoproteins. In order to promote gene expression in target cells, in addition to ensuring the specific uptake via the receptor–ligand engagement, lentiviral vectors with a tissue-specific promoter can be used to target specific cell type. Lentiviral vectors made defective for recombination with host chromosome can promote short-term transgene expression in nondividing cells while reducing insertional mutagenesis. These vectors were used in preclinical animal studies for correction of hemophilia B, b-thalassaemia, Parkinson’s disease (PD), cystic fibrosis, sickle cell anemia, and spinal muscular atrophy.

Nonviral Vectors Despite being highly carcinogenic and immunogenic, in
70% of gene therapy clinical trials viral vectors have been used. On the other hand, although much inefficient compared to the viral counterpart, nonviral vectors are currently being sought considering

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their increased safety profile and capacity to deliver larger genetic material. Nonviral physical methods including gene gun, electroporation, hydrodynamic delivery, sonoporation, and magnetofection are generally less applicable to systemic delivery and therefore, a range of synthetic or semisynthetic nonviral vectors have been developed.

Strategies for long-term and tissue-specific expression of nonviral DNA Following nuclear transport, plasmids that are routinely used as nonviral expression vectors carrying the gene(s) of interest remain episomal, thus reducing the risk of insertional mutagenesis. To control the intensity and the duration of transgene expression, viral enhancers and promoters derived from cytomegalovirus (CMV), respiratory syncytial virus, and simian virus 40 (SV40), and constitutive mammalian promoters, such as the human ubiquitin C and the eukaryotic translation elongation factor 1 alpha 1 (EEF1A1), are frequently used as the regulatory sequences of the therapeutic genes. Additionally, there are numerous cis-acting sequences including various polyadenylation signals introns and scaffold/matrix attachment regions (S/MARs) with ability to increase the level and the duration of transgene expression. DNA size and topology can also influence the gene expression level with the small covalently closed circular plasmid DNA causing greater levels of transgene expression than the large or linearized plasmids. A variety of transposition systems, such as recombinases phiC, PiggyBac, and Sleeping Beauty, can be included to prolong the gene expression by integrating the transgene into the chromosome. Finally, usage of tissue-specific promoters or enhancers, such as the alphafetoprotein (AFP) enhancer or albumin (ALB) promoter, which regulates gene expression exclusively in the liver, can minimize the unwanted transgene expression in nontarget tissues, and thus increase the overall delivery efficacy and reduce the off-target effects (Yin et al., 2014).

Lipid-based DNA Cationic liposomes are the most commonly used nonviral vectors for plasmid DNA delivery. Various cholesterol or PEG-modified cationic liposomal formulations were tested clinically, such as DOTAP–cholesterol for delivery of fus1 tumor suppressor gene and GL67A–DOPE–DMPE–PEG for delivery of CFTR (pGM169) gene in patients with non-small-cell lung cancer and cystic fibrosis, respectively (Yin et al., 2014).

Polymeric DNA vectors Cationic polymers constitute another attractive class of nonviral DNA vectors partly owing to their chemical diversity and potential for surface functionalization. PEI and its variants having a high charge density at reduced pH values are effective in condensation and endosomal escape of DNA, leading to efficient intracellular transgene delivery. The transgene delivery efficacy and cytotoxicity of PEI strongly depend on its molecular weight and structural complexity (linear vs. branched forms). A range of modifications were undertaken for better performance of PEI; for example, block co-polymers of PEG and PEI for improved stability and biocompatibility, degradable disulfide-crosslinked PEIs for reduced toxicity, and alkylated PEI for increased potency (Yin et al., 2014). Intravenous injection of PEI–DNA polyplexes was shown to enhance transgene delivery into the lungs of mice. PEI was studied for local gene therapy of various cancers, including bladder, ovarian and pancreatic cancers, multiple myeloma, B cell lymphoma, and pancreatic ductal adenocarcinoma in humans. A PEG–PEI–cholesterol lipopolymer carrying interleukin-12 (IL-12) plasmid is under clinical investigation for immunotherapy of ovarian and colorectal cancers.

Determinants of polyplex stability Cationic polymers are commonly used to complex with nucleic acids having negatively charges phosphate backbone, forming polyplexes, the stability of which depends on the ratio of positive to negative charges (N/P). For example, linear PEI/plasmid complexes were less stable than branched PEI/plasmid complexes (Madani et al., 2011). In order for systemic delivery, polyplex nanomicelles can be formed through self-assembly of amphiphilic block copolymers carrying a cationic polymer segment which forms the core by condensing with nucleic acid, and a hydrophilic polymer chain, such as PEG on the surface facing the aqueous solution to prevent destabilization of the resultant complexes by hindering nonspecific interactions with blood components, such as nuclease and opsonins.

Determinants of stability of lipid-based complexes The stability of lipoplexes, which are formed by ionic interactions between cationic lipids and anionic nucleic acids, usually depends on the ionic strength under which the lipoplexes are prepared. Thus, a lipoplex formulated under low ionic strength solutions will be destabilized releasing the nucleic acid payloads, after being exposed to physiological saline and serum. Although the existence of PEG on the surface of lipoplexes confers a steric barrier at the lipoplex surface, blocking a variety of interactions with molecular and cellular components in blood and tissues, the ultimate transfection activities of the surface-modified lipoplexes could be hampered. The stability of lipo-polyplexes which are formed by addition of liposomes to preformed polyplexes is mainly determined by the outer liposome shell independent of pDNA and cationic polymer in the polyplexes, whereas the stability of stable nucleic-acid-lipid particle, which consists of a lipid bilayer including a mixture of cationic and neutral lipids (cholesterol and fusogenic lipids) and a PEG-lipid coating, is mainly dependent on the outer liposomal shell.

Passive targeting and endothelial escape Due to the enhanced permeability and retention effect (EPR), nanoparticles with appropriate size could escape the tumor blood capillaries composed underdeveloped, leaky endothelium and be retained in the tumor tissues for days and even weeks due to

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the lack of lymphatic drainage. In addition, acting as vasodilators, nitric oxide (NO), prostaglandins, and bradykinin were reported to enhance the EPR effect in the tumor by increasing its vascular permeability. In fact, the surface properties and diameter of protein corona-decorated particles have notable influence on the outcome of passive targeting. Particles circulating in blood should be larger than 40 kDa with half-lives to be sufficiently high so as to exert the EPR effect. In addition, particles with high positive charges can bind nonspecifically to the luminal surface which is negatively charged due to the presence of sulfated and carboxylate sugar moieties, resulting in rapid clearance from the circulation.

Active targeting and uptake by target cells Covalent or noncovalent attachment of a targeting ligand on the surface of nanoparticles enables them to recognize the specific antigens or receptors on target cells which subsequently engulf the particles through endocytosis. The targetability and efficiency of the uptake are governed by the interactions between the nanoparticle surface groups and the plasma membrane antigens/receptors, which in turn depend on density of the ligands and the antigens/receptors present on a nanoparticle and a cell, respectively. The diverse moieties investigated so far as targeting ligands include carbohydrates (e.g., galactose), monoclonal antibodies (e.g., antiHer2, anti-EGFR), peptides (e.g., Arg-Gly-Asp or RGD), proteins (e.g., lectins, transferrin), vitamins (e.g., vitamin D), and aptamers (e.g., RNA aptamers against HIV glycoprotein).

Endosomal escape The superior transgene expression efficacy of viral vectors over the nonviral vectors is partly due to the ability of viral proteins to facilitate the endosomal escape, following cellular uptake of the viral particle. Thus, in case of adenoviruses, acidic pH of endosome induces conformational changes in viral capsid, leading to endosomal membrane lysis. There are two major strategies undertaken for nonviral vectors to facilitate the endosomal escape: inclusion of fusogenic peptides within the vectors as a virus-mimicking strategy to induce endosomal low pH-induced conformation change of the peptide, leading to fusion with the endosomal membrane; and implementation of “proton sponge” effect with PEI, for instance, by neutralizing the endosomal acidic pH with its excess uncharged amines, eventually bringing in chloride into endosomes, causing osmotic swelling and finally, rupturing the endosomal membrane. In addition, in case of cationic liposomes, the interactions between positive charges of the liposomes and anionic phospholipids of the endosomal membrane cause some of the anionic lipid to displace in a “flip-flop” mechanism and diffuse into the lipoplex, forming a charge neutral ion pair with the cationic lipids and resulting in release of the liposomal contents into the cytosol. The inclusion of neutral helper lipids as a constituent of liposome can accelerate the fusion of lipid layers by switching from a lamellar to a hexagonal phase as the pH of endosome drops. The presence of PEG in the lipoplex surface could impair the phase transition process, ending up with an incomplete endosomal escape (Chowdhury, 2016).

Nuclear targeting Most eukaryotic DNA viruses and some RNA viruses can gain access to the nucleus of their host cells via nuclear membrane pore (NMP) having an internal channel diameter of
25 nm. Intact virions with a diameter smaller than that of the NMP can either translocate through the NMP by recruiting appropriate nuclear import receptors, or undergo conformational changes to allow their outer surface to interact with channel components, thus resulting in the release of their genome in the host nucleus. However, virions significantly larger than the maximum diameter of the NMP are required to partially disassemble or uncoat themselves either in the cytosol or after docking at the NMP before transferring their genomes into the nucleus. Since NMP permits passive transfer of linear DNA fragments up to
300 bp, although dividing cells (e.g., cancer cells) permit entry of plasmid DNA into the nucleus once the nuclear membrane is disrupted during mitosis, nondividing cells do not allow trafficking of molecules larger than 40–45 kDa through NMP unless they possess nuclear localization signals (NLSs) (Chowdhury, 2009). Nanoparticles could be modified with the NLSs originally derived from the Simian virus 40 large T antigen to enter the nucleus following cellular uptake. However, even though a nanocarrier with bound DNA can successfully reach the nucleus, the DNA must be released from the carrier for undergoing transcription in the nucleus. The plasmid that reaches the cytosol after cellular uptake and release from both the endosome and the carrier will be prone to nuclease-mediated degradation during its waiting for the breakdown of nuclear membrane in case of dividing cells (Chowdhury, 2009). For nuclear translocation of a plasmid DNA, the DNA can be covalently linked to a NLS peptide, or specific DNA sequences, such as a 366-bp DNA containing the SV40 origin of replication and early promoter can be incorporated into the DNA by recombinant DNA technology, prior to the complexation of the DNA with a nanocarrier that ensures the cytosolic release of the DNA cargo (Chowdhury, 2009).

Gene Therapy Applications For treating both genetic and acquired human diseases, plasmid DNA carrying gene(s) of interest can be administered either in vivo via different routes, or ex vivo by genetically modifying the autologous cells derived from a patient and reintroducing them in the patient’s body. Ex vivo gene therapy was particularly attempted for the monogenic diseases of blood cells, such as sickle cell anemia or b-thalassemia.

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Gene Therapy of Hereditary Diseases X-linked severe combined immune deficiency X-linked severe combined immune deficiency (X-SCID), a combined cellular and humoral immunodeficiency caused by mutations in interleukin 2 receptor, gamma (IL2RG) gene with almost complete absence of T and natural killer lymphocytes, is fatal in the first 2 years of life without reconstitution of the immune system via bone marrow transplantation or gene therapy. Although autologous transplantation of retrovirally transduced bone marrow cells expressing the IL2RG gene resulted in a functional immune system in the children with X-SCID, 5 out of the 20 patients subsequently developed leukemia as a result of integration of the retroviral DNA into the patient’s genome and the eventual activation of the LIM domain only 2 proto-oncogene in those hematopoietic cell-derived cells (Chowdhury, 2016).

Hemophilia B A mutation in factor IX gene causes a deficiency in factor IX clotting factor, consequentially prolonging bleeding after an injury in patients with hemophilia B. Although administration of a single-dose of AAV2 vectors carrying the factor IX gene in order for its liver-based delivery and expression, led to a long-term correction of hemophilia B in mice and dogs, preexisting antibodies against the capsid of the AAV2 serotype, or a CD8þ T cell response to the capsid led to short-term expression of factor IX with poor therapeutic outcome in a patient with hemophilia. However, peripheral infusion of an AAV8 vector expressing a codon-optimized human factor IX transgene led to long-term expression of factor IX, thus improving the bleeding phenotype in patients with severe hemophilia B (Chowdhury, 2016).

X-linked adrenoleukodystrophy Mutations in the ABCD1 gene encoding peroxisomal membrane protein, ALDP X-linked adrenoleukodystrophy (X-ALD gene product) disrupt the transmembrane transport of very long-chain fatty acids and thereby cause a severe lipid storage disorder with brain demyelination in patients with X-ALD. Lentiviral vector-mediated transfer of ABCD1 gene in the autologous hematopoietic stem cells (HSCs) from the patents with X-ALD showed the outcome apparently comparable to the allogeneic hematopoietic cell transplantation (Chowdhury, 2016).

b-thalassemia

Mutations in beta-globin chains result in the reduced production of hemoglobin in patients with b-thalassemia. Transfer of the b-globin gene with the help of a lentiviral vector into the HSCs of an 18-year-old patient of b-thalassemia led to 10% of the patient’s blood cells having the normal hemoglobin, due to the integration of the b-globin gene without having any insertional mutagenesis (Chowdhury 2016).

Gene Therapy for Cardiovascular Diseases Renin–angiotensin system (RAS) plays a pivotal role in development of atherosclerosis and hypertension, leading to congestive heart failure, myocardial infarction, and kidney damage. In RAS, angiotensinogen (AGT) acts as the precursor in liver for synthesis of angiotensin II which increases blood pressure (BP), promotes progression of atherosclerosis, generates a larger amount of superoxides and reactive oxygen species, and impairs endothelium-dependent dilation by reducing NO production. Silencing of AGT mRNA transcript by nanoparticle-facilitated delivery of specific shRNA was shown to markedly reduce expression of AGT and Ang II in rats, resulting in the decline in BP with the atherosclerotic lesions markedly attenuated. Furthermore, an effective lowering of high BP and attenuation of the pathophysiology were observed in different experimental models of hypertension by overexpressing vasodilators including atrial natriuretic peptide, endothelial nitric oxide synthase, kallikrein, and adrenomedullin (Chowdhury, 2016).

Gene Therapy for Neurodegenerative Diseases Progressive loss of neurons and gradual appearance of disable neurological symptoms characterize the neurodegenerative diseases that predominantly include Huntington’s disease (HD), Alzheimer’s disease (AD), and PD. Silencing of mutant human huntingtin in the striatum and cerebellum of HD mice with a specific shRNA carried by an AAV caused a significant pathological and behavioral improvement as well as a reduction in the size and number of neuronal inclusions. On the other hand, lentiviral delivery of shRNA targeting BACE1 in a transgenic mouse model of AD significantly reduced Ab production, amyloid plaques, and neuronal death, with improved learning and memory. Knockdown of a-synuclein, a principle protein component of Lewy bodies, has been proposed as a potential therapeutic option for the treatment of PD (Chowdhury, 2016).

Gene Therapy for Human Immunodeficiency Virus HIV decreases and weakens the immune system, paving the way for opportunistic infections and development of acquired immunodeficiency syndrome, by binding to and destroying CD4 þ T lymphocytes. Targeting the cellular components involved in the viral

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infection process, such as CD4 (the primary receptor) by which HIV enters the cells, could be a promising strategy. The first clinical trial (NCT00569985; 04047) was based on a lentiviral vector (rHIV7-shI-TAR-CCR5RZ) that was used to transduce autologous, CD34 þ hematopoietic progenitor cells, expressing a shRNA targeted to an exon of the HIV-1 genes tat/rev (shI) to destroy the viral mRNA, a decoy for the HIV TAT-activated RNA (TAR) to antagonize the viral transactivation, and a ribozyme targeting the host T cell CCR5 cytokine receptor (CCR5RZ) (a coreceptor) to block viral entry (Deng et al., 2014).

Gene Therapy for Viral Hepatitis HCV infection is a major cause of chronic liver disease. Single-stranded HCV genome could be an attractive target for potential RNAi-based therapeutics. Specifically designed shRNAs (or siRNAs) can be delivered by either viral or nonviral vectors to target the 50 UTR and 30 UTR, the most conserved regions of the HCV RNA with significant functional importance in the viral life cycle, NS3 helicase which unwinds the viral genomic RNA during replication, NS3 serine protease which contributes to HCV polyprotein maturation, and NS5A, a pleiotropic protein having key roles in both viral RNA replication and modulation of the host cell’s physiology (Motavaf et al., 2012).

Gene Therapy for Cancer Cancer, a leading cause of mortality worldwide, is caused by uncontrolled cell division owing to mutational and epigenetic changes that lead to overexpression or suppression of cellular genes. Metastasis of cancer cells happens due to the downregulation of cell adhesion receptors, upregulation of cell motility-enhancing receptors and activation of membrane metalloproteases. Gene therapy for cancer could be classified into three major categories: tumor suppressor gene replacement therapy, immune gene therapy, and enzyme- or prodrug-based therapy. On the other hand, shRNAs (and siRNAs) have been extensively used to silence antiapoptotic genes, such as bcl-2, proangiogenic growth factor genes, such as fibroblast growth factor (bFGF), VEGF, growth factor receptor genes, such as epidermal growth factor receptor (EGFR), estrogen receptor, human epidermal growth factor receptor-2 (HER2), and VEGFR receptor, multidrug transporter genes, such as MDR1 and MRP1 genes in many preclinical studies, to inhibit tumor cell growth, angiogenesis, metastasis, and chemo-resistance.

p53 gene p53 belongs to the class of tumor suppressors, having crucial roles in preventing tumor development by ensuring cell cycle progression, DNA repair, and induction of apoptosis against cellular stress and damage, and is therefore known as “guardian of genome.” In addition, p53 plays pivotal roles in enhancing the therapeutic effects of antiangiogenesis therapy, chemotherapy, and radiotherapy. Systemic delivery of liposome-p53 DNA complex was reported to reduce the size of the primary tumors and prevent the relapse and metastases of a malignant human breast cancer in nude mice. In addition, intratumoral injection of biodegradable poly(b-amino ester) polymer-carrying p53 gene in xenograft model of SCLC resulted in significant inhibition of tumor growth. GendicineÔ is a p53 adenovirus approved for clinical use in China for the treatment of head and neck squamous cell cancer in combination with radiotherapy (Bakhtiar et al., 2014).

TK gene Herpes simplex virus thymidine kinase (HSVtk), the most commonly used TK gene for cancer gene therapy, was successfully delivered individually as well as in combination with ganciclovir (GCV) (a prodrug) in vivo. Enhanced antitumor efficacy of HSV-1tk/ GCV prodrug system was observed in baby hamster kidney (BHK) tumor grown as xenografts in severe combined immunodeficiency disease (SCID) mice, following delivery with sindbis viral vector. Conversion of prodrug GCV into the active form by the expressed HSVtk promoted the bystander effects responsible for killing surrounding untransducted tumor cells. In a similar approach, repeated delivery of HSVtk gene using hemagglutinating virus of Japan (HVJ)-liposomes in the induced liver tumors, followed by GCV treatment significantly inhibited the tumor growth and markedly improving the survival of animals. In clinical trials, intratumoral injection of adenovirus-associated HSV-tk gene and subsequent intravenous administration of GCV led to an overall mean survival of 70.6 weeks without significant safety issues, compared to 39.0 weeks for the standard therapy group in glioblastoma multiforme (Bakhtiar et al., 2014).

TNF-related apoptosis-inducing ligand gene TNF-related apoptosis-inducing ligand (TRAIL), a cytokine belonging to the TNF superfamily, binds to its receptors, R1 and R2, forming the death-inducing signaling complex which in turn directs cleavage and activation of caspases for inducing apoptosis. TRAIL is considered to be a promising gene therapeutic in cancer treatment considering its specificity in inducing apoptosis and toxicity in cancer cells. A nonviral vector, PEI600-Cyd, prepared by linking low-molecular-weight polyethylenimine (PEI) with b-cyclodextrin (b-CD) was used to introduce the TRIAL gene to mesenchymal stem cells (MSCs) which was later utilized as a delivering agent in vivo to significantly reduce the tumor size. In another study, a low-molecular-weight PEI10 k modified with myristic acid (MC) was complexed with TRAIL gene to form MC-PE10 k/DNA nanoparticles which were subsequently used to deliver the TRAIL gene to the brain, with an improvement in median survival time (Bakhtiar et al., 2014). TRAIL gene-carrying MSCs (MSCsTRAIL) were delivered along with 5-flourouracil to xenografts of HCT116 colon cancer cells, resulting in significant tumor regression compared to the single-agent treatment (Bakhtiar et al., 2014). The combined delivery of doxorubicin and pTRAIL

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gene using PEI-CD (cyclodextrin) was demonstrated to be efficient with higher retention time for drugs, achieving good therapeutic effects in inhibiting the tumor growth with significantly prolonged survival of tumor-bearing mice (Fan et al., 2012). c(RGDyK)– poly(ethylene glycol)–polyethyleneimine (RGD–PEG–PEI) nanoparticles were used in a similar study for the codelivery of TRAIL gene and paclitaxel, producing better antiglioblastoma effects in vivo by overcoming blood–brain barrier and blood–tumor barrier (Zhan et al., 2012).

p21 gene A vital regulator of cell cycle progression, p21, can block cell-cycle progression from the G1 to S phase through inactivation of cyclin/ CDK activity. Retroviral vector-mediated delivery of p21 gene was shown to reduce breast tumor growth in vivo by reducing expression of cyclins D1 and E and Cdks 2, 4, and 6 (Bakhtiar et al., 2014).

Proinflammatory cytokine genes Survival and sustainable growth of a tumor is supported by lack of expression of recognizable tumor antigens, inability of the expressed tumor antigens to adequately stimulate the immune system or downregulation of the immune response by the tumor itself. Tumor expression of proinflammatory cytokine(s), such as GM-CSF or fms-like tyrosine kinase 3 receptor ligand (Flt3L) and subsequent immunization with the tumor lysate containing tumor-associated antigens, could dramatically increase the number of potent antigen-presenting dendritic cells (DCs) in blood. For example, vaccination with the CT26 colon carcinoma cell lysate and the Flt3L-encoding adenoviral vector (pAdFlt3L) injected subcutaneously prevented the tumor growth in a BALB/c mouse model through significant expansion of DCs (Riediger et al., 2013).

Angiotensin gene An octapeptide hormone, angiotensin II, plays a key role in the rennin–angiotensin system by binding with angiotensin II type 1 or type 2 receptors. Angiotensin II type 2 (AT2R) is known to inhibit cell proliferation and stimulate apoptosis. A marked reduction in tumor growth was observed following bolus administration of HIV-1 TAT peptide-based nanoparticles with encapsulated AT2R or TNF-related apoptosis-inducing ligand (TRAIL) pDNA. A single intratumoral injection of defective adenovirus expressing a secretable angiostatin K3 molecule from the CMV promoter (AdK3) into rat C6 glioma or human MDA-MB-231 breast carcinoma (established in athymic mice) resulted in a significant arrest of tumor growth with suppression of neovascularization in the tumors (Bakhtiar et al., 2014). In addition, adenoviral delivery of the genes for selective interleukin (IL), such as IL-2 and IL-12, interferon (IFN), such as IFN-a, IFN-b, and IFNg, and Fas ligand was shown to exert potent antitumor effects in different tumor models.

References Bakhtiar, A., Sayyad, M., Rosli, R., Maruyama, A., & Chowdhury, E. H. (2014). Intracellular delivery of potential therapeutic genes: Prospects in cancer gene therapy. Current Gene Therapy, 14(4), 247–257. Chowdhury, E. H. (2009). Nuclear targeting of viral and non-viral DNA. Expert Opinion on Drug Delivery, 6(7), 697–703. Chowdhury, E. H. (2016). Nanotherapeutics: From laboratory to clinic. Boca Raton, FL: CRC Press. Deng, Y., Wang, C. C., Choy, K. W., et al. (2014). Therapeutic potentials of gene silencing by RNA interference: Principles, challenges, and new strategies. Gene, 538(2), 217–227. Fan, H., Hu, Q. D., Xu, F. J., et al. (2012). In vivo treatment of tumors using host-guest conjugated nanoparticles functionalized with doxorubicin and therapeutic gene pTRAIL. Biomaterials, 3, 1428–1436. Kotterman, M. A., & Schaffer, D. V. (2014). Engineering adeno-associated viruses for clinical gene therapy. Nature Reviews. Genetics, 15(7), 445–451. Madani, S. Y., Naderi, N., Dissanayake, O., Tan, A., & Seifalian, A. M. (2011). A new era of cancer treatment: Carbon nanotubes as drug delivery tools. International Journal of Nanomedicine, 6, 2963–2979. Motavaf, M., Safari, S., & Alavian, S. M. (2012). Therapeutic potential of RNA interference: A new molecular approach to antiviral treatment for hepatitis C. Journal of Viral Hepatitis, 19(11), 757–765. Riediger, C., Wingender, G., Knolle, P., et al. (2013). Fms-like tyrosine kinase 3 receptor ligand (Flt3L)-based vaccination administered with an adenoviral vector prevents tumor growth of colorectal cancer in a BALB/c mouse model. Journal of Cancer Research and Clinical Oncology, 139(12), 2097–2110. Roth, J. A., Nguyen, D., Lawrence, D. D., et al. (1996). Retrovirus-mediated wild-type p53 gene transfer to tumors of patients with lung cancer. Nature Medicine, 2(9), 985–991. Wold, W. S., & Toth, K. (2013). Adenovirus vectors for gene therapy, vaccination and cancer gene therapy. Current Gene Therapy, 13(6), 421–433. Yin, H., Kanasty, R. L., Eltoukhy, A. A., et al. (2014). Non-viral vectors for gene-based therapy. Nature Reviews. Genetics, 15(8), 541–555. Zhan, C., Wei, X., Qian, J., et al. (2012). Co-delivery of TRAIL gene enhances the anti-glioblastoma effect of paclitaxel in vitro and in vivo. Journal of Controlled Release, 160, 630–636.

Materials for Exoskeletal Orthotic and Prosthetic Systems Man Sang Wong, Babak Hassan Beygi, and Yu Zheng, The Hong Kong Polytechnic University of Hong Kong, Hung Hom, Hong Kong © 2019 Elsevier Inc. All rights reserved.

Background of Orthotic and Prosthetic Systems Ankle-Foot Assembly Shank Socket and Socket Interface Suspension Knee Units Upper Limb Prostheses Materials Casting Materials and Fabrication Methods Gypsum plaster or plaster of paris (POP) Fiber glass bandage Foam box impression Alginate Leather Textile Fibers and Fabrics Natural Fibers Synthetic Fibers Metals Steel Aluminum Titanium Wood Synthetic Polymers Plastics Elastomers Adhesives Surface Finishing 3D Printing Summary and Possible Future Direction Further Reading List of Relevant Websites

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Glossary Cranial remodeling orthosis An orthosis prescribed to correct the asymmetry of the skull in infants. Ionomer A copolymer made by bonding of electrically neutral components and the ionized units. Maxillofacial prostheses Any prosthesis fabricated to replace the missing structure of face such as eye, ear and nose due to congenital or acquired impairment. Polycaprolactone A biodegradable polyester with a low melting temperature around 60
C. Polymer Any large molecule, natural or synthetic which is composed of repeated subunits. Polyolefins Any group of polymers produced from a simple Olefin (alkene) with the formula of CnH2n as subunit.

Background of Orthotic and Prosthetic Systems Orthopedic technology targets all devices applied externally to the user’s body to substitute a missing body part (exoprosthesis) or to replace a missing function (orthosis). For clarification, the discussion on endoprosthesis, an implanted internal prosthesis is beyond the topic of this article and the term prosthesis in following pages will stand for exoprostheses. In general, the goal of the appliance is to improve the user’s mobility and ability to perform daily functional activities. Orthotics and prosthetics (O&P) refer to the science and the field of knowledge that deals with orthoses and prostheses design and

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application. The orthotist is a person who designs and fabricates the orthotic devices. Similarly, a prosthetist’s profession is related to the manufacturing and fitting the artificial limb as well as the education of an amputee, a person with missing limb to restore his or her functional abilities. Unique expertise of orthotists and prosthetists in patient assessment, design, and materials offers patients an increased level of mobility and independence. Exoskeletons, also called exoskeletal suits are wearable frameworks alongside of human limbs which are the integrated products of orthotics/prosthetics and robotics. Exosuits are powered by a system of electronic servomotors, pneumatics, levers, hydraulics, and sensors that allow the limb to extend, enhance or even substitute the human function, the functions that might be helpful for disabled people (Fig. 1). Orthotic prescription is developed to achieve specific goals by applying biomechanical principles through orthotic design and material and component selection. Orthotic goals are classified into four groups: B B B B

To protect a joint To assist a joint movement To stabilize a joint by stopping or limiting motion To help manage a skeletal deformity by preventing, supporting or correcting of abnormality (e.g., scoliosis orthoses, cranial remodeling orthoses)

The application of word orthosis is preferred to phrases like brace and splint, although they might be used interchangeably. A standardized nomenclature including an acronym composed of the first letter of the joints encompassed within the orthosis is used for naming and easier communication (Figs. 2–4). The general classifications of orthoses and prostheses have been summarized in Tables 1 and 2, respectively. Orthotic devices can be classified according to the type of materials used as rigid, semi rigid or soft in prefabricated or custommade designs in a variety of procedures. Previously, orthoses were fabricated from metal bands and bars, called conventional

Fig. 1

Application of exoskeleton in disabled people.

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Fig. 2

Spinal orthoses made from plastics, metals and foamed plastics.

Fig. 3

AFOs made from plastics, composites with or without hinges.

Fig. 4

Combination of HKAFOs and upper structures such as TLSOs.

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General classification of orthoses

Lower limb orthoses

Upper limb orthoses

Spinal orthoses

FO (foot orthosis) AFO (ankle foot orthosis) KAFO (knee ankle foot orthosis) HKAFO (hip knee ankle foot orthosis) KO (knee orthosis) HO (hip orthosis)

HO (hand orthosis) WHO (wrist hand orthosis) EWHO (elbow wrist hand orthosis) SEWHO (shoulder elbow wrist hand orthosis) EO (elbow orthosis) SO (shoulder orthosis)

SO (sacral orthosis) LSO (lumbo sacral orthosis) TLSO (thoraco lumbo sacral orthosis) CTLSO (cervico thoraco lumbo sacral orthosis) CO (cervical orthosis)

Table 2

General classification of amputation levels and corresponding prostheses

Lower extremity

Upper extremity

Hip disarticulation Transfemoral (above knee) Knee disarticulation Transtibial (below knee) Ankle disarticulation Partial foot

Shoulder disarticulation Transhumeral (above elbow) Elbow disarticulation Trans radial (below elbow) Wrist disarticulation Partial hand

orthoses; However with introduction of new materials, the majority of orthoses were replaced by plastics/composites or a combination of both metal and plastic/composite materials as it can be represented in current prosthetic designs. Any lower extremity prosthesis includes a socket with a liner or interface, a method of suspension, and a foot. Transfemoral prostheses also incorporate a knee mechanism. Materials application in prosthetic technology is subdivided into main following domains of lower limb and upper limb components:

Ankle-Foot Assembly The distal foundation of the lower limb prosthesis is the foot-ankle assembly. All prosthetic feet should provide a stable base of support for the amputee and absorb the shocks imposed during the walking.

Shank The linking component located between the foot-ankle assembly and the distal portion of the socket is the shank segment. Shank transmits the wearer’s weight from the superstructure to the foot. The type of prosthesis is referred to as either an endoskeletal or an exoskeletal; Exoskeletal types include those prostheses in which a rigid shank characterizes the structure; On the other hand, most prostheses, today use an endoskeletal type or modular system with a central tube structure (pylon) that provides the slight adjustment of prosthesis alignment. Modular pylons are covered with resilient cosmetic foam that enhances the appearance of the prosthesis similar to sound-side limb. Exoskeletal shanks are made of rigid expandable foam which is laminated externally to match the wearer’s contralateral leg.

Socket and Socket Interface The socket is the part that contacts the wearer’s skin. It allows the transmission of forces and moments through the prosthesis to the floor. Most transtibial prostheses are furnished with a combination of insert (soft socket) or liner and prosthetic socks to distribute the stress around the limb and increase the comfort and cushioning. Transfemoral sockets are generally made of a combination of a flexible socket encompassed in a rigid laminated plastic frame.

Suspension Suspension is a component that prevents the prosthesis from slipping off the amputation limb during walking. Old-fashion suspension designs (e.g., a leather thigh corset with side metal joints) are rarely used in the age of advanced technology. There are a variety of belts, strap (e.g., supracondylar strap) and knee sleeves that are also implemented for suspension purpose.

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Fig. 5 (Left) Below knee endoskeletal prosthetic designs with different suspension methods, (right) details of suspension with metal pin locking system.

A metal pin at the end of prosthetic liner which engages into a locking mechanism in socket or a lanyard method (e.g., application of a nylon fabric cord) is considered as suspension methods through the distal end of the socket (Fig. 5).

Knee Units Knee units are mass-produced components and have several functions including the body weight support and provision of controlled movement. Knee mechanism may be simply mechanical in single axis or polycentric design or may function with the use of hydraulic or pneumatic mechanism. By using a liquid medium (usually silicone oil or magnetized rheological fluid), the hydraulic controlled knee mechanisms will facilitate the patient to experience variable walking speed. Some of these hydraulic systems can be driven by incorporating the microprocessors to possess a computer-driven control (Fig. 6). There are some knees and ankle designs including electromechanical motors interlinked with microprocessors which are able to move knee and ankle, a mechanism which might be quite helpful in stair ascending.

Upper Limb Prostheses The upper extremity prosthesis basically consists of a socket, a method of suspension, a power source, plus elbow and shoulder joints depending on level of amputation. There are two basic types of upper extremity prostheses: body-powered and myoelectric.

Fig. 6

Transfemoral prostheses with fluid-control knee units.

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Essentially, a body-powered prosthesis uses body movements, harnessed with control straps and cables to mobilize the terminal device. The terminal device may be a hook or a hand. In Myoelectric prostheses, the implanted electrodes in the prosthetic socket play a key role in detection of signals of contracted muscles in residual limb. The signal will be interpreted in controller to open and close the terminal device. To design an orthotic or prosthetic device, many features should be kept into the consideration such as specific goals and requirements of the patient, weight of device, cosmetics, durability, ease of putting on (donning) and taking off (doffing) and ventilation to prevent skin maceration. Unfortunately, no single material has an ideal set of properties for application in O&P. The type of material influences the design characteristics as well as the effectiveness of the orthosis for a given patient. As an example, the skin exerts water and gases; if they get collected at the surface of an orthotic or prosthetic appliance, they may cause dermatological problems. Selection of appropriate lining materials with enough perforation properties and special consideration such as silver treatment as well as ease of cleanliness may enhance the quality of final products. Similarly, choosing the right material in active patients or those who will involve in sport activities will decrease the risk of appliance failure and potential damages to the user. Orthotic and prosthetic services will address a wide range of complications including the neurological and musculoskeletal disorders occurring in children, adults and elderly people. Following, a description of the most common materials will be provided.

Materials O&P have always derived advantage from advances in technology and materials. The partial substitution of traditional prosthetic and orthotic materials, such as wood, aluminum, and leather by modern materials, such as thermoplastics and advanced composites, has fostered design innovation and resulted in significant improvements in function, durability and appearance of modern prostheses and orthoses. Despite the development of new materials, traditional ones are still in wide use. Metals, wood, leather, fabrics, thermoplastics, thermosetting composites, foamed plastics, and elastomers are the principle materials commonly used in current orthopedic industry. Most orthoses contain several materials with miscellaneous mechanical properties. The materials may be sewn, riveted, or glued to one another. To understand recommended design and fabrication procedures, it is important to have a basic knowledge of the properties of materials. Some of the material properties are listed and defined in Table 3. In addition to materials’ inherent mechanical properties, certain physical characteristics of the materials also affect the material behavior and appliance function such as thickness and shape of the material used in appliance. This article presents an overview of the common materials used in O&P.

Casting Materials and Fabrication Methods There are a variety of manufacturing technologies to fabricate and obtain an orthopedic appliance. Many individual patient solutions begin with conventional fabrication of plaster casts (negative impression) and the manufacture of positive model. Following, some of the most common used materials for casting are discussed.

Gypsum plaster or plaster of paris (POP) POP has two functions in orthotic and prosthetic field; POP bandage is used to generate a negative cast of body part. POP powder is then blended with water to make a liquid plaster poured into the negative cast to create a positive mold resembling the patient’s limb. Table 3

Materials properties

Properties

Definition

Strength Ductility Stiffness

The ability of a material to sustain the external force before the breakage The ability of a material to undergo permanent change of shape without rupture The amount of deformation (e.g., bending or compression) that happens when the load is applied to a material The ability of a material to recover the original dimension after unloading The resistance of a material to penetration in response to a compression load. The measure of material hardness is referred to as its durometer A material weight per volume The ability of a material to withstand a shock force The ability of a material to resist degradation in response to exposure to chemicals The ability of a material to resist the cyclic loading without breaking

Elasticity (resilience) Hardness Density Toughness Corrosion Durability (fatigue resistance)

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POP bandage consists of calcium sulfate hemihydrate generally on a cotton bandage. Addition of water allows the POP to conform to the limb before the completion of crystallization process and plaster setting. In case of big molds, cast fillers in forms of sand might be mixed into the plaster to lower the positive cast weight.

Fiber glass bandage Fiber glass bandage includes knitted glass fibers impregnated with polyurethane (PUR) resin; similar to POP, Fiberglass bandages need to be immersed in water to be activated. This bandage can be used for light supporting limb casting; however they cannot make the same formability as POP bandages do.

Foam box impression The foam box is considered as an alternative to plaster casting for accurately capturing the plantar surface of patient’s in semiweight bearing impressions. Closed-cell and moisture-proof Phenolic foam (made by curing a foaming phenolic resin) will be served for replication of the foot imprint. This foot shape later will be duplicated by pouring the liquid plaster into the foam box.

Alginate Alginate (a natural polysaccharide) is a powder used for taking impression with good surface details. Mixing with water, it makes a creamy gelatinous form used for detailed cast taking of hands, fingers, feet, and maxillofacial prostheses. It will provide greater accuracy than plaster and remains elastic once cured. Alternative to physical negative casting, Computer-Assisted Design and Computer-Assisted Manufacturing (CAD-CAM) techniques are used. With the technology, an optical or laser scanner captures the surface geometry of the body segment. This digital impression is then rectified virtually and used to direct milling machinery to fabricate a rectified positive model from a carving block of PUR. Once the positive model is created, either via casting and hand fabrication or digital scanning/carving, the device is then fabricated in a conventional manner of vacuum forming or lamination over the model. In case of orthoses with metal uprights, factory-made uprights are individually conformed to body contour and then assembled to custom-made segment which is produced by either vacuum-forming or lamination. In an attempt to eliminate the role of positive cast in the procedure of orthotic fabrication, low temperature plastics have been used for many years to be molded directly on the limb. In some technologies, the concept of positive model as a foundation for fabrication of orthoses and prostheses has been totally replaced by direct manufacture of the product from the scanning digital file of the segment. CNC-controlled milling machines can carve materials such as Ethylene Vinyl Acetate (EVA) and cork with different hardness to create a custom-design foot insole. 3D printing is another technology which possesses the direct production of wide range of orthopedic products.

Leather Leather is fabricated from animal skin and hides processed tanning treatment. The final properties of the leather rely on the tanning procedure and the type of hide used in the fabrication of a product. Strength, stretch, formability, and water vapor permeability are the characteristics of leather which makes it valuable for O&P. Leather is used for components such as suspension straps, belts, and thigh corset in conventional hinged lower limb prostheses. Thinner leather can be used as lining material to cover the metal bands and bars in conventional lower limb orthoses, as well as insoles. Cowhide is served as very strong leather and therefore it can be used for straps and the upper portion of shoes.

Textile Fibers and Fabrics Natural and synthetic polymer fabric materials are commonly-used fabrics in orthoses and prostheses.

Natural Fibers Cotton fiber which is taken from the capsules of the cotton plant is perspiration absorbent and tear-resistant. Wool has excellent resilience and high heat-retaining capacity and can absorb the humidity.

Synthetic Fibers Synthetic Fibers might be polymeric-base or mineral-one. Mineral fibers include fiberglass and Carbon which will be explained later. Synthetic polymer fabrics used in orthoses include polyester and polyamide (Nylon and Aramid). Fabrics are widely used for prosthetic socks, tubular stockinette, straps, lining along with other materials, harnesses for upperlimb prostheses as well as fastening and base material in less rigid orthotic supports such as fabric stabilizing spinal corsets. The greatest use is for prosthetic socks and stockinette. They are commonly made of wool, cotton, or blend of these natural fibers combined with nylon, acrylics or other synthetic materials. Fabrics may incorporate rubber to create an elastic cloth applicable for straps used in harness system of upper limb prostheses.

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A very popular use of nylon is hook and loop (Velcro) which are used for frequent engage in strapping to secure an appliance in place in an easier way to buckles and laces. Webbing straps which are strong flat strips of woven fabric might be made from nylon or polyester. Felt, a nonwoven fabric made from textile fibers such as polyester or wool can create a fine surface, and serve as a filling material in lamination or be used for padding purpose. Nylon fabrics nowadays use as a covers material for neoprene products as well as outer layer of many prosthetic liners. Nylon in the form of cord and cable is applied in lanyard suspension system and body-powered prosthetic control, respectively. Cosmetic nylon stockinette will improve the final appearance of covered modular prosthesis.

Metals Metals are used for manufacturing of prosthetic modular components such as adaptors, clamps, and joints as well as uprights and joints in orthotic designs. To bond the materials, steel, aluminum and copper rivets as well as tubular rivets, buckles and loops (made from nickel plated steel) and brass eyelet are implemented to show a minor but important procedure in constructing of an appliance. However, metals mostly used in the fabrication of orthoses and prostheses can be categorized into three groups: steel, aluminum, and titanium.

Steel Iron is never used in pure state. Any iron-based alloy material is generally called as steel. Stainless steel is a steel alloy that contains nickel and chromium to enhance the resistance of alloy to corrosion and oxidation. Because of properties such as durability, corrosion, fatigue-resistance and high strength, stainless steels are used frequently within O&P. The obvious disadvantage is the weight of product. Stainless steel is widely used in factory-made orthotic and prosthetic joints, support uprights, band material, springs and bearings of lower limb orthoses, shoe attachments (stirrup) in conventional lower limb orthoses and spring bands for corset construction in metal frame spinal orthoses. Also, testing screws which are used for connection of orthoses during the trial phase are made from steel.

Aluminum Aluminum alloyed with copper and manganese is well suited for O&P due to low weight and high resistance to corrosion. The primary alloy used in O&P is aluminum alloy 2024 which provides a shiny well-finished component. Two main types of aluminum include the wrought and cast alloys. Cast aluminum is usually used for prefabricated prosthetic components such as finger assembly in hand prostheses. In case of structural purposes such as prosthetic pylon tubes, orthotic uprights, and upper extremity joints, wrought aluminum alloys are preferred to cast ones. In general, aluminum is used in applications in which the device is subjected to lower stresses. To reduce the fatigue failure at attachment site, normally the component needs to be designed from titanium or stainless steel rather than aluminum.

Titanium Titanium alloys are very strong, lighter in weight than steel, and corrosion resistant. Titanium is used in some prefabricated orthotic and prosthetic components when simultaneous strength and light weight are required. In conjunction with light weight composite shells and plates, metal components of exoskeletons are preferably chosen by titanium alloy to decrease the total weight of device.

Wood Unique characteristics of wood make it an appropriate material in lower limb prosthetics application. It is lightweight and, strong. The wooden keel of solid-ankle-cushion-heel (SACH) prosthetic foot is fabricated of maple and encased in a microcellular PUR shell in shape of normal foot. Cork is the most common wood used in orthoses. It is exceptionally lightweight and resilient, making it appropriate for shoe lifts. Sometimes, cork is mixed with rubber or EVA to achieve greater flexibility, cushioning or thermoformability. Except cork, some other woods might be used for shoe elevation or designing the shoe last. A last is a mechanical form in shape of a human foot. It is used by shoemakers as a working surface on which the leather forms and it is made from hardwoods or some high density plastics.

Synthetic Polymers Polymers are typically classified into three groups: plastics including thermoplastics and thermosets, polyurethanes (PUR) and elastomers. Thermoplastics comprise the majority of available polymers.

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Plastics Development of plastic polymers revolutionized orthotic and prosthetic prescription because these materials offered enhanced mechanical properties such as formability which introduced them as predominant materials in this technology. The growth of the plastic industry has been evolving into today’s routine to sophisticated high performance products. There are two major categories of plastics: thermoplastics and thermosetting plastics. Plastics are usually favored because of their superior appearance, uniform color, and the ease at which they can be casted, molded or extruded. Thermoplastics Thermoplastics are frequently used in O&P. Thermoplastic materials get malleable when heated but will retain the new shape when cooled. The most useful physical property of a thermoplastic is its glass transition temperature at which it begins to soften. Thermoplastics are divided into two groups; low-temperature or high-temperature materials, according to the temperature range at which they become formable, whereas high-temperature materials require higher temperatures and must be molded over a positive plaster replica of the patient’s limb. One advantage of thermoplastic materials is that they can be reheated and reshaped, making possible minor adjustments of an appliance during fittings. Low-temperature thermoplastics, those moldable in a warm water bath, can be molded directly onto the body. The basic ingredient of many thermoplastics (low-temperature) is polycaprolactone which helps to make the plastic highly conforming. In contrary, some others include polyisoprene, a synthetic rubber, Low-temperature thermoplastics such as Orfit are most often reserved for orthotic devices that are designed to provide temporary low-stress support and protections such as short-time spinal, upper extremity and fracture orthoses (Fig. 7). Development of low-temperature knitted textile materials can facilitate the fabrication of finger and other small immobilization orthoses. The production of most of orthopedic appliances is accomplished by application of high-temperature thermoplastics. Some of the commonly used materials comprise polyolefins (including polyethylene and polypropylene), acrylic, polycarbonate, acrylonitrile butadiene styrene (ABS), polyamide, saturated polyester, vinyls and acetals) .In spite of versatility of thermoplastic sheets, polyolefins are still considered as most widely used plastics. As a matter of fact, there is no unique plastic which can meet all the requirements. Therefore, when selecting a sheet of plastic, several factors should be kept into consideration including desired finished thickness, shrinkage and rigidity. The characteristics of some of the commonly-used thermoplastics and their typical orthotic and prosthetic application are summarized in Table 4. Thermoforming To fabricate the components of a device such as prosthetic socket and AFO, most of the orthotic and prosthetic laboratories take the advantage of a common technique called thermoforming (drape forming). In this technique, a heated thermoplastic sheet of plastic is drawn on a positive mold; the technique is completed once the plastic sheet is vacuumed and conforms to the shape of the mold (Fig. 8).

Fig. 7

(Top) Low temperature thermoplastics to design different orthoses for upper extremity, (bottom) application of a WHO.

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Thermoplastics used in orthotic and prosthetic fabrication

Generic name

Characteristics

Application

Polyolefins Low-density Polyethylene (LDPE)

• Good toughness and flexibility

Polyolefins High-density Polyethylene (HDPE)

• Thermoforming • Increased rigidity and tensile strength • Good fatigue and wear resistance

• Spinal and upper extremity orthosis • Arch support • Flexible prosthetic sockets • Spinal orthoses • Lower extremity cuff • Resting passive upper limbs orthosis such as

Polyolefins

• Durable and stiff but more difficult to form • Decreased impact strength • Maximum rigidity among olefins • Polypropylene/polyethylene copolymer • Very good thermoforming • Rigid yet flexible • Offers more durability than PP homopolymer • Good endurance to million cycles of repetitive flexes • Resilient • Lightweight • Available in various durometers • Moldable • Copolymer of ethylene with 14%–40% vinyl acetate • Light weight • Excellent shock absorbency • Increased flexibility compared to LDPE • Denser and stronger than PE foam alone • Good shock absorbing capabilities • Moldable • Transparent • High stiffness • Easy to bond to alignment component by using

• Applications require greater strength • Upper and lower extremity • Spinal orthotics • Upper and lower extremity • Nonarticulated dynamic AFOS • Spinal orthotics

• Rubbery thermoplastic • Formable • A more rigid thermoplastic • Good draping (molding) • Semi rigid plastic

• Upper limb orthoses with low force load

• Polypropylene Homopolymer (PP H) Polyolefins • Polypropylene Copolymer (PPC) Polyethylene (PE) foam

Ethylene-vinyl acetate (EVA)

PE foam plus EVA Acrylic PMMA (polymethyl methacrylate) Low-temperature thermoplastic Polyisoprene-based (ezoform) Low-temperature thermoplastic Polycaprolactone–based (Orfit) Polyamides

• Nylon • Aramids (Kevlar)

acrylic resin

WHOs

• Padding and lining straps • Shock absorbing accommodative FOs • Cosmetic cover in endoskeletal prostheses • In foam shape for padding and cushioning in footwear and shoes

• Semi rigid foot orthosis • Posting material • Check socket • Ocular prosthesis

• Upper limb orthoses • Trunk orthoses • Orthotic joints, which are lighter in weight than metal ones

• Reinforcement fiber • In family of saturated (thermoplastic) polyesters • Transparent • Superior impact strength and resistance over acrylic

• Prosthetic check socket • Face mask for sport and burns

Polyacetal (POM ¼ polyoxymethylene)

• Appropriate for vacuum forming • Tough • Chemical resistant • High strength and stiffness • Low coefficient of friction

• Upper limb orthoses • Spine and cervical orthoses • Components in orthopedic devices that require

Acrylonitrile butadiene styrene (ABS)

• A modified polystyrene with improved impact

Polyethylene terephthalate Glycol (PETG)

thermoplastic

PVC (polyvinyl chloride)

Polycarbonate (PC) Ionomer (surlyn)

strength

• Bondable • Transparent • High durability and impact resistance • Copolymer of ethylene and methacrylic acid • Ionically cross-linked • Very high tensile strength • Transparent • Tough and durable • Soft • Excellent thermoforming

a great deal of stiffness such as plastic screws and nuts • Spinal body supports • Seat inserts

• Check socket • Upper extremity orthosis • Spinal orthosis • Partially flexible inner prosthetic socket • Check sockets • Face mask for sport and burns • Cranial helmet (Continued)

362 Table 4

Biomaterials: Biomaterial Applications and Advanced Medical Technologies j Materials for Exoskeletal Orthotic Thermoplastics used in orthotic and prosthetic fabricationdcont'd

Generic name

Characteristics

Application

Polyvinyl chloride (PVC)

• Good stretchability • Increased perspiration

• Cosmetic glove • Top covering material (impregnated with cloth

Polyvinyl alcohol (PVA) Polyether ether ketone (PEEK)

• Soluble in water and alcohol • Rod form • Tough • Spring function • Composite of acrylic and polyvinyl chloride (PVC) • Reinforcing material • Thermoforming • Outstanding toughness

• PVA bags for lamination • Dynamic AFOS

Kydex

Fig. 8

backing)

• Supporting material in Cervical collars (neck orthosis)

• Spinal orthotic such as TLSO body jackets

Cranial remodeling orthoses and transparent face masks made by thermoforming techniques.

Thermosetting materials The cured thermosettings degrade rather than melt upon heating. Thermosetting plastics (Resins) are plastics that are applied over a positive model in liquid form and then chemically cured to maintain the shape of underlying contour. Thermosets are often impregnated into various fabrics by a process of lamination. Epoxy resin (polyepoxides), unsaturated polyester resin and acrylic resins are considered as thermosetting polymers. Epoxy has the highest performance of all thermosetting plastics as the cured resin represents high strength in tension and compression. Acrylic resin is available in different blends each having their own characteristics. Examples include a standard blend of 80% rigid and 20% flexible resin to be used for vacuum laminations. The skin color pigment is an active plastics softener, and as a result it should not be mixed more than 2%–3% to the lamination resin. Due to the transparency and biocompatibility, acrylic resin also can be used in fabrication of Prosthetic eye (Ocular prosthesis) as well as finger nail and toe nail in partial hand and foot prostheses. Polyester resin does not bond well to the fiber. Due to the poor rinsing properties of this resin, it is not recommended for high performance composite manufacture. Forearm segment in above elbow prosthesis is usually a prefabricated component which is laminated with polyester resin while encompassing the steel side bars of a mechanical elbow joint. Composites Fiber-reinforced plastics (FRPs) also called composites, have reformed the O&P by producing a material with enhanced strength and quality yet capable of tolerating compressive and flexural stresses. The most common fiber reinforcements used in polymer composites are fiberglass, carbon, and Kevlar aramid fibers. In general, the volume of fibers should be higher than the volume of resin matrix. Carbon fiber reinforced plastics (CFRP) are widely used in

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orthoses and prostheses including energy storing keels in prosthetic feet, external supporting frame of hydraulic cylinder in prosthetic knee units, pylon tubes in heavy duty cases and orthotic uprights. According to degree of graphitization, this fiber may manifest high strength properties while also being lightweight. Carbon fiber has superior stiffness properties in both compression and tension but, it has relatively low impact strength. In spite of concern about their biocompatibility, aramid fibers have a very high tensile strength. Kevlar fibers might be used in some prosthetic feet as a reinforcement structure to protect the PUR foam from fatigue. Currently, there are two techniques for processing FRP composites in orthotic and prosthetic laboratories namely hand layup with vacuum impregnation-consolidation, and heat curing of resin-impregnated fabrics (Pre-preg). The most common method called vacuum bag lamination includes a hand layup of tubular nylon or cotton stockinette fabrics and reinforcement fibers in shape of cloth, braid or roving placed on the positive mold. This complex is sealed between two layers of polyvinyl alcohol (PVA) plastic bags which can accommodate the wet liquid resin to be poured into the enclosure while the applied vacuum removes the entrapped air from the laminated parts, pressing the resin through the fibers and create a thin-walled structure (Fig. 9). Preimpregnated (Pre-preg) materials are composites of one of those three reinforcement fibers as a foundation material impregnated with a predetermined amount of resins, preferably epoxy. Pre-pregs are easier to work with than current wet-lamination technique while light weight device can attenuate the forces, successfully. These materials are best option for the manufacture of orthoses in a frame construction such as rigid knee orthoses. As a limitation, they cannot be stored in room or higher temperature for long time. Recently, there is an interest to combine the strength and energy storage of carbon fiber with the workability and versatility of the thermoplastics by extruding the carbon fiber or glass fiber compounds inside the thermoplastics. Nylon, polypropylene and PETG are the thermoplastics that have been used in large volume. For partial reinforcement in fabrication of orthoses these fiber reinforced thermoplastic profiles can be formed and bonded to the underlying thermoplastic for enhancement of structure strength. Expanding foams (foamed plastics) Both thermoplastic and thermosetting resins can be used to produce a rigid, semi-rigid, or flexible foam structure. Expanding foams can be used both in orthotic and prosthetic devices, as an interface to protect the skin from the appliance. Bubble gases like hydrogen or nitrogen can be forced into the plastic solid matrix to create resilient materials. These foams are categorized into two groups: open and closed-cell. The cellular structure can reduce the shear forces. In open-cell foam, the cells are interlinked while in closed-cell foam, the cells are completely isolated from each other. Because are liquid-proof, the absorption ratio of body perspiration or odor in closed-cell foams is significantly limited since they are liquid-proof. These closed-cell polyethylene (PE) foams are available in a wide array of durometer hardness PE foams (e.g., Plastazote, Pe-lite) can be used in the manufacture of thermoforming soft socket in lower limb prostheses as well as padding in orthoses and cervical collars. The base material or top-layer cover of accommodative foot orthoses which are highly prescribed for patients with neuropathies could be made from materials such as EVA, neoprene rubber, soft PUR foams, and low density PE foams to provide comfort and shock attenuation. Sometimes the foamed plastic benefit from some treatments such as a layer of silver bonded to material to possess an antimicrobial and antifungal effect. The sandwich foams of PE-EVA with different hardness might be helpful in some patients to absorb the shock while providing a more resilient combination.

Fig. 9

Materials for vacuum bag lamination.

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3D knitted spacer fabrics can be used for padding since they have good compression properties. They might be combined of polyester and nylon in different combinations. Good air permeability and compression behavior are unique properties propose these fabrics as good alternatives for padding.

Elastomers Elastomer comprise a large family of elastic polymers which can snap back to approximately original size and form once the load causing the deformation is released. Because of their properties to absorb and dissipate the loads, many of elastomers are used in shock-absorbing shoe inserts. The classification of four main groups and subgroups of elastomers (e.g., natural and synthetic rubber, PUR, silicone, and thermoplastic elastomers (TPE) has been presented in Table 5. Polyurethane (PUR) The mixture of polyol (specific alcohol) and isocyanate will result in an instant expansion of a foam or rubber-like product called PUR which can be found in block or sheet form in versatile stiffness (flexible to rigid). While soft PUR elastomers are primarily used in body cushioning such as insoles, cosmetic cover in endoskeletal prostheses and prosthetic liners, rigid PUR foams are especially suitable for load parts of the orthopedic technology range (e.g., length difference compensating insoles, shank in exoskeletal prosthetic design and carving blocks in milling machines). PUR elastomers with moderate durometer are generally used to encapsulate the keel of foot unit through a process of injection molding. Also, outer sole of the shoe and flexural ankle joint in AFO can be made from. Thermoplastic elastomers (TPE) TPEs are a class of copolymers which show the advantageous properties of both elasticity and easy processability. The use of TPE in recent decades has dramatically increased. As an example, TPE gels enriched with vitamins and mineral oil can be ideal option for use in cushioning and shock absorption products such as prosthetic gel liners, insoles, heel cups and hypertrophic scar management while they can continuously moisturize the skin. Rubber In general, rubber has considerable elasticity, shock absorbency, and toughness. Synthetic rubber has more resistant to corrosion than natural rubber. Rubber strands may be woven with cotton or other fabrics to create elastic straps. Polychloroprene also called neoprene is available in various densities, making the low-durometer versions suitable as lining material for orthoses, whereas the firmer samples are used for sport and support orthoses. The nylon (polyamide)-covered neoprene acts as a shock absorber while also reducing friction on the foot’s plantar surface. Silicone Dimethyl polysiloxanes is the largest group of commercial silicone polymers used in medicine (including prosthetics). Silicone may form the following polymers: silicone oils (for decreasing the viscosity of silicones), silicone resins (liquid silicone), and silicone rubber (high consistency solid silicone). Depending on whether the vulcanizing process uses heat or not, silicones are available as heat vulcanized (HTV) or room temperature vulcanized (RTV). In general, HTV Silicones show higher strength.

Table 5

Classification of elastomers

Name

Name in details

Comment/application

Natural Rubber Synthetic Rubber

Latex Polyisoprene rubber (IR) Polybutadiene (BR) Styrene-butadiene rubber (SBR) Polychloroprene (CR) also called neoprene

Obtained from the tree Synthetic version of Natural rubber Mostly used in tire production Most important sort of synthetic rubber First synthetic rubber maintain flexibility in a wide temperature range

• Base material of contact

Polyurethane (PUR)

Polyester-urethane Polyether-urethane

Better mechanical properties Better cushioning

Silicone rubber

Dimethyl polysiloxanes

Thermoplastic elastomer (TPE)

Styrene based (TPE-S) Olefin based (TPE-O) Urethane based (TPE-U)

• Cosmetic prostheses • Gel liner • Shoe outsole and footwear • Prosthetic Gel liners

• Insoles • PUR shoe outsoles • Flexural ankle joints • Prosthetic liner

adhesives

• Shock attenuation padding insole

• Rubber shoe outsole • Rubber band in some WHOs and hooks

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Fig. 10

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(Top and bottom) Silicone hands including artistic designs.

Silicone cushioning and elasticity provide excellent wearing comfort, regardless of whether they are used in a socket for prostheses or an orthosis (Fig. 10). The wide range of medical-grade silicones with different degree of stiffness possesses a variety of options to make appliances such as: B B B B B B B B B

End bearing/cushioning pads Cosmetic finger and partial hand prosthetics Face mask for scar compression treatment Cosmetic and functional partial feet Cosmetic glove for upper limb prosthesis External mammary prostheses Maxillofacial prostheses Foot orthotics Socket liners

To facilitate the donning and doffing of liners, different techniques have been implemented in industry. Poly-para-xylylene (Parylene) is a polymer used to create an ultra-smooth surface coating with low coefficient of friction. It bonds to the outer surface of silicone liner at molecular level while this process does not affect the elasticity of substrate silicone. Fabric to gel bonding is another approach to improve the sliding of the gel surface while the fabric cover is in contact with the socket.

Adhesives The adhesives used for orthopedics applications must fulfill a maximum amount of adhesion power, decrease curing speed and be compatibility with a greater range of materials for surface bonding. Examples of the most common glues used in orthopedic and footwear industry include one-part, rubber-based contact adhesive (e.g., polychloroprene, styrene-butadiene rubber) and PURbased contact adhesive. These materials are mostly air drying as the solvent will be evaporated to provide excellent final bond strength.

Surface Finishing Surface coating will improve the acceptance of orthoses and prostheses to the patient. Water transfer printing (hydrographics) and thermopapers are two techniques to match the device with patient’s expectations. Water transfer painting is a coating technology to print and decorate any design on orthopedic devices by using a printed PVA hydrographic film which is placed on the surface of water. Due to the surface tension of water, floating ink pattern will curve and bond around the component. Colorful transfer papers are widely used to cover the thermoplastic material with a specific pattern. Usually an ink-jet printer is used to print an image or pattern on the paper. Then a heat press can transfer image onto the plastic sheets. The final product might be more appealing to the patient and especially in case of pediatric devices, the compliance might be improved.

3D Printing 3D printing, as a process of creating solid objects from digital files has become the medium of new technological revolution in the medical science, including O&P. Due to fabrication ability of customized sophisticated devices, 3D printing has the potential to

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facilitate the design of the orthopedic devices. At the moment this technology is being implemented to fabricate the prosthetic sockets, hand prostheses, AFOs and spinal orthoses. There is a wide variety of material types that are supplied in different states (powder or filament forms) to address the 3D printing technologies including the selective laser sintering (SLS) and fused deposition modeling (FDM) processes. ABS, nylon, and composite form of materials with improved mechanical stiffness such as blend of carbon fiber and polylactic acid (PLA), a biodegradable thermoplastic polyester derived from renewable resources, carbon fiber filled PETG and glass filled nylon are some examples of those materials nowadays used to print the orthopedic devices. There is a potential to use the metal powder of Stainless steel and Titanium for 3D printing of Orthotic and prosthetic components, as well.

Summary and Possible Future Direction There is no doubt that O&P field has developed dramatically and technology advancement has offered outstanding improvement in application of materials. A continually increasing variety of new materials is being used for orthotic and prosthetic fabrication. Using new materials have largely supplanted the traditional methods using leather and steel. Application of Thermoplastics, composites and improved metal alloys in orthopedic devices has offered progress in terms of weight, strength and energy return to customize for individual patient. Materials with better physical properties will be emerging from plastic industry. We are passing from PP, PE and copolymers, wet lamination of sockets and pre-preg sheets to now extruding the plastic with carbon fiber in it and in this way, technology is going to change the field. Next generation of composites might be more evolutionary such as placement of piezoelectric fibers in laminated layers of appliance. Since the piezo crystals can generate electricity when subjected to physical stress, there would be a potential to implement them as a mechanism for battery recharging during walking with prostheses or a method to change the stiffness of a component when electrified. Elastic actuators such as pneumatic artificial muscle (PAM) are used in the field of exoskeletons due to their ability to generate linear forces and motions with a simple mechanism, while remaining lightweight. In most cases, these actuators are composed of elastomeric cylindrical bladders constrained by flexible but inextensible mechanisms such as meshes and embedded fibers. In close future, artificial muscles might be made of fabrics to be integrated into clothing to aid mobility. By using fabrics and textiles, soft wearable exosuits will probably provide a more comfortable interface without any mechanical constrain on wearer’s joints as the hard exoskeletal joints do. Growing experience with osseous integrated prostheses suggest that some patients may achieve greater comfort and function with direct attachment of prosthetic adaptor to implanted metal shaft that protrudes from the limb and by doing so, application of prosthetic sockets may gradually decrease. As a result, in future, the investment might lead to further development of active prosthetic components with new materials and enhancement of control systems including neural control. 3D printing revolutionizes the fabrication techniques of a product and moves the industry forward by making the orthopedic devices more accurate, innovative and improved appearance. There is a very good potential for the application of 3D printing in fabrication of different orthotic and prosthetic components as well as exoskeletons. The concept of natural fibers is being tested to low-cost orthotic and prosthetic devices. Application of composite materials for 3D printing can improve the strength of the final product significantly without affecting the flexibility. Development of recycling and long-lasting materials as well as decreasing the total time of printing will further enhance the feasibility of this technology to serve in orthopedic devices. In summary, the final choice for material selection will rely on local availability of materials, technical expertise and preference of orthotist/prosthetist, expected function and patient’s previous experience; However the interplay of materials science and processing techniques will reveal new insights, and therefore consistent development and enhancement of materials guarantees the achievements of dreams and forthcoming expectations in O&P industry.

Further Reading Coppard, B. M., & Lohman, H. (2014). Introduction to orthotics: A clinical reasoning and problem-solving approach. ElsevierdHealth Sciences Division. Faustini, M. C., Neptune, R. R., Crawford, R. H., & Stanhope, S. J. (2008). Manufacture of passive dynamic ankle–foot orthoses using selective laser sintering. IEEE Transactions on Biomedical Engineering, 55, 2. Herr, H. (2009). Exoskeletons and orthoses: classification, design challenges and future directions. Journal of Neuroengineering and Rehabilitation, 6, 21. Hsu, J. D., Michael, J., & Fisk, J. (2008). AAOS atlas of orthoses and assistive devices E-book. Elsevier Health Sciences. Jacobs, M. L. A., & Austin, N. M. (2013). Orthotic intervention for the hand and upper extremity: Splinting principles and process, 2nd edn. Lippincott Williams & Wilkins. Krajbich, J. I., Pinzur, M. S., Potter, B. K., & Stevens, P. M. (2016). Atlas of amputations and limb deficiencies (4th edn.). American Academy of Orthopaedic Surgeons. Lusardi, M. M., Jorge, M., & Nielsen, C. C. (2013). Orthotics and prosthetics in rehabilitationdE-book. Elsevier Health Sciences. Walbran, M., Turner, K., & McDaid, A. J. (2016). Customized 3D printed ankle-foot orthosis with adaptable carbon fibre composite spring joint. Cogent Engineering, 3, 1227022. Wise, D. L., Trantolo, D. J., Altobelli, D. E., Yaszemski, M. J., & Gresser, J. D. (2013). Human biomaterials applications. New York: Humana Press.

Biomaterials: Biomaterial Applications and Advanced Medical Technologies j Materials for Exoskeletal Orthotic

List of Relevant Websites http://enablingthefuture.org/d“Enabling the future”. https://www.orfit.com/d“Orfit”. https://www.ossur.com/d“Ossur”. http://www.ottobock.com/en/d“Ottobock”. https://wyss.harvard.edu/technology/soft-exosuit/d“Wyss Institute”.

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Microfluidics for Biomedical Applications Shiyu Cheng*, Beijing Engineering Research Center for BioNanotechnology and CAS Key Laboratory for Biomedical Effects of Nanomaterials and Nanosafety, CAS Center for Excellence in Nanoscience, National Center for NanoScience and Technology, Beijing, P. R. China Jinqi Deng*, Beijing Engineering Research Center for BioNanotechnology and CAS Key Laboratory for Biomedical Effects of Nanomaterials and Nanosafety, CAS Center for Excellence in Nanoscience, National Center for NanoScience and Technology, Beijing, P. R. China; and Sino-Danish College, University of Chinese Academy of Sciences, Beijing, P. R. China Wenfu Zheng, Beijing Engineering Research Center for BioNanotechnology and CAS Key Laboratory for Biomedical Effects of Nanomaterials and Nanosafety, CAS Center for Excellence in Nanoscience, National Center for NanoScience and Technology, Beijing, P. R. China Xingyu Jiang, Beijing Engineering Research Center for BioNanotechnology and CAS Key Laboratory for Biomedical Effects of Nanomaterials and Nanosafety, CAS Center for Excellence in Nanoscience, National Center for NanoScience and Technology, Beijing, P. R. China; and Sino-Danish College, University of Chinese Academy of Sciences, Beijing, P. R. China © 2019 Elsevier Inc. All rights reserved.

Introduction Microfluidics for Manipulating Cells Microfluidics for Cells Adhesion on Surfaces Adhesion of single type of cells Adhesion of multiple types of cells Dynamic control of cell adhesion Microfluidics for Cells Migration on Surfaces Topographic properties Shape of cells Cell–cell interactions Wound healing Microfluidics for 3D Patterning Spatial control Soluble molecular gradients Microfluidic simulation of mechanical stimuli Microfluidic-Based Biochemical Analysis Nucleic Acid Detection Protein Analysis Barcode-Based Multiplex Analysis Cell Analysis Further Reading

369 369 369 369 369 369 369 370 371 372 372 372 372 375 375 375 376 378 379 381 383

Glossary Antibodies Proteins that are produced by plasma cells and can bind tightly to their targets, which are called antigens. Cell adhesion Cells interact and attach to a surface, substrate, or another cell by interactions between molecules of cell surfaces. Cell migration Movement of cells in particular directions to specific locations in response to specific external signals, including chemical and mechanical signals. Cell patterning A process to position cells on a substrate with defined spatial selection and in turn to keep stable cell activities on the substrate. Circulating tumor cells Tumor cells that shed from primary lesions and circulate in the bloodstream. The existence of circulating tumor cells can cause tumor recurrence and cancer metastasis. Loop-mediated isothermal amplification An isothermal method for nucleic acids amplification. Microfluidics Precise control and manipulation of fluids that are geometrically constrained to a sub-millimeter scale with specific nano-/micro-structures.

*

These authors contribute equally to this work.

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Encyclopedia of Biomedical Engineering, Volume 1

https://doi.org/10.1016/B978-0-12-801238-3.11029-3

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Introduction Microfluidics typically refers to the precise control of minute amounts of fluids with micro-scale structures. Soft lithographic techniques, based on self-assembly and replica molding, enable the micro-/nano-fabrication and advance the field of microfluidics in a convenient, effective, and low-cost manner. Microcontact printing (mCP) is one of the key soft lithographic techniques, a stamp with patterned relief structures on its surface could generate patterns with feature sizes ranging from 30 nm to 100 mm. Based on soft lithography, several representative advances emerged and have become useful for biomedical applications. In this article, we will introduce recent developments in microfluidics for manipulating cells and microfluidic-based biochemical analysis. We will focus on microfluidic interfacial methods and their extensive applications in fundamental biological research.

Microfluidics for Manipulating Cells Precise control of cell behaviors has been increasingly useful with the development of cell biology and regenerative medicine. The polydimethylsiloxane (PDMS)-based microfluidic chips, a transparent soft material with good oxygen permeability and high level of compatibility, could enable suitable micro-environment for cell culture and real-time observation, thus comprising a great tool for manipulating cells and studying cell-cell interactions.

Microfluidics for Cells Adhesion on Surfaces Most cells require adhesion to solid substrates to carry out biological activities. Mammalian cells could obtain good attachment on surfaces only if suitable extracellular matrix (ECM) proteins are present. ECM, populated by extensive proteins and other molecules, is essential for the regulation of cellular and organismal physiology. It provides a variety of physical and chemical cues to regulate cell behaviors. In particular, physical cues such as topography and stiffness provide mechanical support to cells while chemical cues such as cytokines, ionic strength, and pH values provide chemical stimulations for cells. By employing microfluidic chips, it is straightforward to realize the specific modification of the surface and precisely pattern cells in microchannels with micrometer precision. We will introduce methods for controlling cell adhesion in a confined area, releasing cells from the surface, and dynamic control of cell adhesion.

Adhesion of single type of cells We could control cell adhesion in a specific area by exerting self-assembled monolayers (SAMs) and electrochemical desorption. SAMs are one of the most intensively studied methods to modify the surface. SAMs allow us to control the properties of a surface on the molecular scale with the adsorption of u-substituted alkanethiols on the surface of gold (Au) and palladium (Pd). SAMs can be easily prepared by immersion of a substrate in the solution containing a ligand (Y(CH2)nX) reactive toward the surface, or by exposure of the substrate to the vapor of a reactive species. The thickness of a SAM can be controlled by changing the number (n) of methylene groups in the alkyl chain. EG3-terminated SAMs could resist the absorption of proteins, and thus resist cells attachment and spreading. We confined single human umbilical artery endothelial cells to different micropatterns with SAMs on Au and Pd (Fig. 1A). We could also easily remove the confinement by electrochemical desorption of the (EG)3-terminated SAM. We confined bovine capillary endothelial (BCE) cells into patterns by using mCP and thiols-HS(CH2)11(OCH2OCH2)3OH (C11EG3) and HS(CH2)17CH3(C18). By exerting a cathodic voltage pulse (
1.2 V for 30 s), BCE cells can attach to, and spread across previously protein-inert areas and release patterned cells from the constraints of these patterns (Fig. 1B).

Adhesion of multiple types of cells Situations would be more sophisticated for real tissues, which are usually composed of more than one type of cells. Combining SAMs and microfluidic channels, we could easily confine two or more types of cells to specific locations on surfaces without physical constraints and we could also control the motility of these different types of cells. We confined NIH 3T3 cells and Hela cells in the middle and two sides of stripes (Fig. 1C). Cells moved freely when exerting a cathodic potential of 1.2 V for 30 s (Fig. 1C). This approach is extensively applicable for studying fundamental biomedical problems based on cell–cell interactions.

Dynamic control of cell adhesion Dynamic control of cell adhesion on surfaces allows us to control the mobility of a single cell, or a group of cells at will. We realized reversible cell adhesion using azobenzene-embedded self-assembled monolayers. The azobenzene moiety can be converted photochemically between E and Z configuration to either present or mask RGD ligand and hence modulate cell adhesion (Fig. 1D). The two states of the surface are reversible at different wavelengths of light by the mercury lamp of a standard fluorescence microscope.

Microfluidics for Cells Migration on Surfaces Cell migration plays a fundamental role in many physiological activities, such as embryogenesis, morphogenesis, wound healing, and disease progression. The migration of mammalian cells typically includes several steps: (1) morphological polarization; (2) extension toward the direction of motility; (3) attachments between the leading membranes and the substrates; (4) movement

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Fig. 1 Microfluidics/surface engineering for cells adhesion. (A) Single human umbilical artery endothelial cells with different micropatterns confined by SAMs on Au (left) and Pd (right). Fibronectin, actin, and the nucleus were stained by anti-fibronectin (green), phalloidin-Texas red (red), and DAPI (blue), respectively. (B) BCE cells were allowed to attach to a surface via SAMs. Application of a cathodic voltage pulse released the cells from the microislands. The numbers indicate the time elapsed (in minutes) after the voltage pulse. (C) Top: Phase-contrast and fluorescence micrographs of NIH 3T3 cells and Hela cells patterned on a substrate modified with EG6 by using selective desorption of SAMs. Hela and NIH 3T3 were stained with celltrace calcein green AM and celltrace calcein redorange AM, respectively. Bottom: Application of a cathodic potential on the substrates allowed all cells to move freely on the surface. The numbers indicate the time (in hours) after application of the pulse. (D) Upper panel: The azobenzene moiety can be converted photochemically between the E and Z configurations to either present or mask the RGD ligand. Lower panel: Cells adhered onto SAMs with the azobenzene group in the E configuration. Few cells adhered to the same SAMs with azobenzene in the Z configuration. Cells adhered to the SAMs again when the conformation of azobenzene changed from Z to E. Scale bars are 200 mm. (A) Jiang, X., Bruzewicz, D. A., Thant, M. M., and Whitesides, G. M. (2004). Palladium as a substrate for self-assembled monolayers used in biotechnology. Analytical Chemistry 76, 6116–6121, with permission from American Chemical Society. (B) Jiang, X., Ferrigno, R., Mrksich, M., Whitesides, G. M. (2003). Electrochemical desorption of self-assembled monolayers noninvasively releases patterned cells from geometrical confinements. Journal of American Chemical Society 125, 2366– 2367, with permission from American Chemical Society. (C) Li, Y., Yuan, B., Ji, H., et al. (2007). A method for patterning multiple types of cells by using electrochemical desorption of self-assembled monolayers within microfluidic channels. Angewandte Chemie International Edition 46, 1094–1096, with permission from Wiley-VCH. (D) Liu, D., Xie, Y., Shao, H., Jiang, X. (2009). Using azobenzene-embedded self-assembled monolayers to photochemically control cell adhesion reversibly. Angewandte Chemie International Edition 48, 4406–4408, with permission from Wiley-VCH.

of the bulk of cell body; (5) release the attachment at the sharp end. These processes are influenced by substrate characteristics such as topographic properties, the shape of the cells, and cell–cell interactions.

Topographic properties Most types of cells can polarize and move directionally with topographic stimulus. Researchers studied the role of groove/elevation (ridge) dimensions at the micrometer scale on fibroblast cell migration by correlating cell shape, migration angle, cell orientation and velocity with these dimensions. They found that surface structures significantly influenced migration direction, cell orientation and mean velocity, as well as migration speed in the directions parallel and perpendicular to the grooves/elevations in a surface structure dependent way. Cell migration velocity parallel and perpendicular to the structures was significantly affected by the geometries and dimensions of the substrates. To study influences of the topographic properties, we fabricated microgroove and microgroove-composed structures by applying electrospun (ES) fibers as the template and replica molding the fibers with

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PDMS. We combined microfluidics with the replica-molded substrate from ES fibers to guide neurites and control the morphology of neurons (Fig. 2A).

Shape of cells We identified that apart from the environmental stimulus, the shape of a cell could also determine the direction of mobility. We designed a teardrop-like asymmetric pattern and restricted cells by employing SAMs, then released the constraint on the shape and assessed the direction of motility for individual cells. The cells tend to move toward blunt ends after release from the defined area (Fig. 2B).

Fig. 2 Microfluidics/surface engineering for cells migration. (A) Integration of microfluidic channels with replica mold of aligned ES fibers for guiding neurites. The green fluorescence was neuron-specific marker Tuj1. (B) Left: Migration of a typical mammalian cell on a flat surface. Right: Time-lapse images (in minutes) show the motility of an initially polarized 3T3 fibroblast/COS-7 after the constraint is released. (C) Left: The strategy for patterning multiple types of cells on the same substrate that allows three types of naturally occurring cell–cell interactions. Right: time-lapse phase fluorescence micrographs for the three types of cell–cell interactions between 3T6 and NIH3T3 cells. (D) Patterning different types of cells in a matrix with (left) or without (right) wound at the “wound edge” according to different properties of the surface of PDMS membrane. (E) Microfluidic chip for investigating the neuron-cancer cell interaction in vitro. Neurons, cancer cells and nuclei are stained with Tuj1 (green), panCK (red) and DAPI (blue), respectively. (A) Liu, Y., Sun, Y., Yan, H., et al. (2012). Electrospun fiber template for replica molding of microtopographical neural growth guidance. Small 5, 676–681, with permission from Wiley-VCH. (B) Jiang, X., Bruzewicz, D. A., Wong, A. P., Piel M., and George M. Whitesides G. M. (2005). Directing cell migration with asymmetric micropatterns. Proceedings of the National Academy of Sciences USA 102, 975–978, with permission from National Academy of Sciences. (C) Chen, Z., Li, Y., Liu, W., et al. (2009). Patterning mammalian cells for modeling three types of naturally occurring cell-cell interactions. Angewandte Chemie International Edition 48, 8303–8305, with permission from Wiley-VCH. (D) Yuan, B., Li, Y., Wang, D., et al. (2010). A general approach for patterning multiple types of cells using holey PDMS membranes and microfluidic channels. Advanced Functional Materials 20, 3715–3720, with permission from Wiley-VCH. (E) Lei, Y., Li, J., Wang, N., et al. (2016). An on-chip model for investigating the interaction between neurons and cancer cells. Integrative Biology 8, 359–367, with permission from Royal Society of Chemistry.

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Cell–cell interactions Apart from the physical and chemical cues, cell–cell interactions also influence cell behavior. Microfluidic system is a great tool for simulation of cell–cells interaction in vitro by selectively modifying the surface on the same substrates in micrometer-scale precision. In vivo cell interactions between different type of cells can be divided into three types: (1) both types of cells are immobilized and confined to isolated areas, such as epithelial cells and fibroblasts during ovarian development; (2) one type is immobile and another moves freely, for example, glial cells and neurons during the development of nervous system; (3) both types of cells could move freely, for instance, hepatocytes and fibroblasts in the liver. We used an alkanethiol and a kind of ECM fibronectin to resist and promote cells adhesion, respectively, thus created areas with different surface properties. We transported different types of cells to designed locations on the same surface by employing microchannels and realized patterning multiple types of cells on the same substrate to simulate three types of cell–cell interactions (Fig. 2C). Patterning different types of cells on diverse substrates is crucial for studies of many processes in vitro such as cancer formation, angiogenesis, and metastasis. We circumvented the substrates limitations of previous patterning techniques by employing a thin PDMS membranes as the physical barrier to achieve desired surface properties and to deliver cells. We patterned three types of cells: NIH 3T3 cells, Mardin Darby canine kidney (MDCK) cells, and JEC cells on substrates with micro-scaled grooves (Fig. 2D). After cells adhesion, we removed the microfluidic channels and cells were free to migrate under the influence of each other and eventually contacted each other. We patterned MDCK and NIH 3T3 cells onto PDMS substrates with microgrooves. Although MDCK cells migrate much more slowly than 3T3 on flat substrate, the velocity of migration of MDCK parallel to the grooves is higher than that of 3T3 perpendicular to the grooves. This experiment shows that both cell–cell and cell–substrate interactions simultaneously influent cell group behaviors. For cancer development, we designed a microfluidic compartmentalized device to simulate the interaction between neurons and cancer cells (Fig. 2E) and demonstrated that nerves provided biophysical support for cancer cells and guidance for their directional migration, which was consistent with clinical metastatic behaviors of multiple types of cancers.

Wound healing Wound healing is a complex process, including hemostasis, inflammation, proliferation, and remodeling, which is involved in multicellular activities. Microfluidics is an important tool to study these activities in vitro. For wound healing, we established a microchip model comprising three parallel microfluidic channels for co-culture and wounding assay to study epithelial collective migration. To mimic the scenario of wound tissue, we filled distilled water to the middle channel, causing wounded and normal epithelial cells co-exist on the same substrate. We co-cultured MDCK and NIH 3T3 cells and confirmed the lysed epithelial cells and fibroblasts could enhance epithelial collective migration without the influence of free physical space (Fig. 3A). The microfluidic chip also allows us to in vitro screen wound dressing candidates that can minimize the use of animals for developing better method for wound care. We also employed a cell-on-a-chip model to simulate the cutaneous wound and to in vitro screen the performances of several ES fibrous wound dressings in enhancing wound healing. NIH-3T3 cells and MDCK cells were patterned on the Petri dish using the microchip and then covered by the PCL/gelatin ES mats and bacterial cellulose (BC) mats (Fig. 3B). After peeling off the covered mats, we could observe the migration of cell patterns and measure the distance of cell patterns to evaluate the effect of different wound dressing on cell migration.

Microfluidics for 3D Patterning Patterning cells on 2D substrates is significant in fundamental cell biology research, but is limited in mimicking the physiological structure of tissues in vivo. 3D cell culture is rapidly gaining its popularity since embedment of cells in a 3D extracellular matrix is associated with more relevant physiological behaviors. Microfluidics allow special control over fluids in micrometer-scale channels and hence extend the physiological relevance of 3D culture models. By employing microfluidic techniques in 3D cell culture, we could co-culture cells in a spatially controlled manner, control over signaling gradients, and mimic the microenvironments of in vivo tissues.

Spatial control Spatial control, as the basis of microfluidic 3D cell culture, allows patterning of cells and extracellular microenvironment. A recent trend is to use hydrogels to offer cells more physiologically relevant 3D matrix. Hydrogels enable cells cluster together without need for surface adhesion. If confined in the appropriate 3D microarchitecture, cells have the intrinsic potential to form aligned tissues in vivo will self-organize into functional tissues in vitro. Microvascular networks support metallic activity and define microenvironment conditions within tissues in health and pathology. Recapitulation of functional microvascular structure in vitro could provide a platform for the study of complex vascular phenomena, including angiogenesis and thrombosis. We engineer 3D vascular network in hydrogel by utilizing fibrillogenesis of collagen and a liquid mold (Fig. 4A). The well controlled vascular network exhibited both mechanical stability for perfusing solutions and biocompatibility for cell adhesion and coverage. This approach could be used for modeling of mass transfer-involved physiology in vasculature-rich tissues and organs for regeneration and drug screening. Blood vessel, composed of intima, media, and adventitia, is a typical complex 3D structure. By using a stress-induced rolling membrane (SIRM) technique, we could transform simple patterns on 2D membranes into complex patterns in 3D tubes. These tubes could mimic blood vessels in which different types of cells constitute different layers of the tubular wall, and cells on certain

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Fig. 3 Microfluidics for wound healing. (A) Left: The schematic illustration of a microchip model for co-culture and selective wounding. Right: MDCK cells alone or co-cultured with NIH 3T3 were patterned on a surface. MDCK cells in the middle channel were selectively lysed by distilled water. (B) Left: The schematic illustration of cutaneous wound model in a cell-on-a-chip device. Cells were patterned on the bottom of a Petri dish using the microchip. Right: NIH-3T3 alone, MDCK cells alone, co-cultured NIH-3T3 and MDCK cells were delivered to two sides to study the effects of the wound dressing. (A) Xie, Y., Zhang, W., Wang, L., et al. (2011). A microchip-based model wound with multiple types of cells. Lab Chip 11, 2819–2822, with permission from Royal Society of Chemistry. (B) Li, Y., Wang, S., Huang, R., et al. (2015). Evaluation of the effect of the structure of bacterial cellulose on full thickness skin wound repair on a microfluidic chip, Biomacromolecules 16, 780–789, with permission from American Chemical Society.

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Fig. 4 Microfluidics for 3D patterning. (A) Left: Schematic illustration of engineering of the vascular network in hydrogel. Middle: The vascular network in serpentine and tree-like shapes. Right (top): The monolayer of HUVECs on the wall of branched hydrogel microchannels, stained by Calcein-AM. Right (bottom): The left three images show one microchannel at different depths. The right image shows the cross-section. (B) Left: Schematic illustration of a SIRM and tubes with multiple types of cells. Right: Three types of cells on a SIRM before and after rolling. Red, endothelial cells; green, smooth muscle cells; blue, fibroblasts. (C) Schematic illustration of fabrication of the cell-laden multilayered PCL–PLGA tubes and morphology changes during long-term incubation. (D) 3D neuronal networks on assembled microspheres. The left three are SEM images of neurons on borosilicate microspheres. Middle: Spatial distribution of neurons. Different colors represent the normalized height to the substrate. Right: Polarity of axons (smi312, arrows) and somatodendrites (MAP2, arrow heads). (A) Mu, X., Zheng, W., Xiao, L., Zhang, W., and Jiang, X. (2013). Engineering a 3D vascular network in hydrogel for mimicking a nephron. Lab Chip 13, 1612–1618, with permission from Royal Society of Chemistry. (B) Yuan, B., Jin, Y., Sun, Y., et al. (2012). A strategy for depositing different types of cells in three dimensions to mimic tubular structures in tissues. Advanced Materials 24, 890–896, with permission from Wiley-VCH. (C) Cheng, S., Jin, Y., Wang, N., et al. (2017). Self-adjusting, polymeric multilayered roll that can keep the shapes of the blood vessel scaffolds during biodegradation. Advanced Materials 29, 1700171, with permission from WileyVCH. (D) Huang, Z., Sun, Y., Liu, W., et al. (2014). Assembly of functional three-dimensional neuronal networks on a microchip. Small 10, 2530–2536, with permission from Wiley-VCH.

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layers align either longitudinally or circumferentially around the tube, mimicking their natural arrangement in vivo (Fig. 4B). Based on SIRM technique, we designed structures, such as tubes with wrinkled walls, tubes-in-a-tube structures and spirals that are difficult to fabricate with existing approaches. We also fabricated SIRM using non-elastic materials whereas reported structures of this type usually require elastic materials. Based on the SIRM technique and electrospinning technique, we used PCL film as inner layer and PLGA film as outer layers to form multicell-laden tubular structure (in vitro) or multilayered, acellular tubular structure (in vivo) (Fig. 4C). The inner (PCL) layer of the tube can expand whereas the outer (PLGA) layers will shrink to maintain the stability of the shape and the inner space of the tubular shape both in vitro and in vivo over months. This approach can be generally useful for making scaffolds that require the maintenance of a defined shape, based on FDA-approved materials. To develop 3D neuronal networks, we used microchamber-assisted assembly of borosilicate glass microspheres to fabricate scaffolds to support neurons (Fig. 4D). These neuron-carrying microspheric building blocks have many advantages, including a relatively large surface to volume ratio, a highly ordered manner with the geometrical confinement of microchambers, sufficient nutritional exchange through the spaces between the building blocks, good long-term stability of the 3D structure. By contrast, the polymeric/hydrogel systems tend to deform and collapse.

Soluble molecular gradients Soluble molecular gradients exist in various biological activities in vivo, such as angiogenesis, invasion and migration. The gradients can be precisely controlled since microfluidics enables spatial control over fluids by altering the channel geometry and flow rates. We fabricated gradients that range from pure laminin to pure bovine serum albumin (BSA) by using laminar flows in microchannels. We cultivated rat hippocampal neurons on these substrate-bound gradients and found that axon specification is oriented in the direction of increasing surface density of laminin. To generate temporally and spatially soluble gradients with more complex geometries, researchers patterned ported microchannels within a 3D ECM and examined how geometry of endothelial vasculature alters the special patterning of diffusive gradients and in turn influences the invasion of endothelial cells during angiogenic sprouting.

Microfluidic simulation of mechanical stimuli Mechanical stimuli play important roles in cell regulate microenvironment of cells and development of many tissues. It is straightforward to establish dynamic microenvironment to mimic the in vivo condition of cells by employing microfluidic chips. To realize real-time imaging of molecular dynamics of live cells, we fabricated a device for stretching cells adhering to elastic membranes in equiaxial or uniaxial mode and obtained high-resolution images of stress fibers during the entire process of the stress (Fig. 5A). We observed that several adjacent stress fibers reassembled into a single one after stretching, which indicated that mechanical stretching played important roles in the rearrangement of actin filaments. To clarify specific mechanical properties of different cells and tissues, we built an in vitro model to mimic the cellular microenvironment by combining both mechanical stretch and geometrical control (Fig. 5B). We found that mechanical stretch determined the optimal geometry of myoblast C2C12 cells, whereas vascular endothelial cells and fibroblasts had no such dependency. In the cardiovascular system, endothelial cells (ECs) are constantly subjected to haemodynamic forces in the form of fluid shear stress (FSS) and cyclic stretch (CS) resulting from blood flow and blood pressure. The two mechanical stimulations on ECs are important for the maintenance of the physiological condition of the cardiovascular system and the development of cardiovascular and cerebral diseases. To study the blood vessel biomechanics, we developed a microfluidic flow-stretch chip, which integrated FSS and CS (Fig. 5C). By imposing FSS-only, CS-only, and FSS þ CS stimulation on rat mesenchymal stem cells and human umbilical vein endothelial cells, we found that the alignment of the cellular stress fibers varied with cell type and the type of stimulation. This flow-stretch chip is an ideal model for simulating the haemodynamic microenvironment. The extravasation of tumor cells is critical in tumor metastasis. To study the mechanism underlying tumor cell extravasation, we fabricated a microfluidic chip to simulate both mechanical and biochemical environments of human vascular systems and analyze their synergistic effects on tumor extravasation (Fig. 5D). The Hela cells exhibited highest viability and adhesion activity in the environment of capillary. The arterial microenvironment plays a vital role in the pathology of atherosclerosis (AS). To address the interplay between arterial microenvironment and atherogenesis, we employed a microfluidic AS model to recapture the atherogenic responses of endothelial cells in ways that Petri dish could not (Fig. 5E). We discovered significant cytotoxicity of a clinical anti-atherosclerotic drug probucol on the model, which does not appear on a Petri dish but is supported by previous clinical evidence. In this part, we introduced recent advances of microfluidics for the manipulation of cell adhesion, migration, and surface engineering. We also focused on microfluidic chip models for cell–cell interactions and wound healing. Apart from that, we discussed microfluidic models for 3D patterning, which could mimic the complex in vivo microenvironments and precisely image dynamic behaviors of cells. The microfluidics allows us to integrate multiple cell types, ECM, and various physical and chemical cues to address basic problems in cell biology and provide robust platforms for fundamental biological research.

Microfluidic-Based Biochemical Analysis Biochemical analysis is important for diagnosis and treatment of diseases. Many different methods enable biochemical analysis. However, most of these methods rely on expensive equipment, trained technicians and complex operations, making them unsuitable for resource-limited areas. Microfluidic platform is a good choice for biochemical analysis as most microfluidic devices are

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inexpensive, portable and easy to operate. In this part, we will show several typical methods for biochemical analysis based on microfluidics, including nucleic acid detection, protein analysis and cell analysis.

Nucleic Acid Detection Nucleic acids are the carriers of genetic information and exist in nearly all organisms. The detection of nucleic acid is thus quite important for disease diagnosis and treatment. In some cases, the nucleic acid analysis can provide more accurate results than the protein analysis since the genetic information of organisms is more stable than proteins, whose expression can be affected by the environment. Besides, nucleic acid detection can be quite highly sensitive because nucleic acids can be amplified easily by using different technologies, such as polymerase chain reaction (PCR) and loop-mediated isothermal amplification (LAMP).

Fig. 5 Microfluidics for simulation. (A) Left: Design of the device and stretch modes: equiaxial stretch; uniaxial stretch. Black arrow heads represent the force applied on elastic membranes. Right: Stress fibers of live NIH 3T3 fibroblast cells, transfected with EGFP-actin. Cells were stretched to different extents in the uniaxial mode 16%; 26%; 32%, and relaxed. (B) Left: Top and side views of the stretching device. Right: Single C2C12 cells were successfully confined in different shapes (blue, nuclear stained by DAPI; red, F-actin stained by Rhodamine Phalloidin). (C) Left: Schematic illustration of the microfluidic flow-stretch chip. Right: The stress fibers tend to align in the direction parallel to that of the FSS, CS and the resultant force of the FSS and CS. (D) Left: Schematic illustration of the vascular model chip. Middle: The adhesion of the tumor cells (HeLa, green) on the ECs monolayer (HUVEC, orange) under different mechanical microenvironments of main artery, medium-sized artery, and capillary respectively. Right: The adhesion of HeLa cells on the wall of the simulated major artery, medium-sized artery, and capillary respectively. (E) Left: Schematic illustration of the early-stage AS model. Right: The morphology change of HUVECs on the Petri dish and the AS model before and after probucol treatment. The white arrowhead points to contracted cells, the white arrow points to space between contracted cells. (A) Wang, D., Xie, Y., Yuan, B., et al. (2010). A stretching device for imaging real-time molecular dynamics of live cells adhering to elastic membranes on inverted microscopes during the entire process of the stretch. Integrative Biology 2, 288–293, with permission from Royal Society of Chemistry. (B) Wang, D., Zheng, W., Xie, Y., et al. (2015). Tissue-specific mechanical and geometrical control of cell viability and actin cytoskeleton alignment. Scientific Reports 4, 6106, with permission from Springer Nature. (C) Zheng, W., Jiang, B., Wang, D., et al. (2012). A microfluidic flow-stretch chip for investigating blood vessel biomechanics. Lab Chip 12, 3441–3450, with permission from Royal Society of Chemistry. (D) Huang, R., Zheng, W., Liu, W., et al. (2016). Investigation of tumor cell behaviors on a vascular microenvironment-mimicking microfluidic chip. Scientific Reports, 5, 17,768, with permission from Springer Nature. (E) Zheng, W., Huang, R., Jiang, B., et al. (2016). An early-stage atherosclerosis research model based on microfluidics. Small 12, 2022– 2034, with permission from Wiley-VCH.

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(continued).

Combining the amplification methods with microfluidic chips, the nucleic acids can be detected accurately with small amount of samples. PCR is a non-isothermal method for DNA amplification. A typical cycle of the PCR process includes three steps: denaturing (90– 95 C), annealing (60–65 C) and extension (70–75 C). So for a PCR-based microfluidic chip, the temperature should be controlled carefully, which makes the microchip more complex. Researchers have made more efforts on isothermal amplifications, such as LAMP. LAMP is an isothermal nucleic acid amplification method. Besides DNA polymerase and deoxy-ribonucleoside triphosphates (dNTPs), six different primers are also required for LAMP. Unlike PCR, whose products are DNA duplex exactly the same as their parental chain, the products of LAMP have a loop structure. Compared with other isothermal amplification methods, LAMP has high amplification efficiency and selectivity. We integrated the LAMP on an eight-channel microchip for the detection of Pseudorabies virus (PRV) successfully (Fig. 6A). After the injection of sample and the reaction mixture of LAMP, the microchannels were sealed with uncured PDMS to form isolated microchambers. The microfluidic chip was incubated at 63 C for 1 h for the LAMP process. After amplification, the results could be analyzed either by the naked eyes or by optical absorbance. The total process could be accomplished within 2 h, and 10 fg mL
1 DNA could be detected successfully. We designed an octopus-like multiplex microfluidic loop-mediated isothermal amplification (mmLAMP) assay for the multiplex gene detection (Fig. 6B). With this mmLAMP, we could identify three human influenza A substrains simultaneously, and the limit of detection (LOD) was less than 10 copies mL
1 in 2 mL quantity of sample. In order to reduce the complexity of device fabrication, we used microcapillary to replace PDMS chip for LAMP process. We reported a microcapillary-based LAMP for the detection of two RNA sequences of human immunodeficiency virus (HIV) (Fig. 6C). The samples and reagents were introduced into the microcapillaries, and microcapillaries are sealed with epoxy glue. We put the microcapillaries in two pieces of pocket warmers for 45 min for amplification. After that, fluorescence signal could be caught under excitation. This device was totally equipment-free and easy to operate, and it was quite convenient for the areas with limited resources. To simplify the operation of nucleic acid extraction, we designed an integrated microcapillary for sample-to-answer nucleic acid detection (Fig. 6D). In this device, a Flinders Technology Associates (FTA) membrane was inserted into the microcapillary for the nucleic acid extraction and the relevant reagents were preloaded in the microcapillary for the LAMP process. When we carried out the detection, we introduced the sample into the microcapillary through negative pressure and the nucleic acids were extracted when the sample flow through the FTA membrane. We added a positive pressure at the other end of the microcapillary to push the washing reagents and LAMP mix flow through the FTA membrane for the LAMP process. During the process, the two ends of the microcapillary were sealed with the epoxy glue to avoid contamination. After amplification, the fluorescence signal could be captured using a hand-held flashlight. We also reported a micro-pipette tip-based nucleic acid test (MTNT) (Fig. 6E). Combined

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Fig. 6 Microfluidic-based nucleic acid detection. (A) Photograph (left) and schematic drawing (right) of the eight-channel microfluidic chip for LAMP. (B) Photograph of the octopus-like multiplex microfluidic LAMP system. (C) Microcapillary-based loop-mediated isothermal amplification for nucleic acids detection. The samples and reagents were introduced into the microcapillaries and the microcapillaries were sealed and put in two pieces of pocket warmers for 45 min for amplification. Fluorescence signals indicated results. (D) Integrated microcapillary for nucleic acids detection. The FTA card and relevant reagents were preloaded into microcapillaries. The samples were introduced into microcapillaries through negative force for nucleic acid extraction. The other reagents were pushed to the FTA card through positive pressure. After amplification, the fluorescence signal could be caught be naked eyes. (E) Schematic illustration of the high throughput micro-pipette tip-based nucleic acid test for sample-toanswer analysis. (A) Reproduced from Fang, X., Liu, Y., Kong, J. and Jiang, X. (2010). Loop-mediated isothermal amplification integrated on microfluidic chips for point-of-care quantitative detection of pathogens. Analytical Chemistry 82, 3002–3006, with permission from American Chemical Society. (B) Reproduced from Fang, X., Chen, H., Yu, S., Jiang, X. and Kong, J. (2011). Predicting viruses accurately by a multiplex microfluidic loop-mediated isothermal amplification chip. Analytical Chemistry, 83, 690–695, with permission from American Chemical Society. (C) Reproduced from Zhang, Y., Zhang, L., and Sun, J. et al. (2014). Point-of-care multiplexed assays of nucleic acids using microcapillary-based loop-mediated isothermal amplification. Analytical Chemistry 86, 7057–7062, with permission from American Chemical Society. (D) Reproduced from Zhang, L., Zhang, Y., and Wang, C. et al. (2014). Integrated microcapillary for sample-to-answer nucleic acid pretreatment, amplification, and detection. Analytical Chemistry 86, 10,461–10,466, with permission from American Chemical Society. (E) Reproduced from Lu, W., Wang, J., and Wu, Q. et al. (2016). High-throughput sample-to-answer detection of DNA/RNA in crude samples within functionalized micro-pipette tips. Biosensors and Bioelectronics 75, 38–33, with permission from Elsevier.

with a multichannel pipette, high throughput nucleic acid test could be achieved, and two copies of DNA could be detected successfully. Besides, this MTNT could be also used for different crude samples, including bacteria, cells and even solid plants.

Protein Analysis Many proteins are biomarkers of diseases or physical conditions. So detecting proteins with fast and accurate methods is quite meaningful for clinical use, as well as fundamental research. One of the biggest challenges for protein detection is that proteins

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cannot amplify, unlike nucleic acids. So the methods for proteins detection should be highly sensitive and highly specific. The most commonly used methods for protein detection are antibody-based immunoassays. Antibodies are immunoglobulin that can recognize and bind tightly to their targets, which are known as antigens. Antibodybased protein detection has both excellent sensitivity and selectivity because of the great affinities between either antibodies and antibodies or antigens and antibodies. Enzyme-linked immunosorbent assays (ELISA) are the most commonly used method for protein detection. There are several different kinds of ELISA, including direct ELISA, indirect ELISA, sandwich ELISA and competitive ELISA. Although there are different protocols according to different ELISA types, they all need the solid substrate to immobilize the antigens or antibodies. So the substrate with higher immobilization efficiency is beneficial for detection. Many efforts have been put on this field, including searching for better substrate, modifying the surface of the substrate, and changing the morphology of the substrate. We compared electrospun polycarbonate (ESPC) membranes and track-etched polycarbonate (TEPC) membranes as substrates for microfluidic immunoassays (Fig. 7A). Due to the larger specific surface area, the immunoassay using ESPC exhibited higher sensitivity for HIV detection. Blocking is the second step of conventional ELISA. We reported a microfluidic assay without blocking (MAWB) for HIV detection (Fig. 7B). For the MAWB method, we incubated antigens in the microchannels to form protein stripes on the PDMS slab. Then the microchannels as well as PDMS slab were washed with 0.05% PBST (phosphate buffered saline with 0.05% of Tween-20). The microchannels were peeled off and the PDMS slab was rinsed with water and dried in the air. This process helped to block the activated sites with Teewn-20 quickly. Another microchip with parallel microchannels was utilized to combine with the PDMS slab with the channels vertical to the stripes of antigens. The samples and the labeled secondary antibodies were introduced into the microchannels for the immune-reaction. The LOD of this method was comparable to the conventional ELISA but the total assay time was shortened about 25%. To increase the detection sensitivity, many nanomaterials have been used as alternatives of enzyme for the signal readout. We reported a nano- and micro-integrated protein chip (NMIPC) that uses water-soluble CdTe/CdS core-shell quantum dots (QDs) as signal probes (Fig. 7C). We used this NIMPC to detect carcinoembryonic antigen (CEA). The dynamic range covered six orders of magnitude, and the LOD was 500 fmol L
1, much lower than the conventional fluorophore-based methods. We also reported a label free platform for protein analysis (Fig. 7D). We embedded gold nanoparticles (AuNPs) on a glass substrate through microwave plasma methods to form Au substrate. The Au substrate was blocked with poly (ethylene glycol) (PEG) and functionalized with biotin moiety. Streptavidin and glutaraldehyde were used to capture the antigens on the surface of the substrate. After the capture of antigens, the interaction between AuNPs was affected and the local surface plasmon resonance (LSPR) peak of AuNPs shifted. There was a positive correlation between the antigen concentration and peak shift. With a portable optical bench, the peak shift could be measured by the UV–Vis spectrophotometer easily and the concentration of targets could be determined. Conventional ELISA process needs several hours and researchers have made great efforts to shorten the reaction period. We described a vacuum-accelerated microfluidic immunoassay that can shorten the reaction time significantly (Fig. 7E). This device was consisted with two PDMS layers and a polycarbonate filter membrane that was sandwiched between the two layers. We introduced the antigen solution in the top layers and applied a vacuum to the whole device. Because of the pressure difference, the solution flowed to the bottom layer, and the antigens were immobilized on the filter membrane. We introduced the labeled antibodies to the bottom layer and applied a vacuum again for the immune reaction. We used this device to detect Sudan Red, an illegal food additive. The total time required was less than 15 min and a LOD of 1 ng mL
1 was achieved. Multiple protein detection is quite meaningful for biological applications. Western blot is a commonly used method for protein detection. But normally just one kind of protein can be detected at one time. We reported a microfluidic-based Western blot (Fig. 7F). With this microfluidic-based method, at least 10 proteins could be detected simultaneously, and the amount of antibody needed was just 1% compared with the conventional Western blot.

Barcode-Based Multiplex Analysis Recently, advanced microfabrication techniques make it available to detect multiplex samples or multiplex targets simultaneously, but the data acquisition and analysis process still limit the throughput of these microdevices. We reported several barcode-related devices for high-throughput multiplex analysis. One of the limitations of data analysis is the lack of “function patterns,” so the region of interest (ROI) needs to be defined manually. We designed a 33-channel microfluidic chip for multiplex analysis (Fig. 8A). A piece of microchannels was assembled on a PDMS slab for the antigen immobilization. After immobilization, that piece of microchannels was peeled off and a second one was assembled on the PDMS slab, perpendicular to the previous one, so that a 33  33 matrix was generated, just like a two-dimensional barcode. During the process of immune reaction, we introduced positive control at the edges of the matrix, which could act as “function pattern.” Within a computer program, the ROI could be performed automatically. The total time required for the data acquisition and analysis was 10 s, more than 500 times faster than the conventional process. The barcode readers and smart phones can make the process of data analysis easier. We designed a one-dimensional barcodebased microchip that used the microchannels with different widths to imitate the common one-dimensional barcode (Fig. 8B). The applicability of this design was demonstrated by the simultaneous detection of three HIV-related proteins and seven pathogenspecific oligonucleotides. Recently, we reported a barcoded paper-based assay (BPA) for multiplex analysis (Fig. 8C). For this BPA system, the constant regions which were not involved in the immune reaction were printed by the inkjet printer, and the

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Fig. 7 Microfluidic-based protein analysis. (A) Top: A microfluidic device for bioassays with the ESPC membrane as solid substrate. Bottom: Schematic illustration of the process of the immunoassay. (B) Schematic illustration of the microfluidic assay without blocking. Step-1: antigens immobilization. Step-2: plasma introduction for immune reaction. Step-3: secondary antibodies introduction for signal readout. (C) Schematic illustration of nano- and micro-integrated protein chip with water-soluble QDs as signal probe. (D) Schematic illustration of the label-free microchip for protein analysis. (E) Schematic illustration of the vacuum-accelerated microfluidic immunoassay. (F) Schematic illustration of microfluidic-based Western Blotting. Left: Proteins were transferred to polyvinylidene fluoride (PVDF) membrane by electroblotting and form protein bands. Right: Antibodies were introduced into the microchannels that are perpendicular to the protein bands. (A) Reproduced from Yang, D., Niu, X., and Liu, Y., et al. (2008). Electrospun nanofibrous membranes: A novel solid substrate for microfluidic immunoassays for HIV. Advanced Materials 20, 4770–4775, with permission from Weily-VCH. (B) Reproduced from Song, L., Zhang, Y., and Wang, W. et al. (2012). Microfluidic assay without blocking for rapid HIV screening and confirmation. Biomedical Microdevices 14, 631–640, with permission from Springer. (C) Reproduced from Yan, J., Hu, M., and Li, D. et al. (2008). A nano- and micro- integrated protein chip based on quantum dot probes and a microfluidic network. Nano Research 1, 490–496, with permission from Springer. (D) Reproduced from Zhang, Y., Tang, Y. and Hsieh, Y. et al. (2012). Towards a high-throughput label-free detection system combining localized-surface plasmon resonance and microfluidics. Lab on a Chip 12, 3012–3015, with permission from Royal Society of Chemistry. (E) Reproduced from Liu, Y., Yu, J. and Du, M. et al. (2012). Accelerating microfluidic immunoassays on filter membranes by applying vacuum. Biomedical Microdevices 14, 17–23, with permission from Springer. (F) Reproduced from Pan, W., Chen, W. and Jiang, X. (2010). Microfluidic western blot. Analytical Chemistry 82, 3974–3976, with permission from American Chemical Society.

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Fig. 8 Barcode-based multiplex analysis. (A) The layout of the barcode-based two-dimensional microfluidic chip. (B) A barcoded microchip for multiplex biomolecules analysis using barcode readers or smart phones. (C) Schematic illustration of the barcoded paper-based assay for multiplex proteins and nucleic acids detection. (A) Reproduced from Zhang, Y., Qiao, L., and Ren Y. et al. (2013). Two-dimensional barcode-inspired automatic analysis for arrayed microfluidic immunoassays. Biomicrofluidics 7, 034110, with permission from American Institute of Physics. (B) Reproduced from Zhang, Y., Sun, J., and Zou,Y. et al. (2015). Barcoded microchips for biomolecular assays. Analytical Chemistry 87, 900–906, with permission from American Chemical Society. (C) Reproduced from Yang, M., Zhang, W., Zheng, W, Cao, F. and Jiang, X. (2017). Inkjet-printed barcodes for a rapid and multiplexed paper-based assay compatible with mobile devices. Lab on a Chip 17, 3874–3882, with permission from Royal Society of Chemistry.

variable regions which have capture probes (CPs) were dispensed by the XYZ dispensing platform. We also designed a new group of barcodes and programmed a new application (APP) for the readout of the barcodes. This BPA could be used to detect both proteins and nucleic acids. Using this assay, the whole process of reaction and results analysis could be completed within 10 min.

Cell Analysis Cells are the basic units of organisms for life activities. Isolation and analysis of certain cells are quite powerful methods for biomedical research and applications. However, conventional approaches cannot manipulate cells precisely as the size of cells is quite small, just around 10 mm. In this field, microfluidic methods have excellent performance as the micro-fabricated structures allow better spatial and temporal control of biological samples. Circulating tumor cells (CTCs) are tumor cells that shed from primary lesions and circulate in the bloodstream. CTCs are the main reason for tumor recurrence and cancer metastasis. And CTCs detection can be employed to cancer diagnosis and therapy monitoring. The main challenge for CTCs detection is that the abundance of CTCs in the bloodstream is very low, compared with other blood cells. Many researchers have made efforts to develop effective methods to isolate CTCs with high efficiency, either by physical force difference or by the affinity difference. We designed a label-free cell sorter for CTCs isolation (Fig. 9A). The cell sorter was composed of a double-spiral microchannel with one inlet and three outlets. When the blood sample flowed through the microchannel, different cells could be isolated due to

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Fig. 9 Microfluidic-based cell analysis. (A) Top: Schematic illustration of the double-spiral microfluidic cell sorter for particle/cell separation. Bottom: Schematic illustration of the two counter-rotating Dean vortices forming in the top and bottom halves of the microchannel. The height of the microchannal was 50 mm. The CTCs flowed out from the middle outlet. (B) Schematic illustration of the microfluidic cell sorter with double-spiral microchannel for cell separation. The height was 85 mm. The large CTCs flowed out through the inner outlet. (C) Schematic illustration of the integrated microfluidic cell sorter for size-based CTCs isolation and enrichment. The height of the microchannel was 40 mm. After isolation, small lysed red blood cells (RBCs) and white blood cells (WBCs) were removed from the side outlets, while the large CTCs were deflected into the middle outlet and enriched by the membrane filter for further nucleic acid analysis. (A) Reproduced from Sun, J., Li, M. and Liu, C. et al. (2012). Double spiral microchannel for label-free tumor cell separation and enrichment. Lab on a Chip 12, 3652–3660, with permission from Royal Society of Chemistry. (B) Reproduced from Sun, J., Liu, C. and Li, M. et al. (2013). Size-based hydrodynamic rare tumor cell separation in curved microfluidic channels. Biomicrofluidics 7, 011802, with permission from American Institute of Physics. (C) Reproduced from Wang, J., Lu, W. and Tang, C. et al. (2015). Label-free isolation and mRNA detection of circulating tumor cells from patients with metastatic lung cancer for disease diagnosis and monitoring therapeutic efficacy. Analytical Chemistry 87, 11,893–11,900, with permission from American Chemical Society.

their size difference. After separation, the small blood cells flowed out from the inner outlet, while the large CTCs flowed out from the middle outlet. The recovery rate of tumor cells could be 88.5%. When the height of microchannel increased from 50 to 85 mm, the small cells flowed out from the middle outlet, and the large cells flowed out from the inner outlet (Fig. 9B). Spiked HeLa cells could be separated successfully from the 20  diluted blood sample using this cell sorter. For the further analysis of CTCs, we integrated an 8 mm filter on the outlet for cell enrichment (Fig. 9C). After separation, the CTCs were enriched by the filter. The capture efficiency of this cell sorter was determined to be 74.5  6.1%, and the enrichment factor reached 3.75  106. We used LAMP for CTCs analysis with CK-19 mRNA as target nucleic acid. The results showed correlation with the assessment gotten from X-ray computed tomography (CT). Compared with FDA-approval CellSearch system, this method showed higher detection efficiency, especially for the epithelial mesenchymal transition (EMT) phenotype. Microfluidic-based method has been widely used in biochemical analysis, both on molecule level and cell level. Compared with conventional methods, microfluidic-based methods are time-saving, cost-saving and labor-saving. Besides, microfluidic-based

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devices have been broadly used in point-of-care testing area, as it is convenient to make integrated platform by using microfluidicbased devices.

Further Reading Chen, Z., Li, Y., Liu, W., et al. (2009). Patterning mammalian cells for modeling three types of naturally occurring cell-cell interactions. Angewandte Chemie International Edition, 48, 8303–8305. Chen, Y., Xianyu, Y., & Jiang, X. (2017). Surface modification of gold nanoparticles with small molecules for biochemical analysis. Accounts of Chemical Research, 50, 310–319. Cheng, S., Jin, Y., Wang, N., et al. (2017). Self-adjusting, polymeric multilayered roll that can keep the shapes of the blood vessel scaffolds during biodegradation. Advanced Materials, 29, 1700171. Fang, X., Chen, H., Yu, S., Jiang, X., & Kong, J. (2011). Predicting viruses accurately by a multiplex microfluidic loop-mediated isothermal amplification chip. Analytical Chemistry, 83, 690–695. Gong, P., Zheng, W., Huang, Z., et al. (2013). A strategy for the construction of controlled, three-dimensional, multilayered, tissue-like structures. Advanced Functional Materials, 23, 42–46. Huang, Z., & Jiang, X. (2013). Micro/nano-scale materials and structures for constructing neuronal networks and addressing neurons. Journal of Materials Chemistry C, 1, 7652–7662. Jiang, X., Ferrigno, R., Mrksich, M., & Whitesides, G. M. (2003). Electrochemical desorption of self-assembled monolayers noninvasively releases patterned cells from geometrical confinements. Journal of the American Chemical Society, 125, 2366–2367. Jiang, X., Bruzewicz, D. A., Wong, A. P., Piel, M., George, M., & Whitesides, G. M. (2005). Directing cell migration with asymmetric micropatterns. Proceedings of the National Academy of Sciences of the United States of America, 102, 975–978. Jiang, B., Zheng, W., Zhang, W., & Jiang, X. (2014). Organs on microfluidic chips: A mini review. SCIENCE CHINA Chemistry, 57, 356–364. Li, X. J., & Zhou, Y. (2013). Microfluidic devices for biomedical applications. The United Kingdom: Woodhead Publishing Ltd. Li, Y., Yuan, B., Ji, H., et al. (2007). A method for patterning multiple types of cells by using electrochemical desorption of self-assembled monolayers within microfluidic channels. Angewandte Chemie International Edition, 46, 1094–1096. Liu, D., Xie, Y., Shao, H., & Jiang, X. (2009). Using azobenzene-embedded self-assembled monolayers to photochemically control cell adhesion reversibly. Angewandte Chemie International Edition, 48, 4406–4408. Mu, X., Zheng, W., Sun, J., Zhang, W., & Jiang, X. (2012). Microfluidics for manipulating cells. Small, 9, 9–21. Sun, J., Li, M., Liu, C., et al. (2012). Double spiral microchannel for label-free tumor cell separation and enrichment. Lab on a Chip, 12, 3652–3660. Sun, J., Xianyu, Y., & Jiang, X. (2014). Point-of-care biochemical assays using gold. nanoparticle-implemented microfluidics. Chemical Society Reviews, 43, 6239–6253. Wang, J., Lu, W., Tang, C., et al. (2015). Label-free isolation and mRNA detection of circulating tumor cells from patients with metastatic lung cancer for disease diagnosis and monitoring therapeutic efficacy. Analytical Chemistry, 87, 11893–11900. Wu, J., He, Z., Chen, Q., & Lin, J. (2016). Biochemical analysis on microfluidic chips. Trends in Analytical Chemistry, 80, 213–231. Yang, D., Niu, X., Liu, Y., et al. (2008). Electrospun nanofibrous membranes: A novel solid substrate for microfluidic immunoassays for HIV. Advanced Materials, 20, 4770–4775. Yang, M., Zhang, W., Zheng, W., Cao, F., & Jiang, X. (2017). Inkjet-printed barcodes for a rapid and multiplexed paper-based assay compatible with mobile devices. Lab on a Chip. https://doi.org/10.1039/c7lc00780a. Yuan, B., Li, Y., Wang, D., et al. (2010). A general approach for patterning multiple types of cells using holey PDMS membranes and microfluidic channels. Advanced Functional Materials, 20, 3715–3720. Yuan, B., Jin, Y., Sun, Y., et al. (2012). A strategy for depositing different types of cells in three dimensions to mimic tubular structures in tissues. Advanced Materials, 24, 890–896. Zhang, Y., & Jiang, X. (2013). Microfluidic tools for DNA analysis. In C. Fan (Ed.), DNA nanotechnology (pp. 113–153). Heidelberg: Springer. Zhang, L., Zhang, Y., Wang, C., et al. (2014). Integrated microcapillary for sample-to-answer nucleic acid pretreatment, amplification, and detection. Analytical Chemistry, 86, 10461–10466. Zhang, Y., Zhang, L., Sun, J., et al. (2014). Point-of-care multiplexed assays of nucleic acids using microcapillary-based loop-mediated isothermal amplification. Analytical Chemistry, 86, 7057–7062. Zhang, Y., Sun, J., Zou, Y., et al. (2015). Barcoded microchips for biomolecular assays. Analytical Chemistry, 87, 900–906. Zheng, W., & Jiang, X. (2014). Precise manipulation of cell behaviors on surfaces for construction of tissue/organs. Colloids and Surfaces B: Biointerfaces, 124, 97–110.

Organs-on-Chips Yunki Lee, Song Ih Ahn, and YongTae Kim, George W. Woodruff School of Mechanical Engineering, Wallace H. Coulter Department of Biomedical Engineering, Institute for Electronics and Nanotechnology, Parker H. Petit Institute for Bioengineering and Bioscience, Georgia Institute of Technology, Atlanta, GA, United States © 2019 Elsevier Inc. All rights reserved.

Introduction Vascular System BraindThe Blood–Brain Barrier Heart Liver Gut (Intestine) Muscle Lung Kidney Current Challenges and Future Prospect Further Reading

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Glossary Clinical translation Clinical translation involves the application of discoveries made in the laboratory to diagnostic tools, medicines, procedures, policies and education, in order to improve the health of individuals and the community. Microfluidics The technology for the manipulation of nanoscale amounts of fluids in microscale fluidic channels for a wide range of applications that include chemical synthesis, and biological analysis and engineering. Organs-on-chips A microscale device that mimics biological systems and is used to probe complex human problems.

Introduction The drug development has been strained by lengthy time lines and poor predictive power of preclinical studies, leading to increased development costs. It is estimated that only one in nine drug candidates that enter clinical trials can reach the market and that two thirds of the development costs are spent in the clinical testing stage. The low success rate and the high expense in the later development stage highlight the importance of earlier stage predictive power using more versatile and rapid preclinical models that more accurately predict drug safety and efficacy. However, current in vivo studies using animal models may bear little relation to whole human physiology, and conventional in vitro models are based on twodimensional (2D) culture of transformed cell lines. Recent advanced three-dimensional (3D) cell culture models using custom hydrogels remain to pose technical challenges in recapitulating physiologically relevant microenvironments of human organs or tissues. Meanwhile, microfluidic systems have shown great potential to mimic the physiology and functionality of an organ or tissue interface. These microsystems are engineered with 3D cell culture and controlled mechanical and chemical cues to investigate cellular dynamics and molecular pathways, which are extremely difficult to observe using current in vivo models. This so-called “organs-on-chips” technology aims to recapitulate the chemical and mechanical microenvironment within human organs. The underlying microtechnology of organs-on-chips overcomes several practical limitations including the large number of cells required to build organ structure and function, the large volume of drug and other reagents needed to study, and the low throughput of existing studies needed to drive the economics of the preclinical drug evaluation process. Recent years have reported new approaches for organs-on-chips that reconstitute the structure and function of human organs and their interactions. Successful development of organs-on-chips will enable scientists to predict more accurately the efficacy of therapeutic drug candidates, revolutionizing drug development, disease modeling, and personalized medicine. This article provides the fundamental understanding of the structure and function of key human organs, describes the cellular features and microenvironments, and highlights current technologies to microengineer the key functional units in organs-on-chips systems. It also includes brief discussion on current challenges and future prospects for these technologies to decode its potential to practice.

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Vascular System The vascular system or the circulatory system carries blood and lymph throughout the body. The blood vessels like arteries and veins carry blood throughout the body, transporting oxygen and nutrients to the pheripheral tissues and removing the wastes from target tissues. The lymph vessels filter and drain lymph from the body to maintain the fluid environment. The dysfunction of the vascular system therefore causes critical pathologies including cardiovascular diseases (e.g., hypertension, atherosclerosis, and restenosis), tumor angiogenesis, and cancer metastasis. The endothelium lining the inner layers of blood vessels in the cardiovascular system experiences pulsatile flow-induced wall shear stress and transmural pressure, and pericytes and smooth muscle cells wrapping the outer walls of blood vessels contribute to the stabilization of the vascular system (Fig. 1A). Intercellular communication between these cells is regulated by signaling pathways such as through gap junctions that allow the exchange of metabolites, ions, and other essential molecules. Direct interaction between endothelial cells and smooth muscle cells faciliates the synchronization of their behaviors along the vascular wall. The microengineered vascular system represents an undisputed advance in developing physiologically relevant vasculature models. The microfluidic device integrated with micropost structures allows for the investigation of microvessel formation, organization, and function under highly controlled conditions including biochemical gradients and flow-mediated dynamics. One representative approach is to replicate vasculogenic formation and angiogenic sprouting morphogenesis at the intraluminal side of the microvascular network (Fig. 1B). The extracellular matrix-mimicking hydrogels are being developed to provide more physiologically relevant 3D microenvironment to vascular endothelial cells, pericytes, and smooth muscle cells. Recent advances in 3D bioprinting technology allow high-precision construction of microarchitechture with hydrogel bioinks containing vascular cells and extracellular matrix components. This recapitulation of the complex microarchitechture of a human vascularized tissue combined with precisely controlled flow conditions with physiological context enables a perfusable capillary network within 3D extracelluar matrix and maintains the endothelial barrier function. Despite progress in engineering vascular systems, reproduction of the key biophysical and biochemical features of the basement membrane and interstitial matrix in the microengineered vascular systems with the desired luminal structure and size requires more physiologically relevant biomaterials and advanced microfabrication approaches. Co-culture with stromal cells and immune cells in the developed vascular models is being studied with better maintenance of phenotypic cell stability and intercommunication. Modeling of complicated vascular functions including regulation of blood pressure, vasoactivity, hemostatic balance, permeability, and immunity and complex pathological conditions including thrombus formation and ischemia/reperfusion injury remains challenging.

BraindThe Blood–Brain Barrier The brain is a highly specialized organ serving as the centeral nervous system that controls dynamic and coordinated responses to the environmental changes in other organs. These highly coordinated biological activities are maintained by the blood–brain barrier, a selectively permeable structure that restricts the passage of nonlipophilic species or large substances (> 400 Da). Dysfunction of the blood-brain barrier leads to disruption of brain homeostasis, ultimately causing a variety of neurological disorders including Alzheimer’s disease, Parkinson’s disease, and amyotrophic lateral sclerosis. The blood-brain barrier is a highly organized structure that consists of specialized endothelial cells that line the blood vessels with the basal lamina, pericytes that wrap around these vessels, and end-feet of astrocytes extending to the vesselsdcollectively referred to as the blood–brain barrier contributing to the brain homestasis (Fig. 2A). The crosstalks between these cells maintain the integrity of the blood-brain barrier. Pericytes and astrocytes act on the specialized endothelial cells to maintain an extensive

Fig. 1 (A) Schematic illustration of vascular system. (B) Microengineered vascular system to investigate microvascular network organization by mechanical and biochemical stimuli. Reproduced from Kim, S., Chung, M., Ahn, J. et al. (2016). Interstitial flow regulates the angiogenic response and phenotype of endothelial cells in a 3D culture model. Lab on a Chip 16(21), 4189–4199, with permission from Royal Society of Chemistry.

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Fig. 2 (A) Schematic illustration of neurovascular functional unit consisting of neurons, blood–brain barrier including endothelial cells, pericytes, astrocytes and glial cells. (B) Engineered neurovascular unit on a chip. (C) 3D human blood-brain barrier on a chip. (A and B) Reproduced from Alcendor, D. J., Block, F. E., Cliffel, D. E. et al. (2013). Neurovascular unit on a chip: Implications for translational applications. Stem Cell Research & Therapy 4(Suppl 1), S18, with permission from BioMed Central Ltd. Reproduced with permission from Herland, A.,van der Meer, A. D., FitzGerald, E. A. et al. (2016). Distinct contributions of astrocytes and pericytes to neuroinflammation identified in a 3D human blood-brain barrier on a chip. PLOS One 11(3), e0150360.

network of tight junctions, promote the expression of specialized transporters, and limit the rate of transcellular pinocytosis. In addition, the extracellular matrix as well as the basement membrane of the endothelium plays a key role in the brain microenvironment homeostasis. While the basement membrane is predominantly made up of collagen, laminin, and fibronectin, the neural extracellular matrix comprises a dense network of proteoglycans, hyaluronan, tenascins, and small amounts of fibrous proteins and adhesive glycoproteins. In vitro systems of the blood–brain barrier that incorporate all those key components with its capability for spatiotemporal regulation of the microenvironment may provide a promising tool for prescreening and optimization of new drugs for neurological diseases prior to animal and clinical trials. However, due to the complexity, no existing in vitro models have closely recapitulated the blood–brain barrier with physiological context yet. The current state-of-the art model is a 3D multicompartment microfluidic device that includes the central nervous system and cerebral spinal fluid, which is used to investigate the role of both the blood– brain barrier and the blood-cerebral spinal fluid barrier in modulating chemical body–brain interactions, and to characterize the interactions of pericytes, astrocytes, microglia, and neuronal and endothelial cells in the brain and its barriers (Fig. 2B). The critical aspect of the in vitro blood–brain barrier model development is to mimic a 3D microenvironment of a brain tissue. Several multicomponent hydrogel scaffolds are widely utilized with manipulation of physicochemical properties such as matrix stiffness and surface chemistry. The in vitro 3D models with an endothelial lumen structure have been primarily developed in a collagen type I gel where the endothelial cells arranged in a cylindrical monolayer inside the microfludic platforms construct similar geometry to brain microvessel (Fig. 2C). No physiologically relevant in vitro neurovascular unit systems including the blood–brain barrier and 3D glial and neuronal networks exist due to the complexity. Mimicking brain microenvironments for better maintenance of brain cell phenotypic stability remains challenging. In addition to the recent focus on the development of blood–brain barrier models, 3D neural cultures designed to manipulate the complex neuronal networks, long processes, and synapses are required for better modeling neurological diseases, as well as drug screening.

Heart The heart is a muscular organ that pumps blood through the circulatory system. The heart wall consists of three cellular layers: the outer epicardium, middle myocardium, and inner endocardium. The epicardium, a thin membrane that covers the myocardium, functions as a source for multipotent progenitor cells, and regulates cardiomyocyte proliferation, differentiation, and regeneration. The myocardium, a layer located between the epicardium and the endocardium, comprises the primary contractile unit of the heart (Fig. 3A). A cardiomyocyte is the primary cell type in the myocardium that terminally differentiate into a muscular cell, which accounts for the transmission of electrical signals (e.g., action potential) through the gap junctions, coordinating the contractile activity of the muscle. These physiological functions of the cardiac tissue are maintained with other important integrated components: fibroblasts, extracellular matrix, and vascular systems. The components for the extracellular matrix in the heart tissue include collagen (type I and III), proteoglycans, fibronectin, and basement membrane proteins. The cell–cell interactions have a significant impact on the tissue responses via the secretion of a combination of different factors (e.g., growth factor, cytokine) that modulate the molecular composition of the cardiac tissue microenvironment. Such an organized orchestra of cells and tissue factors regulates, stabilizes, and reinforces the cardiac phenotype. It is the concerted interactions between the cells, extracellumar matrix, and signaling molecules all together that contribute to the development and maintenance of the functions of the heart.

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Fig. 3 (A) Schematic illustration of anatomic structure and organization of cardiac-muscle tissue. (B) Engineered heart-on-a-chip device based on muscular thin film technique. Reproduced from Agarwal, A., Goss, J. A., Cho, A. et al. (2013). Microfluidic heart on a chip for higher throughput pharmacological studies. Lab on a Chip 13(18), 3599–3608, with permission from Royal Society of Chemistry.

To study cardiac physiology, pathology, and pharmacology, engineered heart-on-a-chip platforms have to provide precise control over these essential factors; for example, combination of topographical and electrical cues to engineer the cardiac tissue with matured neonatal rat cardiomyocytes in microscale devices allows for increased cellular maturation and elongation with gap junction formation. The current state-of-the-art of heart-on-a-chip is to use a muscular thin film technique to recapitulate the structure and function of the myocardium (Fig. 3B). This system enables the real-time measurement of the cardiac muscle tissue contraction in response to pathophysiological conditions combining mechanical, electrical, and chemical stimuli. The advance of this technology has demonstrated not only microcontact printing of fibronectin on the thin flexible PDMS films to electrically stimulate anisotropically organized cardiomyocytes, but also, multiple engineered cardiac tissues on a chip for drug testing applications in a robust and high-throughput manner using 3D printing technology. While most heart-on-a-chip platforms use animal-derived primary cells, more focus on integration of human origin and humanderived iPSCs with these models will be required for better mimicking human physiological response. In addition, development of 3D engineered myocardium microenvironment and integration of these platforms with vasculature remains challenging. Engineered cardiac tissue constructs that include both physiologically relevant mechanical and electrical properties and provide realtime measurement of contractility and action potential signals will impact the field for the applications of in vitro drug testing, disease modeling, and biological mechanistic studies.

Liver The liver, the largest organ in the body, performs essential metabolic functions that include synthesis of proteins (e.g., albumin), detoxification of endogenous and exogenous substances ingested through the gastrointestinal tract, and production of biochemicals necessary for digestion. The internal structure of the liver has approximately 100,000 small hexagonal functional units known as lobules, where the oxygen consumption creates the functional heterogeneity between periportal and perivenous zones (i.e., liver zonation) during methabolism and detoxification process. The hepatocyte, a majority of cells in the liver that constitute 70%–85% of the liver’s mass, performs primary functions including metabolism, detoxification, storage, digestion, and bile production. The walls of hepatic sinusoid are lined by three different cell types: sinusoidal endothelial cells, Kupffer cells, and hepatic stellate cells (Fig. 4A). The liver sinusoids are lined with sinusoidal endothelial cells, separating hepatocytes from flowing blood. A thin extracellular matrix in the space of Disse between hepatocytes and endothelial cell layer is composed of fibronectin, collagen, laminin, and other basement membrane proteins, and retains important substances including glycosoaminoglycans, growth factors (e.g., basic fibroblast growth factor, hepatic growth factor, and vascular endothelial growth factor) and cytokines. Kupffer cells preferentially reside in the sinusoid of the periportal zone, and form the macrophase population. Quiescent stellate cells located in the space of Disse store vitamin A, and regulate regeneration, vascular remodeling, differentiation, and inflammation. To mimic the liver microenvironment, long-term maintenance of hepatocyte viability and functionality is required for the metabolism and liver toxicity studies. Incorporation of a perfusable flow provides continuous nutrient/oxygen supply and physiological shear stress required to recapitulate liver tissue microenvironment. These approaches are enhanced with efforts to simulate the physiologically relevant morphology of the liver functional units; endothelial barrier of liver sinusoid and hepatic lobule structure, highlighting the importance of structure-function relationship of these modeling approaches and combining physiological cues such as dynamic flow and vascularization within the liver-like structure (Fig. 4B). To predict drug and biothreat agent toxicity under physiological conditions in real-time instead of conventional end-point assays, microengineered liver-on-a-chip platforms have to integrate electrochemical measurement technology. Incorporation of oxygen sensors in a liver-on-a-chip bioreactor device allows for the detection of minute changes of oxygen concentrations across the tissue and the real-time tracking of the dynamics of

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Fig. 4 (A) Schematic illustration of the basic liver tissue unit, hepatic lobule and sinusoid. Microengineered (B) in vitro lobule-mimetic and (C) liver sinusoid-like bioreactor models consisting of hepatocytes and endothelial cells. Reproduced from Ho, C. T., Lin, R. Z., Chen, R. J. et al. (2013). Liver-cell patterning lab chip: Mimicking the morphology of liver lobule tssue. Lab on a Chip 13(18), 3578–3587; Domansky, K., Inman, W., Serdy, J. et al. (2010). Perfused multiwell plate for 3D liver tissue engineering. Lab on a Chip 10(1), 51–58, with permission from Royal Society of Chemistry.

metabolic adaptation to mitochondrial dysfunction (Fig. 4C). In addition, combining this monitoring technology with 3D bioprinting platforms, researchers have encapsulated hepatic spheroids in hydrogels printed directly into the microfluidic bioreactor device, allowing for the evaluation of drug-induced toxicity through dynamic monitoring of metabolite production. Current challenges include the development and long-term validation of reliable 3D disease models that mimic vascularization and dynamic flow in the liver and the availiability of human cell sources such as iPSC-derived human liver cells, which may extend the potential for the personalized drug screening applications.

Gut (Intestine) As a vital organ in the gastrointestinal tract of digestive system, the intestine operates several essential functions such as digestion, nutrient transport into blood circulation, metabolism, and elimination of toxic substances. The intestinal wall composed of various differentiated cell types with a tight junction and covered by a mucus layer forms a highly selective barrier whose physiological functions are closely associated with the peristalsis, villi of the mucus layer, and microbiome activity (Fig. 5A). The intestinal epithelium regulates these fundamental immune-regulartory functions through the secretion of mucins and antimicrobial peptide while gut microbiome prevents infection and controls immune system through the balanced host-gut microbiome crosstalk. Infections, inflammation, and drug toxicity can cause epithelial dysfunction and gut microbiome imbalance, such as inflammatory bowel disease.

Fig. 5 (A) Schematic illustration of native intestinal wall with villi structure and tissue interface with blood vessel. (B) Design of human gut-on-achip microdevice composed of the gut lumen and blood compartment. Cyclic suction to both side channels exerts peristalsis-like motion, ultimately producing relevant host-microbe interaction. Reproduced from Kim, H. J., Huh, D., Hamilton, G. et al. (2012). Human gut-on-a-chip inhabited by microbial flora that experiences intestinal peristalsis-like motions and flow. Lab on a Chip 12(12), 2165–2174, with permission from Royal Society of Chemistry.

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Most substances and drugs absorbed through the intestine into the blood stream diffuse across the mucus layer, the epithelial cell layer of the intestine wall, basement membrane, and the endothelial cell layer lining the capillaries. The intestinal epithelium is composed of different types of epithelial cells mainly including enterocytes and goblet cells. The enterocytes covered with surfaceenhancing microvilli produce various catabolic enzymes on their exterior luminar surface for the absorption and methabolism functions. In most areas of the intestine, the goblet cells secrete the mucus layer which protects the epithelial layer from the luminal contents, and play critical roles in regulating innate immune responses and homeostasis. The basement membrane, consistsing of laminin, collagen type IV, and proteoglycans, supports the intestinal epithelium for tissue development, function, and repair. The current state-of-the art model of gut-on-a-chip is based on double-layered perfusable channels separated by an extracellular matrix-coated porous flexible membrane in order to simulate the intestine lumen and the blood compartment (Fig. 5B). The intestinal epithelial cells form a high integrity epithelium-barrier on one side of membrane with contiuous fluid shear stress while the cyclic mechanical deformation triggered by stretching and relaxing the porous membrane mimics physiological peristaltic motions. This mechanically active platform enables the cell polarization and differentiation into a columnar epithelium and the subsequent intestinal villi formation. When co-cultured with a normal intestinal microbe on the luminal surface, the intestinal microenvironment engineered to mimic peristalsis-like motion and fluid flow improves the intestinal barrier function. This physiological organlevel function reconstitution has not been possible in 2D static culture system due to contamination and overgrowth of the microbe. This platform is being further adapted to study inflammation response at tissue and organ levels, and antibiotic and probiotic therapeutic effects toward injured intestinal epithelium, thereby verifying host-microbiome crosstalks in intestinal pathophysiology and dissect disease mechanisms. These in vitro models provide new insights of the interplay between multiple different types of cells and tissues (e.g., intestinal epithelial/endothelial cells, immune cells, commensal microbe and pathogenic bacterial), molecular components, and physical cues in an organ-relevant microenvironment. Some other recent modeling approaches include hostmicrobe interface model under aerobic-anaerobic conditions that highlight the significant technical advance in manipulating metabolically mismatched microenvironments. Current challenge in this technology is to obtain healthy primary human cells for improved extrapolation of results to humans and to contain a host of “non-GI” cells including mast cells which play a key part in the barrier integrity and physiology of the intestinal epithelium.

Muscle Skeletal muscle tissue accounts for 30%–40% of total body weight, and is responsible for gross movements. Skeletal muscle consists of myofilaments, sarcomeres, myofibrils and muscle fiber. The sarcomeres, the muscle contractile units, result from the assembly of myosin and actin myofilaments. The repetition of sarcomeres generates myofibrils, which assemble into myofibers, long cylindrical multinucleated muscle cells. Muscle fibers are the basic unit of muscles and organize into fascicles, which are surrounded by a network of blood vessels and nerve fibers (Fig. 6A). Pathological conditions include muscular genetic disease such as Duchenne

Fig. 6 (A) Anatomic structure of organized skeletal muscle fiber surrounded by capillary blood vessels and neuronal network. (B) In vitro model of neuromuscular junction cocultured with motoneuron cell bodies and muscle bundles in the microfluidic device. (i–ii) Representative images of neuron-muscle coculture and neurite extension. Axons and myocytes are immunoreactive for NFM and desmin, respectively. (iii) Visualization of fibroblast (green), motor axon (red) and myotubes (blue) in the distal chamber. Reproduced from Southam, K. A., King, A. E., Blizzard, C. A. et al. (2013). Microfluidic primary culture model of the lower motor neuron-neuromuscular junction circuit. Journal of Neuroscience Methods 218(2), 164– 169, with permission from Elsevier.

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muscular dystrophy and neuromuscular degenerative disease such as amyotrophic lateral sclerosis. In case of the mutation in the dystrophin gene, the abnormal transmission of contraction alters the muscle contractility, leading to muscle atrophy. The microenvironment of the skeletal muscle is composed of myofiber, extracellular matrix, and cellular components including muscle satellite cells, fibroadipogenic progenitors, and endothelial cells in the microvalsculature and neuronal network. The satellite cells, which make up approximately 2%–10% of the total number of myonuclei, are quiescent in physiological condition, but are activated for proliferation and differentiation to form new myofibers consisting of multinucleated myotubes in response to mechanical and chemical stimuli resulting from injurious events. These myofibers are covered by a thin layer of connective tissues mostly composed of laminin and collagen type IV. The cells in the muscle microenvironment are anatomically adjacent to each other, and their behaviors are strictly regulated through the dynamic interplay with vasculature. Vascular inflow into muscle takes place through the arteries that are distributed along muscle fibers. Vessel branching occurs obliquely or perpendicularly to the main vessels thereby surrounding muscle fibers, and allow for their perfusion. The contractitility of aligned muscle fibers is modulated by electrical stimulation. Motoneurons located in the spinal cord innervate skeletal myofibers by forming neuromuscular junctions and trigger muscle contraction by releasing transmitter acetylcholine. Generation of functional muscle fibers requires fine control of topographical (e.g., surface micropatterning/microgrooves), biochemical (e.g., soluble factors and matrix composistion), mechanical (e.g., cyclic strain and matrix stiffness), and electrical ques in the microdevices. Incorporation of these stimulations increases cellular actitivites including myogenic differentiation and myotube formation, maturation and alignment, thus effectively improving the contractile behavior of engineered muscle tissue units. Another important characteristic in the muscle microenvironment, the neuromuscular junctions are constructed to investigate cell–cell interaction between motor neurons and muscle cells, which is crucial to understand mechanisms leading to muscle degeneration or diseases (Fig. 6B). Precise spatiotemporal control of local chemical environment allows for investigation of the effect of chemoattactant gradients or neuron-secreted factors on myotube functions in co-culture condition. Incorporation of compartmentalization system with microgrooves allows axonal outgrowth from motoneuron cell bodies and formation of neuromuscular junctions with muscle cells. More recent approaches mimic more realistic morphology, and increase physiological relevance with 3D functional neuromuscular junction platforms engineered by coculturing with myoblast-derived 3D muscle strips and motor neurons within a hydrogel matrix. These 3D muscle tissue models compartmentalize muscle fibers and motor meurons, mimicking the geometry of the spinal cord-limb physical separation and facilitating real-time monitor axonal outgrowth and muscle innervation. The combination strategy of microfluidic and optogenetic technologies enables development of highly controllable and physiologically relevant tissue model for motor units and ultimately for drug screening assay applications. The major challenges lie in the manipulation of cues that induce differentiation toward contractile lineage, the understanding of how contraction is impaired in the muscular disorders and how to rescue it, and the engineering of de novo surrogates for tissue repair or nonmedical applications. Current models still lack important features that would enable them to fully characterize skeletal muscle pathologies. To better characterize the response of pathological muscle models, the readout of important parameters such as muscle contractility, indicative of muscle health, needs to be added to the systems on the basis of similar apparatuses reported for muscle models. Lastly, novel findings indicate how the pathological muscular endothelium is an important element to be considered, although vascularized pathological muscle model is not currently available.

Lung The lung is a primary organ of the respiratory system that performs gas (oxygen–carbon dioxide) exchange through passive diffusion in the basic functional unit, alveoli, which are hollow cavities surrounded by numerous pulmonary capillaries. The alveolar wall is a physical barrier that is composed of two cell layers of the alveolar epithelium and the vascular endothelium, and that contributes to the regulation of lung tissue homestasis (Fig. 7A). The alveolar epithelial cells are exposed to mechanical stresses that have critical roles in the regulation of the key pulmonary functions, including their response to environment-induced damage by releasing biologically active factors. The defense mechanisms of lung tissue against various environmental insults (e.g., pathogens and tissue damaging agents) include secreting antimicrobial peptides and inflammatory mediators, and initiating immune cell recruitment to the damaged or infected site. Through the epithelial-endothelial crosstalk, various mediators including chemokines, cytokines, growth factors and other small molecules such as nitric oxide or reactive oxygen species are released to recruit innate and adaptive immune cells. The infection and chronic inflammation are responsible for most lung diseases, particularly chronic obstructive pulmonary diseases including chronic bronchitis, emphysema, and asthma. The extracellular matrix of the pulmonary alveoli is composed of collagen (type I and III), elastin, glycosaminoglycans and proteoglycans. In the lung tissue, the pulmonary extracellular matrix determines lung-tissue strucuture, and provides mechanical stability, elastic recoil and maintenance of normal interstitial fluid dynamics, which are critical for physiological lung function. Moreover, biochemical/mechanical signals initiated by the extracellular matrix regulate cellular behavior, and play key roles in development, remodeling and homeostasis of the lung tissue. To replicate the strcutrual and functional complexity of the alveolar-capillary interface in the lung, the state-of-the art lung-on-achip system leverages compartmentalized microfluidic chambers separated by a porous membrane to create the microarchitecture and dynamic microenvironment of the alceolar-capillary unit (Fig. 7B). The epithelial and endothelial cells are co-cultured on the opposite sides of extracellular matrix-coated membrane, and are respectively exposed to air and flowing media, ultimately resulting in air–liquid barrier formation in microdevices. The cyclic stretching implemented in the system allows for the simulation of the

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Fig. 7 (A) Schematic illustration of the lung functional unit. (B) (i) Cyclic stretchable lung-on-a-chip microdevice consisting of epithelial and endothelial cell layers for reproducing the alveolar–capillary barrier in a microfluidic device. (ii) A tissue–tissue interface consisting of a single layer of the alveolar epithelium closely apposed to a monolayer of the microvascular endothelium by long-term microfluidic co-culture. Reproduced from Huh, D., Kim, H. J., Fraser, J. P. et al. (2013). Microfabrication of human organs-on-chips. Nature Protocols 8(11), 2135–2157, with permission from Nature Publishing Group.

lung physiological breathing motion. This mechanically active microdevice reproduces organ-level responses to bacteria and proinflammatory cytokines as well as nanoparticles, which provides more realistic in vitro models to study pathogen-triggered pathological mechanism and body’s response to drugs. This platform is continuously used to simulate human lung diseases such as pulmonary inflammation by perfusing proinflammatory factors. Chronic obstructive pulmonary diseases are also modeled by lining the airway with epithelial cells and stimulating with viral mimic polyinosinic–polycytidylic acid or lipopolysaccharide endotoxin. Current challenge in the microengineered lung-on-a-chip models include the applications of human primary cells or specific patient-derived cells, which would enable creating more relevant and complex hallmarks of human lung diseases. In addition, more cell types such as pneumocyte II or macrophage immune cells would be integrated in the well-estabilised model to assess some mechanisms in certain aspects of lung response to inflammation and infection.

Kidney The kidney is a vital organ that maintains hemeostasis in the body and controls blood pressue through blood filtration, waste excretion, and essential substance reabsorption. It also plays a primary role in the elimination of toxins and their metabolites, therefore is susceptible to injury during the toxin removal process. The kidney tissue consists of an inner medulla, an outer cortex and millions of nephrons, the basic structural and functional unit of the kidney. The nephron is composed of two main structures: the glomerulus that is responsible for renal ultrafiltration and the renal tubule for reabsorption of useful substances such as water, glucose, and amino acid (Fig. 8A). The glomerular filtration barrier consists of endothelial cells, podocytes, mesangial cells, and epithelial cells. The endothelial and epithelial monolayers share the extracellular matrix known as the glomerular basement membrane where these three distinctive layers form the glomerular filtration barrier. The filtration and excretion mechanism include size, charge, and shape selectivity at the glomerular barrier that has size-selective slit nanopores in foot processes and surface negative charge of glomerular endothelium. In contrast, the renal tubular reabsorption is the process where the removed water and solutes from the glomerular capillaries transport into the blood circulatory system to maintain homeostasis, which mostly occurs in the proximal tubule by osmotic pressure and active transport of the tubular epithelial cells. Impairment of the renal filtration by disrupted epithelial cell junctions causes failure of homeostasis, leading to kidney diseases. Acute renal injury is caused by decreased blood flow to the kidneys or injury from drugs and viral infection. Chronic renal failure includes inflammation such as glomerulonephritis or pyelonephritis, polycystic kidney disease, renal fibrosis and kidney stone. The kidney-on-a-chip platforms recapitulate tissue- and organ-level physiology of the nephron functional unit, focusing on the renal tubule and capillary microenvironment of the kidney-specific functional unit on a chip. The state-of-art proximal tubule model is a multilayer microfluidic device system consisting of luminal flow channel and interstitial space compartmented by a porous membrane. The integration of relevant fluidic shear stress and transepithelial osmotic gradients increases the physiological relevance, demonstrating polarization and cytoskeletal reorganization of the renal tubular epithelial cells. This approach is enhanced to construct the kidney glomerulus as well as disease model. The diabetic nephropathy disease model of the glomerular endothelial barrier was estabilished in the compartmentalized microfluidic device consisting of tissue-specific cellular components and 3D extracellular matrix (Fig. 8B). Under fluid flow condition, 3D hydrogel membrane was lined by endothelial cells and podocytes, and served as the glomerular filtration barrier, showing selective permeability for large proteins. This device was also employed to verify high glucose-induced pathological responses where hyperglycermia proved to cause glomerular dysfunction

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Fig. 8 (A) Schematic illustration of nephron functional unit (left) and glomerular structure (right) in kidney tissue. (B) In vitro glomerular model consisting of glomerular endothelial cells, podocytes and hydrogel membrane. This design mimics the glomerular filteration barrier. Reproduced from Wang, L., Tao, T., Su, W. et al. (2017). A disease model of diabetic nephropathy in a glomerulus-on-a-chip microdevice. Lab on a Chip 17(10), 1749– 1760, with permission from Royal Society of Chemistry.

with increased barrier permeability. In addition, to better recapitulate in vivo microenvironment, with combining bioprinting technique, perfusable 3D human renal proximal tubule was created on a chip. This engineered 3D tubule model promoted tissue-like epithelium formation with enhanced epithelial morphology and functional properties, compared to the 2D system. Most work leverages double-layered microfluidic compartmentalization to develop the renal tubular microenvironment of the nephron. More studies are now focusing on mimicking a glomerular filtration barrier structure and function, which will contribute to the modeling of many renal diseases that are related to impaired activity of glomerulus. Glomerular dysfunction leads to the alteration of podocyte foot processes, which is responsible for inefficient blood ultrafiltration and ultimately proteinuria and other pathological diseases. Introduction of primary human cells into this microengineered nephron functional unit would provide more realistic physiological and pathological in vitro models for studying drug therapeutic efficacy and toxicology to kidney tissues.

Current Challenges and Future Prospect The overall challenge of this field is to ultimately scale up these microengineered organ constructs to the volume required for pharmaceutical applications and to achieve reliable high-throughput systems for efficient drug screening. More practically, the central question lies in how to create the simplest yet physiologically relevant (and also useful) models that reproduce the critical functions of interest organs with reliable human primary cell sources. To further recapitulate the structure and function of human organs, the current organs-on-chips technologies need to maintain a healthy physiological state where nutrients and waste products do not reach unacceptably low or high levels and establish precision fluid flow on organ model platforms for controlling drug, nutrient, and metabolite concentrations and applying fluid shear to cultured cell populations. Moreover, automation technologies should be incorporated for monitoring cell culture conditions in multiwell plates as well as real-time assessment of morphology, cell growth, and protein expression. Real-time measurement of basic cell culture parameters such as oxygen, temperature, pH, and molecular measurements will further expand this innovative technologies for the rapid translation of drug candidates to the clinic. Lastly, due to the complexity of each organ-on-a-chip system, multiorgan system integration and scaling remain challenging in the context of the application to drug screening. Successful development of organs-on-chips with these advances will be able to improve the cost effectiveness and predictive power of preclinical studies assessing the safety and efficacy of compounds (e.g., absorption, distribution, metabolism, and excretion), and eliminate ineffective drug candidates as early as possible. Furthermore, this innovative approach may allow us to be able to respond to newly evolving demands including personalized or precision medicine, rising rates for many chronic diseases, and threats from emerging infectious diseases.

Further Reading Benam, K. H., Villenave, R., Lucchesi, C., et al. (2016). Small airway-on-a-Chip enables analysis of human lung inflammation and drug responses in vitro. Nature Methods, 13(2), 151–157. Bersini, S., Arrigoni, C., Lopa, S., et al. (2016). Engineered miniaturized models of musculoskeletal diseases. Drug Discovery Today, 21(9), 1429–1436.

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Huh, D., Matthews, B. D., Mammoto, A., et al. (2010). Reconstituting organ-level lung functions on a chip. Science, 328(5986), 1662–1668. Khetani, S. R., & Bhatia, S. N. (2008). Microscale culture of human liver cells for drug development. Nature Biotechnology, 26(1), 120–126. Kim, H. J., Li, H., Collins, J. J., et al. (2016). Contributions of microbiome and mechanical deformation to intestinal bacterial overgrowth and inflammation in a human gut-on-a-chip. Proceedings of the National Academy of Sciences of the United States of America, 113(1), E7–E15. Kolesky, D. B., Homan, K. A., Skylar-Scott, M. A., & Lewis, J. A. (2016). Three-dimensional bioprinting of thick vascularized tissues. Proceedings of the National Academy of Sciences of the United States of America, 113(12), 3179–3184. Lind, J. U., Busbee, T. A., Valentine, A. D., et al. (2017). Instrumented cardiac microphysiological devices via multimaterial three-dimensional printing. Nature Materials, 16(3), 303–308. McCain, M. L., Sheehy, S. P., Grosberg, A., Goss, J. A., & Parker, K. K. (2013). Recapitulating maladaptive, multiscale remodeling of failing myocardium on a Chip. Proceedings of the National Academy of Sciences of the United States of America, 110(24), 9770–9775. Park, G. S., Park, M. H., Shin, W., et al. (2017). Emulating host-microbiome ecosystem of human gastrointestinal tract in vitro. Stem Cell Reviews and Reports, 13(3), 321–334. Uzel, S. G., Platt, R. J., Subramanian, V., et al. (2016). Microfluidic device for the formation of optically excitable, three-dimensional, compartmentalized motor units. Science Advances, 2(8), e1501429. Wilmer, M. J., Ng, C. P., Lanz, H. L., et al. (2016). Kidney-on-a-Chip Technology for drug-induced nephrotoxicity screening. Trends in Biotechnology, 34(2), 156–170. Wong, K. H., Chan, J. M., Kamm, R. D., & Tien, J. (2012). Microfluidic models of vascular functions. Annual Review of Biomedical Engineering, 14, 205–230. Yi, Y., Park, J., Lim, J., Lee, C. J., & Lee, S. H. (2015). Central nervous system and its disease models on a chip. Trends in Biotechnology, 33(12), 762–776. Yoon No, D., Lee, K. H., Lee, J., & Lee, S. H. (2015). 3D liver models on a microplatform: Well-defined culture, engineering of liver tissue and liver-on-a-chip. Lab on a Chip, 15(19), 3822–3837.

Shape-Memory Polymer Medical Devices Muhammad Y Razzaq, Markus Reinthaler, Mark Schro¨der, Christian Wischke, and Andreas Lendlein, Institute of Biomaterial Science and Berlin Brandenburg Center for Regenerative Therapies, Teltow, Germany © 2019 Elsevier Inc. All rights reserved.

Medical Devices Characteristics and General Requirements Upcoming Clinical Demands for Smart Medical Devices Shape-Memory Polymers Working Principle Polymer Network Structure Molecular Switches Respond to External Stimulation Programming Processes Are Mandatory for the SME From Dual-Shape Effects to Complex Shape Switching Quantitative Analysis of the SME Tailoring Shape-Memory Polymers for Use in Medical Devices Switching Temperature Cell and Tissue Compatibility Degradation Rates Sterilization Medical Devices Based on Shape-Memory Polymers Overview Surgical Devices Sutures Surgical fasteners Vascular Devices Intravascular stents Clot removal Aneurysm therapy Orthopedic Devices Filling of bone cavities Drug Delivery Devices Miniaturization of SMP Outlook Further Reading

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Glossary Medical device An instrument, apparatus, implant, in vitro reagent, or similar or related article that is used to diagnose, prevent, or treat disease or other conditions, and does not achieve its purposes in or on the human body primarily by pharmacological, immunological, or metabolic meansdbut typically by physical principles. Minimally invasive intervention Surgical technique that limits the size of incisions needed and so lessens wound healing time, associated pain and risk of infection. Polymer network A structure, in which all polymer chains are interconnected by crosslinks to form a single macroscopic entity. Programming An external mechanical manipulation of the SMP to achieve a temporarily fixed, second shape. Shape-memory polymer (SMP) A stimuli-sensitive polymeric material that has the ability to return from a deformed state (temporary shape) to its original (permanent) shape by exposure to an external stimulus (trigger), such as heat, light etc. Stent Biomaterial-based tubulous construct inserted into the lumen of an anatomic vessel or duct to keep the passage way open.

Nomenclature PCL Poly(ε-caprolactone) PCLDMA Poly(ε-caprolactone)dimethacrylate

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Rf Shape fixity ratio Rr Shape recovery ratio SME Shape-memory effect SMP Shape-memory polymer Tg Glass transition temperature Tm Melting transition temperature Tsw Switching temperature Ttrans Thermal transition temperature Ts,max Temperature, at which the maximum stress is generated

Medical Devices Characteristics and General Requirements Medical devices are systems dedicated to diagnose, prevent, or treat diseases or other conditions. Examples include man-made implants, which are employed to replace a missing biological structure, support a damaged biological structure, or enhance an existing biological structure, for example, by physical principles. These devices are usually implanted by surgical procedures. As they are directly exposed to tissue, they should be nontoxic, nonimmunogenic and sterile. From a historical point of view, dental implants are important medical devices, as this technology has been used and developed since the earliest civilizations. For example, ancient Egypts used ivory to replace single teeth, whereas metals and more recently polymers are the present materials of choice. Beside mechanical properties, as illustrated by dental devices, modern implants also need to exhibit additional features, which are highly dependent on the application area of a device. For example, hemocompatibility and low thrombogenicity are essential characteristics for devices, which are applied in the vascular system and therefore directly interact with the blood stream. Stents, artificial heart valves as well as extracorporal circulatory systems may be mentioned in this context. A group of medical devices with increasing relevance are temporary implants, which degrade after a certain period. In this way, they are able to support organ functions just as long as necessary without leaving behind foreign materials. This concept may avoid long-term side effects like chronic inflammation (foreign body response). Furthermore, the diffusion characteristics or the erosion behavior of polymeric biomaterials were prerequisites for their application as matrix materials in controlled drug-release systems. Importantly, the degradation products need to be either metabolized or excreted and should, again, be nontoxic.

Upcoming Clinical Demands for Smart Medical Devices Over the past few years, minimally invasive interventions have increasingly replaced standard surgical procedures. This trend continues and imposes specific requirements on materials and upcoming technologies. One of the key questions arising is how to adapt bulky devices in order to precisely deliver them to certain areas inside the body through small incisions. Furthermore, a general problem in device based procedures is the high anatomic variability of every individual patient, raising the need for “personalized designs.” For example, heart valves and valve dysfunctions exist in numerous anatomic variabilities and the optimal percutaneous treatment will require individually adapted devices. Here, materials that can change their shape in a predefined way on demand could potentially play an important role. An implant could be inserted into the body through a small incision in a temporary stabilized compressed or elongated shape. As soon as the implant is placed in the body, it is heated to body temperature, which may trigger its change into the desired bulky applicationrelevant shape. Furthermore, the type of shape switch might be adjusted to the individual needs of the patient. It may also be of interest to combine the shape switch with other functions, such as degradability or drug release. In this article, we discuss shape-memory polymers (SMPs) as stimuli-sensitive matrix materials that allow spatially directed shape switches and their possible applications in medical devices. This includes an overview of fields of applications including general surgery, the cardiovascular system or applications in orthopedic surgery. Furthermore, the general principle of the thermally induced SME is described, and selected examples for biomaterials with SME are introduced.

Shape-Memory Polymers Working Principle The shape-memory effect (SME) of a polymer describes the capacity to switch from a temporary shape to a permanent shape. In order to allow a SME in polymers, a number of preconditions need to be fulfilled: the material needs to exhibit a polymer network structure, suitable elasticity, a process to transfer the material to the temporary shape by elastic deformation, and molecular switches with the ability to form reversible crosslinks for fixation of the temporary shape. The permanent shape of the sample is determined

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Fig. 1 Principle of the thermally induced SME in polymers. (A) Schematic demonstration of a thermomechanical cycle of programming a SMP and its shape recovery. (B) Photo series demonstrating the macroscopic SME for a poly(3 -caprolactone) based shape-memory polymer network with a Ttrans ¼ 51 C (length of sample is 6 cm).

by the polymer network structure, while the temporary shape is achieved by a programming process that may involve stretching, compression, bending and/or twisting deformations. Recovery to the original shape is carried out by application of an external stimulus such as heat, light or moisture to remove the temporary crosslinks. A schematic demonstration of the programming and recovery of a thermally induced SMP is provided in Fig. 1A. The process of programming and recovery to the original shape can be repeated several times depending on the application.

Polymer Network Structure SMPs exhibit a network architecture consisting of netpoints, chain segments in between the netpoints, and molecular switches. The netpoints can be either of physical or chemical nature and determine the permanent shape to the polymer network. Multiblock copolymers, whose morphology consists of at least two types of phase-segregated domains, are examples for physically crosslinked polymers. Here, the domains providing the highest thermal transition temperature act as netpoints. The domains associated with the second highest thermal transition can serve as the molecular switch. Such materials are examples of temperatureinduced SMPs. Shape-memory polymer networks with covalent netpoints can be obtained either by in situ crosslinking during polymer synthesis or by postpolymerization, which includes radiation or chemical methods. Due to the covalent nature of netpoints, covalent networks are often capable of a more quantitative switching toward the original shape. Besides temperature-sensitive SMPs with covalent network structure, also other types of molecular switches can be introduced.

Molecular Switches Respond to External Stimulation In principle, molecular switches can be selected from a variety of chemical components as long as they are able to form crosslinks that are reversible in nature. The most frequently reported SMPs are thermo-sensitive and can be activated by heat as an external stimulus. Here, the temporary physical crosslinks base on domains, which are formed from switching segments and are associated with distinct thermal transition temperatures Ttrans (e.g., melting transition Tm or glass transition Tg) (Fig. 2). The fixation of a temporary shape can in this case be achieved by vitrification or crystallization of switching domains upon cooling, while shape recovery can subsequently be induced by heating. The macroscopic shape recovery of a SMP with crystallizable switching domain is shown in Fig. 1B. Heat can be applied directly, but also indirectly in a non-contact mode. Indirect heating can be realized by incorporated dyes or magnetic nanoparticles, which can transform energy provided by light or alternating magnetic fields, respectively, into heat locally. Other types of molecular switches base on reversible covalent bonds, which are formed between two functional groups incorporated in the polymer network at a given condition, and can be cleaved at a different condition. Examples include the lightinduced SME, which can be controlled independently from heat and is obtained by incorporating light-sensitive reversibly reacting

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(B)

Fig. 2 Schematic representation of molecular architecture of covalently crosslinked SMPs (black lines: switching segments, red circles: netpoints) with (A) amorphous and (B) crystallizable switching segments.

molecular switches into the polymer networks. Another example is SME induced by ultrasound cavitation, which is based on reversible dissociation of metal–ligand coordination bonds under the effect of ultrasonic waves. Furthermore, redox active moieties can also be utilized to create reversible covalent bonds in SMPs. For instance, cellulose derivatives having a crosslinkable mercapto group enabled a redox sensitive SME. However, the shape-recovery ratio (Rr), a value that would be 100% for full recovery to the original shape, decreased with the repetition of the redox treatment because side reactions reduced the number of available mercapto groups.

Programming Processes Are Mandatory for the SME Along with the polymer network structure, a suitable elasticity is required for SMPs as mentioned above. This elasticity is essential to allow the deformation of the polymer during a programming step and, importantly, to drive the shape recovery to the permanent shape. Based on their mandatory role to realize a SME, the programming technologies are also called shape-memory creation procedures (SMCP). In case of temperature-sensitive SMP, the programming of the polymer network is carried out by the action of an external stress applied at a temperature T > Ttrans (Fig. 1). At this temperature, the chain segments are flexible and can be deformed to a less coiled conformation, resulting in a loss of entropy. The overall deformation capability can be adjusted by the length and flexibility of the polymer chain segments, from which the network has been constructed. Cooling the sample to T < Ttrans causes the crystallization or the vitrification of the switching segments, preventing the immediate recoiling of the polymer chains and thus fixing the temporary shape of the polymer sample. The macroscopic recovery to the permanent shape, in this case induced by heating to T > Ttrans, is enabled by entropy elasticity of the switching segments that move toward an entropically favored conformation.

From Dual-Shape Effects to Complex Shape Switching By introducing a second type of switching segments into the polymer network, the capability of memorizing two distinguishable shapes can be implemented, thus extending the dual-shape capability to a triple-shape effect. Such a triple-shape polymer network consists of at least two segregated domains with two transition temperatures Ttrans,A and Ttrans,B and can change on demand from a first shape (A) to a second shape (B) and from there to a third shape (C) when stimulated by two subsequent temperature increases. An example of such a multiphase polymer network is an AB copolymer network based on crystallizable poly(ε-caprolactone) (PCL, Ttrans,A ¼ Tm,PCL z 50 C) and amorphous poly(cyclohexyl methacrylate) (PCHMA, Ttrans,B ¼ Tg,PCHMA z 140 C) segments. By adding further distinct phase transitions in a polymer or by using a polymer with a broad thermal transition associated with the switching domain, the number of temporary shapes can be increased. In this way, a multishape effect (MSE) can be enabled in polymers. These polymers are capable to memorize and recover more than two predefined shapes. Such an effect has been realized in covalently crosslinked copolymer networks, interpenetrating polymer networks, and in multimaterial polymer systems. However, the dual-shape and MSEs are categorized as “one-way” phenomena as they are not reversible and an external manipulation is always required to again program the samples. A reversible actuation can be observed in polymer networks containing at least one crystalline switching domain. Here, a constant stress (sc) has to be maintained on the sample at all times, which induces crystallization along the direction of the load. On the molecular level such a reversible SME (rSME) is based on the crystallization-induced elongation (CIE) of preoriented switching chain segments, which occurs during cooling to temperatures Tlow < Tm under a specific sc, while heating to temperatures above Tm results in melt-induced contraction (MIC). By using a polymer network with two crystallizable switching domains, an rSME has been demonstrated, which is possible under stress-free conditions. This actuation effect requires a modified thermomechanical programming procedure and is named reversible bidirectional SME (rbSME). In principle, a SME is not limited to large samples but also works for microstructures, which are useful for miniaturized biomedical devices and microfluidic systems. A variety of shape-memory micro-objects has been explored and their one-way and rSMEs are

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investigated. However, specific thermomechanical programming and characterization procedures at microscale are mandatory to observe the SME at micro-level.

Quantitative Analysis of the SME The thermally induced SME of macroscopic samples can be quantified in cyclic, thermomechanical tests that consist of the programming as well as the recovery step. A variety of test protocols has been developed, which are based on recovery under stress-free or strain-controlled conditions. They allow determining the switching temperature (Tsw) of thermo-sensitive SMPs, at which the maximum recovery takes place during stress-free recovery conditions, or the temperature, at which the maximum stress is generated (Ts,max) during strain-controlled recovery process. Other important shape-memory properties include the extent, to which the temporary shape can be fixed (shape fixity ratio Rf) and the permanent shape is recovered (shape recovery ratio Rr). These shape-memory characteristics strongly depend on the polymer’s structural design as well as on the parameters (e.g., heating or cooling rates or the type of mechanical deformation) applied during the SMCP.

Tailoring Shape-Memory Polymers for Use in Medical Devices Switching Temperature Heat can be considered as the most commonly evaluated switching mechanism in the body, which, however, must be carried out at a physiological temperature at or slightly above 37 C. For SMPs with a Tsw lower than body temperature, a premature recovery before implantation can take place, whereas hyperthermia to induce SMPs with Tsw well above 37 C may cause cell and tissue damage. General principles to tailor Tsw include modification of the respective Ttrans by changing the segment length between netpoints and / or by copolymerizing different types of monomers. Some of the common monomers used for the synthesis of degradable SMPs with tailorable Tsw are provided in Fig. 3A. Copolymerization of these cyclic monomers has also enabled controlling the hydrolytic degradation rate of the resulting polymers.

(A)

(B) Glycolide

L,L-Dilactide

oligo[(rac-lactide)-coglycolide]tetrol

Diglycolide

D,D-Dilactide

I-[OH]4

Caprolactone oligo[(rac-lactide)-co-

(ε -caprolactone)]tetrol

rac-Dilactide

p-Dioxanone

ε -Caprolactone

Morpholino-2,5-dione pentaerythriteI =

p-Dioxanone

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(D) (C)

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ε (%)

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75

MW Loss

50 25 0 0

10

20

30 40 T (°C)

50

60

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Homogeneous Hydrolysis

70 Random PLGA

Heterogeneous Hydrolysis

Time

Fig. 3 (A) Common cyclic monomers used for synthesis of biodegradable SMPs with different Tsw, (B) synthesis of star-shaped rac-dilactide-based macrotetrols used for the synthesis of polyesterurethanes with different Tsw, (C) strain recovery process of different copolyesterurethanes with Tsw ranging from 14 C to 56 C (Tsw ¼ inflection point of recovery curve), (D) effect of monomer sequence on degradation behavior of PLGA copolymer. Reprinted with permission from Biomacromolecules 2009, 10, 975–982, Copyright 2009, American Chemical Society and J. Am. Chem. Soc., 2011, 133 (18), 6910–6913, Copyright 2011, American Chemical Society.

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The Tm of semicrystalline polymer networks prepared by photo-crosslinking of oligo(3 -caprolactone) dimethacrylate could be varied in a range from 30 C to 50 C depending on the molecular weights of the oligomeric precursor. Similarly, the terpolymer poly(L-lactide-ran-glycolide-ran-trimethylene carbonate) presented Tgs in the range of 38 C–42 C, which could be tailored by using different weight contents of L-lactide. Furthermore, the Tgs of amorphous polymer networks from star-shaped rac-dilactide-based macrotetrols and a diisocyanate can be controlled systematically by incorporation of p-dioxanone, diglycolide, or ε-caprolactone as comonomer (Fig. 3B and C). The resulting Tsw could be adjusted between 14 C and 56 C by selection of comonomer type and ratio without affecting the advantageous elastic properties of the polymer networks.

Cell and Tissue Compatibility SMPs designed for biomedical applications are expected to be nontoxic. Therefore, systematic studies both in vitro and in vivo are required to better address the concern of biocompatibility. The use of natural polymers such as polypeptides or polysaccharides is an important approach to design biocompatible SMPs. Another strategy is to use synthetic macromolecules built from compounds naturally occurring in the human body to minimize potential toxic side effects. For example, polyesters can be made based on molecules abundant in human metabolism, like glycerol and lactic, citric or sebacic acids. To create such shape-memory polyesters, polycondensation of polyols and diacids endogenous to human metabolism is carried out. A variety of biocompatible polymers based on poly(glycerol-sebacate), poly(glycerol-dodecanoate), and polydiol citrate (using 1,8-octanediol, 1,12-dodecanediol, and 1,16hexadecanediol together with citric acid) have been synthesized and their shape-memory performance was investigated. Furthermore, a bio-based poly(propylene sebacate) was synthesized from 1,3-propanediol, sebacic acid, and itaconic acid produced via fermentation or extraction. The Tsw could be tuned in a range from 12 C to 54 C by a copolymerization with diethylene glycol to tailor the chain flexibility. In vitro fibroblast response demonstrated that these polymers are potentially cell compatible and are promising candidates for medical devices. Along with biopolymers, the compatibility of commercially available polyurethane based SMPs (Mitsubishi, Japan) is investigated. The absence of both cytotoxic effects on L929 fibroblast cultures and mutagenicity with strains of Salmonella typhimurium were first relevant results to allow further development of these materials to be used for endovascular interventions. Also, the behavior of L929 mouse fibroblasts and other cells from human and rat in contact with biodegradable oligo(ε-caprolactone)dimethacrylate (OCLD) based SMP networks was investigated. The differentiation capacity of mesenchymal stem cells into osteoblasts and adipocytes was supported by the polymer network. The shape-memory effect did not affect the majority of the adherent cells. While in vitro studies may give a first hint on cellular response for selected types of cells, often based on immortalized cell lines, the exploration of in vivo behavior of implanted materials provides insights in complex cellular responses of multiple involved cell types. As the place of application (e.g., subcutaneous or intravascular), the disease and surgery-associated trauma (e.g., bone fractures or intervention-induced tissue response based on surgical skills), the type and properties of SMP material including potential impurities (e.g., chemical composition, surface properties, sterility, degradation products), the mode of action of the SMP device (e.g., pressure on adjacent tissues) and individual factors associated to the treated person/animal (e.g., immunological state) all can affect in vivo response, a generalized statement on histocompatibility of SMP cannot be given. However, several preclinical studies demonstrated a promising in vivo behavior. For instance, tissue integration and angiogenesis was observed in rats along with slow degradation of multiblock copolymers composed of oligo(p-dioxanone)- and oligo(ε-caprolactone)-segments. Photocrosslinked SMP networks from poly(ε-caprolactone)-co-(a-allyl carboxylate ε-caprolactone) showed no inflammatory response when implanted adjacent to femoral artery ligations in mice, a model for ischemia. Polymer networks from poly(D,L-lactide) containing polyhedral oligomeric silsesquioxane (POSS) netpoints demonstrated that different phases of generally mild inflammatory responses can be distinguished after subcutaneous implantation in rats before and after the onset of material degradation. The associated formation of fibrous capsules around the implant depict the cellular response to the foreign body, which is characteristic to most implants and may preferentially be resorbed with time. In specific applications, however, fibrous tissue formation in response to the implant may be desired to permanently alter the implantation site, as long as this response is well controlled, aseptic, and local. A low level chronic inflammatory reaction to a spiral shaped SMP occluder from poly(DL-lactic acid)-based poly(urethane urea) inserted for contraception in rabbit fallopian tubes was reported to result in a permanent tubal occlusion due to an almost complete replacement of mucosa and muscle by fibrous tissue. For occlusion of aneurisms with polyurethane SMP foams (see “Aneurism Therapy” section), the desired ingrowth of connective tissue with only a mild inflammatory reaction was confirmed.

Degradation Rates The degradability of SMP can be achieved by introducing hydrolyzable bonds in the polymer chains, which can be cleaved under physiological conditions. Polymeric materials may be degraded via chemical and enzymatic oxidation upon exposure to body fluids and tissues, via hydrolysis catalyzed by acids, bases, or salts, and via enzyme-catalyzed hydrolysis reactions. The degradation rates of the SMPs can be tailored by copolymerization and crosslinking procedures. For example, the hydrolytic degradation rate of SMP based on oligo[(ε-hydroxycaproate)-co-glycolide] dimethacrylates and butyl acrylate can be tailored by adjusting the weight content of glycolide units and n-butyl acrylate weight content during the crosslinking reaction. The degradation rate could be accelerated by increasing the glycolide content and reduced by increasing amounts of n-butyl acrylate units. The enzymatic degradation of crosslinked poly(ε-caprolactone)-based SMP can also be controlled by adding pendant coumarin groups. Here, the coumarin groups

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hindered the excess of enzyme to the degradable polymer chains resulting in a decreased degradation rate. It should be noted that also the order of how monomers are build in the polymer chain during copolymerization affects the hydrolytic degradation of the polymers. A dramatically different hydrolysis rate of alternating PLGA and random sequence PLGA was observed (Fig. 3D). Along with chemical composition, other factors such as water diffusion, monomer solubility and diffusion, material homogeneity, processing technique, and device geometry and size play a role to control the degradability. Furthermore, polymer blending and surface modification can be used to adjust the degradation of the SMPs.

Sterilization Sterility is a mandatory feature of SMP medical devices for implantation. However, the sterilization method can influence the polymer structure, SME functionality, biocompatibility and thus performance of a device. For temperature-induced SMP, the sterilization technique should ideally be operated below Ttrans, which typically excludes heat (160 C) or steam sterilization (121 C) from the list of possible methods at least for thermoplastic or programmed SMP samples. Thus, low-temperature sterilization (LTS) methods based on g- or e-beam-irradiation, hydrogen peroxide vapor, low-temperature gas plasma and ethylene oxide (EO) may be preferred. Among them, a careful selection is needed as also these techniques may induce changes in polymer structure under certain conditions. For instance, hydrogen peroxide vapor or EO may react with certain chemical groups of the SMP, while g-, e-beam, and plasma sterilization may be associated with changes in the molecular structure of the polymer. Recently, low-temperature plasma (LTP) sterilization was used for SMP polyurethane foams and the effects of sterilization on the chemophysical, thermomechanical, and shape-memory performance of polyurethane were investigated. Plasma sterilization (Sterrad100St, H2O2, 52 min) did not influence the SME of PU foams, but an increase in the open porosity was observed. Furthermore, plasma treatment had no significant effect on the material’s cytotoxicity on L929 cells in vitro. In another study, six sterilization techniques (LTP, steam, ethylene oxide, e-beam radiation, gamma radiation, and nitrogen dioxide) were evaluated for acrylate based SMP networks from poly(ethylene glycol) dimethacrylate copolymerized with tert-butyl acrylate or methyl methacrylate. The samples were shown to be cytotoxic after LTP sterilization as the surface was oxidized, which negatively affected the biocompatibility. Gamma radiation decreased the Tg and significantly increased the rubbery modulus. The other sterilization methods did not significantly alter the thermomechanical properties. The sterilized SMPs exhibited a full shape recovery under free-strain conditions even after a storage of 1 year at 20 C. The Tsw increased by up to 9 C and the speed of the strain recovery by up to 9% due to physical aging. Storage and aging of SMPs would allow for faster activation in vivo if the increase in Tsw did not negatively affect recovery.

Medical Devices Based on Shape-Memory Polymers Overview Owing to a combination of advantageous features such as lightness, low cost, easy processability, high recoverable strains, tailored thermomechanical properties, and shape-recovery temperatures as well as possibly in vivo degradation, SMP could be beneficially used in the fabrication of various medical devices, particular for minimally invasive applications. However, the promise of SMPs in biomedical applications has yet to be fully realized and so far is restricted to conceptional studies in some clinical fields. Potential medical applications of SMPs are provided in Fig. 4. In this section, selected SMP-based medical devices are discussed, with special emphasis on those works where evidence of functionality either in vitro or in vivo has been obtained.

Surgical Devices Sutures One of the cornerstones in conventional surgery is to create stable tissue connections or to secure implants, which has traditionally been performed by surgical sutures. In order to safely adapt tissue margins, the suture has to be tied at an extent providing sufficient stress on the wound lips. Both, inadequately high or low suturing forces may result in poor healing. Especially in minimally invasive (keyhole) surgeries, stitching may become a major limitation due to lack of space. Selftightening sutures may therefore be of particular relevance in these types of operations. The first biodegradable self-tightening suture was based on a multiblock copolymer synthesized from oligo(ε-caprolactone)diol and crystallizable oligo(p-dioxanone)diols. The suture was first stretched (about 200%) to achieve a temporary shape. Modification of the stretching conditions or the molecular structures of the multiblock copolymer were used to adapt the exerted tension of the SMP suture. Series of experiments were conducted to demonstrate the feasibility of these shape-memory sutures. Once the suture was applied for example to close an incision, the SME capability of the polymer was activated by the body temperature, tying together the wound margins. As the adopted polymer may also be degradable, there may be no need to remove the suture after the incision has healed. The sutures proved to be cell compatible when incubated with 3T3 fibroblasts, umbilical endothelial cells and human fibroblasts. For the selected programming conditions, shape-memory based contraction of the polymer suture was 25%, when tested in 0.9% (w/v) saline, blood or air at 38 C–45 C.

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Surgical fastener

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Sutures Stents for intraluminal prothesis

Soft tissue fixation to bones General surgical uses

Archwires

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Intravascular applications

Ocular

Heat packs

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Clot removal devices

Thermally-induced SMP material platforms

Brain Microgripper & retrieval devices

Neuronal electrode Urogenital tract Retrieval devices

Percutaneous vascular access products

Fertility control plug Stents for intraluminal prothesis

Dialysis needle

Anchoring cannula

Catheter for apheresis

Fig. 4 Potential medical applications of shape-memory polymers. Reprinted from Comprehensive Biomaterials 2011, 479–496, Copyright 2011, with permission from Elsevier.

Surgical fasteners Another application of SMP has recently been suggested for the fixation of medical devices during surgery. The concept involves needles of temporary straight shape that are pierced through the implanted device, such as a hernia mesh, and the respective subjacent tissue. Upon heating to body temperature, both ends of the needle recover to a helical shape, thus fixing the implant in the desired place. The temperature-induced self-tightening of a bioabsorbable PLGA based subcuticular staple has been demonstrated. The use of these fasteners reduces the need for staples of different sizes and provides gentle force to close the wound. Furthermore, cannulas containing SMP-based anchoring structures are proposed for minimally invasive procedures. The cannulas are utilized to provide an access port for surgical procedures. The proposed cannula include an elongated longitudinal shaft having a distal opening, a proximal opening for receipt of the operating instruments, an anchoring member including a SMP on the outer surface of the elongated longitudinal shaft, and a sliding tube mounted on the outer surface of the cannula. After inserting the cannula, it can be prevented from migrating or slipping out through the incision using the SMP-based anchoring structures on the distal portion of the cannula.

Vascular Devices The use of medical devices inside of blood vessels is necessary in highly critical medical conditions such as vessel occlusion. High precision of device insertion to the site of placement, control of device expansion and/or fixation in the tissue, and long-term performance are typical key requirements of vascular devices, which make these applications highly relevant for SMP-based implants but at the same time also more challenging than some other applications introduced above.

Intravascular stents Stents are tubular structures, designed to re-open a stenosed/occluded vessel and therefore to establish adequate blood flow as shown in Fig. 5. As demonstrated in coronary interventions, supporting the vessel wall by implanting these devices has become one of the crucial steps to ensure durability of the treatment as they prevent elastic recoil. Traditionally, metallic stents were mounted on a balloon catheter and delivered from a peripheral access. Beside balloon-expandable stent technologies, selfexpandable designs have been developed, which are typically used to treat peripheral vascular disease. Balloon-expandable metal stents are commonly made of stainless steel, cobalt or titanium, whereas nitinol-alloys are mainly used in self-expandable stents.

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Vessel

Stent Plaque

Fig. 5

Reopening of a stenosed vessel by using a stent.

Polymer based scaffolds, which offer the option of biodegradability, have recently been introduced in the interventional treatment of coronary artery disease. The idea behind this technology is to support the vessel wall just as long as necessary and therefore avoid long-term complications like late stent thrombosis. This problem has especially been observed in drug eluting stents. In this context, degradation time seems to be an important parameter and according to current opinion should be a period of 6–9 months. Recently, the ABSORB (Abbott vascular) and the DESolve (Elixir Medical corp.) devices based on degradable polymers have been granted CE marks and are commercially available. However, so far metallic stents are still the gold standard in cardiovascular medicine and the field of polymer stents is still under development. SMP-based stent technologies have not been in clinical use so far, however due to a strong clinical trend toward minimally invasive techniques, SMP-based stent technologies have been continuously explored. These stents can be preprogrammed to a compact shape, which on the one hand may be activated at body temperature, resulting in natural deployment without need for auxiliary devices. Proposed stent designs include spiral shaped structures based on copolymer networks of star-shaped poly(ε-caprolactone) (acting as switching segment) and polyester poly[(R)-3-hydroxybutyrate-co-(R)-3-hydroxyvalerate] (PHBV) (acting as hard segment). These spirals would partially uncoil upon shape recovery, a process resulting in a twisting of each end of the spiral and possibly a reduction in device length. In another study, either solid or faster expanding, perforated tubular stents based on tert-butyl acrylate and poly(ethylene glycol)dimethacrylate (PEGDMA) crosslinkers were programmed to a small temporary shape by furling and rolling, which would have to be reversed at the site of placement. In both cases, practical concerns in the precision of placement would need to be addressed. More advanced concepts have considered a highly perforated tubular design with small struts, which should be programmed in situ by aid of balloon catheters. This approach would allow stent removal by diameter reduction via the shape-memory effect. In addition, SMPU based neurovascular stent prototypes were fabricated. These tubular devices were laser etched for more flexible stents that also could navigate through small tortuous vesselsdhowever, it has been revealed that resistance to withstand the critical collapse pressure in the vessel at elevated body temperature may not be given in all cases.

Clot removal Micro-devices for mechanical clot removal have recently attracted significant attention in minimally invasive surgery as an alternative to thrombolytic agents, as these drugs are associated with bleeding complications. First experimental devices have already been used and offer a transcatheter approach, guided from a peripheral access. After crossing the clot, the SMP-based device is activated by an external stimulus and converted in a grabber-like shape, which allows extraction of the embolus. For example, a corkscrew temporary shape had been proved to be effective under physiological flow and pressure conditions (Fig. 6). Controllable shape switching can be achieved by indirect heating using a light-absorbing dye incorporated into the SMP or by electro resistive heating of hybrid devices with a nitinol core covered with a polymeric shell. This hybrid device enabled a higher recovery and therefore retraction force in the blood flow. The functionality was demonstrated in vitro and in vivo in a water-filled silicone neurovascular model and a rabbit carotid occlusion model, respectively.

Aneurysm therapy An aneurysm is a localized, blood-filled balloon-like bulge in the wall of a blood vessel. Especially in the brain, aneurysms can cause symptoms by compressing adjacent structures or can rupture, which is a life threatening complication. Based on the aneurysm anatomy and the clinical circumstances, endovascular or surgical treatment-strategies are available. Both treatment modalities attempt to disconnect the aneurysm from the blood stream either by clipping, coiling, or stenting. The idea of coiling is to insert a filling material into the vascular bulge (usually platinum coils), inducing thrombosis and therefore blocking off the aneurysm

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Microcatheter

Clot Straightened coil

1. Blood vessel with a clot

Fig. 6

Helical coil

2. Insertion of a microcatheter with straightened coil to the site of clot

3. Coil returning back to its original helical shape by laser heating

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Procedure of mechanical removal of embolus using a SMP-based clot removal device.

cavity. In 2013, 28,000 people in Western Europe received a coil based minimally invasive aneurysm treatment (Clin. Neuroradiol. 2015, 25 (2), 317–324). However, coils are prone to incomplete cavity filling or to rupture the aneurysm wall by coil puncture. The utilization of SMPU spherical foams as filling materials has shown to be a good alternative to metallic coils. These devices are softer compared to metallic coils, reducing the risk of rupture, and may also improve the closure performance due to enhanced alignment to irregular aneurysm shapes. From a procedure point of view, the device can be programmed to its temporary shape for delivery by crimping onto the catheter. The device should revert back into its permanent shape inside the aneurysm when exposed to the body temperature (Fig. 7). Furthermore, by tuning the pore size of a SMP foam by the composition and synthesis condition, cellular infiltration can be controlled to support the blood clotting and anchoring of the clot in the aneurysm. An important requirement is radio-opacity of devices as the procedure is fluoroscopy guided, for which tungsten particulate filler were incorporated into the SMP. In vitro models of shape-memory polymers for embolization of aneurisms were explored to predict the

Aneurysm

Catheter

Stent

compressed SMP-Foam

1. Vessel with Aneurysm Fig. 7

2. Stenting and delivery of the compressed Foam

expanded SMP-Foam

3. Foam expansion

4. Remainingfilled aneurysm

Procedure of aneurysm treatment by using a SMP-based occluder device and additional stent.

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stresses caused by the SMP, thermal and fluid dynamic changes, and changes in vascular dynamics. Also in vivo studies have been conducted with SMP foams versus metal coils in a vein pouch aneurism model, which provides similar aneurism sizes to those found in the human intracranial vasculature. These studies suggested a much higher occlusion of aneurism volume by SMP foams compared to bare metal coils when examined 180 days after implantation. While significant research has gone into this application of SMPs, FDA clearance for this application is still required. In addition to neurovascular applications, SMP occluder systems including foams as well as fibrous SMP coils are explored in vivo in peripheral vasculature to treat, for example, aneurisms or hemorrhage, or to block the blood supply to tumors.

Orthopedic Devices Filling of bone cavities Bone tissue typically allows spontaneous healing of defects up to a critical size. Larger defects above the critical size are currently preferentially treated by using bone autografts from sources like the iliac crest or the ribs (gold standard), which are not always available and/or bring additional burden to patients during harvesting. An unmet need are devices that would enable the simple access and optimal adaption to irregular shapes of defect cavities as well as eventually the regeneration of bone. Other than synthetic grafts that have to be tailored manually to the defect for the respective patient, SMP foams can be fixed in a compressed temporary shape, which allows simple insertion of the implant to the defect. By shape recovery, the foam expands until it fits the cavity boundaries and remains anchored by applying its recovery force to the cavity boundaries during constraint recovery. This anchoring method works like a press-fit fixation. Often, materials with a temperature-induced shape recovery, in some cases at critically high temperatures, have been proposed. An alternative may be materials with a water-induced shape-memory effect, which can be realized by water-induced plasticization of glassy switching domains. An example can be found in gelatin-based porous hydrogels stabilized via covalent crosslinks. When water was added to compressed dry-state samples, the original shape was recovered (Rr of 88%  5% to 95%  5%) along with a reduction of the Tg of the switching domains from 50 C–60 C to 0 C. The inherent cell-binding motives of gelatin-based material and the merging of smaller pores to larger sized cavities during degradation were beneficial as concluded from cell proliferation, support of cell differentiation, and a purely material-induced regeneration in different animal models with outcomes comparable to less easily accessible cancellous bone graft. Beyond filling of bone defects, orthopedic applications of SMP expand also to the fixation of tendon and ligaments. For a biostable polyether ether ketone SMP medical device, which is designed for nonrotational insertion in (drilled) bone cavities and prevents rotational force on the tissue to be fixated, an FDA clearance has been granted.

Drug Delivery Devices In recent years, degradable SMPs have been equipped with drug-release functionality by drug incorporation, which leads to multifunctional SMPs. A detailed analysis including examinations of the kinetics of drug incorporation and in vitro release as well as tissue compatibility, in vivo release and degradation are pivotal for these devices. Furthermore, the incorporation of drugs may result in a disturbance of polymer morphology or thermomechanical properties, and thus may lead to an impaired SME of a temperature-induced SMP. Therefore, investigations concerning the interplay between drug payload, SME and polymer degradation are required in order to obtain tailored properties of a device for a specific application. In recent years, a variety of SMP-based materials for drug delivery systems has been proposed. Degradable SMP networks prepared by UV-curing of oligo[(ε-caprolactone)-co-glycolide]-dimethacrylates precursors were loaded with ethacridine lactate or enoxacin as examples of hydrophilic and hydrophobic test drugs, respectively, by swelling and by incorporation during network synthesis. The investigation of the SME and drug delivery for this set of materials showed that the shape-memory functionality at relevant temperatures (28 C–42 C) was maintained after drug incorporation and the diffusion-controlled release was timely independent of the polymer degradation. Similar systems were further proposed as implantable devices with body temperature-induced shape switch to avoid implant migration by anchoring in the tissue, enabling local or systemic drug delivery. Further covalent polymer network systems were shown to have controlled drug-release capacity and, in animal studies, slowly degraded over the period of weeks when implanted in rats. Although there are not yet massive publications in this field, drug loaded SMPs hold substantial promise for applications in the field of controlled drug release.

Miniaturization of SMP SMPs are promising candidates to generate switchable micro-objects or miniaturized biomedical devices such as drug delivery systems. Recently, shape-memory particles composed of biodegradable multiblock copolymers of poly(u-pentadecalactone) and poly(ε-caprolactone) have demonstrated the ability to be switched from prolate ellipsoids to spheroid. Recently, SMP microparticles based on poly(ε-caprolactone) and poly(ethylene glycol) have shown a reversible switch from spherical to ellipsoidal shape by cyclic heating and cooling between 0 C and 43 C. In another study, a soft lithographic technique is used to fabricate SMP micro cuboids. Resulting microcuboids enabled a nano (surface roughness) and micro-level (cuboidal geometry) shape recovery process after programming by different compression ratios.

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Miniaturization of SMP-based features has also been used to design responsive micro-structured surfaces with switchable material properties such as wetting, adhesion and friction. For example, micro-tipped SMP surfaces were used to fabricate reversible dry adhesives. The micro-tips in the deformed state achieved a tight contact and bonding with the substrate. Upon heating, the microtips returned to their protruded state, causing de-bonding. In another example, surface wetting was controlled by using micro-pillars developed on the surface of SMP. The deformed and original/recovered SMP pillar array displayed distinct wettability, which was characterized by the static/dynamic water contact angles and the droplet sliding angle. Furthermore, as cell growth is considered to be sensitive to the topographic features of a substrate, different cell behaviors were observed in response to changes of the surface micro-structure changes in SMPs.

Outlook In summary, the unique capability of shape-memory polymers to undergo predefined shape switches between two or more desired shapes as well as the recent demonstration of reversible bidirectional actuation by SMP materials underline the high potential of SMP in biomedical applications. This is reflected also in the high attention that SMP experience in academic research. Some first SMP-based medical devices have progressed far towards clinical application including successful examination in approval processes. These systems are pioneers for other SMP-based medical devices that will follow.

Further Reading Behl, M., Kratz, K., Noechel, U., Sauter, T., & Lendlein, A. (2013). Temperature-memory polymer actuators. Proceedings of the National Academy of Sciences of the United States of America, 110(31), 12555–12559. Chen, M. C., Tsai, H. W., Chang, Y., Lai, W. Y., Mi, F. L., Liu, C. T., Wong, H. S., & Sung, H. W. (2007). Rapidly self-expandable polymeric stents with a shape-memory property. Biomacromolecules, 8(9), 2774–2780. Filion, T. M., Xu, J., Prasad, M. L., & Song, J. (2011). In vivo tissue responses to thermal-responsive shape memory polymer nanocomposites. Biomaterials, 32(4), 985–991. Hardy, J. G., Palma, M., Wind, S. J., & Biggs, M. J. (2016). Responsive biomaterials: Advances in materials based on shape-memory polymers. Advanced Materials, 28, 5717–5724. Horn, J., Hwang, W., Jessen, S. L., Keller, B. K., Miller, M. W., Tuzun, E., Hartman, J., Clubb, F. J., & Maitland, D. J. (2017). Comparison of shape memory polymer foam versus bare metal coil treatments in an in vivo porcine sidewall aneurysm model. Journal of Biomedical Materials Research Part B Applied Biomaterials, 105(7), 1892–1905. Jung, F., Wischke, C., & Lendlein, A. (2010). Degradable, multifunctional cardiovascular implants: Challenges and hurdles. MRS Bulletin, 35, 607–613. Lendlein, A., & Langer, R. (2002). Biodegradable, elastic shape-memory polymers for potential biomedical applications. Science, 296, 1673–1676. Lendlein, A., Behl, M., Hiebl, B., & Wischke, C. (2010). Shape-memory polymers as a technology platform for biomedical applications. Expert Review of Medical Devices, 7, 357–379. Neffe, A. T., Pierce, B. F., Tronci, G., Ma, N., Pittermann, E., Gebauer, T., Frank, O., Schossig, M., Xu, X., Willie, B. M., Forner, M., Ellinghaus, A., Lienau, J., Duda, G. N., & Lendlein, A. (2015). One step creation of multifunctional 3D architectured hydrogels inducing bone regeneration. Advanced Materials, 27(10), 1738–1744. Rodriguez, J. N., Clubb, F. J., Wilson, T. S., Miller, M. W., Fossum, T. W., Hartman, J., Tuzun, E., Singhal, P., & Maitland, D. J. (2014). In vivo response to an implanted shape memory polyurethane foam in a porcine aneurysm model. Journal of Biomedial Materials Research Part A, 102(5), 1231–1242. Serrano, M. C., & Ameer, G. A. (2012). Recent insights into the biomedical applications of shape-memory polymers. Macromolecular Bioscience, 12, 1156–1171. Small, W., Singhal, P., Wilson, T. S., & Maitland, D. J. (2010). Biomedical applications of thermally activated shape memory polymers. Journal of Materials Chemistry, 20, 3356–3366. Wischke, C., Neffe, A. T., & Lendlein, A. (2010). Controlled drug release from biodegradable shape-memory polymers. Advances in Polymer Science, 226, 177–205. Wischke, C., Schossig, M., & Lendlein, A. (2014). Shape-memory effect of micro-/nanoparticles from thermoplastic multiblock copolymers. Small, 10(1), 83–87. Yakacki, C. M., Shandas, R., Lanning, C., Rech, B., Eckstein, A., & Gall, K. (2007). Unconstrained recovery characterization of shape-memory polymer networks for cardiovascular applications. Biomaterials, 28, 2255–2263.

REGENERATIVE ENGINEERING Adult Bone Marrow-Derived Stem Cells: Immunomodulation in the Context of Disease and Injury AE Ting and SA Busch, Athersys, Inc., Cleveland, OH, USA © 2019 Elsevier Inc. All rights reserved.

Introduction Immunomodulatory Properties of BMSCs Graft-versus-Host Disease (GvHD) Type 1 Diabetes Spinal Cord Injury Multiple Sclerosis Stroke Potential Safety Issues Associated with BMSCs Future Perspectives References

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Glossary Allogeneic from an organism of the same species but dissimilar in genotype. Angiogenesis formation of new blood vessels. Autologous from the same organism. Cytokine small cell-signaling molecules important in cell–cell communication and immune modulation. Immunomodulatory the ability to suppress or enhance an immune system response. Xenogeneic cells originate from a donor of a different species than the recipient.

Introduction Adult bone marrow-derived stem cells, or BMSCs, are a heterogeneous population of multipotent progenitor cells. BMSCs are adherent and form colonies when grown in culture and are capable of multilineage differentiation as they have the capacity to produce multiple tissue types within the mesenchymal lineage including bone, adipose, and cartilage. These cells are positive for the cell surface markers CD73, CD90, and CD105, but negative for CD11b, CD45, CD79a, and HLA-DR; however, the phenotype of these cells is dependent on passage, cell density, and culture media (Ankrum and Karp, 2010). Under the umbrella of BMSCs are multiple characterized cell types including mesenchymal stem cells, marrow isolated adult multilineage inducible cells, and multipotent adult progenitor cells (Mays et al., 2007; Le Blanc, 2002). While these cells were first isolated from bone marrow, they have since been discovered in fat and several other adult tissues, and it has been determined that the source also dictates their phenotype and differentiation capacity. BMSCs were initially considered for therapeutic applications based on their multilineage differentiation capacity and their potential to replace tissue, but it was later determined that these cells primarily exert their therapeutic effects through the secretion of paracrine factors that can significantly modulate both the immune response and inflammatory processes in the context of disease and injury. Furthermore, BMSCs home to sites of injury using the CXCR4-SDF1 chemotactic axis and the secretion of stromal cellderived factor 1 (SDF1) by BMSCs can lead to the recruitment of endogenous stem cells to aid in repair (Sordi, 2009). The shift in scientific thought from considering BMSCs as replacement cells that would differentiate and regenerate injured tissue to the current hypothesis that BMSCs exert benefit through the production of trophic factors and modulation of inflammation has also led researchers to change their therapeutic approach from local to systemic administration. The ability to isolate, culture, expand, and characterize the properties of BMSC have led to an ever-increasing exploration of the clinical utility of BMSCs. Currently, more than 100 clinical trials are being conducted using BMSCs to treat a variety of diseases including: graft-versus-host disease (GvHD), acute myocardial infarction, diabetes, multiple sclerosis (MS), spinal cord injury (SCI), and stroke (Ankrum and Karp, 2010).

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Immunomodulatory Properties of BMSCs As a biological therapeutic, BMSCs have a number of important properties. First, the cells do not require immunosuppressive drugs for host acceptance (as is required for a bone marrow or hematopoietic stem cell transplant) and second, the cells actually suppress the host T-cell response. Unlike most cells, BMSCs do not elicit an allogeneic response when presented to T cells, in vivo or in vitro (Di Nicola et al., 2002; Bartholomew et al., 2002; Liechty et al., 2000). Major histocompatibility complex (MHC) molecules are expressed on all mature cells and used to distinguish self from nonself. While BMSCs express low to medium levels of MHC class I and escape lysis by natural killer (NK) cells (Rasmusson et al., 2003), there is no detectable MHC class II present to activate T cells. Furthermore, there is no expression of the costimulatory molecules CD40, CD80, and CD86 that are required for T-cell activation (Pittenger et al., 1999). Therefore, in contrast to other forms of cell transplantation (e.g., hematopoietic stem cell transplant), no immunosuppressive agents are required when using allogenic BMSCs. In addition to being able to avoid recognition by the immune system, BMSCs are also able to suppress the allogeneic T-cell response. This has been demonstrated to be a dose-dependent response and is observed both in vitro and in vivo and shown to occur for both naive and memory T cells (Di Nicola et al., 2002; Dazzi et al., 2002). A number of studies have been performed to examine the mechanism by which BMSCs suppress T-cell function (Table 1). While the mechanisms underlying the immunosuppressive effects are not clearly understood, there are several hypotheses that have been explored. Because the majority of these effects do not require cell–cell contact (Yagi et al., 2010), these hypotheses have focused on the inhibition of T-cell proliferation by secreted factors such as hepatocyte growth factor, transforming growth factor beta (TGF-beta) 1, prostaglandin E2 (PGE2), IL-2, IL-10, indoleamine 2,3 deoxygenase, iNOS, soluble HLA-G, and soluble IL-1 receptor. While the precise mechanism has not been determined, a number of factors have been identified while others are more controversial. What is clear is that BMSCs secrete a variety of factors that can modulate the immune system and that there are probably undiscovered additional factors involved. Importantly, it also appears that BMSCs require interaction with T cells in order to generate immunosuppressive trophic factors, in a process called licensing. When conditioned media from naive BMSCs alone is tested, no inhibition of T-cell activation is observed. It is only after exposure of BMSCs to T cells that the BMSC-conditioned media is found to inhibit T-cell activation. Thus, the coincubation of BMSCs with T cells is a requirement for the production of inhibitory factors. The secretion of the inflammatory cytokine interferon gamma (IFNg) by T cells has been implicated as one of the major factors required for licensing. When BMSCs are stimulated with IFNg, the conditioned media from these BMSC is then capable of inhibiting T-cell activation (Ren et al., 2008; Polchert et al., 2008). Other cytokines that have been suggested to be involved in licensing include tumor necrosis factor alpha (TNFa) and IL-1b. BMSCs have also been observed to have effects on other cells from both the innate and adaptive immune system including helper T cells, regulatory T cells (Tregs), macrophages, antigen-presenting cells (APC), NK cells, and B cells (Figure 1). When BMSCs are incubated with monocytes, they inhibit their differentiation into dendritic cells and thus reduce the ability to create APC. BMSCs can alter the cytokine profile of macrophages, DCs, naive and effector T cells (TH1 and TH2), and NK cells from a proinflammatory to antiinflammatory state. This has been observed primarily by the types of cytokines secreted by the cells, either proinflammatory (i.e., IFNg, TNFa, IL-1b) or antiinflammatory (i.e., IL-4, IL-10) (Le Blanc and Ringden, 2007; English et al., 2010; Tolar et al., 2010). BMSCs decrease dendritic cell production of TNFa and increase production of IL-10. Similarly, helper T cells decrease IFNg and increase IL-4 production (Aggarwal and Pittenger, 2005). BMSC not only drive immune cells toward an antiinflammatory phenotype, but also increase the presence of Tregs that are important for regulating the tolerance of the host immune system and may be involved in reducing the alloreactive T-cell response in the presence of BMSC. The production of Tregs requires cell–cell contact with

Table 1

Potential mechanisms of allogeneic T-cell inhibition

Factor

Mechanism

Evidence

Prostaglandin E2 (PGE2)

Secretion of PGE2, a known immune regulator, by MSCs is upregulated in the presence of T cells IFNg stimulated MSCs express IDO, an enzyme that breaks down tryptophan into kynurenine, which has been demonstrated to block T-cell proliferation Cytokines secreted by MSCs are known to inhibit T cells

PGE2 inhibitors reduce the ability of MSCs to inhibit T-cell proliferation An antagonist of IDO, 1-methyl-L-tryptophan blocks the effects of MSCs on T-cell proliferation

Indoleamine 2,3 deoxygenase (IDO) Hepatocyte growth factor (HGF), transforming growth factor beta (TGF-beta) Nitric oxide (NO) HLA-G

MSCs produce NO in the presence of T cells and NO suppresses Stat-5 phosphorylation, which is required for T-cell proliferation HLA-G is a nonclassical HLA Class I molecule that inhibits T-cell proliferation and is secreted by MSCs in an IL-10 dependent manner

MSC, mesenchymal stem cells; GvHD, graft-versus-host disease.

Antibodies against HGF and TGF-beta block MSCs ability to inhibit T-cell proliferation iNOS deficient MSCs, which cannot express NO, are unable to inhibit T-cell proliferation and unable to provide benefit in a mouse GvHD model Antibodies against HLA-G block the ability of MSC to inhibit T-cell proliferation

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Figure 1 Diagrammatic summary of the major CD4þ and CD8þ T-cell effector subtypes (as derived from naïve CD4þ and CD8þ T cells), the reported effects of adult bone marrow-derived stem cells (BMSCs) on these effectors, and some clinically important disease associations for each. Bidirectional arrows indicate reported interconversion (plasticity) between Th1/Th17 phenotypes and Th17/iTreg phenotypes that may be of relevance to BMSC immune modulatory effects. CTL, cytotoxic T lymphocyte; DC, dendritic cell; DTH, delayed-type hypersensitivity; FOXP3, forkhead box P3 transcription factor; GvHD, graft-versus-host disease; IFNg, interferon-gamma; IL, interleukin; iTreg, induced regulatory T cell; nTreg, natural regulatory T cell; Th1, T helper type 1 cell; Th2, T helper type 2 cell; Th17, T helper type 17 cell; Treg, regulatory T cell. Duffy, M.M., Ritter, T., Ceredig, R., Griffin, M.D., 2011. Mesenchymal stem cell effects on T-cell effector pathways. Stem Cell Res. Ther. 2, 34.

BMSC and PGE2, and TGF-beta that induces FoxP3, a transcription factor critical for the differentiation of Tregs (English et al., 2009). Although this process is not nearly as well characterized, BMSCs have also been demonstrated to have effects on B-cell proliferation upon stimulation (Asari et al., 2009). In vivo studies have demonstrated that BMSCs can reduce antibody production in a heart transplant model (Ge et al., 2009).

Graft-versus-Host Disease (GvHD) Bone marrow transplants or hematopoietic stem cell transplantations (HSCT) have been used for over 40 years to treat hematological malignancies. One of the major complications that can occur with an allogeneic HSCT is GvHD, a severe inflammatory condition that occurs when the donor allogeneic T cells are activated by the host APC resulting in further stimulation of cellular and proinflammatory responses that result in tissue injury (Krensky et al., 1990). While GvHD is a multiorgan disease, the most clinically relevant organ damage occurs in the epithelial cell layer of the skin, gastrointestinal tract, and liver. Based on in vitro studies that demonstrated the immunomodulatory properties of BMSC, a number of preclinical and clinical studies were performed to determine if BMSCs could be used to treat GvHD (Tolar et al., 2011). The first clinical study was performed by Le Blanc and colleagues in a child suffering from GvHD who was treated with BMSCs isolated from the mother (Le Blanc et al., 2004). Subsequently, both the gut and liver GvHD were reversed after two infusions of BMSC. Based on this initial finding, eight additional patients with steroid-refractory GvHD were treated with BMSC, of which six patients had complete resolution of their GvHD (Ringden et al., 2007). The significant response observed in these patients led to a number of clinical trials using HLAidentical allograft donors, haploidentical donors, or unrelated HLA-mismatched donors. One of the first studies ever done was to

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give patients BMSCs from either HLA-identical haploidentical or third-party donors. Of the 55 patients in this study (Le Blanc et al., 2008), 27 exhibited a complete response with 24 of these receiving third-party BMSC. These results demonstrated the utility of using allogeneic BMSCs to treat GvHD. The ability to use allogeneic or third-party BMSCs provides many benefits. First, allogeneic BMSCs can be generated in advance from healthy donors, thus providing a more economical and uniform source of donor cells. Furthermore, allogeneic BMSCs are available on demand for the treatment. In fact, an allogeneic BMSC product has recently been approved for commercial use in Canada and New Zealand for the treatment of steroid-refractory GvHD in pediatric patients (www.osiris.com). In addition to the treatment of GvHD, there has been a recent study examining the potential use of BMSC as a prophylactic treatment for GvHD. In a Phase I clinical trial, BMSCs were administered to individuals undergoing allogeneic HSCT for the treatment of leukemia and related conditions, the BMSCs were well tolerated in both the single infusion and repeat infusion arms and the data also suggested that the therapy may provide benefit to recipients of allogeneic HSCT, by reducing the incidence and severity of GvHD as compared to historical clinical experience (Maziarz et al., 2012).

Type 1 Diabetes Type 1 diabetes is a T-cell-mediated, autoimmune disorder that leads to destruction of pancreatic b islet cells, and is characterized by the presence of antiislet cell antibodies and severe insulitis (Vija et al., 2009). BMSCs have been shown to suppress autoreactive Tcell responses in models of autoimmunity, making them an ideal candidate for use in Type 1 diabetes mellitus (T1DM). The efficacy of BMSCs in the treatment of T1DM has been tested or proposed at several time points of intervention: (1) as a preventative treatment or as a means to delay the full development of the disease, (2) as a therapy to treat complications arising from T1DM, (3) as a supportive treatment during the isolation of donor pancreatic islet cells for an islet transplant, (4) as an adjunct therapy with islet transplantation, and (5) as a rescue therapy upon onset of islet graft failure. Several studies have attempted to drive BMSCs to differentiate into insulin-secreting cells, with mixed results, but the field remains focused on the use of BMSCs as immunomodulators in the context of diabetes. BMSCs have been shown to significantly delay disease onset in diabetic mouse models (Busch et al., 2011a). In a streptozotocininduced model of diabetes, green fluorescent protein (GFP) donor BMSCs significantly reduced blood glucose levels as compared to controls. Transplantation of bone marrow resulted in an increase in insulin positive islets, but no GFPþ/insulinþ cells were observed, suggesting that the results were due to proliferation of host islet cells. It has also been demonstrated that upon islet transplantation, recipient islet precursor cells do not divide and contribute to graft function, so understanding the BMSCs-induced stimulation of host cell division or insulin production may be particularly important. These studies have led to the working hypothesis that BMSCs exert their beneficial effects through indirect mechanisms such as protection of remaining b-cells or stimulation of endogenous b-cell replacement. The early success of BMSCs in delaying the onset of diabetes led to studies examining the ability of BMSCs to prevent initial islet loss and promote engraftment of islets after pancreatic islet transplantation. It is thought that an islet transplant has the potential to cure Type 1 diabetes; however, there are currently major limitations to widespread implementation, including loss of large numbers of islets in the immediate setting posttransplant, lack of revascularization, and transplant rejection (Korsgren et al., 2005). The effect of multiple doses of BMSCs, in combination with immunosuppressive therapy, on islet graft rejection has been examined in diabetic rats (Solari et al., 2009). The results demonstrated that both intraportal and IV-administered BMSCs prolonged graft function through prevention of acute rejection in a dose-dependent fashion, and no difference was seen between syngeneic or allogeneic BMSCs. The ability of BMSC transplants to prevent rejection was similar to immunosuppressive therapy, but the combination of BMSC and immunosuppressive therapy together were not more efficacious. Islet grafts can deteriorate over time due to chronic allograft rejection, local islet toxicity as a result of the immuno-suppressive regimen, recurrent autoimmunity, and/or failure of islet regeneration (Ding et al., 2010). Drugs used to prevent islet allograft loss adversely effect b-cell function and glycemic control. BMSC therapy may have the potential to eliminate maintenance of some systemic immunosuppressive drug therapies, thus relieving negative effects of these drugs on the graft itself in addition to removing the risks of continued immunosuppression. The failure of the islet graft and the loss of insulin independence can involve rejection due to the activation of alloreactive T cells and a reoccurrence of the original autoimmune disease. The breakdown of immunologic tolerance may result in a cross-reactive memory response against the transplanted islets, resulting in loss of b-cell mass. Leukopenia, a decrease in the number of white blood cells in circulation, can result from immunosuppressive therapy and favors the generation of islet-reactive T cells, leading to islet destruction (Monti et al., 2008). It is thought that BMSCs may protect transplanted allogeneic islets by negatively regulating persistent T-cell autoimmunity and controlling the activation and effector function of alloreactive T cells. BMSCs may also suppress activation and proliferation of B cells, and prevent the differentiation and maturation of dendritic cells, effectively preventing islet destruction. Further studies are necessary to determine the ability of BMSCs or other cell therapies to prolong graft function or reverse graft failure. Multiple clinical studies have examined the ability of bone marrow-derived cells to improve clinical outcome in both Type 1 and Type 2 diabetes, and have established a safety profile for the use of these cells in diabetic patients. A Phase II, multicenter, randomized, double-blind, placebo-controlled study has been initiated to evaluate the safety and efficacy of adult human BMSCs for the treatment of recently diagnosed Type 1 diabetes. The use of BMSCs as an adjunct therapy for islet transplantation warrants further exploration and advancement toward clinical use.

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Spinal Cord Injury Spinal cord injury (SCI) results in disruption of the blood–brain barrier and initiates a cascade of inflammatory processes leading to infiltration of immune cells and secondary cell death that extends beyond the site of initial injury (Silver and Miller, 2004). This reactive process of secondary injury takes place in the days and weeks following SCI, and can result in exacerbation of neurological dysfunction. An additional cause of spinal cord regeneration failure is the formation of the glial scar, which involves activation of astrocytes in an attempt to restore the blood–brain barrier. These astrocytes produce inhibitory chondroitin sulfate proteoglycan, a major barrier to regenerating axons. Other potential inhibitors in the glial scar are myelin-associated proteins, which also inhibit neurite outgrowth and hinder repair. BMSC therapy has been proposed to treat both acute and chronic SCI through multiple potential mechanisms of action (Wright et al., 2011). The immunosuppressive properties of BMSCs may reduce the acute inflammatory response to SCI, reducing secondary injury and cavitation. Direct transplantation of BMSCs after injury has been shown to increase preservation of spinal cord tissue and decrease neuropathic pain. Direct transplantation of BMSCs into the cord may modify activation of astrocytes or encourage axonal regeneration through downregulation of inhibitory components in the glial scar. BMSCs produce a number of growth factors, including nerve growth factor (NGF) and vascular endothelial growth factor (VEGF), which could promote axonal outgrowth even in an inhibitory environment. Additionally, some reports suggest that BMSCs could act as bridges, guiding regenerating axons across the injury cavity. BMSCs produce molecules including laminin, fibronectin, and collagen, which could decrease cavitation and provide a permissive environment for growing axons (Ankeny et al., 2004). As in most injury situations, some degree of inflammation is a necessary and beneficial component of the recovery process after SCI (David and Kroner, 2011). Macrophages are known to phagocytose the myelin debris, clearing the way for functional regeneration, and to produce some protective cytokines and growth factors, which could enhance regeneration. Two subtypes of macrophages have been described with regards to their phenotype and activity: classically activated macrophages (M1) and alternatively activated (M2). M1 macrophages are typically considered to be the product of activation with proinflammatory cytokines IFNg and TNFa. Alternatively, activated macrophages are the product of activation with the cytokines interleukin-4 (IL-4) and IL-13, and possess enhanced phagocytic capabilities and antiinflammatory activities, which are thought to contribute to their beneficial effects after SCI. BMSCs and related cell types have been shown to drive macrophages toward the alternatively activated M2 phenotype and concurrently promote white matter sparing and reduce the effects of the inhibitory glial scar (Busch et al., 2011b; Nakajima et al., 2012). To date, the majority of cell therapy-focused clinical trials for SCI have utilized whole mononuclear cell preparations (MCPs) from bone marrow, not cultured adherent BMSCs (Wright et al., 2011). MCPs are generally administered alongside granulocytemacrophage colony stimulating factor to mobilize the migration of these cells into the lesioned spinal cord and induce activation resulting in the secretion of neurotrophic cytokines at the site of injury. Modest increases in neurological function have been reported, but as relatively few patients have been treated thus far, it is difficult to determine if these results are due to an intrinsic recovery process or directly attributable to the treatment itself.

Multiple Sclerosis Multiple sclerosis (MS) is an immune-mediated disorder of the nervous system, which is characterized by inflammation and axonal demyelination and degeneration (Odinak et al., 2011). The etiology of MS is uncertain, but it is likely that both genetic and environmental factors contribute to disease onset and progression. Two aspects of the disease have the potential to be addressed by stem cell therapy: prevention of further CNS damage via immunomodulation, and promotion of remyelination and repair. The basis for the use of BMSCs in the treatment of MS comes largely from an animal model of MS, the experimental autoimmune encephalomyelitis (EAE) (Auletta et al., 2012). In this model, the combination of injected myelin immunogenic peptide and an adjuvant generates widespread inflammation and demyelination in many regions of the brain and spinal cord. It has been demonstrated that IV-infused BMSCs improve the clinical course and pathology scores in an EAE model (Zappia et al., 2005). This decreased severity of the disease severity was paralleled by suppression of inflammation as indicated by T-cell anergy and decreasing demyelination and this response was associated with induction of tolerance toward the immunizing antigen. Intraventricularly injected BMSCs have been observed to migrate to white matter lesions and induce upregulation of growth factors and proliferation of endogenous oligodendrocyte progenitor cells (Kassis et al., 2008). As a whole, these preclinical results suggest that not only do BMSCs inhibit autoimmune attack, but they also provide significant neuroprotection despite their limited infiltration into the CNS. Several studies have been undertaken to determine the safety and efficacy of BMSCs for the treatment of MS in the clinical setting (Freedman et al., 2010). In a small study, 10 patients were given BMSCs via intrathecal administration and the only conclusion was that the approach was clinically feasible (Connick et al., 2011). The preliminary results of a phase I/II study reported that a combination of intravenous and intrathecal administration of BMSCs given to 15 patients with MS resulted in no significant side effects and revealed no unexpected pathology 1 year following injection (Slavin et al., 2008). This preliminary safety data suggest that BMSCs can be considered relatively safe in the context of severe disease; however, additional studies are ongoing to determine efficacy.

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Stroke Therapeutic options for patients suffering an ischemic stroke are extremely limited and agents which can exert neuroprotective effects, induce neural plasticity, or encourage remodeling and neovascularization are of great interest. As in other injury paradigms, the use of BMSCs was initially proposed as a cell replacement therapy; however, significant and reproducible transdifferentiation of BMSCs into neural lineages in vivo have not been observed (Honmou et al., 2012). Following an ischemic stroke, some level of spontaneous functional recovery is generally seen in both stroke patients and animal models, and it is possible that BMSCs can harness and enhance this compensatory neural plasticity or remodeling in the days and weeks postinjury. Transplantation of BMSCs in the hours to days following ischemia can reduce the size of the infarct and improve functional outcome measures in rodent models of ischemic stroke. BMSCs are known to release growth factors and stimulate the release of beneficial factors from endogenous tissue. Intravenously delivered BMSCs also influence the ischemic tissue itself, reducing apoptosis of cells at the lesion boundary and promoting proliferation of endogenous cells at the site of injury. BMSCs have been demonstrated to release brain-derived neurotrophic factor (BDNF) and levels of this growth factor have been shown to increase at the site of ischemia in rats receiving BMSC treatment (Nomura et al., 2005). BMSCs can also be engineered to express growth factors such as BDNF or VEGF to be released over time at the injury site, providing a local and sustained neuroprotective, neurostimulatory, and/or angiogenic therapy. The proinflammatory environment after stroke leads to breakdown of the blood–brain barrier, which worsens the deficit associated with the initial insult. Ischemic stroke stimulates adrenergic activation that promotes the release of immune cells from the spleen, ultimately leading to a loss of splenic mass (Ajmo et al., 2009). Previous work has shown ischemic stroke to be associated with the loss of splenocytes in conjunction with an increase in terminal deoxynucleotidyl transferase dUTP nick end labeling (TUNEL)- þ apoptotic cells within the spleen, leading to decreased T-cell proliferation and cytokine production culminating in a state of immunosuppression after stroke (Offner et al., 2006a,b). Separate experiments have shown a peripheral increase in production of the proinflammatory cytokines TNFa, IFNg, IL-6, monocyte chemotactic protein-1 (MCP-1), and IL-2 after ischemic stroke (Offner et al., 2006a,b). After intravenous injection, BMSCs come into contact with many organ systems including the spleen. Intravenous BMSC injection preserves the lost splenic mass (via inhibition of CD8þ T-cell release) and potentially increases splenocyte proliferation resulting in production of antiinflammatory cytokines such as IL-4 and IL-10 (Walker et al., 2010; Vendrame et al., 2006). The production of these cytokines could modulate the proinflammatory response after injury in the lesion itself and in the penumbral regions of the stroke. These results support the emerging concept that intravenous administration of bone marrow-derived cells can enhance stroke recovery through direct effects on peripheral organs. The safety, feasibility, and efficacy of autologous BMSCs administered intravenously have been examined in multiple Phase I stroke clinical trials. MRIs have shown no tumor or abnormal cell growth in any patient to date and no severe adverse cellrelated effects have been reported. Bang and colleagues reported success in a 30-patient trial in patients suffering from severe middle cerebral artery stroke in which 5 patients received 1  108 autologous BMSCs (Bang et al., 2005). While improvements were observed in patients who received BMSCs, the treatment group was small and treatment procedures were not blinded. Additional studies are necessary to conclusively determine if therapeutic intervention with BMSCs can improve clinical outcome in ischemic stroke.

Potential Safety Issues Associated with BMSCs Despite the success of BMSCs in preclinical models, several potential safety issues are associated with the use of BMSCs, particularly in immunosuppressed patients. Given the strong capacity of BMSCs for immunomodulation, concerns have arisen that BMSCs might interfere with immune responses against pathogens and therefore increase the risk of infection (Nauta and Fibbe, 2007). Fortunately, no increases in infection have been observed in any of the clinical trials to date, in particular in those patients who are already on immunosuppressive drugs. In fact, it was recently demonstrated that stimulated human BMSCs have potent antimicrobial effector function against bacteria, protozoal parasites, and viruses (Meisel et al., 2011). In preclinical models, infused adherent stem cells can home to the stroma bed of preexisting tumors, with the implication that through trophic support or immunomodulation these cells can promote tumor growth or block tumor clearance (Kuhn and Tuan, 2010). Conflicting data exist in the literature for this hypothesis without clarification of whether exogenously provided adherent stem cells impact the endogenous endothelial and stromal populations involved in tumor initiation or growth; however, the majority of reports do not reflect increased tumorigenic risk (Burt et al., 2008; Sensebe et al., 2012). Although human BMSCs do not appear to be readily susceptible to chromosomal aberration in culture (Nauta and Fibbe, 2007), expanded BMSCs should be tested for chromosomal stability and purity prior to clinical administration. A recent review article reports on completed clinical studies with BMSCs covering more than 5000 patients treated in over 100 clinical studies bracketing 15 therapeutic areas including both acute and chronic diseases (Ankrum and Karp, 2010). To date, with studies initiated over 16 years ago, no reports of tumorigenicity by donor product, or increased frequency of host tumorigenesis has been reported. While it is clear that only long-term patient follow-up will provide a statistically valid evaluation of this association, near-term therapeutic decisions should be driven by a careful analysis of patient risk/benefit considerations in specific disease settings.

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Future Perspectives The use of adult stem cells to treat inflammatory and autoimmune diseases offers novel and exciting prospects for the future. BMSCs have immunomodulatory properties that make them uniquely suited to treat autoimmune disease and inflammation. Multiple clinical studies have examined the ability of bone marrow-derived cells to improve clinical outcome and have established a safety profile for the use of these cells in patients. Optimization of several parameters regarding BMSC therapy must be determined, including the ideal time of administration, dosing regimen, route of administration, and identification of biomarkers that will identify the optimal time for intervention and will likely vary by indication. With the recent approval of BMSCs for the treatment of GvHD, the field of adult stem cell therapy looks forward to clinical use in other areas. Based on the Phase II clinical trials underway to establish the potency of BMSCs in MS, stroke, SCI, and diabetes, additional approvals are expected in the near future.

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Assessment of Cellular Responses of Tissue Constructs in vitro in Regenerative Engineering Margaret AT Freeberg, Jacob G Kallenbach, and Hani A Awad, University of Rochester, Rochester, NY, United States © 2019 Elsevier Inc. All rights reserved.

Biochemical and Biophysical Regulation of Cellular Responses Characterizing Cellular Proliferation and Senescence Quantifying Cellular Proliferation Detection of Cellular Senescence Characterizing Cellular Viability and Death Live/Dead Staining ATP/Metabolic Activity Tetrazolium reduction Resazurin reduction Intracellular ATP quantification Intracellular protease activity Caspase activity Characterizing Cellular Differentiation Microscopy-Based Techniques Confocal and multiphoton microscopy Electron microscopy FRET microscopy Quantitative Gene Expression Analysis DNA microarray RNA sequencing Quantitative Protein Analysis Western blot Enzyme-linked immunosorbent assays (ELISA) Flow cytometry Förster resonance energy transfer assays Assessment of Mechanical Properties Emerging Technologies Further Reading

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Glossary Apoptosis A process of programmed cell death or suicide triggered by intrinsic (mitochondria-associated proteins) or extrinsic (death receptor activation) pathways, characterized by an ordered sequence of events leading to chromatin degradation, mitochondria disintegration, morphological changes and fragmentation, and ultimately inflammation-free phagocytosis of the apoptotic bodies or fragments of the dying cell. Cell cycle The ordered sequence of biological processes within a cell leading to the duplication of its genetic material (DNA) and the production of two daughter cells; conventionally divided into a first gap (G1), a DNA synthesis (S) phase, a second gap (G2), and a cell division or mitosis (M) phase. Cell differentiation A multi-step process through which a stem cell loses its pluripotency and alters its phenotype to assume a specialized function, characterized by overt changes such as cell shape and cell size, and covert changes such as altered gene expression, metabolic activity, secretory phenotype, and responsiveness to external stimuli. Cellular proliferation A process that results in a net increase of the number of cells through the balance of cell division and cell death. Necrosis A process of accidental or premature cell death by autolysis due to injurious events such as infection or trauma, leading to the activation of an inflammatory response to eliminate the dead cell debris by phagocytosis. Senescence A state of permanent and irreversible cell-cycle arrest associated with replicative aging in metabolically active and viable cells, which is characterized by DNA damage, upregulation of cyclin-dependent kinase inhibitors, and an abnormal secretory phenotype.

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Abbreviations 3D Three-dimensional AFM Atomic force microscopy ATP Adenosine triphosphate BrdU bromodeoxyuridine or 5-bromo-20 -deoxyuridine Cas9 CRISPR-associated protein-9 nuclease CRISPR Clustered regularly interspaced short palindromic repeats DNA Deoxyribonucleic acid ECM Extracellular matrix ELISA Enzyme-linked immunosorbent assay FRET Förster (fluorescence) resonance energy transfer IHC Immunohistochemistry ISH In situ hybridization MSC Mesenchymal stem cells MTS 3-(4,5-Dimethylthiazol-2-yl)-5-(3-carboxymethoxyphenyl)-2-(4-sulfophenyl)-2H-tetrazolium MTT 3-(4,5-Dimethylthiazol- 2-yl)-2,5-diphenyltetrazolium bromide NAD Nicotinamide adenine dinucleotide NADP Nicotinamide adenine dinucleotide phosphate NGS Next generation sequencing PCNA Proliferating cell nuclear antigen PCR Polymerase chain reaction PDMS Polydimethylsiloxane RNA Ribonucleic acid RNA-seq RNA sequencing RNAi RNA interference ROS Reactive oxygen species RT-PCR Reverse transcription-polymerase chain reaction SASP Senescence-associated secretory phenotype SDS-PAGE Sodium dodecyl sulfate polyacrylamide gel electrophoresis WST-1 2-(2-Methoxy-4-nitrophenyl)-3-(4-nitrophenyl)-5-(2,4-disulfophenyl)-2H-tetrazolium X-gal 5-Bromo-4-chloro-3-indolyl-b-D-galactopyranoside XTT 2,3-Bis-(2-methoxy-4-nitro-5-sulfophenyl)-2H-tetrazolium-5-carboxanilide

The replacement of terminally-diseased organs or massive tissue loss have been and continue to be critical needs in medicine. Traditionally, these surgical needs have been fulfilled using donated organs or tissue allografts. However, viable donor organs are limited in their availability, carry numerous health risks, and require immunosuppressants for the life of the recipient. Processed allograft tissues are readily available through tissue banks and can provide satisfactory short-term outcomes, but generally do not integrate or mediate regeneration. Over the past 3 decades, tissue engineering and regenerative medicine, or regenerative engineering, have emerged to design replacement tissues and organs that overcome the limitations of traditional sources. In the classical tissue engineering paradigm, cells are seeded onto biomaterial scaffolds and stimulated biochemically and physically to grow a functioning tissue or organ in vitro, in a simulated physiological environment such as a dish in an incubator or a more sophisticated bioreactor system, to eventually be surgically implanted into a patient. The design of engineered tissues and organs must take into account the different choices available for the three main building blocks of cells, signals, and scaffolds, which are beyond the scope of this chapter. Regardless of these choices, this bottom-up approach inherently has a number of response variables that must be monitored, optimized, and reproducibly-controlled; most importantly perhaps are related to cellular responses to their engineered environment. When cells are dissociated from their physiological context (their native ECM) to be cultured in vitro, they can be induced to assume one of four fates or phenotypes: proliferative, differentiated, senescent, or apoptotic. Likewise, when dissociated cells are incorporated into a 3D biomaterial matrix in an engineered tissue construct, they can alter their phenotype depending on their interactions with their engineered ECM substrate and the abundance of biological and biochemical signaling factors in their in vitro culture environment. Monitoring the cellular responses to these biophysical and biochemical cues can be predictive of the function of the engineered tissue. Assessment of cellular responses can be accomplished using invasive or destructive assays that require sacrificial constructs in cross sectional studies, or using minimally- or non-invasive assays that can enable longitudinal, real-time monitoring of the functional maturation of the engineered tissue.

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Biochemical and Biophysical Regulation of Cellular Responses To create complex functional tissues, it is important to mimic the native tissue environment to restore function. Tissues and organs are generally anisotropic, with spatially heterogeneous composition, morphology, nano- and micro-topographies, and biomechanical properties that are optimized for function. The native and engineered tissue environments are composed of both biophysical and biochemical cues that orchestrate cellular phenotype and drive altered responses. There is extensive research to understand how biophysical and biochemical signals regulate cell fate and direct differentiation. Built-in features of scaffold design such as substrate architecture and topography (surface patterning and fiber orientation) and substrate stiffness can alter how cells perceive applied mechanical forces (tensile, compression, shear), interact with neighboring cells, and respond to biophysical and biochemical cues to alter their phenotype (Fig. 1). Externally applied physical cues can include mechanical, electrical, electromagnetic, acoustic and ultrasound, and photo stimulation. The effects of these biophysical cues on stem cells have been extensively investigated in vitro. Surface chemistry and topography affect protein adsorption from cell culture media, which consequently affect cell attachment mechanisms and impact cellular behaviors. In addition, substrate or scaffold stiffness is an important biophysical cue that can direct cell differentiation. Cells adjust their cytoskeletal tension and stiffness to match their environmental substrate, therefore reduced cytoskeletal organization is typically observed in cells grown on low stiffness substrates. This effect has mechanobiological consequences. For example, a substrate stiffness of 0.1–1 KPa induces neurogenic differentiation, while a stiffness of 8–17 KPa induces myogenic differentiation, and a 25–40 KPa stiffness induces osteogenic differentiation of stem cells. On the other hand, substrate stiffness < 0.1 KPa maintain pluripotency of stem cells. Likewise, external mechanical forces of tension, compression or shear can either maintain cellular

Fig. 1 Biochemical and biophysical regulation of cellular responses. Biochemical cues include soluble molecules added to the culture media or can include surface functionalization with proteins, peptides, or integrins. Physical cues include surface properties (topography, stiffness, alignment, etc.), mechanical forces, electrical stimulation, application of ultrasound, and photostimulation. These cues can drive cellular responses, including differentiation, proliferation, apoptosis, and senescence. Adapted from Ding, S. Kingshott, P., Thissen, H., Pera, M. and Wang, P. Y. (2017). “Modulation of human mesenchymal and pluripotent stem cell behavior using biophysical and biochemical cues: A review.” Biotechnology and Bioengineering 114(2), 260–280.

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pluripotency by enhancing proliferation and inhibiting differentiation or stimulating cell lineage differentiation. For example, cyclic compression of MSC can promote cartilage formation or chondrogenesis, while shear stress from fluid flow can stimulate MSC differentiation into endothelial cells to form vessel networks, whereas high intracellular tension can encourage MSC toward osteoblasts. Even though specific mechanical forces are correlated with differentiated cell types, the mechanical environment needs to be maintained in order to maintain the differentiated cell phenotype. Electromagnetic stimulation is of particular importance for cardiac and neural lineage cells that function to communicate via electrical pulses. Neurogenesis and cardiomyogenesis depend greatly on intensity, duration, and application method of the electromagnetic stimulus. Overall, electromagnetic stimuli have been applied successfully to differentiate neurons and cardiomyocytes, and it remains an area of interesting research to evaluate their effects on other cell types. Ultrasound is a mechanical wave (low intensity, high frequency) that can modulate and pattern cellular behavior and tissue morphology. Besides its clinical imaging diagnostic value, ultrasound has been shown to stimulate osteogenesis, increase proliferation, and increase differentiation of stem cells, and as such offers exciting potential as a safe tissue engineering and regeneration approach. Photostimulation is likely wavelength- and cell type-dependent as it has been shown to increase epidermal stem cell proliferation and migration, while decreasing proliferation of bone marrow-derived MSC. Biochemical cues are primarily soluble factors that are delivered as signals to guide cellular responses and include cytokines and growth factors, genetic editors (RNAi, CRISPR/Cas9), proteins and proteases, and exosomes. In general, biochemical cues are easy to deliver in culture media. However, it is often advantageous to tether biochemical cues to scaffolds to control their extended release and their bioavailability for their target cells. This has been an area of intensive research in tissue engineering, where optimization studies are typically performed in vitro. For example, osteogenic differentiation of MSC requires ascorbate, beta-glycerophosphate, and bone morphogenetic proteins (BMPs) to stimulate the expression of the osteogenic genes Runx2, osteocalcin, and osteopontin, whereas adipogenesis requires dexamethasone, indomethacin, and insulin. Transforming growth factor beta (TGF-b) is typically added at varying doses to induce mesodermal lineages of bone, fat, cartilage, tendon, ligament, and cardiac differentiation. Furthermore, cardiomyocyte differentiation from pluripotent stem cells within embryoid bodies is promoted by culturing with nonessential amino acids, L-glutamine, b-mercaptoethanol, and high concentrations (20%) of fetal bovine serum. Moreover, functional insulin producing pancreatic b-cells were generated from stem cells through systematic addition and removal of activin, Wnt, fibroblast growth factor 10 (FGF10), retinoic acid, insulin-like growth factor 1 (IGF) and smoothened antagonists (SANT). Extracellular matrix soluble and insoluble proteins are also important in controlling cellular behavior. For example, Matrigel, a gelatinous ECM mixture from mouse sarcoma cells, supposedly maintains stem cell self-renewal in cell culture applications, but is an exogenous murine mixture limited in its clinical applications. Fetal bovine serum (FBS) is ubiquitously used in cell culture as a biochemical supplement because it contains copious biomolecules such as vitronectin, fibronectin, and albumin, as well as amino acids, sugars, lipids and growth factors and cytokines. The adsorption of serum proteins on engineered materials critically influences cell adhesion and cell migration. Engineered surface modifications, whether that be protein or polymer coatings on materials, have been shown to affect MSC behavior. For example, protein coatings on polydimethylsiloxane (PDMS) surfaces primarily utilizing fibronectin and Type I collagen have been shown to promote cell adhesion and growth in cell culture applications. Immobilized bioactive peptides can directly influence cell phenotype through juxtacrine signaling. Since biophysical and biochemical stimuli combine in native tissues, similar combinations need to be recapitulated for growing artificially engineered tissues. Therefore, they systematically need optimization, and this bottom up approach often employs empirical experimental strategies. This can be expensive, time-consuming, and highly variable due not only to the need for sequential addition and subtraction of growth factors, which varies for different cell phenotypes, but also due to the variability in biologically-derived culture media supplements such as sera. Furthermore, exogenous soluble factors and animal derived products can elicit problems of contamination especially for clinically translatable cellular therapies. Thus, a major emphasis of tissue engineering and regenerative medicine research has been, and continues to be focused on characterizing the cellular responses to biophysical and biochemical stimuli in vitro prior to in vivo testing and clinical translation. These cellular responses induce overt and covert alterations in cell phenotypes, which require specialized characterization techniques. Overt phenotypic changes such as changes in cell numbers and morphological changes including cytoskeletal organization and contractility can be assessed using microscopy. Covert changes, as with cell differentiation for example, typically require biological and biochemical assays to measure the associated transcriptional and translational phenomena. Some standard techniques measure these changes by lysing the cells to measure altered gene expression, protein content and phosphorylation, or protein-DNA interactions, to name a few examples. Other techniques can often rely non-destructive assays to characterize real time measurable changes in cellular responses such as activatable ion channels (e.g., Ca2 þ flux) and receptor-ligand binding or changes induced by the cells such as the cell-mediated unfolding of cryptic peptide sequences in structural proteins, as well as cell signaling (e.g., using Förster resonance energy transfer (FRET)-based approaches). In practice, there are numerous techniques that can be employed to assess cellular proliferation and metabolic activity, multipotency or differentiation state, and whether the cells are being driven to pathologic phenotypes such as senescence or apoptosis. Typically, a comprehensive assessment of cellular responses in engineered tissue constructs uses a combination of outcomes based on gene expression, protein content and phosphorylation, enzymatic activity, imaging, and mechanical characterization. This manuscript offers a survey of the state-of-the-art tools and assays for the assessment of cellular responses within engineered tissues in vitro.

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Characterizing Cellular Proliferation and Senescence The cell cycle is an orchestrated sequence of events during which a cell duplicates its contents and divides into two daughter cells. There are 4 phases of the cell cycle: G1, S, G2 and M phases (Fig. 2). The G1 phase, or Gap 1, is an interphase during which the cell synthesizes mRNA and proteins necessary for transitioning into the subsequent step of DNA synthesis. The G1/S checkpoint in this interphase orchestrates cell fate decisions, instructing the cells to become quiescent, differentiate, or proceed to DNA replication and chromosome duplication in the S phase. Dysregulated G1/S transitioning can lead to transformation and cancer. Once DNA replication in the S phase is complete and DNA damage is detected and repaired, the cell cycle enters into a second interphase known as G2 or Gap 2, in which the cells experience rapid growth and protein synthesis in preparation for mitosis in the M phase. Another critical DNA damage control takes place within the G2/M checkpoint, which ensures that mitosis is not initiated before damage to DNA during replication is repaired. All major checkpoint transitions in the cell cycle are orchestrated by cyclins, cyclin dependent kinases (CDK) and a large family of cyclin and CDK regulators. When cells exit the cell cycle they enter a phase known as G0 or the resting phase, in which they can acquire a quiescent phenotype or a non-cycling state. Tissue resident stem cells mostly exist in this state. Often cells in G0 can reversibly differentiate without losing the ability to re-enter cell cycle such as the case with hepatocytes. However, if the cells irreversibly differentiate into a postmitotic phenotype such as osteocytes, myocytes, and mature neurons they cannot re-enter the cell cycle. Cells in G0 can sometime assume another irreversible phenotype known as senescence, a degenerative and often aging-associated change in the cells accumulating over many cycles of cell divisions, which often contributes to a pathology in the tissue. A third phenotype relates to a process of programmed cell death known as apoptosis, which is characterized by overt morphological changes preceding death, including membrane blebbing, cell shrinkage, nuclear fragmentation, chromatin condensation and chromosomal DNA fragmentation. Apoptosis is a highly regulated process and has been linked to cell cycle since a number of tumor suppressor genes have been shown to sensitize cells to programmed death.

Quantifying Cellular Proliferation Detection of gene expression or antigens present exclusively in proliferating cells represent standard techniques for assessment of cell proliferation. The nuclear protein Ki-67 is a bona fide proliferation marker, which is expressed during all phases of the cell cycle but not during quiescence or in differentiated cells (Fig. 2). While its function remains largely unknown, recent discoveries suggested a role for Ki-67 in organizing heterochromatin including compaction and long-range genomic interactions. Other active cell cycle markers include the proliferating cell nuclear antigen (PCNA), which is a homotrimer scaffolding protein involved in DNA replication and repair and chromatin remodeling. PCNA nuclear localization is particularly increased during the S phase of the cell cycle. Both Ki-67 and PCNA can be detected in cell-seeded engineered tissue constructs using standard immunohistochemistry techniques with a variety of antigen- and species-specific, commercially available antibodies. Other less commonly reported nuclear antigen markers of cell proliferation, sometimes used as indices for cancer diagnosis, include type II DNA topoisomerase and phosphohistone H3, both of which are essential for several proliferation events in DNA replication and chromatin modification and can be detected using immunohistochemistry.

Fig. 2 The cell cycle and cellular phenotypes. There are 4 phases of the cell cycle: G1 Phase, S phase, G2 phase, and M Phase. Ki-67 and PCNA are nuclear antigens exclusively present in proliferating cells. Cells can exit the cell cycle to enter quiescence, or can differentiate or undergo senescence or apoptosis. Differentiation can be assessed histologically by examining morphology or using IHC for phenotypic antigen markers. Senescence can be detected by increased b-galactosidase. Apoptosis can be detected by probing for caspase activity.

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Cell proliferation can also be measured by quantifying newly synthesized DNA by treating cells in cultures with the synthetic thymidine analog bromodeoxyuridine (5-bromo-20 -deoxyuridine or BrdU). BrdU signal can be simply detected as it gets incorporated into newly synthesized DNA during the S phase using immunohistochemistry and visualized by a colorimetric, chemiluminescent, or fluorescent reporter signal. More sophisticated analysis of the cell cycle phases using fluorescent-BrDU labeling can be performed using flow cytometry.

Detection of Cellular Senescence Cellular senescence is a phenotype associated with DNA damage, in which cells cease to divide but remain metabolically active. Senescence often occurs when cells in culture have reached their maximum number of divisions, the Hayflick limit, and is often associated with telomere shortening leading to DNA damage due to reduced telomerase activity. Environmental stress factors in culture such as reactive oxygen species (ROS) can also induce senescence, sometimes downstream of commonly-used growth factors, including TGF-b, or under supraphysiological oxygen conditions common in many tissue engineering scenarios. In vivo, senescence is associated with aging-related pathologies, in which senescent cells produce a pro-inflammatory secretome. This senescence associated secretory phenotype (SASP), which consists of inflammatory cytokines, growth factors, and proteases is a bona fide markers panel for senescent activity, which remains understudied in tissue engineering. SASP can be quantified at the gene expression level in the cells or at the secreted protein level using enzyme-linked immunosorbent assays (ELISA), where innovations in multianalyte detection assays can provide comprehensive phenotypic assessments. Morphologically, senescent cells grow larger with flattened bodies and express b-galactosidase, an enzyme whose activity in catalyzing the hydrolysis of b-galactosides can be detected using chromogenic substrates. The most popular substrate for detecting b-galactosidase is the glycoside X-gal (5-bromo-4-chloro-3-indolyl-b-D-galactopyranoside). The X-gal assay relies on the activity of b-galactosidase to induce release of a soluble indolyl group from X-gal, which can be subsequently oxidized to form a blue precipitate that can be visualized by eye in whole-mount constructs or using light microscopy at a higher magnification. Furthermore, nuclear antigens belonging to a large family of CDK inhibitors, such as p21 and p16, have been associated with the induction of senescence without detectable DNA damage, and therefore have also been used as bona fide markers of senescence, which can be detected immunohistochemically. Moreover, senescent cells display components of the DNA damage response (DDR) including phosphorylated gH2AX in their nuclei, which can be used as a means to assess the senescent state of the cells in vitro using immunohistochemistry.

Characterizing Cellular Viability and Death In assessing the function of a tissue engineered construct in vitro, it is critical to determine if cells are viable within their engineered environment. There are numerous methods to evaluate cellular viability, which include assessment of cell and DNA damage, ATP concentration, and metabolic activity. Cells in engineered tissue can experience death by apoptosis or necrosis. Apoptosis is programed cell death, which plays an indispensable role in morphogenesis and organogenesis during embryonic development, and is one of the key cellular responses to stress within adult tissues or organs, often in the context of pathology. During apoptosis cells begin to retract and condense as they begin to degrade their internal components. Apoptosis happens very rapidly, on the order of 20–60 minutes, making it very difficult to measure. There a number of biochemical processes that occur between the onset of apoptosis and the end point of cell and nuclear fragmentation. These processes include cytoskeletal reorganization, alterations (blebbing) in the membrane, proteolytic cleavage, DNA degradation, and fragmentation into apoptotic bodies. These biochemical processes can be probed to detect apoptosis. Necrosis is a distinct mode of cell death characterized as “accidental” or “uncontrolled” destruction of the cell and release of its contents, which invokes an inflammatory response. Necrosis lacks the ordered sequence of events observed in apoptosis. Regardless of the mechanism, cell death in engineered tissues in vitro can result from environmental stress factors including extreme hypoxia or diffusion-limited nutrient bioavailability, and must be assessed as endpoint outcomes using sacrificial constructs or using real-time, non-invasive assays. There are numerous assays and techniques available to measure cell viability. Each assay has advantages and disadvantages. Therefore, it is important to identify what best suits the experimental needs. Common viability assays include live/dead staining, ATP/metabolic assays, and detection of protease activity. Additional assays, which will not be discussed herein, examine individual cells to assess cell morphology, chromatin condensation, and detection of fragmented DNA. These are typically measured by light or electron microscopy, flow cytometry, or using time-lapse microscopy.

Live/Dead Staining This approach introduces a combination of two distinct fluorophores (green/red) to identify live and dead cells, respectively, within an engineered construct. Typically, a membrane permeant, optically-silenced fluorophore, such as calcein-acetoxymethyl ester (AM) can enter into the cytosol of a live cells, where AM is hydrolyzed by intracellular esterases to liberate the green fluorescent calcein

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dye. The calcein fluorescence intensity is proportional to the number of viable cells because esterase activity is absent in dying cells. The second fluorophore, typically a red-fluorescent ethidium homodimer, is not membrane permeant but has the ability to bind DNA in dead cells where the cell and nuclear membranes are disrupted. Cell viability (green) or death (red) can then be visualized using fluorescent or confocal microscopy, or quantitatively assessed using a spectrophotometer (plate reader) or a molecular imaging instrument capable of measuring fluorescence in tissues.

ATP/Metabolic Activity Methods that measure aspects of general cellular metabolism or enzymatic activity include: tetrazolium reduction, resazurin reduction, protease markers, and ATP detection. Metabolism based methods require incubation of a reagent with cells, which in the presence of metabolically active cells will react with a substrate to generate a colorimetric or fluorescent signal for detection. The signal is proportional to the number of viable cells present, since dead cells lose their ability to convert the substrate to a measurable signal.

Tetrazolium reduction Tetrazolium salt compounds are commonly used to detect viability. There are two basic categories of tetrazolium salts: (1) Cationic salts that can permeate viable eukaryotic cells through electrostatic interactions with the anionic plasma membrane. Cationic tetrazolium salts include MTT (3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide). (2) Anionic tetrazolium salts require an electron-coupling reagent to permeate live cells. These include MTS (3-(4,5-dimethylthiazol-2-yl)-5-(3-carboxymethoxyphenyl)-2(4-sulfophenyl)-2H-tetrazolium), XTT (2,3-bis-(2-methoxy-4-nitro-5-sulfophenyl)-2H-tetrazolium-5-carboxanilide), and WST-1 (2-(2-methoxy-4-nitrophenyl)-3-(4-nitrophenyl)-5-(2,4-disulfophenyl)-2H-tetrazolium). All of these assays depend upon the action of dehydrogenase enzymes in metabolically active cells to reduce anabolic cofactors such as nicotinamide adenine dinucleotide (NAD) and nicotinamide adenine dinucleotide phosphate (NADP). In the presence of active dehydrogenase enzymes, the colorimetric salt permeating into the cytosol of a metabolically active cell is reduced producing an intracellular purple precipitate (formazan), which typically requires additional steps (such as solubilization) before it can be quantified by absorbance using a spectrophotometer or a plate reader. Tetrazolium reduction is toxic to cells and is an endpoint measure of viability, which can only be performed in sacrificial constructs.

Resazurin reduction Resazurin (7-hydroxy-3H-phenoxazin-3-one-10-oxide) is a cell permeable redox indicator. Similar to tetrazolium reduction, viable cells with active metabolism can reduce resazurin to produce a pink and fluorescent precipitate (resorufin) which can be measured colorimetrically or fluorescently using a spectrophotometer or a spectrofluorometer. As with tetrazolium salts, the reduction of resazurin salts is induced by dehydrogenase enzymes in metabolically active and viable cells. Common resazurin reduction assays include Alamar Blue.

Intracellular ATP quantification Another cell viability assay utilizes of the critical role of intracellular ATP in respiration of living cells. Measurement of ATP concentration employs bioluminescent detection using firefly luciferase. ATP detection methods require lysing the cells to release ATP, which is subsequently reacted with luciferase to generate photons. When cells begin to die, they lose their membrane integrity and ability to synthesize ATP, and as such dying cells contain little to no ATP. Additionally, endogenous ATPases rapidly deplete any remaining ATP from the cytoplasm of a dying cell. Therefore, there is a tight linear correlation between cell number and the concentration of ATP. The bioluminescence intensity can be measured by any plate reader or other instruments capable of detecting luminescent signals.

Intracellular protease activity Protease activity within living cells, including cytoplasmic aminopeptidase, can be exploited as markers for viability by introducing a cell permeant fluorogenic protease substrate such as glycyl-phenylalanyl-aminofuorocoumarin (GF-AFC). The aminopeptidase-mediated cleavage of the GF amino acids liberates the fluorescent AFC, producing a signal proportional to cell viability, which can be measured using a spectrofluorometer. Intracellular protease activity can also be exploited as tools to assess cytotoxicity, such as lactate dehydrogenase, which is released by necrotic and apoptotic cells. In general, assessment of intracellular protease activity is non-toxic to cells, and therefore, has the benefit of enabling longitudinal, non-destructive assessments.

Caspase activity A defining feature of apoptosis is the activation of caspase enzymes, which play critical roles in the programmed cleavage of protein substrates and in the subsequent fragmentation of the cell. Caspase enzymes can be detected in cells in the early stages of apoptosis using immunohistochemistry. Alternatively, cell-permeant, caspase-susceptible fluorogenic substrates can be used with intact cells in engineered tissues or with tissue/cell lysates, and caspase activity as a measure of cell apoptosis can be quantified with a spectrofluorometer.

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Characterizing Cellular Differentiation Microscopy-Based Techniques The naked eye can view objects around 200 mm, but cannot resolve objects at the scale of typical animal cells (10–20 mm in diameter) and intracellular organelles (500–5000 nm). Microscopy is a useful tool that allows for visualization of cells within viable or fixed engineered tissues in vitro. The resolving power of a microscope is defined as the smallest detectable separation distance between two points on a specimen, which depends on several factors including the wavelength of light (or electrons in an electron microscopy), the optical and geometric features of the lenses, and sample preparation. Light microscopy can resolve features down to 0.2 mm. The resolving power of scanning and transmission electron microscopy (SEM and TEM) is about 2 and 0.1 nm, respectively (Fig. 3). Microscopy techniques have advanced over the past decades to enable detection, measurement, and time-lapse monitoring of numerous phenomena within living cells or engineered tissues at multiple scales. These techniques can be utilized to assess cellular differentiation, by probing specific cellular phenotypes using immunohistochemistry techniques, or as a tool to monitor cellular morphology and dynamics of cellular interactions within the scaffolds such as adhesion, spreading, and migration. The use of light or fluorescence microscopy in assessing cellular responses in tissue engineered constructs is limited by the size of the construct, its increasing opacity with maturation and compaction, and the refractive and diffractive features of the scaffold architecture. Dynamic observations on living cells within scaffolds are therefore restricted to superficial regions of interest, and often utilize fluorescent reporters to tag cells or intracellular proteins. However, if the objective is to capture detailed mechanistic information on cellular responses within the scaffolds, it is advantageous to perform cross-sectional, time-course studies on cryogenically or chemically fixed engineered tissues using routine histology with the aid of a specific stains, antibodies for immunohistochemistry, or complementary nucleic acid probes for in situ hybridization. Brightfield histology is commonly performed by microscopic examination of thinly cut sections of fixed tissue mounted on glass slides and subsequently stained with special chemical dyes to give contrast to differentially visualize the various classes of extracellular matrix proteins and cells, as well as intracellular organelles. There are numerous dyes with selective binding affinity to DNA, carbohydrates, lipids and proteins that can be used for primary, counter, or differential staining. The choice of which stain to use can be made depending on the types of cells and extracellular matrix expected and the questions asked. Immunohistochemistry and in situ hybridization offer a level of molecular resolution not possible with histology. In immunohistochemistry, primary antibodies specific to intra- or extracellular antigens (proteins) are incubated with the tissue sections, and then immunolabeled with appropriate secondary antibodies tethered to a chromogen or a fluorophore to visualize the antigen of interest. When unbound antibodies are washed, the sections can be viewed with the aid of a microscope to give a binary determination of whether the antigen of interest is present in the tissue sample. Immunohistochemistry is a powerful diagnostic tool in clinical pathology, and can be equally powerful in assessment of cellular responses and molecular evolution of engineered tissues. In situ hybridization localizes the expression of genes of interest in populations of cells within engineered tissues. The principle of

Fig. 3

The resolving power of microscopy.

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in situ hybridization takes advantage of the complementary nature of base pairing (A-T(U), C-G) in DNA and RNA. In this technique, a labeled (fluorescent, radioactive, or chromogenic) short sequence of nucleic acid bases (probe) designed to be complementary to a known sequence of DNA or RNA (gene of interest) is incubated with the tissue or section, and allowed to recombine with the gene of interest. When unbound probe is washed away, the detection of a signal or lack thereof gives a binary determination of whether the gene is expressed in the cells within the tissue samples, which can be done in whole mount preparations or with the aid of microscopic examination of a tissue section. While histology-based techniques including immunohistochemistry and in situ hybridization are valuable tools, whose use is commonplace in biomedical sciences including tissue engineering, they are not without limitations. These imaging-based techniques are for the most part non-quantitative. They are also sensitive to sampling, in the sense that only small tissue samples or biopsies in thinly sliced sections are examined to assess cellular responses in thick 3D tissues. Furthermore, histology-based techniques are highly empirical recipes, where nuanced procedural deviations in fixation, demineralization, or staining steps can lead to difficulties or misinterpretations. Another limitation is that there are no universally accepted predictive marker or panel of markers that would be conclusive regarding the functional maturity of engineered tissues. In that sense, histological-based assessments can be biased to common paradigms because they target a priori selected antigens or genes of interest, and therefore can be an effective tool for hypothesis testing, but a less effective discovery instrument. The cross-sectional nature of these destructive techniques gives snapshots of discrete points in time, and therefore the functional assessment of the cellular responses in engineered tissues should take into account the histological history (morphology, immunohistochemistry, and in situ hybridization), along with other relevant outcomes. A word of caution is that histology imaging produces more information than meets the untrained eye, and therefore the reading of histology slides should ideally be performed by trained pathologist.

Confocal and multiphoton microscopy Visualizing tissue engineered constructs using the histology-based techniques mostly uses standard phase-contrast (bright-field) microscopy or fluorescence microscopy. Both techniques are effective in imaging cells in monolayer cell cultures or thin tissue sections on glass slides, but have very limited applications in 3D engineered constructs. While fluorescence microscopy can be employed to detect the general location, number, morphology or viability of cells aided by genetic or chemical introduction of fluorophores, the quality of images weakens with depth, because the entire sample is evenly excited resulting in a blurred signal and unfocused background. Confocal laser scanning microscopy (CLSM) is a specialized fluorescent microscopy technique that increases the optical resolution by using a pinhole aperture to block out-of-focus light within the image. Then a composite image is created by scanning across a field of view. This technique increases the resolution to about 200 nm (0.2 mm) and allows for 3D imaging (z-stacks) to a depth of about 150–300 mm within the tissue construct, depending on the density and optical properties of the tissue. However, CLSM can activate photo-sensitive transients within the cells resulting in cell and DNA damage or cell death, and can therefore affect the cellular responses and functional outcomes. As such, CLSM should not be broadly considered a noninvasive imaging technique and perhaps should be used in conjunction with culture media additives such as antioxidants, provided that they do not influence the maturation of the tissue engineered constructs. Multiphoton-excited fluorescence microscopy overcomes limitations of confocal microscopy to enable deeper light penetration up to several hundred microns depth with reduced photodamage and exquisite resolution, which enables sharp 3D imaging of intra- and extracellular structures and assessment of dynamic interactions of the cells with their engineered extracellular matrix. Label-free multiphoton fluorescence microscopy can also take advantage of a nonlinear optical phenomenon known as second harmonic generation (SHG), which allows discrimination of signals from specific structures (such as fibrous collagen) from the bulk tissue-engineered constructs. SHG multiphoton microscopy can be particularly useful in imaging structural maturity and assessing the compositional anisotropy of tissue engineered extracellular matrix.

Electron microscopy Electron microscopy involves the use of accelerated electron beams, instead of light, to visualize the sample. This technique allows for much greater magnification, resolving details at the nanoscale, which is useful to observe very small structures within the cell. It is an endpoint imaging technique, which requires attention to specimen fixation and subsequent preparation protocols. In scanning electron microscopy (SEM), the electron beam projected at a non-normal angle scans a region of interest, losing energy by a variety of mechanisms, ultimately resolving surface features of the sample. In SEM, an ultrathin conductive coating (e.g., with gold), typically applied by low vacuum sputter coating, is necessary for imaging non-conductive samples such as tissue engineered scaffolds. In transmission electron microscopy (TEM) the electrons are transmitted through the specimen, and as such the specimens should be sectioned to thin slices. TEM can achieve magnifications approaching 50 million times and resolutions at the sub-nanoscale (< 1 Å or 0.1 nm). SEM has lower resolution compared to TEM, however, because the information retrieved in SEM are limited to 3D surface features, it has the advantage of being performed on bulk engineered tissues without the need for sectioning. Furthermore, the advent of environmental scanning electron microscopy (ESEM) facilitates imaging unfixed hydrated biological samples, producing images of good quality in low vacuum without the need for conductive coating.

FRET microscopy Despite the unsurpassed high-resolution imaging of the electron microscope, one of its main limitations is the lack of compatible labeling techniques necessary to interrogate the dynamic interactions between intracellular proteins in signal transduction pathways and numerous vital cellular processes. FRET microscopy is an emerging microscopy technique that enables precise localization and

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probing of the interactions between proteins in living cells at the molecular level, and could be of interest in mechanistic studies in the area of tissue engineering. FRET microscopy is based on combining the principles of Förster (Fluorescence) Resonance Energy Transfer phenomenon with high resolution fluorescence microscopy. This enables high resolution spatial and temporal assessment of the weak and transient molecular interactions between proteins inside the living cell.

Quantitative Gene Expression Analysis The central molecular biology dogma, which states that in the cell, DNA is transcribed into RNA, which is subsequently translated into proteins, offers a powerful toolbox for the interrogation of cellular functions. RNA can be extracted from tissue engineered constructs to assess changes in gene expression over time during the course of in vitro culture and in response to external stimuli. Extraction of high quality RNA in sufficient quantities from bulk tissue constructs can be challenging. The extraction techniques are nuanced and empirical, wherein subtle changes in various steps or reagents can lead to dramatic effects on purity and yield. This purity and yield challenge arises from specific and non-specific physical and chemical interactions between the nucleic acids and the biomaterials. In determining which extraction protocol to employ it is critical to think about the nature of these interactions and always check for RNA quality and yield. The reverse transcription of RNA into complementary DNA and its subsequent amplification offer some of the most powerful techniques for the assessment of gene expression, which can be used to identify differentiation state of cells or offer a mechanistic explanation of the cellular response to stimuli. Polymerase chain reaction (PCR) is used to amplify a single copy or a few copies of a segment DNA generated from RNA by reverse transcription techniques. Reverse transcription PCR or RT-PCR takes advantage of the complementary nature of deoxyribonucleotide-base pairing and is catalyzed by two enzymes, reverse transcriptase and DNA polymerase. Reverse transcriptase, first discovered in RNA viruses, catalyzes the synthesis of complementary DNA (cDNA) from RNA templates. DNA polymerase, an essential enzyme for DNA replication in mammalian cells, catalyzes the copying of new DNA strands from existing ones through multiple rounds of replication. In RT-PCR, the amplification selectively amplifies a gene transcript using a specific primer sequence unique to the gene being probed. The essential steps of PCR include thermal denaturation of the double-stranded DNA, annealing of the primer sequence to the complementary sites on the denatured DNA strands, followed by extension of the annealed primers by DNA polymerase with the aid of mixed-in oligonucleotides. This results in doubling the DNA product with every cycle of PCR. The amplified DNA product can then be visualized using gel electrophoresis and DNA-binding dyes. Quantification of the PCR product on a gel can be done by imaging the gel and quantifying the intensity of the gel band normalized to a housekeeping gene band (ubiquitously expressed gene across all treatment groups), which allows for assessment of relative changes in gene expression across experimental groups (Fig. 4). In comparison to classical RT-PCR methods, reverse transcription quantitative PCR (RT-qPCR), also known as quantitative realtime PCR, uses probes made of sequence-specific oligonucleotides labeled with a fluorescent reporter to allow the real-time detection of probe hybridization with its complementary sequence on the DNA strand with every cycle of the amplification. This allows for real-time quantification of the relative expression of the gene of interest, normalized to a housekeeping gene based on the fluorescence intensity of the hybridized probe. PCR-based techniques for assessment of relative gene expression are effective tools for hypothesis testing because they measure a priori selected genes of interest. The same principles used in PCR can be used for discovery instruments such as DNA microarrays and RNA sequencing.

DNA microarray A DNA microarray is made of large numbers of microscopic spots, to which probes (DNA sequences corresponding to a specific gene per spot) is attached. The probe at a spot can hybridize a target cDNA from an experimental sample if the gene it was designed to detect is expressed, and the probe-target hybridization can be detected because the target cDNA is a priori labeled (e.g., with fluorescence or chemiluminescence). The strength of the fluorescence or chemiluminescence signal at a spot on the microarray reflects the amount of target sample hybridizing to the probe, which is representative of the relative expression of the gene identified by the address of the spot on the microarray.

RNA sequencing In tissue engineering, RNA sequencing (RNA-seq), which utilizes next-generation sequencing (NGS) technology for unbiased reading of the whole transcriptome, can be used to assess the differential expression of genes (DEG) within purified RNA sample from different engineered constructs or experimental conditions. While the utility of RNA-seq in tissue engineering might be focused on the assessment of DEG, it is nevertheless a very powerful technique that can assess other important features of the transcriptome. RNA-seq overcomes many of the practical limitations of microarrays. Most notably, RNA-seq does not require a priori knowledge of the sequences of the gene of interest. Rather, once the sequences are determined using NGS, they can be aligned to mapped genomes of different species to identify the genes significantly expressed and the levels of differential gene expression between experimental conditions. This technique generates large sets of data whose analysis requires access to numerous knowledge databases of gene functions, pathways, and networks and sophisticated statistical analysis techniques for unbiased interpretation of the information. It is, therefore, a powerful discovery and hypothesis generating tool. It should be noted that RNA extraction from bulk tissue and the aforementioned techniques cannot discriminate between the different cell types contributing to differential gene expression in engineered composite tissues. Single cell sequencing, which is technically feasible using NGS, can be enabled by additional sophisticated techniques such as laser capture microdissection, which

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Fig. 4 Gene expression analysis methods. There are three methods of gene expression analysis (1) reverse transcription-polymerase chain reaction (RT-PCR), (2) DNA microarray, and (3) RNA sequencing (RNA-seq). mRNA is isolated from a population of cells, which can be directly sequenced or reverse transcribed to generate cDNA. RNA-seq allows for detection of transcripts of thousands of genes from a given sample. PCR methods result in quantification of expression of a single gene within a sample. The PCR amplification reaction involves three steps: heating to separate DNA strands, hybridization of primer sequences, and DNA synthesis with DNA polymerase. Each cycle doubles the number of DNA transcriptions. PCR products can be quantified at the end of the amplification or in real-time (qPCR). Lastly, multiple primers can be bound to a microarray that allows for detection of hundreds of genes. Fluorescence signal detection and normalization to a control sample allows for quantification of changes in gene expression.

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allows the isolation of single cell or a uniform population of cells in a tissue (e.g., from a histology slide) for gene expression analysis under direct microscopic visualization.

Quantitative Protein Analysis Proteins perform many processes and functions within the cell including providing structural support and catalyzing enzymatic reactions, and can be probed to assess cell phenotype. Tissue engineers can measure intracellular proteins, proteins expressed on the surface of cells, and proteins secreted by cells into the extracellular matrix. Most of the methods utilize immunological detection of proteins by introducing antibodies to probe for specific antigens on the target protein. These methods require lysing the tissue and cells and are therefore used as endpoint destructive assays.

Western blot A Western blot is a semiquantitative immunological method for detecting electrophoretically separated proteins. Isolated proteins are first separated using a sodium dodecyl sulfate polyacrylamide (SDS-PAGE) gel electrophoresis (SDS-PAGE). After separation using SDS-PAGE, the proteins are transferred to a nitrocellulose paper and incubated with a specific antibody that has been coupled to a label (e.g., a radioactive isotope or a fluorescent dye), which can then be developed for detection. This method of protein detection provides qualitative and semi-quantitative data about the protein of interest relative to a ubiquitous housekeeping or loading control protein not affected by the experimental conditions. The size of the protein band can be normalized by a loading control band using image analysis software to provide semi-quantitative results.

Enzyme-linked immunosorbent assays (ELISA) Enzyme-linked immunosorbent assays or ELISA is a common quantitative immunoassay used to measure the concentration of intracellular or secreted peptides and proteins from tissue lysates and culture media using antibodies typically labeled with an enzyme marker. The key step of ELISA is the immobilization of the antigens from the sample to a surface of a multiwell plate. Protein detection is accomplished by measuring enzyme activity on a colorimetric substrate following an enzyme–antibody– antigen reaction. The enzyme activity on the substrate causes a colorimetric change that is proportional to the concentration of antigen (protein) and can be measured spectrophotometrically. In Direct ELISA, the antigen can be captured by direct adsorption to the assay plate and labeled with a primary antibody. The second type of ELISA is a “sandwich” assay in which the antigen is bound between two primary antibodiesdthe capture and the detection antibody. The third type is known as Competitive ELISA, in which an unlabeled antibody is incubated with the sample presumably containing the antigen of interest, and then added to an antigen-coated well. The next step after washing unbound antibodies is the addition of an enzyme-linked detection antibody, specific to the primary antibody and detection is enabled by adding a colorimetric enzyme substrate.

Flow cytometry Flow cytometry is a powerful laser-based tool used to sort cell populations based on surface or intracellular protein expression. The most common use of flow cytometry in tissue engineering is the characterization of cell types (e.g., stem cells) prior to seeding on biomaterial scaffolds. Protein detection is completed by fluorescently labeling the cellular proteins, commonly labeled using antibodies coupled to fluorophores, and passing the cell suspension through a laser which excites the fluorophore to emit light at varying wavelengths. The fluorescence counts can be measured to determine the number of cells expressing the protein(s) of interest.

Fo¨rster resonance energy transfer assays Förster or fluorescence resonance energy transfer or FRET-based assays are based on the principle of measuring the proximity of two fluorophores and is commonly used to detect molecular interactions intramolecular distances. The principle mechanism of FRET is energy transfer between the two fluorophores. A donor fluorophore is initially excited and can transfer energy to an acceptor fluorophore. FRET can provide information about distances between domains of the protein to determine its conformational state and protein interactions. Additionally, FRET-based assays can be used to measure proteolytic enzyme activity, interactions between proteins such as integrins and extracellular proteins, and metabolic or signaling pathways in engineered tissues.

Assessment of Mechanical Properties The mechanical properties of cells play a significant role in many biological processes. Cellular stiffness can be an indicator of actin organization and differentiation status. Micro-indentation using an Atomic Force Microscope (AFM) is a common method used to measure the stiffness of living cells. The general methodology is to indent a cell with an AFM tip of a specified geometry and measure the applied force from the bending of the AFM cantilever. The small tip is made of silicon or silicon nitride and is made by nanofabrication techniques. The displacement of the cantilever can be measured with a laser. Mathematical modeling is used to fit the force-displacement curve to determine cellular stiffness. In addition to measuring cellular stiffness, AFM can be used to scan the surface of a tissue engineered construct to determine the nanotopography. While the assessment of cellular mechanics can be informative of cellular phenotype, the assessment of construct biomechanics can provide information about the functional properties of engineered tissues. In creating tissue constructs in vitro it is not only

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important to mimic biological responses but also the mechanical phenotype of the native tissue. Typical tissue loading patterns include: tension, compression, and shear. Simple tests assess mechanical properties in biologically-relevant modalities. Appropriate biomechanical assessment of an in vitro constructs depends on the engineered tissue’s function. For example, tendon constructs can be tested in uniaxial tension while engineered cartilage is tested in confined or unconfined compression or shear and engineered heart valves can be tested in flexion. Engineered skin on the other hand can be stretched in biaxial loading configurations. These types of biomechanical tests provide information about structural and material properties (force-deformation and stress–strain constitutive laws, respectively), which can be critical to the function of the engineered tissue or organ.

Emerging Technologies The ability to make real-time measurements of key responses of cells within an engineered tissue using non-invasive imaging techniques in a high throughput system is a critical need for the heuristic science of tissue engineering. Genome editing using CRISPR/ Cas9 is an emerging technology that not only has therapeutic value, but can also enable the introduction of multiplex reporter signals (e.g., fluorescence) as readouts of critical cellular functions or phenotypic gene expression and secretome markers, which in theory can be detected using label-free, time-lapse fluorescence microscopy. Human microphysiological systems (tissue- and organ-on-chip) can also provide a high throughput platform for screening the myriad combinations of stimuli that might need to be optimized, investigate the interactions with other tissues and organs, and curtail the reliance on animal testing in the early phases of development. The combination of the two technologies, genome editing and microphysiological systems, can be transformative for the practice of tissue engineering.

Further Reading Alberts, B. (2015). Molecular biology of the cell. New York: Garland Science, Taylor and Francis Group. Berthiaume, F., Maguire, T. J., & Yarmush, M. L. (2011). Tissue engineering and regenerative medicine: History, progress, and challenges. Annual Review of Chemical and Biomolecular Engineering, 2, 403–430. Ding, S., Kingshott, P., Thissen, H., Pera, M., & Wang, P. Y. (2017). Modulation of human mesenchymal and pluripotent stem cell behavior using biophysical and biochemical cues: A review. Biotechnology and Bioengineering, 114(2), 260–280. Guilak, F., Butler, D. L., Goldstein, S. A., & Baaijens, F. P. T. (2014). Biomechanics and mechanobiology in functional tissue engineering. Journal of Biomechanics, 47(9), 1933– 1940. https://doi.org/10.1016/j.jbiomech.2014.04.019. Ireland, R. G., & Simmons, C. A. (2015). Human pluripotent stem cell Mechanobiology: Manipulating the biophysical microenvironment for regenerative medicine and tissue engineering applications. Stem Cells, 33(11), 3187–3196. Sittampalam, G. S., Coussens, N. P., Brimacombe, K., et al. (Eds.). (2004). Assay Guidance Manual. Bethesda, MD: Eli Lilly & Company and the National Center for Advancing Translational Sciences. https://www.ncbi.nlm.nih.gov/books/NBK53196/. Vielreicher, M., Schürmann, S., Detsch, R., Schmidt, M. A., Buttgereit, A., Boccaccini, A., & Friedrich, O. (2013). Taking a deep look: modern microscopy technologies to optimize the design and functionality of biocompatible scaffolds for tissue engineering in regenerative medicine. Journal of the Royal Society Interface, 10, 20130263. https://doi.org/ 10.1098/rsif.2013.0263.

Assessment of Tissue Constructs In Vivo in Regenerative Engineering Anuradha Subramanian and Swaminathan Sethuraman, SASTRA Deemed University, Thanjavur, India © 2019 Elsevier Inc. All rights reserved.

Introduction Animal Models Cardiovascular Tissue Construct Nerve Tissue Constructs Bone Tissue Construct Osteochondral Interfacial Tissue Construct Challenges in Assessment of In Vivo References

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Introduction Transplantation procedures are clinically required due to the increasing prevalence of end-stage organ failure and loss of tissues due to disease, disorder, and trauma. Shortage of donor organs and risk of immunosuppressing agents administered are the major limitations of organ transplantation. Regenerative engineering is an interdisciplinary approach that combines trophic factors, cells and extracellular matrix analog for the functional restoration of complex tissues and organ systems (Gu et al., 2014; Furth et al., 2007). The basic requirement in tissue engineering medical products development is to characterize the safety and regenerating efficacy of tissue construct using both in vitro and in vivo models. Though in vitro techniques are useful they are limited due to the nonphysiological culture condition, depletion of media with the accumulation of by-products, sudden exchange of media, diffusion limited oxygen supply and inability of monolayer culture to recapitulate complex in vivo system provoke the importance of in vivo assessment of tissue constructs (Slack et al., 2015). The preclinical assessment of toxicity and efficacy of tissue substitutes is required for the translation of tissue constructs into clinics. Although various animal models are available and are used to envisage human response, each animal model has its own limitations due to the difference in anatomy, physiology, and pathophysiology from species to species (Ahern et al., 2009). Animal models such as rodents, pigs, rabbits, dogs, sheep, and primates are widely used in various tissue-engineering applications (Fan et al., 2014; Fricain et al., 2013; Hu et al., 2013). Successful regeneration of any tissue mandates the fabrication of ideal tissue substitute that is equivalent to structure, function, and mechanical strength of native tissue (Fan et al., 2009; Yu et al., 2008; Millera et al., 2012). However, none of the animal models are suitable to assess all the desirable characteristics of ideal tissue constructs while many animal models exist to evaluate the specific properties (Ahern et al., 2009; Angius et al., 2013). Hence, the choice of animal models for the assessment of tissue constructs in regenerative engineering plays a vital role as the results from preclinical models are often extrapolated to human responses.

Animal Models Animals such as guinea pigs, goats, dogs, pigs, rodents, rabbits, and primates such as apes and monkeys are most widely used to test the toxicity and efficacy in tissue engineering. An appropriate animal model should be determined based on the tissue constructs and its applications (Cui et al., 2007). The choice of the animal model needs to reflect the appropriate aspect of human physiology and pathological response (Rashid et al., 2004). For example, rodents are used for many tissue engineering applications to evaluate the efficacy in regenerating skin, sciatic nerve, vascular conduit due to the ease of care, availability and the physiology of rats is more similar to humans (Allen et al., 2014; Krupnick et al., 2002). However, the results obtained from rodent model require further testing and evaluation in larger animal models as these results cannot be easily correlated to humans due to the difference in biomechanics and anatomical features with humans (Liebschner, 2004; Milano et al., 2006). Each animal model has its own merits and limitations. Rabbits have been chosen for some studies due to its availability, costeffectiveness, and ease of maintenance. The anatomy, physiology, and immunological characteristics of pigs are equivalent to humans (Summerfield et al., 2015; Guo and Yang, 2015). However, rapid growth rate, expensive animal maintenance including care and feeding, complex surgical and anesthetic procedures with a mismatch in engineered constructs due to rapid growth of animals are the potential drawbacks of pig models (Rashid et al., 2004). Dogs are much preferable animal models than pigs as it has modest growth rate (Lim and Temenoff, 2009). Dog breeds such as beagles, foxhounds have been used as animal models for various tissue applications. Short-term studies have been carried out in small animal models such as rodents, guinea pigs, and rabbits up to 12 weeks whereas long-term studies have been performed in longer life-expectancy animals such as dogs, rabbits, sheep, goats, pigs, etc. (Li et al., 2006; Chang et al., 2006). Selection of appropriate animal models to evaluate the specific properties

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of tissue construct is critical for the thorough investigation of the biomaterial substitutes. Hence, substitutes for cardiovascular, nerve, bone, and osteochondral tissues require the search of appropriate animal models as follows.

Cardiovascular Tissue Construct Cardiovascular disease is one of the leading causes of death worldwide. Biomaterials intended for cardiac or vascular applications have to be extensively evaluated using in vivo preclinical animal models based on International Organization for Standardization (ISO). The general requirements for the evaluation have been found in ISO 5840-1:2015 (Cardiovascular implants-cardiac valve prosthesis) and ISO 7198:2016 (Cardiovascular implants and extracorporeal systems-vascular prosthesis-tubular vascular graft and vascular patches). Prior to clinical evaluation, ISO requires that the tissue constructs should be tested in envisioned anatomical sites using preclinical animal models. Larger animal models such as dog, pig, sheep, and primates are preferable model for testing the therapeutic efficacy since it shows slower heart rate that closely resemblance the cardiac tissue physiology to humans (Chong and Murry, 2014). Preclinical examination of cardiovascular implants such as heart valve, vascular grafts, and cardiac assist device should be based on the expected complications of device-on-host and host-on-device. For instance, implantation of heart valve prostheses may lead to paravalvular leak, thrombosis, emboli formation, calcification, hemorrhage, infection, hemolytic anemia, etc., whereas the grafting of vascular tissue substitutes results in anastomotic hyperplasia, disintegration, false aneurysm, infection, thrombosis, embolism, etc. (Zilla et al., 2007). Complications of vascular stents are thrombosis, incomplete expansion, strut-based inflammation, proliferative restenosis, and infections (Scott, 2006). Hence, blood compatibility and blood–material interaction should be evaluated for testing the safety aspects of the cardiovascular implants. Canine model is advantageous for assessing the cardiovascular regeneration due to the resemblance of cardiovascular physiology to humans (Byrom et al., 2010). In addition, metabolic and hematological differences, white blood cells differential count and platelet counts are also identical to humans (Levine et al., 2015). This model is mainly used to evaluate the thrombogenesis of stent due to its hypercoagulation property compared humans (Bauer and Moritz, 2013). Fufukoshi et al., have fabricated robust type C biotubes and implanted into the femoral arteries of beagles by end–end anastomosis. The implanted modified biotubes showed higher external pressure resistance, excellent burst strength and 100% patency rate at 7 days of postimplantation compared to original thinner wall biotubes (Furukoshi et al., 2016). Similarly, sheep is the choice of animal model for evaluating the cardiac valves due to the close resemblance to the mechanical, hemodynamic flow property and calcification propensity of human (Driessen-Mol et al., 2014; Schmidt et al., 2007). Specifically juvenile sheep < 6 months remain the best model of choice for the evaluation of calcification of vascular grafts (Herijgers et al., 1999). However, sheep may not be a sensitive model for the assessment of valve thrombosis due to the higher platelet count and its adhesiveness with respect to humans (Ratner, 2013). Pigs are more ideal models for evaluating the pacemakers due to the similarities of electrophysiological parameters such as P and R wave amplitude, P-R intervals, Q-T interval, QRS duration with humans (Ratner, 2013). Although larger animal models are ideal for preclinical evaluation of vascular grafts, rat carotid and abdominal aorta replacement models have been used for the evaluation of in vivo performance of vascular grafts due to ease of surgical exposure without temporary removal of visceral organs (Allen et al., 2014; Katsimpoulas et al., 2015). Chan et al., had recently evaluated the efficacy of polycaprolactone electrospun vascular grafts in mouse carotid grafting model using C57BL/6 transgenic mice. Use of transgenic mice for vascular grafting procedure showed higher neointimal hyperplasia at proximal anastomosis and found to decrease towards the mid portion comparable to that of physiological response of humans to vascular replacement (Chan et al., 2017). Since cardiovascular implants have direct contact with blood and soft tissues, evaluating the consequence of blood–material and soft tissue–material interactions are necessary to extrapolate the human physiological and pathological response of cardiovascular grafts.

Nerve Tissue Constructs Regeneration of critical-sized defect (5–30 cm) in peripheral nerve repair demands the exogenous nerve conduits to reestablish the outgrown neurites to the distal end of the axons (Kaplan et al., 2015). On the contrary, spinal cord injury remains challenging, as these axons do not regenerate intrinsically (Geoffroy et al., 2017). Current strategies to repair nerve gaps up to 10 cm is nerve autografts whereas vascularized autografts are required for repairing the gaps of > 10 cm (Trehan et al., 2016; Palispis and Gupta, 2017). Though autografts are gold standards, these grafts are not ideal due to the shortcomings such as donor site morbidity, size mismatch and poor availability (Rinker and Vyas, 2014). Hence, the search for new biomaterials and novel strategies to repair the larger size defects demands in vivo testing using animal models. Animal models for critical-sized defects have been chosen based on the relevance to neurobiological regeneration profile of the humans since the critical size of the defect may vary from species to species (Angius et al., 2013). Critical nerve defect of rat and rabbit, is 1.5 and 3 cm respectively, while pigs and humans have up to 4 cm (Kaplan et al., 2015). Mohammadi et al., bridged 10 mm sciatic nerve defect using cyclosporine-A releasing chitosan conduit in rat sciatic model. Apart from histomorphometric and immunohistochemical analysis to confirm the regenerated axons and Schwann cell-specific biomarker (S-100), functional recovery of regenerated nerves play vital role in the efficacy evaluation (Mohammadi et al., 2014). Walking track analysis is a comprehensive test used to evaluate the sensory input, cortical integration, and motor response by determining functional sciatic nerve index and muscle mass. Peptides have been self-assembled in aligned

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poly(lactide-co-glycolide) nanofibers and the efficacy of the scaffolds to regenerate the sciatic nerve was evaluated in a rat model for 16 weeks (Nune et al., 2017). Similarly electrophysiological assessment has been performed at third and sixth month of the implantation of conductive polypyrrole/poly(D,L-lactic acid) nerve conduits in rat sciatic model by recording nerve conduction velocity on lower leg triceps (Xu et al., 2014). Unlike humans, rodents have shorter nerve gap (< 10 mm), complete recovery after injury and rapid peripheral axonal regeneration that limits the use of rat sciatic model used for the evaluation of biomaterial constructs. For the evaluation of tissue construct in repairing spinal cord injury (SCI), rodents especially rats are widely used since it develops fluidfilled cystic cavity immediately after injury, which mimics the human pathology. Out of various injury models, contusion SCI model are preferable in order to mimic the pathophysiology of the humans (Sharif-Alhoseini et al., 2017). However, this biomechanical force based injury fail to create the lesion at specific axons. Complete transection of nerve is another model that mimic the severe category of injury where no axons are spared between the two nerve stumps (Lee and Lee, 2013). Hence, this model is more advantageous to evaluate the axonal regeneration potential of biomaterial scaffolds. Rabbits are considered as a poor animal model for nerve tissue engineering applications due to hopping and variation in hind limb function than humans (Angius et al., 2013). Despite the use of larger animal models such as dogs, sheep, pigs, and monkeys to evaluate the synthetic constructs with larger gaps, expensive animal care and ethics restrict its potential use for functional evaluation. Among the various larger animal models, axotomy of nonhuman primates is equivalent to humans, which drives better comparison with probable human response (Hu et al., 2007).

Bone Tissue Construct Animal models explored for testing of repair and regeneration of bone substitutes are only boney vertebrates from fish to mammals. Each model has its own merits and limitations in evaluating the specific biological activity of the tissue construct and tissue responses based on the factors such as defect size, anatomy, and physiology, biology of animals (An et al., 2000). To elucidate the surface osteoconductivity of the tissue substitute, small defects of maximum 2–5 mm and minimum of 1 mm (mice, rats, rabbits) is considered whereas for porous scaffolds, the larger defect of 1–5 cm of maximum and minimum of 5–10 mm is used, as the mass transport function is the major variable of porous substitutes (Reichert et al., 2009). Rodent models are generally not preferred for assessing the tissue constructs for bone repair due to the life-long process of skeleton growth, absence of osteoclast tunneling and limited trabecular bone content in comparison with bone biology of human (Miller et al., 1995). Rabbits are most widely used in musculoskeletal and mandibular research studies owing to its ease of handling, larger trabecular bone mass, short skeletal maturity with Haversian remodeling as humans (Jowsey, 1966). Biology and composition of bone for sheep, goat, dog, and pig are equivalent to humans (Bagi et al., 2011). Tissue-engineered bone substitute using hydroxyapatite/tricalcium phosphate scaffold seeded with adipose-derived stem cells demonstrated the reconstruction of bone in alveolar cleft model of mongrel dog (Pourebrahim et al., 2013). Systemic variables such as age, sex, endocrine factors such as estrogen, parathyroid hormone PTH and local variables such as type and size of defects interfere with the outcome of bone repair (Florencio-Silva et al., 2015). Hence, critical-sized bone defects remain the better translational model for the evaluation of bone tissue constructs. Anatomy, morphology, and remodeling with mineral density, mineral composition of bone in pigs show the close resemblance to human bones except the denser trabecular network (Pearce et al., 2007). Further regeneration rate of bone in pigs (1.2–1.5 mm/day) is equivalent to humans (1.0–1.5 mm/day) (Raschke et al., 1999). Among the large animals, multiple defects of 10 mm wide and 15 mm long created in canine femoral multi-defect model aids in the intra-subject assessment of grafts (Cui et al., 2007). Similarly multiple criticalsized cranial defects (20 mm diameter) in Dutch Texel sheep have been filled with different poly(trimethylene carbonate)-based composite materials where one of defect left untreated (Zeng et al., 2017). Sheep is an excellent model to evaluate different bone forming ability of tissue constructs by offering multiple surgical sites for critical-sized cranial defects.

Osteochondral Interfacial Tissue Construct Degeneration of articular cartilage is a common disease in elder patients resulting in dysfunction of joints, long-term disability associated with chronic pain. These defects progress to affect the underlying bone that necessitates the total joint replacement. However, available therapeutic options are not fully curative to regenerate hyaline cartilage, and instead they develop fibrocartilage (Prado et al., 2016). Ease of handling, caging, and maintenance costs of animals are the major reasons for using murine models. However, rats and mice are not ideal models to predict the response of humans due to the smaller and thinner cartilage tissue which heals better healing potential when compared to humans (Bagi et al., 2011). Cartilage thickness of murine, rabbit, canine, caprine, ovine, porcine, equine, and human is 0.1; 0.3; 0.6–1.3; 0.7–1.5; 0.4–0.5; 1.5–2.0; 0.96–3.13, and 2.2–2.5 mm respectively (Chu et al., 2010). Based on the thickness of cartilage, the critical size defect of rabbit (3 mm), dog (4 mm), sheep (6–7 mm), pig (6 mm), horse (9 mm), and human (10 mm) also varies (Frisbie et al., 2006). Despite the advantage of early skeletal maturity in rabbits, spontaneous healing potential with difference in biomechanics is major limitations in correlating the results from rabbit to humans (Hurtig et al., 2011; An and Freidman, 1998). Although canine models share many cartilage pathologies with humans such as osteochondritis, osteoarthritis, canine knee joint may not model the human knee joint due to the anatomical difference (Proffen et al., 2012). Characteristics such as joint size, biomechanics, weight, lack of intrinsic healing, trabecular bone thickness of porcine similar to humans make the pigs as an ideal animal model (Chang et al., 2006). However, expensive animal maintenance and poor availability of skeletally mature animals are the major barriers of porcine model (Vasara et al., 2006). Performing preclinical evaluation

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of osteochondral tissue construct in equine model receives much attention due to the anatomical equivalence to human knee with similar bone mineral density, larger defects with ease of assessment of repair tissues, poor intrinsic healing and fully extended stifle during walk (Colbath et al., 2016; Bertuglia et al., 2016; Harrison et al., 2014).

Challenges in Assessment of In Vivo Animal studies are performed to test tissue constructs to evaluate the response to interventions so that adequate data is available to plan clinical trials. The safety and efficacy data obtained for substitutes in a homogenous animal population does not always prove to be effective in humans also. Hence, the choice of appropriate animal model is essential to translate the research outcome from animal to humans. Unresolved challenges remain in the translation of preclinical animal model to clinics. They are: (i) difference in anatomical, physiological, and pathological response between animals and humans; (ii) lack of randomized or blind studies in animal model; (iii) difficulty in systemic and local toxicity prediction; (iv) surgical limitation; (v) size mismatch; (vi) ethical concerns; and (vii) differential healing response. Further, factors such as device exposure duration, nature of injury, injured tissue; age and species may also influence the outcome. Most of the animal models are ideal for evaluation of complete safety and efficacy of the tissue constructs. For example, animal models from murine to nonhuman primates shows extensive endothelialization in lumen surface as healing response, which is absent in humans. The success rate for translation of tissue substitutes to clinics depends on the choice of the animal models that in turn depends on the end-use application. Hence, the choice of appropriate animal models is important to bridge the preclinical gap for translating the materials from lab to clinics.

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Bioengineered Kidney and Bladder DS Koslov and A Atala, Wake Forest Baptist Medical Center, Winston–Salem, NC, USA © 2019 Elsevier Inc. All rights reserved.

Introduction Cells Used in Regenerative Medicine Adult Stem Cells Embryonic Stem Cells Somatic Cell Nuclear Transfer Altered Nuclear Transfer Single Cell Embryo Biopsy Cells from Arrested Embryos Amniotic Fluid and Placental Stem Cells Induced Pluripotent Stem Cells Biomaterials Acellular Matrices Synthetic Biomaterials Biologic Biomaterials Vascularization and Innervation Bioreactors Renal-Specific Studies Cell-Based Renal Therapy Renal Precursor Therapy Factor-Based Renal Therapy Materials-Based Renal Therapy Organ Decellularization and Repopulation Conclusion References

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Glossary Blastocyst Early developmental stage, typically 70–100 cells. It contains the inner cell mass, which will develop to the trilaminar germ layers, and trophoblast, which will become the placenta. Embryo Zygote derivative that develops into offspring; in humans, this is defined from fertilization to the end of the first 8 weeks gestation. Embryoid Body In vivo aggregation of pluripotent stem cells, which can undergo differentiation to cells of all three germ layers. A source of difficulty in clinical stem cell use. Extracellular Matrix (ECM) Complex array of proteins, polysaccharides, growth factors, and proteoglycans found in animal tissue that provide support, regulate intercellular communication, divide tissues, and guide development as an embryo. Commonly referred to as connective tissue, it is essential in tissue growth and function. Myelomeningocele Congenital defect with open spinal canal with protrusion of meninges and spinal nerves through the defect. Patients may have associated defects and neural damage to structures such as the bladder. Stem Cells Population of immature cells that remain undifferentiated and can become several types of mature cells. Several types exist based on the location and range of mature cells they can become.

Introduction Organ transplantation has long been a therapy for patients suffering from diseased and terminally damaged organs, even dating to antiquity. Initial contributions to transplantation came from the joined efforts of Charles Lindbergh and Alexis Carrel, a Nobel Prize Laureate in Medicine, in ex vivo organ preservation. Transplantation came to the forefront in 1954 when Joseph Murray performed the first whole organ transplant between identical twins. Murray continued his work in the 1960s when he successfully transplanted a kidney between nonrelated patients. However, lack of good immune suppression, inability to monitor and control rejection, and donor shortages spurred physicians and scientists to look for alternative sources of transplantable tissues.

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Soon, synthetic materials were introduced to substitute for whole organ transplantation. Tetrafluoroethylene (Teflon) and silicone, for example, introduced a wide array of devices that could be applied for human use. While these devices provided structural replacement for organs, providing function was the next challenge. As new techniques for cell harvesting, culture, and expansion were developed, studies of the extracellular matrix (ECM) and its interaction with cells led the way to a better understanding of cell and tissue growth and differentiation. The advent of human bone marrow transplant in the 1970s was the first in development of several ways to combine devices and materials with cell biology concepts, ultimately creating a new field called tissue engineering. As scientists from different fields came together with the common goal of tissue replacement, the field of tissue engineering became more formally established. More recently, the fields of cell transplantation and tissue engineering have been combined with stem cell biology, creating what is now known as ‘regenerative medicine.’ While organ transplantation remains a mainstay of treatment for patients with severely compromised organ function, the number of patients in need of treatment far exceeds the organ supply, and this shortfall is expected to worsen as the global population ages. However, recent advances in regenerative medicine suggest that it can provide new alternatives to donor organs. In the past two decades, scientists have attempted to grow native and stem cells, engineer tissues, and design treatment modalities using regenerative medicine techniques for almost every tissue of the human body. The genitourinary system has been at the forefront in the advancement of bioengineering. Both basic science and clinical experience have shown advances in both bladder and kidney bioengineering, two organs that are targeted by congenital, neurological, traumatic, and degenerative processes. This article will discuss the basics of tissue engineering including cell isolation and biomaterial selection, and then specific experiences in both urologic and renal tissue regeneration.

Cells Used in Regenerative Medicine Bioengineering relies on appropriate selection of cells to populate the new tissues. Ideally, the cells would be easily derived in an ethically acceptable manner, expand easily in culture, and be safe to use in the clinical setting. By definition, stem cells regenerate indefinitely in an undifferentiated state, and can mature into one of many types of cells (Hipp and Atala, 2008). They come in three main categories: adult, embryonic, or amniotic fluid and placental. More recently, a new population of induced pluripotent stem cells (iPSCs) has been introduced. Embryonic stem cells (ESCs) are obtained from the inner cell mass of a blastocyst and are pluripotent. Fetal and neonatal amniotic fluid and placenta taken during amniocentesis or chorionic villus sampling contain multipotent cells that may be useful in cell therapy applications. Adult stem cells, on the other hand, are usually isolated from organ or bone marrow biopsies. When cells are used for tissue reconstitution, donor tissue is dissociated into individual cells that are either implanted directly into the host or expanded in culture and then attached to a support matrix or reimplanted after expansion (Hipp and Atala, 2008). Many techniques for generating stem cells have been studied over the past few decades, and the main ones are discussed in detail below. Table 1 summarizes the unique forms of generating pluripotent stem cells.

Adult Stem Cells Adult stem cells are unipotent undifferentiated cells located within specific organs and tissues including muscle (Crisan et al., 2008), gastrointestinal tract (Weiner, 2008), brain (Taupin, 2006), bone marrow (Till and McCulloch, 1964), and skin (Jensen et al., 2008), and replace cells lost in their respective locations only. Apart from those located in bone marrow, these stem cells are small in number and difficult to maintain in culture (Mimeault and Batra, 2008). An advantage of using native adult stem cells is that they can be obtained from the specific organ to be regenerated, and then expanded and used in the same patient without rejection. Major Table 1

Methods of generating pluripotent stem cells

Method

Advantage

Limitation

Somatic cell nuclear transfer

Customized stem cells Has been shown to work in nonhuman primates Patient specific to embryo Does not destroy or create embryos Has been done in humans

Requires oocytes Has not been shown to work in humans Allogeneic cell types Are not known if single cells are totipotent Requires coculturing with a previously established human ESC line Allogeneic cell types Quality of cell lines might be questionable Ethical issues surround embryos with no potential Modified genome Has not been done with human cells Retroviral transduction Oncogenes

Single cell embryo biopsy

Arrested embryos Altered nuclear transfer Reprogramming

Cells obtained from discarded embryos Has been done in humans Customized stem cells Customized stem cells No embryos or oocytes needed Has been done with human cells

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advances in cell culture techniques have been made in the past decade and make autologous adult stem cells attractive targets for clinical applications. In the bladder, autologous cells can be isolated for engineered tissue repopulation. Progenitor cells in the bladder have been identified in larger quantities near the neck of the bladder and trigone (Nguyen et al., 2007), and procedures have been developed to expand urothelial cells in vitro. These cells can subsequently be seeded on to a matrix and reintroduced to the host with varying degrees of success. Even cells derived from a diseased organ can be expanded and reimplanted in a scaffold and will reacquire normal form and function (Lai et al., 2002). Within the adult stem cell group exists a unique subpopulation, the mesenchymal stem cells (MSCs)/multipotent adult progenitor cells, which have the ability to differentiate into several cell types. Originating in the bone marrow, MSCs are multipotent, demonstrating the ability to differentiate developmentally if injected into a blastocyst (da Silva Meirelles et al., 2008), as well as to integrate into pulmonary structures (Nolen-Walston, 2008), spleen (in’t Anker et al., 2003), muscle (Crisan et al., 2008; Luttun et al., 2006), liver (Mimeault and Batra, 2008; Ikeda et al., 2008), and gastrointestinal tract (Jiang et al., 2002). These cells are easier to maintain in culture than other adult progenitor cells (Caplan, 2007). Notably, MSCs have demonstrated renoprotection in several models of acute kidney injury. MSCs were found to be increased in numbers in peripheral circulation after acute kidney ischemia (Kale et al., 2003) and were seen to relocate to the kidney (Togel et al., 2005a), likely in response to increased levels of stromal cell-derived factor-1 (Togel et al., 2005b). The MSCs were traced to the kidney (Lange et al., 2005), and administration of MSCs in postischemic renal injury hastened renal function (Lange et al., 2005; Kunter, 2007; Morigi et al., 2004; Lin et al., 2003). They have also shown to have positive effects on chronic renal failure models (Choi et al., 2009). It has been established that MSCs that move to the kidney after injury do not integrate into the organ, rather they deliver microvesicles in a paracrine manner, which provides renoprotective molecules such as insulinlike growth factor-1 (IGF-1) and vascular endothelial growth factor (VEGF). Furthermore, MSC may reduce inflammatory reactions after acute kidney damage by inhibiting several T-lymphocyte activities, keep dendritic cells in an immature state, reduce interleukin (IL)-2 production, and augment CD4þCD25þ T-regulatory cells (Aggarwal et al., 2005). Therapeutic use of MSC in kidney-related injury may yield an effective manner of cell-based therapy. Beyond adult and MSCs, pluripotent cells also pose a promising source for tissue regeneration.

Embryonic Stem Cells Pluripotent cells were found in the inner cell mass of the human embryo in 1981, and given the name human embryonic stem cells (hESCs). These cells are able to differentiate into all cells of the human body, excluding placental cells, and self-propagate in an undifferentiated state. These cells have great therapeutic potential, but their use is limited by both biological and ethical factors (Landry and Zucker, 2004). Several studies have demonstrated that tissues of all germ layers arise from ESC (Zhang et al., 2001; Reubinoff et al., 2000; Levenberg et al., 2002; Kehat et al., 2001; Assady et al., 2001; Chung et al., 2006), and protocols exist to expand these cells ex vivo (Thompson et al., 1998). Therapeutic value is questionable, however, as these cells form embryoid bodies in vitro and teratomas in vivo and may induce an inflammatory immune response. Additionally, use of discarded embryos to isolate hESC is an area of ethical controversy.

Somatic Cell Nuclear Transfer Somatic cell nuclear transfer (SCNT) is the process of transplanting nuclei from adult cells into oocytes or blastocysts and allowing them to grow and differentiate, producing pluripotent cells. Figure 1 illustrates SCNT. This process has both reproductive and

Figure 1

Somatic Cell Nuclear Transfer (SCNT) to produce replacement tissues using a patient’s cells.

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therapeutic implications. Both have the same initial process, removing an adult cell nucleus, placing into an oocyte, and stimulating it to grow with electricity or chemicals. This will produce an embryo genetically identical to the donated cell nucleus: if planted in a uterus, a clone will develop. If the embryo is used for tissue development, it is considered therapeutic. The resultant cells are immunoidentical to the donor and capable of identical growth to a naturally formed embryo (Hochedlinger et al., 2002; Brambink et al., 2006). The interest of SCNT is to employ the capacity to develop nonimmunogenic cells that can develop into any tissue to replace the damaged organ. In SCNT, a viable embryo is not always developed. A new study demonstrates that leaving the oocyte’s genetic material and simply adding the somatic cell genome will produce a triploid blastocyst that can lead to a viable embryo that can develop all three germ layers (Noggle et al., 2012). Several studies to date have demonstrated the capacity to remove cell nuclei from animal model cells, place in blastocysts and expand ex vivo, replace in the organism, and produce anatomic and functional renal units (Lanza et al., 2002; Atala et al., 1995; Yoo et al., 1996). Reproductive SCNT is banned in most countries due to ethical dilemmas, while therapeutic SCNT is an important form of research.

Altered Nuclear Transfer Altered nuclear transfer takes a genetically modified somatic cell nucleus and transplants it into an oocyte. The mutation allows development to a certain point, and then halts progression. This technique allows development of pluripotent stem cells using a blastocyst without the controversy of hESC, has been shown to work in mice, but needs further studies to understand how the mutation affects development (Meissner et al., 2006; Benahmed et al., 2001).

Single Cell Embryo Biopsy hESC research is an area of ethical controversy in that it results in the destruction of embryos; a method of isolating these cells without destroying the embryo would be advantageous. In 2006, Chung and colleagues were the first to report the generation of mouse ESC lines in this manner (Becker and Chung et al., 2006). Their method was based on obtaining a single cell embryo biopsy for preimplantation genetic diagnosis. Cells were taken from eight-cell blastomeres rather than from blastocysts. The cells differentiated into derivatives of all three embryonic germ layers in vitro, as well as into teratomas in vivo, and were considered pluripotent. In addition, the mouse embryos that resulted from the biopsied blastomeres developed to term without a reduction in their developmental potential.

Cells from Arrested Embryos hESC lines can also be derived from arrested embryos from in vitro fertilization (IVF) clinics (Zhang et al., 2006a). Only a few of the zygotes produced in the process of IVF will develop to the morula and blastocyst stages, and the ones that stop dividing are considered dead embryos and are usually discarded (Hardy, 1993; Geber and Sampaio, 1999; Landry et al., 2004). As not all of these cells in the discarded embryos are abnormal (Martinez et al., 2002), this may prove to be a source of hESCs. More studies are needed to characterize the full proliferation and differentiation potential of ESCs derived from arrested embryos.

Amniotic Fluid and Placental Stem Cells A fourth stem cell population is the amniotic fluid and placental-derived stem cell (AFPS). Aspirates taken from amniotic sac yield a heterogeneous population of cells (von Koskull et al., 1981; Medina-Gomez, 1982), mostly fetal in origin. Yet a unique form of stem cell can be isolated from the amniotic fluid and placenta called AFPS. These cells are less differentiated than adult stem cells but more so than ESCs, can form all three germ layers, and notably do not form teratomas in vivo. They also can differentiate to cells of all three germ layers including myogenic, neurallike, adipogenic, hepatic, and osteogenic (De Coppi et al., 2007b), as well as mature cartilage (Kolambkar et al., 2007), kidney (Perin et al., 2007), lung (Warburton, 2008). They can be obtained via chorionic villus sampling or amniocentesis and have the potential to be banked for self or allogenic use. A major advantage of isolating cells from gestational tissue is that it allows for autologous reimplantation, effectively bypassing the possibility of immune rejection. Several techniques have been developed to isolate and maintain AFPS in vitro, allowing a more thorough understanding of this cell population. Remarkably, AFPS is capable of propagating rapidly while keeping normal karyotypes even in late passages (Atala, 2009). Telomere length, and cell cycle check points (Brown et al., 2002) are normal, and expression of stem cell markers SSEA4 and OCT4 are seen (De Coppi et al., 2007b), which indicate an ESC-like cell, but not exactly identical. These cells do form all germ layers in vitro, but not teratomas in vivo, making them a more appealing therapeutic option for regeneration (Thompson et al., 1998). Compared to ESCs, iPSCs, and MSCs, AFPS grow more rapidly and there are established protocols to induce the less immature cells to differentiate into too many cell types (De Coppi et al., 2007b). ESCs are more plastic than MSC/AFPS, but AFPS express OCT4, which is correlated with ESC plasticity (De Coppi et al., 2007b). There are implications for use in bladder muscle (De Coppi et al.,

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2007a), nerve (De Coppi et al., 2007b), kidney (Perin et al., 2010), lung (Carraro et al., 2008), heart (Bollini et al., 2011) and heart valve (Weber et al., 2011), diaphragm (Fuchs et al., 2004), bone (De Coppi et al., 2007b), and cartilage (Kolambkar et al., 2007). Of note, amniotic fluid stem cells (AFSC) have shown to reduce renal damage postischemia, mediated by released antiapoptotic effects, activated Akt, IL-6, and VEGF release, and this effect can be enhanced by preconditioning AFPS with glial cell line-derived neurotrophic factor (GDNF) (Rota et al., 2012).

Induced Pluripotent Stem Cells iPSCs, a recent development, pose an alternative source of pluripotent cells for use in tissue regeneration. Nobel Prize laureate Shinya Yamanaka created iPSC through somatic cell reprogramming (Takahashi and Yamanaka, 2006). These cells maintain the morphology and growth properties of ESC and express ESC marker genes while avoiding the ethical and immunological problems associated with the use of ESC. A variety of cells can be used to produce iPSCs, including keratinocytes, mesenchymal cells in fat, oral mucosa, dental pulp, peripheral blood, and cord blood (Zhou and Ding, 2010). iPSCs, in turn, have successfully been seeded on to scaffolds and differentiated to other cell types (Xie et al., 2011). iPSCs have typically been produced by viral vectors integrating key transcription factors such as Oct3/4, Sox2, Klf4, and c-Myc into somatic cells. This does, however, potentially increase the risk of tumor formation. Alternative methods of iPSC generation, such as plasmid transfection and direct protein delivery, are under study and efficacy is uncertain (Zhou and Ding, 2010). The establishment of safe methods of iPSC generation for clinical applications is an ongoing process and iPSCs may be applicable for medical treatment in the near future. While selection of appropriate cells for bioengineering is crucial, and an area of development, it is only one branch of designing an organ. Appropriate material structure for the cells to grow on is equally significant.

Biomaterials Cell growth and migration is heavily influenced by their microenvironment, including the ECM, soluble factors, and physical stimuli. The ECM is a dynamic environment that provides structure and guidance cues to cells, a barrier between different compartments or cell types, and intercellular communication regulation. It also aids in embryologic development, tissue repair, and wound healing. The main components of ECM are structural proteins, collagens, elastins, and proteoglycans (Aitken and Bagli, 2009). In the urinary bladder, the composition and architecture of native ECM influences development, and provides function and protection against urine and potential pathogens. Replacement of biomaterials thus must perform several roles. They ideally would guide cell growth, maintain visceral structure and correct cellular alignment, facilitate vascularization and innervation, provide resilience to the stresses of the normal bioenvironment, and be degradable with nontoxic degradation products. In the bladder, the material must coordinate both urothelial and detrusor smooth muscle growth to facilitate stretching and relaxing for normal urination. In the kidney, it must facilitate blood filtration and urine draining. Several types of biomaterials have been explored: biologics including alginate and collagen, decellularized tissue matrices such as bladder or gastrointestinal submucosa, and synthetic materials such as polylactic acid, polyglycolic acid, and poly(lactic-co-glycolic acid). The first type discussed is native acellular matrices, which have been isolated and used in both in vitro and in vivo studies. Figure 2 demonstrates several types of biomaterials.

Acellular Matrices Acellular matrices have shown promise in various settings. Matrices are isolated most often from bladder acellular matrix or small intestine submucosa (SIS), and decellularized by chemical and mechanical techniques (Sutherland et al., 1996). Of note, these matrices do maintain their biological properties, and facilitate cell ingrowth and structure through type 1 collagen, glycosaminoglycans, and other bioactive proteins and factors including VEGF, fibronectin, transforming growth factor-b1, and fibroblast growth factor (FGF) (Chun et al., 2007; Yang et al., 2010; Kanai et al., 1999). Acellular matrices have been successful in guiding cell migration and growth in bladder regeneration models, both with and without preseeding with urothelial and smooth muscle cells (Oberpenning et al., 1999; Kropp et al., 1996; Zhang et al., 2006b; Reddy et al., 2000). Rat trials have shown vascularization with acellular matrices (Sievert et al., 2007), but larger pig and dog models need cellularization in order to develop successful blood supply (Zhang et al., 2006b; Brown et al., 2002). Without cell seeding, cellular matrices have shown normal urothelial growth and abnormal smooth muscle growth (Atala, 1996, 1998). Similar findings were shown in a clinical trial conducted on patients with exstrophy–epispadias complex and bladders with poor capacity and compliance using SIS as the matrix for bladder augmentation (Cainoe et al., 2012). Follow-up studies demonstrated integration of the matrix into the bladder, normal urothelial growth, decreased smooth muscle:collagen ratio, and no significant increase in capacity and compliance. Other biomaterials of interest are synthetics.

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Figure 2 Scanning electron microscopic image of biomaterials used in tissue engineering. Upper left, collagen sponge; upper right, acellular matrix from pig bladder submucosa; lower, polyglycolic acid matrix.

Synthetic Biomaterials Synthetic biomaterial polymers (polyglycolic acid, polylactic acid, poly-lacto-co-glycolic acid, and recently silk fibroin) are also an area of development (Pariente et al., 2001). These materials are quickly formed, thermoplastic, and thus malleable to many shapes, and have biodegradation by-products that can be rapidly cleared from the body. Synthetics are formed into sponges or meshes to achieve the appropriate high surface area to volume and porosity that are required for adequate tissue growth. Usually, permanent synthetic materials used for bladder reconstruction succumb to mechanical failure and urinary stone formation, and use of degradable materials leads to fibroblast deposition, scarring, graft contracture, and a reduced reservoir volume over time. A significant clinical trial comparing collagen and a collagen–Polyglycolic acid (PGA) hybrid in patients with myelomeningocele-related neuropathic bladder needing cystoplasty demonstrated that use of the composite led to improved compliance and capacity compared to collagen alone. Optimal outcomes were seen in patients treated with a scaffold seeded with autologous urothelial and smooth muscle cells, anastomosed to existing bladder, and covered with omentum (Atala et al., 2006). Results of this trial are shown in Figure 3. Beyond acellular matrices and synthetic materials, biologics such as collagen and alginate have been focused on as well.

Figure 3 Construction of engineered bladder. Left, cell-seeded bladder scaffold material; center, the seeded scaffold is anastomosed to native bladder; right, implant covered with fibrin glue and omentum.

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Biologic Biomaterials Biologic materials such as collagen and alginate are advantageous in that they are less immunogenic. Collagen, a structural protein found throughout the body, is minimally inflammatory and has already been approved by the United States Food and Drug Administration (USFDA) for wound care (Cen et al., 2008). Collagen can be manipulated with chemicals and physical methods to increase intermolecular cross-linking, which makes it resilient (Li, 1995). Additionally, collagen has been shown to guide cell activity and phenotype via specific interactions (Silver et al., 1992). Collagen and alginate have been used in bioprinting with inject techniques, and are used as ‘ink’ to print organs into designed three-dimensional structures that allow appropriate cell and tissue growth (Boland et al., 2006; Campbell, 2007).

Vascularization and Innervation Designing a bladder that truly replaces the lost tissues requires adequate blood supply and innervation. The bladder derives its blood source primarily from branches of the internal iliac artery, and this needs to be reestablished in the bioengineered tissue, a formidable task, and there is a limited distance to which cells can be located from a blood supply (Folkman et al., 1992). It has been demonstrated that new vessels develop in two manners: vasculogenesis, which is a de novo development of a vessel and capillaries from mesoderm (Drake et al., 1998), while angiogenesis is a two-stage process where capillaries are synthesized from existing vessels. During angiogenesis, under the guidance of factors VEGF and FGF (Battegay, 1995), endothelial cells separate, the basement membrane of the vessel degrades by enzymes, endothelial cells migrate and proliferate to form the edge of a new capillary, a lumen is formed, a new basement membrane is synthesized, and ultimately, pericytes and vascular smooth muscle form (Folkman et al., 1992; Ferrara et al., 2007; Yamanaka et al., 2002). To engage this process, various methods of integrating VEGF and FGF into the biomaterial and cell constructs have shown increased vascularization and reduced graft shrinkage (Gerber et al., 1998; Askari et al., 2004). Innervation of the bladder is as important and perhaps more complicated than its vasculature. A complex arrangement of afferent and efferent nerves with sympathetic and parasympathetic as well as central and peripheral components allows the bladder to work in synergy with the urethra to fill and void at low pressures. To regain lost function, this component must also be present. Neurotrophic growth factors include nerve growth factor (NGF), GDNF, and even VEGF. Excessive levels of NGF can cause hyperinnervation and hypersensitivity, while low levels will not promote adequate branching and axonal growth (Schnegelsberg et al., 2010; Piquilloud et al., 2007). A combination of NGF with VEGF or GDNF integrated into engineered tissues will likely prove to be the optimal way to lead to regenerative bladder innervation (Madduri, 2009; Kikuno et al., 2009; Nitta et al., 2010), and prove that appropriate neuronal growth will rely on biomaterial physical alignment (Corey et al., 2007).

Bioreactors Cell phenotype and tissue formation depends on physical stimuli to influence cellular alignment, cell proliferation, remodeling and production of ECM components. Bioreactors recreate physiologic conditions ex vivo, which provide these stimuli to engineered tissues in the form of forces including compression, shear stresses, pulsatile flow, and electrical stimulation. This process has been proved to be important in the engineered blood vessels (Seliktar et al., 2000), cartilage (Elder and Athanasiou, 2009), and ligaments (Moreau et al., 2008). In bladder bioengineering, simulating normal physiological conditions of bladder cycling (filling and emptying) may provide the additional mechanical properties required to improve functional outcome after implantation (Hubschmid et al., 2005; Wallis et al., 2008; Haberstroh et al., 2002; Farhart and Yeger, 2008). It has been shown that this stretching effect on bladders positively influenced in vitro cell and tissue growth, cellular alignment, and changes in gene expression (e.g., collagens, uroplakin II). Further investigations are needed to decide on the importance of ex vivo bioreactors in bladder engineering. Another approach is the in vivo bioreactor, where bioengineered tissues are temporarily placed into an organism to promote growth and development, and then transplanted into the true recipient. This in vivo preconditioning may further support tissue development, enhance vascularization of the engineered tissue, and prevent fibrosis and loss of contractility. In an experiment from Baumert et al., SIS was seeded with urothelial and smooth muscle cells and placed into a porcine omentum. Three weeks after placement in this in vivo bioreactor, the scaffolds showed a multilayered urothelium and organized outer layers composed of SMCs and fibroblasts, and rich tissue vascularization. The constructs could also be transferred to the site of bladder replacement without compromising the blood supply (Baumert et al., 2007). While promising, this required a two-stage procedure and the potential benefits over current strategies need to be further investigated before they can be applied in the clinic.

Renal-Specific Studies While significant advances have been made in bladder bioengineering, there is also research primarily applicable to the kidney. The kidney is a complex organ with multiple cell types and an integrative functional anatomy that renders it one of the most difficult organ to reconstruct. It is the target of many congenital and acquired diseases including diabetes, hypertension, polycystic kidney

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disease, as well as ischemia and toxic injury, which leave patients with acute or permanent renal damage. The United States Renal Data System reported 594 374 patients with end stage renal disease (ESRD) in 2010, with 116 946 new patients in that same year (USRDS, 2012). While renal transplantation remains the cornerstone in treating ESRD, many will spend years on dialysis. While effective, dialysis is a temporary solution, and regenerative medicine approaches have been heavily studied to address renal bioengineering. Solutions need to address the myriad of functions the kidney performs, including filtering blood, excreting waste, regulating systemic pH and water balance, and hormone production. The approaches discussed are cell based, renal precursor based, factor based, and materials based.

Cell-Based Renal Therapy As discussed above, MSCs have demonstrated the ability to aid in renal recovery after ischemic and toxic injury. Administration of these cells after acute injury has been renoprotective (Togel et al., 2005), but has no effect on fibrotic changes (Stokman et al., 2008).

Renal Precursor Therapy Many advances in renal regeneration are based on embryology. The kidney derives from intermediate mesoderm of the urogenital ridge in the fetus in a three-stage process. Located along the posterior abdominal wall, the first structures to appear are the pronephros. Developing caudal to cranial from the urogenital ridge starting at day 22 of gestation, these structures are composed of pronephric tubules and duct. Caudal to the pronephros, the mesonephros develop and are embryologically functional, but ultimately degenerate with some remnant cells migrating to the developing adrenogonadal primordia, which become adrenal glands and gonads. The final embryologic structures to develop are metanephros, which develop from metanephric mesenchyme and epithelial ureteric bud interactions, leading to a renal vesicle, and progressively nephrons, podocytes, and capillaries of the glomerulus are developed (Maezawa et al., 2012). Thus the kidney is derived from many sites, leading to its intricate form and function. Regenerative approaches keep this complexity in focus to replace these organs. Several groups have worked on transplanting renal precursors, both allogeneic and xenogeneic, and most prominently have found that the immature structures were less susceptible to host rejection, and recruited vasculature from host structures. These structures were found to produce urine and fulfill other endocrine roles of the kidney. Early studies of renal precursors repeatedly demonstrated that transplanting more premature metanephros were less likely to be rejected, were capable of producing urine, and could do so in extrarenal locations. A series of experiments demonstrated that E13– E16 metanephroi could be transplanted into neonatal mice recipients. They were able to show vascularization, maturation, and tubular extension to host medulla in donated tissues. They were also filtering. These results were not reproduced in adult recipient mice (Woolf et al., 1990). Subsequent studies demonstrated function in metanephroi placed outside the renal parenchyma, including subcapsular (Abrahamson et al., 1994) and omental placement (Rogers et al., 1998), which was able to prolong survival in anephric rats by an average of 58 h compared to controls (controls were those not receiving transplant and those with transplant without ureteroureterostomy to host ureter) (Rogers and Hammerman, 2004). Transplanted metanephroi were shown to produce urine as well. Of note, human (7–8 week) and pig (3.5–4 week) precursors can grow, undergo complete nephrogenesis, and produce urine in immunodeficient mice (Dekel et al., 2003). The same group also indicated that vascularization of these structures was of host origin, and that the less immunogenicity observed may be accounted for by this fact. Recently, animal models with blood loss-induced anemia after being transplanted with xenogeneic metanephroi have shown host MSC recruitment and subsequent erythropoietin (EPO) production. Anemic mice received extrarenal metanephros and anemic cat received extrarenal pig metanephros; both recipients demonstrated increased EPO levels of host MSC origin (Matsumoto, 2012). This study demonstrates that metanephros can serve as nonimmunogenic scaffold for host tissue migration and differentiation to EPO-producing cells, but that these cells are multipotent mesenchymal cells, and not pluripotent, as demonstrated by ESC and induced pluripotent cell failure to integrate and produce EPO. Transplantation of embryonic structures is promising, due to the observed decreased immunogenicity, and possible reconstitution of all renal functions, including filtration, urine production, and EPO production. Additional studies in 2012 indicate that metanephros transplant can even increase blood pressure in an induced hypotension model in anephric rats, used to simulate dialysis-associated hypotension (Yokote et al., 2012). Embryonic renal structures may play a key role in recovering renal loss in patients in the future.

Factor-Based Renal Therapy A separate approach to renal regeneration is administration of specific growth factors that have shown positive effects on renal repair. In animal studies, epidermal growth factor and hepatocyte growth factor have improved survival in acute renal injury (Hammerman and Miller, 1994). In human trials, IGF-1 has shown to improve renal function in patients with ESRD (Vijayan et al., 1999). Potential problems with exogenous administration of growth factors are their possible myriad of effects in other

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tissues, difficulty in maintenance of therapeutic levels, and dose-dependent responses. Very specific delivery systems need to be established for growth factor-based therapy.

Materials-Based Renal Therapy Previous efforts in tissue engineering of the kidney have been directed toward the development of extracorporeal renal support systems made of biologic and synthetic components. One approach has been development of an extracorporeal device with a membrane and a single renal cell type to form a monolayer of cells to perform the endocrine and metabolic renal functions, and it could be used with conventional hemodialysis (Aebischer et al., 1987; Ip and Aebischer, 1989; Humes et al., 1999; Tiranathanagul, 2005). Many researchers have used porcine and canine renal epithelial cell lines in this model. Problems arose with sodium transport, multilayer growth, and necrosis (Fujita et al., 2002). Membrane material and coating strongly influenced monolayer adhesion and function (Zhang et al., 2009; Chun et al., 2007; Sato et al., 2005). A similar study looked at human renal epithelial cells grown on a fiber cartridge for uremic dogs. It was shown to excrete ammonia and convert 1,25-dihydroxyvitamin D3, but not adequately transport electrolytes (Humes et al., 2002). Devices described as above have been used in phase I and II clinical trials with some evidence of assistance in recovery in patients with acute kidney injury (Humes et al., 2004; Tumlin et al., 2008). Primary problems using this method lie in exposure of renal cells to nonphysiologic environments. Unique cell culture techniques, however, are being studied to target this dilemma (Minuth et al., 2005; Schumacher et al., 2002; Minuth and Strehl, 2007). Research into implantable renal replacement devices have been conducted, but this poses a formidable challenge due to the foreign body response generated. Experiments in rats with this approach have been done, but are far from clinical trials at this point (Minuth and Strehl, 2007).

Organ Decellularization and Repopulation While use of entirely manufactured renal replacement devices has been attempted as described, others sought to decellularize ex vivo organs and repopulate the scaffold with new cells to generate transplantable viscera. The kidney has complex anatomy that derives from differing germ cell layers, and this approach would ideally preserve vascular and parenchymal anatomy to see if the residual extracellular matrices would provide niches for cell growth and retain the signaling capacity to guide stem cell differentiation. Optimal decellularization techniques would rapidly remove all native cells while maintaining organ structure and matrix integrity, and several groups have recently made advances in using this approach. Bonandrini et al. (2014) showed with scanning electron microscopy and microcomputerized tomography that structures were conserved after decellulizing rat kidneys with sodium dodecyl sulfate; afterward they were able to seed the matrix with murine embryonic stem cells and the cells showed signs of dedifferentiation. Orlando et al. showed that decellularized porcine kidneys could be reimplanted unseeded in live pigs and maintain renal structure after 2 weeks in vivo (Orlando et al., 2012). Ross et al. were able to demonstrate that residual renal matrices after decellularization from rats could guide xenobiologic precursor cells (mouse ESCs) to differentiate into mature, tissueappropriate cells (Ross et al., 2009). Song et al. were able to remove cells from three species (rat, pig, human), reseed them with human umbilical venous endothelial cells, and rat neonatal kidney cells, and create units that filter and produce urine in recipient rats after growth in a perfusion bioreactor (Song et al., 2013). These exciting advancements in both manufactured and repopulated kidneys show promise as a way to provide renal replacement for the current demand for transplant that far exceeds the supply of donors.

Conclusion The field of regenerative medicine has made advances in kidney and bladder bioengineering. A wide array of studies, both in the laboratory and clinic, have developed techniques to replace bladder tissue when augmentation is needed. For the kidney, studies in the role of MSCs have shown promise in repairing acute and chronic injury, as well as in the fields of factor-based, and materialbased renal therapy. The future of bioengineering and regenerative medicine is promising, with urology and nephrology leading much of the new findings and developments.

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Bioengineering Scaffolds for Regenerative Engineering Zichen Qian, Daniel Radke, Wenkai Jia, Mitch Tahtinen, Guifang Wang, and Feng Zhao, Michigan Technological University, Houghton, MI, United States © 2019 Elsevier Inc. All rights reserved.

Introduction Scaffolding Materials Natural Materials Synthetic Materials Hybrid Materials Scaffold Fabrication Technologies Decellularization Mechanical impact of decellularization protocol Biological impact of decellularization protocol Polymeric Scaffold Fabrication Highly porous scaffold fabrication Electrospinning Bioprinting of 3-D Scaffolds Stereolithography Solid free-form fabrication Bioprinted scaffolds Applications of Scaffolds in Tissue Engineering Introduction Spinal Cord/Neuron Repair Wound Healing Bone Regeneration Vascular Scaffolds Summary and Future Directions References Further Reading

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Introduction Tissue engineering has been a rising field since the mid-1980s when the term was used loosely to cover everything from prosthetics to tissue manipulation. Today, the field of tissue engineering has evolved into applying precise engineering principles to clinical problems (Winterswijk and Nout, 2007). Cells, scaffolds, and signals are often called the tissue engineering triad, which are combined through engineering technologies to achieve the replacement or regeneration of damaged tissues (O’Brien, 2011). The purpose of scaffolding is to mimic the structure and function of extracellular matrix (ECM) in native tissues to fully, or partially replace, damaged or diseased tissues. Ideally, scaffolds should be able to execute physiological functions including mechanical support, cell attachment and survival, and nutrient/waste transportation once implanted. They should also disappear upon the completion of tissue regeneration (Minuth et al., 1997; Chan and Leong, 2008; Stock and Vacanti, 2001; Hutmacher, 2001). Fabrication of scaffolds should meet several criteria regardless of target tissue types. The scaffolds need be biocompatible, biodegradable, mechanically compatible, and bioactive. Metals, ceramics, and polymers are the three fundamental biomaterials, in which metals and ceramics are often utilized for the manufacturing of stiff tissues such as bone scaffolds and dental scaffolds (Graney et al.,  2016; Chróscicka et al., 2016; Denry and Holloway, 2014; Kim et al., 2016a; Zhao et al., 2017; Capek et al., 2016; Yu et al., 2017; Wang et al., 2016). For scaffold engineering, polymeric materials have gained popularity since it is easier to produce synthetic polymers with specific properties and shape them into desired architectures. Drawbacks of synthetic polymers include the lack of bioactivity and tissue integration capability, compared with natural proteins, which could result in the rejection of the scaffold. In addition, local pH can be distorted by some acid polymer degradation by-products and result in cell necrosis (O’Brien, 2011). On the other hand, fabrication of natural polymers into scaffolds with desired mechanical properties is challenging. To overcome the disadvantages of synthetic and natural polymers, a combination of the two materials is widely accepted as a strategy to balance scaffold properties (Grayson et al., 2004). Numerous fabrication methods have been developed to tailor “tissue specificity” of scaffolds to fulfill diverse needs. The most straightforward method of fabricating such scaffolds would be decellularizing similar tissues from another source to remove most of the antigens while preserving the mechanical and biochemical properties (Liao et al., 2008; Rieder et al., 2005). Studies are trying to manipulate antigen-removed xenogeneic tissues for human use to widen the tissue donor availability; however, the sources of

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tissues are still quite limited. Fortunately, tissue decellularization is not the only option. Fabrication technologies including salt leaching, gas foaming, phase separation, and electrospinning can generate three-dimensional (3-D) variable architectures that mimic the nature of different tissue microenvironments. Nevertheless, it remains difficult to recreate complex microarchitectures such as the interface between different tissues (Hollister, 2005). With the development of 3-D printing technologies, computeraided bioprinting of specialty tissues with precise geometry has become possible (Nowicki et al., 2016). By careful selection of printing techniques and materials, it is possible to fabricate functional scaffolds with tailored architecture and mechanical properties to replace a variety of tissues. This article will provide a broad overview of scaffolding materials and fabrication techniques. Beyond providing a review of general principles, specific examples of regenerative applications that utilize scaffolds fabricated through different techniques are also examined and discussed.

Scaffolding Materials Material selection is critical for scaffold design due to the bioscaffold material having to be nontoxic and supportive to cell adhesion, proliferation, and migration (Chan and Leong, 2008). The scaffold should also provide proper mechanical support to included cells and match the mechanical properties of native tissues and organs. Based on the material sources and compositions, scaffold materials can be classified into natural materials, synthetic materials, and hybrid materials. Hybrid materials are designed to integrate different classes of materials to balance the scaffold properties to meet various requirements (Grayson et al., 2004).

Natural Materials The natural ECM is a collection of molecules secreted by cells that contains complex components including collagen, laminin, hyaluronic acid (HA), and fibronectin. ECM provides an ideal environment for cells to proliferate and carry out their functions. ECM can mechanically support cells to form an integrated and functional tissue and provide biomolecules to maintain cell motility and viability. Ideally, tissue-engineered scaffolds should play a similar function to natural ECM, which supplies mechanical and molecular support for cell and tissue regeneration. Materials isolated from physiological systems have great potential for scaffold fabrication due to their structural and biochemical similarity to native tissues. These natural materials are biodegradable and can provide biological signaling for cell attachment, proliferation, and differentiation. Proteins derived from ECM are much more popular in the tissue engineering field compared with decellularized ECM. A single ECM component can be extracted from different species, which overcomes the donor limitation. Scaffolds from single component are more reproducible and hold better potential to be manufactured in large scale. Collagen is the most popular protein-based scaffolding material in tissue engineering. Other naturally derived proteins such as fibrin and HA have also been used for engineering different tissues. Collagen is one of the main protein components in ECM, providing support to connective tissues and maintaining tissue integrity. There are 28 types of collagen that have been identified, in which collagen types I, II, III, and V compose the main part of the bone, cartilage, tendon, and skin. Collagen as a scaffold component possesses advantages such as high mechanical strength, biocompatibility, degradability, and broader accessibility. The collagen fibrils in scaffolds carry integrin binding sites, such as integrins a2b1 and a1b2 (Heino, 2000), which are beneficial for cell adhesion and migration (Heino, 2000). The porous structure created in the collagen scaffold allows cells to penetrate and further promote tissue regeneration (Glowacki and Mizuno, 2008). These characteristics make collagen a popular scaffold material in tissue engineering (Cen et al., 2008). For instance, bone and cartilage tissues are supportive tissues that require scaffolds to bear high mechanical loads, which inspired researchers to apply collagen when engineering bone and cartilage tissues. Combining chondrocytes with a collagen scaffold could promote organized cartilage defect healing within 12 months (De Franceschi et al., 2005). Applying a collagen I sponge scaffold as a transient cartilage template showed improved scaffold stiffness and chondrocyte vitality. The chondrocytes deposited collagen I, collagen X, and alkaline phosphatase, indicating the collagen scaffold could be potentially utilized in endochondral bone regeneration (Oliveira et al., 2010). Due to its abundant sources, bioproperties, and mechanical strength, collagen scaffolds have also been widely used in engineering other tissues, such as cardiac (Boland et al., 2004; Boccafoschi et al., 2005; Park et al., 2009), skin (Ma et al., 2003; Gautam et al., 2014), and nerve (Bozkurt et al., 2007) tissues. Fibrin is a nonglobular fibrous protein that contributes to blood clotting. It is also a fibrous network polymer formed by polymerization of fibrinogen in the presence of thrombin, resulting in the porous structure of a fibrin clot that mimics ECM (Barsotti et al., 2011). Fibrinogen can be isolated from blood plasma of the patient, which reduces the risk of immune rejection and disease transmission. Fibrin contains cadherin, integrin, and platelet binding sites (Mosesson, 2005), through which the cell adhesion, migration, and proliferation are promoted (Ehrbar et al., 2005). Fibrin has been widely applied as a carrier for cells and growth factors in tissue engineering applications. With the introduction of thrombin, fibrinogen solution undergoes a cascade of polymerization to form a hydrogel, meanwhile encapsulating cells and absorbing growth factors. Applying fibrin gel as a cell carrier could promote seeding efficiency and increase ECM accumulation in the extracellular space (Ye et al., 2000). In a study that applied fibrin as a myofibroblast carrier for cardiovascular regeneration, the addition of fibrin decreased the loss of collagen and promoted ECM formation, indicating fibrin could promote tissue regeneration by promoting the accumulation of ECM components (Mol et al., 2005). A similar study compared the performance of fibrin with collagen as an arterial media equivalent (Grassl et al., 2002). Results showed that fibrin could promote smooth muscle cells to secrete more than three times the amount of collagen than the collagen gel

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group, resulting in a mechanically stronger arterial media. This study indicates the feasibility of applying fibrin in cardiovascular engineering. Fibrin scaffolds have also been tested for repairing spinal cord injury by carrying the neurotrophic factors, such as neurotrophin-3, which can be incorporated into a fibrin scaffold for sustained release and resulted in significantly enhanced sprouting of neural fibers near the transplantation site. All these studies show that fibrin scaffolds function as excellent cell or growth factor carriers for tissue engineering applications (Johnson et al., 2009). The challenge for applying a fibrin scaffold is that fibrin degrades easily. Adding protease inhibitors to the fibrin scaffold could help alleviate this problem, but the added concentration of protease inhibitors needs to be minimized to avoid interrupting cellular activities (Mol et al., 2005; Shaikh et al., 2008). Beyond natural proteins, polysaccharides are also extensively studied as scaffolding materials. Polysaccharides are long-chain polymeric carbohydrates composed of monosaccharide units, which are bonded together by glycosidic linkages. Polysaccharides can be obtained from abundant and relatively cost-effective sources, including animal, vegetal, and microbial. Polysaccharides are biocompatible and generally have good hemocompatibility. They can be processed into special scaffolds to meet the diverse requirements of different tissue types (Mano et al., 2007). Cellulose, starch, and chitin are representative examples of widely used polysaccharides. However, cellulose requires further modification for use in tissue engineering applications due to its water sensitivity and low cell adhesion on the scaffold surface (Kalia et al., 2011). Also, starch lacks dimensional stability and mechanical strength and is difficult to process, making it not applicable as a pure scaffold (Lu et al., 2009). These limitations confine the applications of cellulose and starch. Chitin can be found in far-ranging sources such as the cuticle of insects, the cell wall of fungi, and the shell of crustacean (Ozdil and Aydin, 2014). Although chitin is insoluble in common solvents, its deacetylated form, chitosan, can be easily dissolved in dilute acids (pH < 6). Moreover, chitosan has been proved to be biologically renewable, biodegradable, biocompatible, and nonantigenic (Di Martino et al., 2005). Due to its structural similarity with various glycosaminoglycans (GAGs) found in articular cartilage, chitosan has been used to modulate chondrocyte morphology and differentiation for cartilage tissue engineering (Suh and Matthew, 2000). Results from in vivo studies showed that chitosan hydrogel could support cartilage matrix accumulation and promote cartilage regeneration (Hoemann et al., 2005; Hao et al., 2010). Moreover, chitosan can promote osteoblast proliferation and mineralized ECM deposition by upregulating associated osteogenesis gene expression (Mathews et al., 2011). Therefore, chitosan has been fabricated into sponge or highly porous composite scaffolds for bone regeneration (Seol et al., 2004; Zhang et al., 2008; Venkatesan and Kim, 2010; Li et al., 2005a). Another naturally derived material that has been extensively used is HA, which is a linear GAG with high molecular weight and biodegradability (Fakhari and Berkland, 2013). HA is a component of most connective, epithelial, and neural tissues. It associates with cell surface receptors. For example, association with integrins modulates cellular behaviors such as cell attachment, motility, and proliferation (Evanko et al., 1999). HA can be massively produced by animal tissue extraction, bacterial synthesis, and enzymatic synthesis from nucleotide sugars (Boeriu et al., 2013). The nonlinear viscoelastic property endues HA the function of lubrication and shock absorption (Zhang and Christopher, 2015). It can also function as space-filling materials and increase the viscoelastic property of bulk materials, which makes it an attractive material in cartilage (Yoo et al., 2005) and bone engineering (Mano et al., 2007). HA-based scaffolds can stimulate matrix deposition when seeded with preadipocytes, indicating that HA scaffolds are able to be applied when generating volume-retaining tissue (Stillaert et al., 2008). Another experiment applied constructed HA scaffolds seeded with preadipocytes as an implant for soft tissue regeneration (Stillaert et al., 2008). The scaffold successfully integrated with native tissue, showing the capability of HA scaffolds in adipose tissue regeneration. The high solubility of HA at room temperature and under acidic pH values is the main obstacle for keeping its structural integrity. Adding cross-linkers could help solve this problem (Allison and Grande-Allen, 2006; Zheng Shu et al., 2004). Besides the several materials introduced earlier, there are many other types of natural material that are widely applied as materials for tissue engineering scaffolds, such as GAGs, alginate, or silk fibroin. Due to the naturally derived nature of these scaffolds, they have reduced immunological reaction and can provide a suitable microenvironment for cell vitality, migration, proliferation, and differentiation. However, the degradation rate of natural scaffolds is difficult to control. Their mechanical strength is usually not strong enough to support certain tissue functions. Optimization of processing and design strategies is required for the application of natural materials, which will be discussed later.

Synthetic Materials Although natural materials provide excellent properties including biocompatibility, biodegradability, and bioactivity, issues such as immunogenicity, reproducibility of fabrication, impaired mechanical properties, and control of degradation rate need to be resolved for practical applications (Lutolf and Hubbell, 2005). Compared with biological materials, synthetic materials are much more controllable and easy to process. Synthetic materials have been widely investigated as tissue engineering scaffolding materials due to their flexibility in physical, chemical, and mechanical properties modulation. They can be modified to have different properties that fulfill the criteria of a variety of implantation sites with decent reproducibility (Bracaglia and Fisher, 2015). Conventional categories of synthetic materials include ceramic, metal, and polymeric materials. Those materials have been extensively studied for tissue engineering applications. Synthetic polymers provide highly tunable mechanical strength and controllable degradation rate, which make them appealing candidates for both soft and stiff tissue engineering scaffolds. Linear aliphatic polyesters, such as poly(lactic acid) (PLA), poly(glycolic acid) (PGA), and their copolymers, poly(lactic acid-co-glycolic acid) (PLGA), are the most frequently used polymers as tissue engineering scaffolds due to their biodegradability and biocompatibility (Armentano et al., 2010). The degradation products are nontoxic components and can be removed by natural metabolism. Besides, they have been approved by US Food and Drug

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Administration (FDA) for clinical use. Another well-accepted aliphatic polyester used in tissue engineering is the poly(3 -caprolactone) (PCL) (Woodruff and Hutmacher, 2010). PCL has a relatively low degradation rate compared with PLA and PLGA, which provides prolonged support for tissues that require a longer regeneration process. In particular, PCL has been proved to carry a degradation rate similar to the rate of new bone formation (Lam et al., 2009), which makes it attractive for bone tissue engineering. Mesenchymal stem cells (MSCs) cultured on PCL supplied with osteogenic factors are observed to form mineralization and collagen I deposition on the scaffold (Yoshimoto et al., 2003). PCL also shows the ability to induce chondrogenesis of MSCs seeded on its scaffold to differentiate toward chondrocytes and enhance secretion of cartilage-specific ECM (Li et al., 2005b), which designates its capability in engineering cartilages. In addition, there are some other synthetic polymers that have been widely applied as tissue engineering scaffolds, such as poly(propylene fumarate) (PPF) and poly(glycerol sebacate) (PGS). PPF is a good candidate for bone engineering scaffolds. It is biodegradable, biocompatible, and osteoconductive and can be reinforced by cross-linking (Shi et al., 2005; Zhu et al., 2010). PGS is biocompatible, biodegradable, and cost-effective. Its flexibility makes it popular for engineering soft tissues (Motlagh et al., 2006; Masoumi et al., 2013; Ravichandran et al., 2012), such as skin and cardiovascular tissues. Broadly used ceramic materials consist of calcium phosphate (CaP) ceramics, CaP cements, and bioactive glasses. These materials have their unique advantages, but are not appearing to be very attractive for engineering most tissue types due to their slow degradation rate. However, they are preferred in constructing bone and dental scaffolds due to their superior load bearing, wear properties, and integration promoting effect on human tissues (Rahaman et al., 2011). Within the variety of ceramics, CaP compounds (LeGeros, 2002) and bioactive glass (Rahaman et al., 2011; Thomas et al., 2005) are exceptional for fabrication of bone engineering scaffolds due to their osteoconductive properties. Hydroxyapatite (HAp), beta-tricalcium phosphate, biphasic CaP, carbonated apatite, and amorphous CaP are the most common CaP ceramics that have been explored in bone tissue engineering (Habraken et al., 2007). Currently, metal scaffolds are mainly applied in treating orthopedic diseases. Typical metal scaffolds contain stainless steel, cobalt-based alloys, titanium-based alloys, and magnesium-based alloys (Yang et al., 2001). Metal scaffolds possess strong mechanical strength and thus can provide robust tissue mechanical support as bone replacement. Moreover, metal scaffolds have unique properties that allow for distinctive scaffold designs. For example, titanium-based alloys are able to be fabricated into shape memory scaffolds, which means that they can regain their original shape after deformation. This shape memory property allows large strain while bearing loading and returning to original position when relaxed. To take advantage of this property, Ni–Ti alloy has been widely used as orthodontic plates and staples for bone fractures (Wang et al., 2016). The controllable degradation property is another benefit of certain metal alloys. Compared with biodegradable polymers, biodegradable metals can provide higher mechanical strength, and they are applicable in supporting tissue architectures (Farraro et al., 2014). However, the degradation by-products of biodegradable metals could potentially cause inflammation. Controlling the degradation rate and by-product removal should be taken into consideration when designing metal-incorporated scaffolds (Lietaert et al., 2013). Some metal scaffolds face mechanical mismatch problems that can damage surrounding tissues and causes foreign body reactions (Jung et al., 2015). To solve these problems, metal scaffolds are fabricated into porous structures to decrease material stiffness, while increasing cell infiltration and tissue regeneration (Yang et al., 2001). Synthetic materials provide better plasticity and stronger mechanical strength than naturally derived materials. Their degradation rate and mechanical strength are more controllable through different synthesis approaches. The reproducibility also makes them promising as tissue engineering scaffolds. However, the synthetic materials often lack biocompatibility and bioactivity, which should be resolved during the scaffold design and fabrication (Seo and Park, 2010).

Hybrid Materials As described earlier, both natural and synthetic materials have their pros and cons that need to be resolved for a good scaffold design. It is usually difficult for one type of material to fulfill the complex requirements for tissue regeneration. Novel design strategies for different combinations of biomaterials with diverse properties to complement the gaps of each other are necessary. The addition of natural materials into synthetic scaffolds would provide molecular signals to the scaffold and increase the biocompatibility, while the addition of synthetic materials into natural scaffolds could significantly increase the mechanical strength (Chan and Leong, 2008). The hybrid of materials can be achieved by blending, coating, or chemical grafting of needed moieties on polymers. For example, a hybrid of PLGA and collagen for cartilage tissue engineering took advantage of PLGA as mechanical supporting skeleton while collagen provided cell binding sites to allow integration (Dai et al., 2010). As the result anticipated, chondrocytes showed even distribution and deposited abundant ECM on the scaffold. In another study, ECM-derived components were combined with PCL that led to favorable biological and mechanical properties for cardiovascular tissue engineering (Heydarkhan-Hagvall et al., 2008). Poly(3-hydroxybutyrate-co-4-hydroxybutyrate) was coated on decellularized porcine aortic heart valves, giving rise to significantly enhanced mechanical strength and well-maintained biological properties (Hong et al., 2009). The coating of fibrin on PLA increased the elasticity of the scaffold, enhanced cell proliferation, cell adhesion, and even cell distribution (Gamboa-Martínez et al., 2011). Collagen-coated co-poly(L-lactide/epsilon-caprolactone) (PLCL) tube was applied to treat urethral defects (Kanatani et al., 2007). After being implanted in an animal model, the tube was functionalized through epithelialization and pericyte wrappings, which benefited from the incorporation of collagen. Polypyrrole (PPy) has applications as neural conduits based on its electric conductivity. The incorporation of HA into PPy improved the biocompatibility of the scaffold while preserving its original conductivity for neural scaffold designs (Cen et al., 2004). Research studied the mechanical and biological properties of collagen–PCL blends that were applied for engineered skin. With a small amount of PCL, the strength of

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acellular collagen increased up to 10% compared with a purely collagen scaffold, while the biocompatibility of the scaffolds was not affected (Powell and Boyce, 2009). The promising next generation of tissue engineering scaffolds should have the integration of desired strength, degradability, and bioactivity. The materials should not only contain growth factors facilitating cell attachment, proliferation, and differentiation but also provide suitable mechanical strength for supporting tissue function. The hybridized materials offer flexibility and controllability for fulfilling these requirements.

Scaffold Fabrication Technologies Different strategies have been explored for creating tissue engineering scaffolds. Decellularization of tissues to generate scaffolds is the most straightforward method to create tissue-specific scaffolds. To successfully obtain a functional scaffold from decellularization, delicate cell removal processes are usually required to preserve the mechanical properties and bioactivities postdecellularization. Nonetheless, the sources of donor tissues are far behind demand (Liao et al., 2008; Gilbert et al., 2006). Various natural and synthetic polymers have been processed with different strategies to recreate the porosity and bioactivity of the target tissues and to widen the selection other than decellularized tissue scaffolds. As aforementioned, polymeric materials provide highly tunable mechanical strength and controllable degradation rate by their chemical composition. Different fabrication techniques, including particulate leaching, gas foaming, phase separation, and electrospinning, have been developed to create pores and microarchitectures that mimic spatial structures of specific tissues (Zhang et al., 2005; Annabi et al., 2009; Zhang et al., 2012; Davidenko et al., 2012; Chew et al., 2008). The fabricated scaffolds must be porous to support nutrient and oxygen transportation and encourage cell penetration, adhesion, and proliferation to promote regeneration. The structure and size of the pores can be roughly controlled by these methods, but recreation of detailed microstructures is not possible. With the development of 3-D printing technologies, more precisely controlled fabrication processes are introduced to manipulate the scaffold microarchitectures through the computerdirected nozzle dispersion (Nowicki et al., 2016).

Decellularization Tissue engineering scaffolds can be fabricated from harvested tissues and organs through the removal of the native cellular components. The remaining ECM provides an ideal environment to accommodate tissue cells. ECM is specialized for the local context of its environment but is composed of the same general components, even across species (Bernard et al., 1983a, 1983b; Constantinou and Jimenez, 1991; Exposito et al., 1992). Therefore, the goal of decellularization is to effectively remove all cellular and nuclear materials, while minimizing the impact on the ECM, to produce a scaffold that will maximize compatibility with native tissues and lessen host immune response. The most common decellularization procedures today typically incorporate a series or combination of treatment methods, but ultimate efficacy depends upon the context of the tissue. Nevertheless, the standard procedure involves treatment with a nonionic detergent to disrupt lipid–lipid and lipid–protein interactions, rupturing the membranes of cells and removing their attachments to the ECM while minimally affecting the protein–protein interactions of the ECM itself. This is followed by a brief enzymatic treatment to disrupt the peptide bonds of genetic materials that would otherwise provoke a host response, usually in combination with mechanical agitation to wash out the resulting cellular debris (Gilbert et al., 2006). This process is illustrated in Fig. 1. It was originally thought that a decellularization protocol with fast lysing and removal of cells would be preferential for reducing the exposure time of the tissue to the detergent and thus minimizing the damage of the detergent on the

Fig. 1 Decellularization of tissues to create scaffolds. The tissue from the donor is treated with detergent to break the lipid bilayer of the cell nucleus. DNA is removed by an extensive wash with DNase treatment if necessary. Detergent is neutralized post decellularization to ensure the biocompatibility of the decellularized scaffolds. Illustration is created based on modification of Servier Medical Art under a creative commons 3.0 unported licenses.

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remaining ECM. In comparing a nonionic detergent-based decellularization protocol (Triton X-100) with faster anionic detergentbased protocol (sodium dodecyl sulfate, SDS, and sodium deoxycholate, SDC) of porcine aortic heart valves and pericardia, it was found that the lessened exposure time of the SDC-based protocol did not result in better preservation of structural ECM components (Roosens et al., 2016). Both protocols led to similar alterations to the ECM, but cell nuclear staining revealed DNA fragments still present in all SDC treated tissues that were not caught with staining of cellular mass. Gel electrophoresis confirmed the presence of these DNA fragments after the SDC protocol and the absence of any such DNA fragments after the Triton X-100 protocol. Despite the preservation of ECM tissue structure under these protocols, porcine heart valves decellularized with the SDC protocol showed a significant reduction in soluble collagen that was conserved in the Triton X-100 protocol. It was also found that the pericardia were less susceptible to the deterioration of soluble collagen following decellularization protocols (Roosens et al., 2016). Triton X-100 remains one of the most popular detergents for decellularization protocols, having proved itself both thorough with cell removal and gentle on ECM. Given that one of the primary advantages of a decellularized ECM scaffold is its mechanical similarity to native tissue, considerable interest was drawn to analyzing how decellularization alters the ECM that is left behind.

Mechanical impact of decellularization protocol The mechanical properties of the scaffold after decellularization depend upon the chemicals and methods used during the decellularization procedure (Liao et al., 2008). The disruption or loss of collagen has been implicated as the major contributor to the alteration of mechanical strength following decellularization, while the disruption of elastin alters mechanical resilience (Liao et al., 2008; Grauss et al., 2005). Ideally, a minimal impact on the ECM structure and mechanics is desired, but this is not always possible. Even the most delicate treatments will alter the mechanical properties slightly due to the absence of the cellular components, though such slight alterations are usually without statistical significance. Generally, it is preferable to sacrifice a small mismatch in mechanics than to risk a host response provoked by lingering genetic materials. The mechanical impact of the decellularization protocol primarily depends on how the protocol alters the collagen and elastin of the ECM, which is in turn dependent upon the context of the tissue to be decellularized. For example, porcine aortic heart valves decellularized with a Triton X-100 based protocol (1% concentration) increased in net extensibility by around 100%, based on circumferential and radial maximum stretch ratios at 60 N/m equibiaxial tension (Liao et al., 2008). This corresponded with a reduction in tissue area and thickness by about one standard deviation of the sample (n ¼ 18) according to morphological analysis using scanning electron microscopy (SEM). Small angle light scattering analysis confirmed the preservation of the collagen fiber structure, but the Triton X-100 treatment did not preserve the regional alignments found in the untreated aortic valves and resulted in a more homogeneous fiber structure. SEM analysis calculated a normalized orientation index and found decellularization caused a decrease in areas that were relatively highly aligned in the native aortic valves, but this decrease was not statistically significant. Decellularization resulted in a major loss of tissue stiffness and a shift from a nearly linear moment–curvature response to a nonlinear relationship, causing a 46%–80% reduction in stiffness. Histological analysis confirmed the absence of any cellular material after decellularization; however, the well-organized collagen crimp structure of the fibrosa was disrupted and the spongiosa was depleted (Liao et al., 2008). Bovine pericardial tissue decellularized with a Triton X-100 treatment (1% concentration) did not exhibit any significant change in thickness or microstructure and did not alter stress-relaxation behavior when treated with Triton X-100 alone (Mendoza-Novelo et al., 2011). The complete decellularization procedure resulted in an increase in maximal stress retention of approximately 20% when compared to untreated tissue and reduced the native tissue modulus without statistical significance. No statistical difference was observed in the ultimate tensile stress between the native and decellularized pericardial tissue. Pellegata et al. observed that porcine aortic and carotid arterial segments ranging from 2 to 11 mm in lumen diameter decellularized by an SDC-based protocol (4% concentration) displayed no significant difference in Young’s modulus, compliance, ultimate circumferential stress, burst pressure, or suture retention strength when compared to native tissues. There was however a significant decrease in ultimate strain and an increase in residual stress after relaxation (Pellegata et al., 2013). HE staining revealed no alterations to elastic laminae structure or organization, but the absence of cellular components between laminae layers was thought to have altered the mechanical interaction between them.

Biological impact of decellularization protocol In addition to altering the structurally derived mechanical properties of the scaffold, the method of decellularization alters both the chemically and physically derived biological interactions with the decellularized ECM scaffold (Rieder et al., 2005). The first heart valves engineered through decellularization of porcine heart valves were disastrous failures. Incomplete decellularization induced a severe inflammatory response that led to the development of a fibrous sheath on the interior and exterior graft surfaces, blocking cell repopulation and endothelialization and ultimately leading to graft failure (Simon et al., 2003). The development of this dense fibrous sheath began as early as 2 days after implantation. Histological analysis of grafts explanted post failure indicated a severe foreign body reaction started at the exterior of the graft with dense neutrophil granulocyte and macrophage reactions present within the fibrous sheath, which infiltrated into the leaflet tissue. Lymphocytes were present after 1 year. Inexplicable calcium deposits were discovered within the wall of the conduit but not the leaflet, and analysis of preimplant samples revealed the presence of similar calcium deposits and dense patches of remnant cells. Evaluation of immunogenicity and thrombogenicity of porcine pulmonary leaflets, decellularized by four different decellularization protocols, revealed significant differences that correlated with alterations to the ECM architecture (Zhou et al., 2010). b-thromboglobulin (b-TG) and thrombin–antithrombin complex (TAT) were used to determine thrombogenicity and revealed no significant differences in and between native and decellularized tissues according to TAT. However, b-TG revealed a significant difference between untreated tissue and those treated with SDC or SDS, which displayed

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similar b-TG levels, and a very significant difference between untreated tissue and those treated with trypsin/EDTA or trypsin/EDTA/ Triton X-100, which displayed similar b-TG levels. Two-photon laser scanning microscopy revealed the SDC protocol left collagen and elastin structures almost completely unaffected, the SDS protocol caused elastin fibers to become compact and microcurled and collagen fibers to become compact and lose structure detail, the trypsin/EDTA protocol revealed much fewer elastin and collagen fibers with the collagen fibers appearing more loose and wavy, and the trypsin/EDTA/Triton X-100 protocol kept elastin fibers intact while the collagen fibers lost their wavelike structure and became misaligned (Zhou et al., 2010). Many studies only test decellularized scaffolds for immediate biocompatibility and successful end-stage regeneration. As such, mechanisms linking altered ECM structure to the development and progression of immunologic response have not been thoroughly investigated, though previous studies have examined ECM structures fabricated from a variety of protocols and noted distinct acute and chronic host response and remodeling (Badylak and Gilbert, 2008). Residual chemicals from decellularization protocols have also recently been implicated in significantly altering immunogenicity, though specific mechanisms have yet to be identified (Poornejad et al., 2016). It is difficult to identify which protocols will have what effect on which tissues. SDC and Triton X-100 remain two of the primary detergents used, SDC requires less treatment time to remove cellular components, but research has shown this does not necessarily translate to better preservation of ECM. On the other hand, Triton X-100 is viewed as the safe alternative, requiring longer exposure times but lessening the impact to the ECM, though again studies have shown this is not always the case and depends on the tissue being decellularized. Literature searches to discover what methods have generated what results with the tissue of interest are key. These existing protocols can then be modified to attain specifically desired results, such as experimentally adjusting concentration or exposure time of a particular treatment for a particular tissue to balance the cell removal, ECM preservation, immunogenicity, and mechanical properties.

Polymeric Scaffold Fabrication Highly porous scaffold fabrication As aforementioned, pore size and porosity are critical for scaffolds to allow mass transportation and cell penetration. Thermally induced phase separation has been developed as a simple and cost-effective fabrication method to generate porous structures in 3-D scaffolds (Fig. 2A). Generally, the polymer solution is frozen at low temperature to achieve liquid–solid phase separation. The polymer-lean phase is then removed by evaporation or extraction, and the polymer-rich phase is left to form a polymer skeleton with porous structures (Zeng et al., 2015; Whang et al., 1995). The cooling temperature and cooling rate are used to control the pore size from several micrometers to around 100 mm (Haugh et al., 2009; Peters et al., 2014). In addition, uniaxial aligned geometry of the pores can also be created by introducing a uniaxial temperature gradient (Zhang et al., 2012; Davidenko et al., 2012). Unfortunately, through this method, homogenous interconnectivity, porosity, and pore size are difficult to be reproduced. The pore-topore variation is impossible to regulate (Matsuba et al., 2003). As an alternative, the porogen leaching method is able to produce more uniform porous structures in scaffolds, which disperse a porogen (usually sodium chloride salt granules) into the water-insoluble polymer through solvent casting followed by porogen leaching in water, leaving a matrix with interconnective porous structure (Fig. 2B) (Tran et al., 2011; Goldstein et al., 2001). Biocompatible solvents are required during the polymer casting process; otherwise, the residue of solvent retained in the scaffolds could cause cytotoxicity. As pore size and structure of the scaffold are dependent on the crystal structure of the porogen, a variety of porogens other than salts have been investigated. Spherical porogens are more favorable since they create well-controlled interconnectivity compared with cubic shapes, due to the contact geometry (Zhang et al., 2005). For example, gelatin microspheres with uniform particle size have been fabricated and adapted as a porogen (chloroform as solvent) for fabricating PCL scaffolds. The generated pore size can be modified by tuning the size of the gelatin microspheres, while porosity can be controlled by the amount of gelatin microsphere added during fabrication. The resulting scaffolds have less creep deformation and better interconnectivity than sodium chloride control (Draghi et al., 2005). Another study used poly(ethylene glycol) (PEG) as a porogen to control the pore properties of a polymer scaffold, such as PCL. The size of the pore was determined by employment of different molecular weight PEG, and the porosity was manipulated through changing the ratio of PEG added into the blended materials (Columbus et al., 2014). Gas foaming has been utilized as another method to create porous scaffolds due to its organic, solvent-free, and inexpensive nature (Fig. 2C). Like salt leaching, a foaming agent is added into the polymer solution to create bubbles, which is then dried to obtain the interconnected porous scaffold. Effervescent salt-based gas foaming is more cost-efficient, but surfactants are usually added to stabilize the foam. Dense gas CO2 foaming is more controllable and surfactant-free, but cosolvents are needed to enhance the solubility of CO2. Through dense gas CO2 foaming, the pore size can be controlled by adjusting the temperature and pressure (Annabi et al., 2010). A sole method introduced previously has its own limitations, and multiple methods are combined to improve the properties of porous scaffolds during the fabrication process. For example, to increase the interconnectivity and also include large pores in the PLLA scaffold fabricated by phase separation, salt particulates (200–250 mm) were added into the polymer solution prior to phase separation. After salt leaching, the microstructure of PLLA was significantly improved, resulting in better nutrient transportation and cell proliferation (Tu et al., 2003). Importantly, the fabrication methods discussed in this section are not highly equipmentdependent and can be processed at low costs. Moreover, the porous structure of the scaffolds is relatively controllable and reproducible that is essential for industrial mass production of scaffolds.

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Fig. 2 Fabrication methods of highly porous polymeric scaffolds. (A) Temperature gradient is introduced to the freezing process of the polymer solution, inducing phase separation. The polymer-lean phase is then extracted with vacuum, resulting in pores after drying. (B) Particulate leaching of polymer/porogen mixture. Water soluble porogen is removed by water leaching, leaving void spaces between insoluble polymer structures. (C) Gas foaming agent is added into polymer solution to create a foamy polymer solution. The foam is then dried to form spongy structures. SEM images were modified from (A) Li, X.-T., Zhang, Y. and Chen, G. Q. (2008) Nanofibrous polyhydroxyalkanoate matrices as cell growth supporting materials. Biomaterials 29(27), 3720–3728. https://doi.org/10.1016/j.biomaterials.2008.06.004, with permission from Elsevier. (B) Chatterjee, K., Hung, S., Kumar, G. and Simon, C. G. (2012). Time-dependent effects of pre-aging 3D polymer scaffolds in cell culture medium on cell proliferation. Journal of Functional Biomaterials 3(2), 372–381. https://doi.org/10.3390/jfb3020372, with permission from MDIP. (C) Martín-de León, J., Bernardo, V. and Rodríguez-Pérez, M. A. (2016). Low density nanocellular polymers based on PMMA produced by gas dissolution foaming: fabrication and cellular structure characterization. Polymers 8(7), 265. doi:10.3390/polym8070265, with permission from MDIP.

Electrospinning Highly porous scaffolds can be created via methods discussed in the earlier section, but those techniques are not sufficient to reproduce the nanofibrous microstructure of ECM in native tissues. Electrospinning has been adapted to fabricate fibrous tissueengineered scaffolds due to the favorable features of fibers such as large surface area, interconnectivity, adequate porosity, and good mechanical stability. Synthetic, natural, and hybrid polymeric materials have been utilized to make fibrous scaffolds. A general setup for electrospinning includes a syringe pump with flow control, a high-voltage power supplier, and a grounded collector. Polymer fibers are drawn from a Taylor cone, formed at the droplet of the flowing polymer, under an electric field, and deposited on the collector to form a fiber layer. The fiber layers can then be folded or glued together to form scaffolds. Various parameters such as voltage, flow rate, and collector distance/structure of electrospinning can be tuned to obtain fibers with diameters from nanoscale to microscale (Liu et al., 2017; Chen et al., 2015; Santos et al., 2013; Paskiabi et al., 2015; Kim and Kim, 2014; Sekiya et al., 2013; Qi et al., 2016). With the standard setting of electrospinning, random and aligned electrospun fibers can be produced (Fig. 3A). Alignment is an attractive feature of scaffolds that facilitate regeneration of highly oriented tissues such as nerve and muscle (Kim et al., 2016b; Aviss et al., 2010). Several strategies including electrostatic and mechanical methods have been used to govern the electrospun fiber alignment (spiral and linear, Fig. 3B and C), by which the collector is either placed in aligned electric field or spinning at a certain rate to regulate alignment of electrospun fibers. Linear alignment of the fibers can be collected by the rotating reel setup, and the fiber size is controllable by the speed of revolution (Chew et al., 2008). Radial alignment can be produced by a novel tethered pyroelectrodynamic spinning system. This method is independent from high-voltage source and nozzle-free, which produces spiral fibrous scaffolds resembling the cochlear morphology (Coppola et al., 2014; Mecozzi et al., 2016). Aligned electrospun fibers (axial/radial) are proved to regulate cell morphology, cell migration, and cellular differentiation both in vitroand in vivo, which can promote tissue regeneration by directing the cell growth with desired orientation (Chew et al., 2008; Kim et al., 2008; Kutikov et al., 2015). For instance, the aligned scaffold could guide the axonal growth where two damaged nerve ends need to be bridged. The neurite extension benefited from the alignment guidance of the scaffolds (Yao et al., 2009).

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Fig. 3 (A) General electrospinning setup with syringe pump flow control, high-voltage power source, grounded fiber collector, and humidity control box. (B–D) Different collector designs with their corresponding collected fiber organization.

3-D fibrous scaffolds mimicking native ECM can be achieved from electrospinning techniques by using a tubular spinning collector or simply wrapping the electrospun scaffold in tubular shape. For example, PLGA–collagen–elastin tubular scaffolds were synthesized for mimicking native arteries with enhanced physical strength and cell ingrowth (Stitzel et al., 2006). By stacking electrospun fibrous scaffolds with different materials and different orientation of the fibers, complex structures can be constructed through multiple different layers in one conduit. As an example, 3-D spiral scaffolds were achieved by simply wrapping an aligned electrospun fibrous scaffold in the direction perpendicular to the fiber alignment (Kutikov et al., 2015). Such morphology mimics the orientation of the cortical bone, which is favorable for fabrication of cortical bone replacement. The space between the fibers inside the created spiral scaffold allows cell penetration and calcium deposition to realize reformation of the bone tissue (Kutikov et al., 2015).

Bioprinting of 3-D Scaffolds Existing industrial and commercial printing technologies have been adapted to a biological medium as a method to reliably fabricate 3-D biological scaffolds. This process has proved to be much more efficient than the tedious growth and manipulation of cell sheets but remains subject to the constraints and dependencies of existing printing technologies. These adapted printing methods can produce either acellular or cellularized scaffolds, though certain methods are unable to accommodate the inclusion of cells by nature of their fabrication process. For example, printers that utilize a sintering laser to melt or cure material may end up inducing cell death through hyperthermia. Rapid prototyping methods follow the same principle as cell-sheet stacking, creating a scaffold from the bottom up by fabricating one layer at a time, and do so with much greater precision and speed (Fig. 4A). Bioprinting methods deposit a biological “ink” following the principle of additive manufacturing to generate a 3-D structure. This “bioink,” as it has come to be known, is a mixture of biological and cellular components suspended within a delivery medium such as gelatin, collagen, HA, alginate, or PEG, though any biocompatible material possessing viscosity suitable for the printing method may be used (Nowicki et al., 2016). These 3-D bioprinting methods primarily differ in their composition of bioink and method of deposition, though a simple extrusion method is beginning to emerge as standard practice. Among these broad categories, a few techniques stand out for their potential in fabricating biological scaffolds.

Stereolithography Stereolithography (SLA) uses a deflected or projected laser to cure and harden exposed areas of a photoreactive polymer at the surface of a reservoir of material. 3-D constructs can be fabricated by curing successive layers of polymer, as illustrated by

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Fig. 4 Demonstration of general 3-D printing workstation with controlled nozzle dispenser (A). (B) Stereolithography is based on photo-initiated polymerization by shining a laser at desired location to polymerize locally to form features inside polymer precursor solution. (C) Solid free-form fabrication (SFF) deposit polymers into desired structures. (D) Combined strategy of SFF with another nozzle depositing bioink protected cells to form cell integrated scaffolds. The printed cellular scaffold is usually immersed in medium to sustain the viability of the cells.

Fig. 4B. This technology has previously been used industrially to create rapid prototypes and functional models, but theoretically, any photopolymer of suitable viscosity can be used as fabrication material. Biodegradable polymers such as diethyl fumarate and PPF, HAp, and even photocurable ceramic suspensions have been used to fabricate tissue engineering scaffolds for bone regeneration (Christenson et al., 2007; Cooke et al., 2003; Langton et al., 1997; Leukers et al., 2005). The limitations of SLA include the geometric fidelity of the fabricated construct, which may be compromised by rehydration (Matsuda and Magoshi, 2002), and its restriction to homogenous materials (Bártolo, 2011). Powder-fusion printing (PFP) is similar to SLA in resolution and principle, though it utilizes a granular material to be bound together through laser sintering (Yang et al., 2002). This broadens the available fabrication material to include metals and plastics, but does not overcome the homogenous material hurdle and severely hinders the ability to include biological materials during fabrication.

Solid free-form fabrication Though SLA and PFP are suitable for the rapid fabrication of tissue engineering scaffolds, they are unsuitable for the fabrication of biologically inclusive materials, requiring that cells and biological components be seeded onto the scaffolds post production. Solid free-form fabrication (SFF) utilizes a precise Cartesian coordinate positioning system similar to those used to guide the sintering lasers of SLA and PFP but directs the position of a nozzle to deposit material as depicted in Fig. 4C (Sears et al., 2016). In the previous industrial applications, this material has been a polymer feedstock forced through a heated nozzle or drawn from a heated reservoir, but this method can also print biologically relevant polymers to produce precision lattice structures (Zein et al., 2002). SFF has been used to print hydrogels that utilized a semi-interpenetrating network of PEG and alginate with silicate nanoplatelets to give the gel zero-shear viscosity above 10 kPa s, granting shape retention after printing and shear-thinning properties that allowed for extrusion (Compton and Lewis, 2014; Hong et al., 2015). Due to the dependency upon material viscosity for the rate of flow and subsequent deposition, the machine must be calibrated specifically to individual materials to allow for proper printing results (Billiet et al., 2014; Fedorovich et al., 2008; Khalil et al., 2005).

Bioprinted scaffolds The most common and affordable method, originating from conversions of commercially available polymer 3-D printers based off of industrial designs, is microextrusion printing (Murphy and Atala, 2014). It is essentially a simplified, scaled down version of SFF, often utilizing only a single deposition nozzle. This provides an advantage over inkjet and laser printing methods in that it can deposit very high cell densities. Laser-assisted printing is capable of much higher resolution; however, it requires rapid gelation kinetics to achieve high shape fidelity, which results in low overall flowrate and relatively long print times (Guillotin and Guillemot, 2011). A subset of extrusion printing known as fused deposition modeling (FDM) allows for the precise fabrication of graded

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microstructures with controlled porosity and layer geometry, owing to its scalability and computer control software (Nowicki et al., 2016). FDM can be expanded upon for even greater control through coaxial extrusion, which simply allows for multiple bioinks or other material reservoirs to be simultaneously deposited through the same nozzle (Jia et al., 2016). This allows for rapid and controlled layering of different materials for the manipulation of micro and bulk structural properties or even integrated blending or encapsulation of materials with distinct properties. As shown in Fig. 4D, this allows cells to be embedded into the scaffold during fabrication. Coaxial extrusion has allowed for the direct 3-D bioprinting of perfusable vascular constructs and endothelialized myocardium (Zhang et al., 2016a). SFF may also be combined with existing scaffold fabrication procedures to aid in the production of specialized devices. Ainola et al. utilized SFF to produce a microfibrous scaffold encapsulated by an electrospun nanofibrous mesh to create a 3-D scaffold for cartilage repair that would confine MSCs and chondrocytes to the damaged site and impede the ingrowth of fibroblasts to prevent fibrous scarring (Ainola et al., 2015). Zhang et al. (2016b) developed an in-house SFF 3D printing method to fabricate a magnetothermal bioceramic scaffold for simultaneous hyperthermal anticancer therapy and bone regeneration. The scaffold could be raised to 42 C through the application of a magnetic field therapy that successfully reduced the viability of local osteosarcoma cells without compromising the regenerative promotion of the scaffold during in vitro testing. Together, those studies demonstrated the potential of 3-D printing in biomanufacturing tissue-engineered scaffolds with high controllability.

Applications of Scaffolds in Tissue Engineering Introduction Applications investigated in this section primarily look to replace the customary practice of autografts. In neural, skin, bone, and cardiovascular tissue autografts are used frequently to repair large injuries. The major downfall of autografts is the requirement for a second surgical site where a donor tissue will be harvested. This donor site is prone to morbidity and infection. Moreover, autografts are not always accessible for all patients such as those who have undergone multiple surgeries or the elderly who do not have viable tissue to use. Thus, engineered tissues are critically needed to replace autografts.

Spinal Cord/Neuron Repair In peripheral neurons, Schwann cells proliferate and align to form bands of Büngner that guide sprouting axonal growths, though the chance of a functional regeneration decreases with the severity of the injury and distance of the resulting gap. Furthermore, injuries to peripheral nerves typically involve scarring of connective tissues, resulting in aberrant sprouts believed to be responsible for spontaneous neuropathic pain symptoms (Deumens et al., 2010). Nerve gaps greater than 2 cm require surgical intervention. The current gold standard of treatment is an autologous nerve graft, but results are limited since the tension generated by suturing nerve segments larger than a few millimeters severely impedes axon regeneration. Limited supply of donor nerves and the requirement of a second surgery site that is prone to infection and morbidity are additional limiting factors (Deumens et al., 2010; Jenkins et al., 2015; Gu et al., 2014). Bioengineered scaffolds offer an alternative intervention able to integrate the multiple mechanisms involved in successful and functional neuron repair. An ideal neural engineered scaffold would incorporate neuroprotective and neuroregenerative drugs (Kabu et al., 2015); biomimetic mechanical properties (Yao et al., 2016); topographical and physical guidance cues (Mobasseri et al., 2015); factors for the recruitment, differentiation, and growth of stem cells to avoid scar formation (Ajioka et al., 2014); angiogenic factors for sustainable growth and effective integration (Álvarez et al., 2014); and functional priming via electric stimulation (Akhavan et al., 2016). Current neural scaffolds focus primarily on providing physical guidance to developing axons through the use of channels or conduits fabricated from polymers or biologically derived materials, often placing emphasis on the surface patterning to promote alignment (Pawar et al., 2015; Archibald et al., 1991; Zhu et al., 2015; Vaysse et al., 2015). More recently, interest has been drawn to the integration of neural scaffolds with neural stem cells (NSCs) (Sharifi et al., 2016). Studies have shown that scaffolds constructed from ECM derived from central nervous system (CNS) tissues demonstrated the ability to differentiate NSCs into neurons while ECM derived from non-CNS tissues did not (Crapo et al., 2013; Duan et al., 2016). Proper differentiation of stem cells, either seeded on the scaffold or recruited endogenously, will speed regeneration and reduce the risk of glial scarring.

Wound Healing Wound healing involves a set of interrelated physiological events including inflammation, proliferation, and remodeling (Stadelmann et al., 1998). However, pathological conditions such as diabetes mellitus can disrupt the wound healing process and result in chronic wounds (Blakytny and Jude, 2006). Tissue-engineered skin grafting scaffolds have been included to treat chronic cutaneous wounds. The optimal scaffold should be nonimmunogenic and biodegradable while incorporating, or being able to recruit, cells and growth factors that can synthesize granulation tissue, allow for reepithelialization, and absorb harmful free radicals that promote inflammation. Furthermore, the scaffold must be able to facilitate the contraction of the wound and reduce scar tissue formation (Zhong et al., 2010). Several skin substitutes have already been approved by FDA to treat diabetic foot ulcers and venous leg ulcers (Hu et al., 2006; Marston et al., 2003; Landsman et al., 2011). These skin substitutes are constructed with matrix scaffolds seeded with fibroblasts

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and/or keratinocytes to resemble the contents of the skin. For example, Demagraft is a cryopreserved fibroblast-derived skin substitute fabricated by seeding fibroblasts on bioabsorbable PGA mesh to constantly provide growth factors and matrix proteins that promote diabetic wound healing (Naughton et al., 1997). Unfortunately, the average time to fabricate these skin substitutes by harvesting the patients’ own cells is long, and the transportation of living cell products is difficult (Zelen et al., 2016). These drawbacks inspired the development of acellular skin grafts from synthetic and natural polymers such as PLGA, collagen, alginate, and chitosan, to avoid the complications of live-cell products (Zong et al., 2003; Singer and Clark, 1999; Mayol et al., 2014; O’Meara et al., 2015). Those scaffolds are helpful in assisting wound closure. Many growth factors play a role in this process, notably vascular endothelial growth factor and basic fibroblast growth factor, which are known to induce cell proliferation and angiogenesis. Embedding these growth factions into scaffolds offers the opportunity to further accelerate the wound healing process in vivo by allowing faster penetration of inflammatory cells and promoting mesenchymal cell recruitment (Breen et al., 2008; Losi et al., 2013; Sun et al., 2014). The incorporation of stem cells directly into scaffolds is another promising method that offers abundant growth factors while reducing scarring. An MSC sheet with built-in vascular networks has been created and was shown to accelerate wound healing, modulate inflammation, and reduce scarring through the therapeutic functions of MSCs and vascularization (Chen et al., 2017). Together, these studies have demonstrated the clinical potential of tissue-engineered scaffolds in promoting wound healing. The in vivo diabetic models have confirmed that those therapeutic scaffolds are able to heal both diabetic and nondiabetic chronic wounds (Dash et al., 2013; Chereddy et al., 2015).

Bone Regeneration The ideal bone tissue engineering scaffold should not only meet the required mechanical properties but also have osteoconductive (stimulate bone cells), osteoinductive (stimulate undifferentiated cells), and osteogenic (stimulate both mechanisms) potentials (Albrektsson and Johansson, 2001; Giannoudis et al., 2011; Khan et al., 2008; Yi et al., 2016; Srouji et al., 2006; El-Ghannam, 2005). Since bone tissue is majorly constructed by collagen I and HAp, a variety of bone tissue-engineered scaffolds have been designed based on the variation of these two materials (Polo-Corrales et al., 2014; Aravamudhan et al., 2013; Schneider et al., 2010; Oh et al., 2011; Niu et al., 2012). While collagen serves as a continuous organic template, CaP, as inorganic constituents, not only strengthen the scaffold but also provide osteoconductivity to the scaffold. Collagen promotes osteoblast adhesion and calcium deposition, which is critical for bone tissue formation and regeneration (Thitiset et al., 2013). Through mineralization and cross-linking, collagen as a scaffold can successfully approach the mechanical strength of cancellous bone that is critical for repairing load-bearing bones (Dhand et al., 2016). HAp or other CaP ceramics are chemically and mechanically compatible with the bones (Kalita et al., 2007; Lc, 2009; Panseri et al., 2012). To promote osteoinductivity, highly porous CaP ceramics have been fabricated to encourage cell integration. Meanwhile, the CaP materials can release Ca2 þ ions to create a suitable microenvironment to promote osteogenic differentiation and mineral deposition (Rahaman et al., 2011). Researchers have investigated the effects of incorporating different cell types and growth factors into the scaffolds to stimulate osteogenesis. For example, incorporation of osteoblasts differentiated from MSCs can enhance the mechanical properties of engineered bone-like tissue (Naito et al., 2011). Beside osteogenic properties, a good bone scaffold should promote angiogenesis to form and maintain a healthy bone during regeneration (He et al., 2013). In summary, to mimic the bone tissue microenvironment and promote regeneration, mechanically qualified scaffolds with incorporated growth factors are needed to promote both osteogenesis and angiogenesis. Most scaffolds fulfilling the criteria are still at proof-of-concept stage, and in vivo testing is on demand in order to apply them clinically.

Vascular Scaffolds Cardiovascular disease remains the leading cause of death worldwide (Pashneh-Tala et al., 2015). Modern treatments such as stents and angioplasty only provide temporary relief to occluded blood vessels (Alfonso et al., 2003; Moussavian et al., 2001). Autografts may suffer from a mismatch of size, vessel structure, or mechanics that limit their suitability as a replacement and carry the complications of donor-site morbidity associated with a second surgery. In some cases, such as in systemic vascular diseases like atherosclerosis, autografts may already be just as unsuitable as the vessel they are intended to replace. Synthetic vascular grafts have shown to be at least as effective as autografts in clinical studies; however, they must be paired with anticoagulants, such as a surface conjugation of heparin, to avoid thrombus and occlusion (Devine et al., 2004; Begovac et al., 2003; Pulli et al., 2010; Gutowska et al., 1995). This makes synthetic materials unsuitable for replacing small-diameter blood vessels, such as those used to bypass coronary blockages, where even a small degree of thrombus or hyperplasia can completely occlude the lumen (Seifu et al., 2013). For this reason, research has shifted toward facilitating integration of the vascular graft with the native tissues of the host. The components of the ECM are generally conserved across species, and so, blood vessels harvested from animal sources can be decellularized to remove the immune-response-triggering cells and genetic material, leaving a scaffold that very closely mimics the native host tissue and provides an ideal environment for vascular reconstruction (Gilbert et al., 2006). However, decellularization procedures are still being refined, and the effects on bulk tissues have not been rigorously investigated, but it has been shown that different tissues respond differently to the various steps of any given decellularization protocol (Badylak and Gilbert, 2008). Each decellularization procedure must be custom-tailored to the tissue at hand. A procedure too intense may alter the ECM structure or organization resulting in a change to its natural function via alteration of the physical or chemical interaction with cells, or it may

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leave behind traces of cytotoxic detergents that impede cellular regrowth and host integration. If the procedure is too weak, it may leave behind foreign cells or genetic materials capable of provoking a host response that could in turn potentially compromise the entire graft. Decellularization procedures can also be used to extract ECM from cultured cells, to establish a more reproducible tissue source, and have demonstrated the ability to integrate with the native host tissue and even remodel and grow with native blood vessels in animal studies (Syedain et al., 2016). However, decellularized ECM is only a scaffold and still requires an appropriate blood-contacting surface that can avoid formation of thrombus but still promote native cell ingrowth. ECM components like elastin have shown to be effective blood-contacting surfaces for maintaining hemocompatibility and regulating cellular regrowth but are structurally weak on their own (Yokoyama et al., 2017). These blood-contacting surfaces may be reinforced with synthetic materials for mechanical strength (McCarthy et al., 2015) or fabricated with controlled architecture to impart more durable mechanics (Xing et al., 2017). The ultimate goal, however, is to replace the diseased blood vessel with as close an approximation to its natural state as possible. For this, biodegradable and naturally derived materials show the best potential (Wu et al., 2012).

Summary and Future Directions The development of tissue-engineered scaffolds has been dramatically accelerated in the last decade. Novel technologies now allow the potential to regenerate most tissues. In this article, new materials were reviewed for creating special scaffolds with unique applications. More prospective composites are being investigated to combine both natural and synthetic materials to achieve a balance between mechanical properties and bioactivities. Emerging technologies such as electrospinning, particulate leaching/phase separation, tissue decellularization, and 3-D printing have been explained in detail. It is generally accepted that the scaffolds need to be vascularized, or otherwise be able to promote vascularization, to further foster tissue regeneration. While more functional scaffolds are being created, a lot of studies have shifted toward developing vascularization strategies for engineered scaffolds. Scaffold vascularization could be realized by coculturing endothelial cells and mural cells to form vasculatures inside the scaffold prior to implantation. Other vascularization strategies including creating microchannels, embedding angiogenic growth factors, and assembling endothelialized tissue constructs are being extensively studied to improve the survival and functions of implanted constructs (Thein-Han and Xu, 2013; Chen et al., 2014; Baldwin et al., 2014; Hasan et al., 2014; Kolesky et al., 2014; Laschke and Menger, 2016). 3-D printing of vascular channels inside scaffolds has been proved to support cell metabolic functions within the constructs (Miller et al., 2012). With the development of 3-D bioprinting technology, creating large scaffolds for whole-organ replacement may become possible. However, the challenges of printing an organ-size scaffold that resembles the complex architecture, mimics the mechanical properties, and restores the function of target organ remain unsolved. To overcome the challenges, highly computer-aided printing system with microsize printing nozzles have been used to recreate the anatomical architecture and mechanical strength of the femurs, branched coronary arteries, trabeculated embryonic hearts, and human brains (Hinton et al., 2015). Another challenge of bioprinting large scaffolds is to sustain the already-printed structures while the whole construct is printing. A sacrificial polymer framework can be printed during the process of scaffold printing. It can then be removed after the printed scaffold is reinforced (Kang et al., 2016). Last but not the least, large organ-size scaffolds have the problem of reseeding cells into the core of the scaffold. Bioreactor system with perfusion functions can help aid cell penetration, but incorporating cells during the printing process might be a better solution, which strongly relies on a superior design of bioink to sustain and protect the cells while printing (Adam et al., 2016; Ott et al., 2008; Hinton et al., 2015). The next generation of tissue-engineered scaffolds of organ size will be based on the 3-D bioprinting technology that enables the clinical possibility of using biomimetic synthetic organs to fulfill the transplantation demands.

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Zhu, X., et al. (2010). Effects of composite formulation on mechanical properties of biodegradable poly (propylene fumarate)/bone fiber scaffolds. International Journal of Polymer Science, 2010. Zhu, W., et al. (2015). Highly aligned nanocomposite scaffolds by electrospinning and electrospraying for neural tissue regeneration. Nanomedicine: Nanotechnology, Biology and Medicine, 11(3), 693–704. Zong, X., et al. (2003). Structure and morphology changes during in vitro degradation of electrospun poly(glycolide-co-lactide) nanofiber membrane. Biomacromolecules, 4(2), 416–423.

Further Reading Chatterjee, K., et al. (2012). Time-dependent effects of pre-aging 3D polymer scaffolds in cell culture medium on cell proliferation. Journal of Functional Biomaterials, 3(2), 372–381. Li, X.-T., Zhang, Y., & Chen, G.-Q. (2008). Nanofibrous polyhydroxyalkanoate matrices as cell growth supporting materials. Biomaterials, 29(27), 3720–3728. Martín-de León, J., Bernardo, V., & Rodríguez-Pérez, M.Á. (2016). Low density Nanocellular polymers based on PMMA produced by gas dissolution foaming: fabrication and cellular structure characterization. Polymer, 8(7), 265.

Biomaterials for Tissue Engineering and Regenerative Medicine Ohan S Manoukian and Naseem Sardashti, University of Connecticut Health, Farmington, CT, United States; and University of Connecticut, Storrs, CT, United States Teagen Stedman, University of Connecticut Health, Farmington, CT, United States Katie Gailiunas, Anurag Ojha, Aura Penalosa, Christopher Mancuso, and Michelle Hobert, University of Connecticut, Storrs, CT, United States Sangamesh G Kumbar, University of Connecticut Health, Farmington, CT, United States; and University of Connecticut, Storrs, CT, United States © 2019 Elsevier Inc. All rights reserved.

Introduction Natural Biomaterials Collagen Alginate Chitosan Silk Cellulose Bacterial Cellulose Fibrin Gelatin Synthetic Biomaterials Polyglycolic Acid Polylactic Acid Poly Lactic-co-Glycolic Acid Polycaprolactone Polyetheretherketone Polyethylene Glycol Polymethyl Methacrylate Conclusion Acknowledgments References Further Reading

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Introduction Biomaterials are natural or synthetic materials intended to interface with biological systems without causing harm to the body. Biomaterials are utilized in many applications that intend to evaluate, treat, or replace any tissue or organ in the body. Depending on the intended application, materials are selected from metals, polymers, ceramics, and composites that are biocompatible. The design must have the ability to mimic properties of tissues and perform the intended biological function. Biomedical engineering is a multidisciplinary field that combines aspects such as biomaterials, bioinformatics, biomechanics, instrumentation, biology, computer science, and medicine to solve medical problems. Biomaterials are essential to biomedical engineering because optimal integration between the body and biomedical device is essential for a successful design. If the immune system rejects the apparatus, it will fail. Regardless of the structural or functional design of the biomedical device, if the material releases toxins or is unstable in the body, it will not be a successful solution. Recent advancements in materials science, such as nanotechnology, have allowed subfields of biomedical engineering including tissue engineering and drug delivery to improve healthcare. There are many applications and scenarios where these biomaterials can be utilized to assist the recovery process for patients. These applications cover a broad range and a single biomaterial is usually not limited to one application. Furthermore, creating composites may enhance the desired properties and minimize the weaknesses of the combined materials. Given the large spectrum of regenerative medicine applications, biomaterials range from full-size polyetheretherketone (PEEK) cranial implants down to the nanostructures such as nanofibers for fabricating scaffolds. Bone screws, vertebral body replacements, micro-scale scaffolds, nanoscale scaffolds, and drug delivery systems are among the most popular applications for biomaterials. New materials and applications are constantly being investigated and researched in this dynamic and rapidly evolving field. This article is meant to serve as an introduction to Biomaterials for Regenerative Medicine, highlighting the properties and applications of various natural and synthetic biomaterials commonly used in tissue engineering and regenerative medicine.

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Natural Biomaterials Natural biomaterials have existed for centuries, and are presently used in a variety of biomedical applications. These natural biomaterials have versatile capabilities due to their optimal biocompatibility, biodegradability, and remodeling abilities. This allows for their application in the repair or replacement of damaged tissues and organs within the body. Additionally, natural biomaterials have the ability to support cell migration, proliferation, differentiation, and adhesion. These characteristics are vital for tissue engineering as the naturally derived biomaterials will promote the attachment and migration of the cells from the surrounding environment, and in turn encourage tissue regeneration (Ha et al., 2013; Ige et al., 2012). Although natural biomaterials have numerous advantages, they still pose potential problems. A limitation of natural biomaterials is the immunogenic response that can result following implantation. Another disadvantage, especially for natural polymers, is the tendency to decompose at temperatures below their melting point. This limits their use in implants as the natural biomaterials lack the versatility to fabricate a range of shapes and sizes. Due to the unpredictability of an in vivo source such as an animal, lot-to-lot variability is an additional concern with these biomaterials (Ige et al., 2012; Bartis and Pongrácz, 2011). Naturally derived biomaterials can be organized into several groups such as protein-based biomaterials, polysaccharide-based biomaterials, and decellularized tissue-derived biomaterials. Protein-based biomaterials include collagen, gelatin, and silk. In addition, polysaccharide-based biomaterials are comprised of cellulose and chitosan. Finally, examples of decellularized tissue-derived biomaterials include decellularized blood vessels and liver (Ha et al., 2013). Several natural biomaterials will be further discussed in the following sections.

Collagen Collagen is the most abundant, natural polymer found in the body, and it accounts for a third of the body’s proteins. This is caused by the abundance of three amino acids: glycine, proline, and hydroxyproline. These amino acids provide a basis for collagen’s triple helical structure to create the macroscopic fibers seen in the extracellular matrix (ECM) of tissue, bone, etc. (see Fig. 1). Collagen’s high resistance to tensile forces maintains and stabilizes structures within tissues (Jenkins et al., 2003). As the most governing protein in the extracellular matrix, collagen accounts for nearly 75% of the dry weight of skin (Mizuno et al., 2003). Collagen’s mechanical properties, biocompatibility, and compatibility with other polymers makes it an ideal material for tissue regeneration (see Fig. 2). As previously stated, collagen is highly biocompatible. Collagen’s ability to conduct electricity makes the protein ideal for many tissue engineering purposes. The more conductive a polymer is, the better cell attachment and cell proliferation on the scaffold (Jenkins et al., 2003). Collagen’s conductivity, approximately 0.3 Siemens/m, can be enhanced by combining the material with other polymers, and thus improve its tissue engineering functionality (MacDonald et al., 2008). The degradation of collagen can be attributed to collagenase, which is an enzyme naturally found in the human body. Collagenase is responsible for the degradation of collagen molecules, which generally occurs in a short period of time. The degradation time is contingent upon the molecular weight of the collagen molecules, which vary throughout the body.

The length scales of collagen type I Nano

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Fig. 1 The hierarchical structure of collagen type I leads to specific biological functions and characteristics across length scales. Reprinted with permission from Walters, B. D. and Stegemann, J. P. (2014) Strategies for directing the structure and function of three-dimensional collagen biomaterials across length scales. Acta Biomaterialia 10 (4): 1488–1501.

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Fig. 2 Scanning electron microscopy (SEM) images of electrospun collagen-based fibers. (A, B) Mesh like and aligned collagen fibers, respectively. (C) A 50–50 collagen–chitosan fiber mesh. Reprinted with permission from Walters, B. D. and Stegemann, J. P. (2014) Strategies for directing the structure and function of three-dimensional collagen biomaterials across length scales. Acta Biomaterialia 10 (4): 1488–1501.

Collagen has been very useful in numerous applications and research studies. In a study conducted by Cho et al., the use of pheochromocytoma (PC12) cells showed that collagen helps nerve growth with its ability to conduct electricity through carbon nanotubes (Cho and Borgens, 2010). The study shows more improvement in cell viability for scaffolds electrically stimulated than the control scaffolds that were not electrically stimulated. The neurite length in electrically stimulated scaffolds was longer and had a higher density of developed neural cells when compared with scaffolds of the control. Studies such as these indicate that collagen can be used as a valuable platform for tissue regeneration.

Alginate Alginate is a natural occurring polymer found in brown seaweed. It is a commonly used biomaterial known for its ability to be made into a gel and retain a structure similar to the ECM. It is commonly used for wound healing, drug delivery, and tissue engineering (Manoukian et al., 2016a). Alginate has great biocompatibility and produces little to no immunogenic response. The studies that did see an immunogenic response were likely caused by impurities present in the polymer. Since alginate is retrieved from a natural source, it acquires many impurities from the environment. Studies that used highly purified alginate saw no immunogenic response, and thus only purified alginate should be introduced to the body. The chemical structure of alginate can be seen in Fig. 3 (Lee and Mooney, 2012). This naturally occurring polymer is not degradable in mammals due to the lack of the enzyme alginase, which degrades alginate. However, one solution to this problem is to make alginate gels that are ionically cross-linked since then they will dissolve. An alternative approach would be to partially oxidize the alginate chains. Gels can also be made out of G-blocks or gene fragments that are retrieved from alginate. The G-blocks are oxidized and cross-linked with adipic acid dihydrazide (AAD) to form gels. This allows for degradable soft gels that dissolve much slower than most gels. The more AAD in the gel, the slower the degradation (Lee and Mooney, 2012). Due to its generally slow degradation, alginate is most suited for applications that will require longer periods of time. For example, alginate has been used as a biomaterial in myocardial repair. In the study, an alginate hydrogel was used to deliver stem cells to an infarcted rat heart. This method of delivery saw much better results regarding scar size and microvascular density than methods that directly injected the cells.

Fig. 3 Alginate structure and the egg-box model of hydrogel formation. Reproduced from Lee, K.Y., Mooney, D.J. (2012). Alginate: Properties and biomedical applications. Progress in Polymer Science 371, 106–126.

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When designing any cell culture scaffold for tissue regeneration, it is important to mimic the meshwork of the native ECM. The ECM is an elaborate cellular environment that is responsible for a variety of cell functions such as cell-to-cell communication, and for providing structural support. Synthesizing a scaffold similar to the native ECM will ensure optimal interaction with the native cells as well as the necessary structural support (Geckil et al., 2010). Alginate’s structure is similar to the ECM of cardiac tissue, which makes it a desirable biomaterial for cardiac tissue engineering. Future research intends to design alginate scaffolds that mimic the native ECM and inspire native cardiac tissue functions. A schematic design of engineering cardiac tissue can be seen in Fig. 4 (Ruvinov and Cohen, 2016). Alginate has also been used with hydroxyapatite to produce a scaffold intended for bone regeneration. The alginate was used in combination with the hydroxyapatite to create porous microspheres. A porous scaffold allows for greater integration of cells and a more stable bone structure as materials begin to degrade (Rossi et al., 2012). In terms of drug delivery, alginate is commonly used due to its gelling ability. Hydrogels are readily dissolvable, allowing them to release drugs in a predictable manner. Alginate was used as a neuro-bridge to treat spinal cord injuries by enriching the polymer with two growth factors known to enhance spinal cord repair. The alginate scaffold released these growth factors and was seen to increase the number of surviving neurons in spinal cord injuries (Grulova et al., 2015).

Chitosan Chitosan is the principal derivate from the polysaccharide polymer chitin, which can be found in the skeleton and internal structure of invertebrates, such as in the exoskeleton of shellfish. It is the second most popular natural polysaccharide after cellulose, but their structure differs by replacing hydroxyl for an acetamido group in the C-2 position (see Fig. 5). Chitosan is obtained by partial deacetylation of chitin, that is, the removal of proteins and the dissolution of calcium carbonate (Dutta et al., 2004). The degree of deacetylation of chitin can vary in order to modify its solubility and solution properties. To obtain chitosan, the typical degree of acetylation is less than 0.35, which contributes to its nontoxic characteristics (Kumar, 2000). In addition to its nontoxic properties, chitosan is a naturally abundant and renewable biomaterial with important properties for regenerative medicine such as biocompatibility and biodegradation. It has good biological performance and it degrades into products that are easily absorbed. Physically, chitosan has poor mechanical properties, which is why it is often used as part of a composite (Dutta et al., 2004). Chitosan is a biodegradable polymer with a semicrystalline structure. Porous chitosan scaffolds have been used for implantation, and the results showed that a connective tissue matrix formed within the implant’s pore spaces. Additionally, there was

Fig. 4 Three-dimensional engineered nanowired cardiac tissue. (A) Isolated cardiomyocytes that were cultured in either alginate or alginatenanowired composites. (B) The cardiomyocytes in the alginate scaffolds (top) only form small clusters with nonsynchronous behavior while the alginate-nanowired scaffold exhibit synchronization throughout the scaffold. Reproduced from Ruvinov, E., Cohen, S. (2016). Alginate biomaterial for the treatment of myocardial infarction: Progress, translational strategies, and clinical outlook: From ocean algae to patient bedside. Advanced Drug Delivery Reviews 96, 54–76.

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CH2OH

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Fig. 5 Chemical structure of chitosan. Reprinted with permission from Dutta, P. K., Duta, J. D., Tripathi, V. S. (2004). Chitin and chitosan: Chemistry, properties and applications. Journal of Scientific & Industrial Research 63, 20–31.

very low incidence of chitosan-specific reaction, including cellular and serological reactions (VandeVord et al., 2002). Several studies showed that chitosan’s degradation was related to the degree of acetylation and the polymer’s molecular mass. In vertebrates, lysozyme and bacterial enzymes typically degrade this biomaterial. In vivo studies have also reported ways in which chitosan degrades with different administration methods. In oral administration, degradation occurs mainly in the gastrointestinal track, while in intravenous administration, it is likely to degrade in the liver and kidney (Kean and Thanou, 2010). In several studies, chitosan has shown promising results as a scaffolding material for tissue engineering. It is easily processed into porous scaffolds, films, and beads. Chitosan’s applications include bone tissue, central nervous system, and articular cartilage. Depending on the application, composites are made to improve and complement the biomaterial properties. For example, in bone tissue engineering, a microporous chitosan/calcium phosphate composite scaffold showed enhanced osteoblast attachment that increased scaffold strength while maintaining biocompatibility. Moreover, a chitosan-glycosaminoglycan (GAG) composite has been successfully used to aid articular cartilage repair (Dutta et al., 2004). In terms of wound healing, the deacetylation of chitin into chitosan showed significant improvement in its interaction with mammalian tissues. The cells involved include osteoblasts, fibroblasts, macrophages, and keratinocytes, which aid the tissue regeneration and repair process (Madihally and Matthew, 1999). Moreover, chitin and its derivatives have also shown to increase extracellular lysozyme activity, inhibit fibroplasia, and promote tissue growth for better healing (Kumar, 2000). In some circumstances, wound healing incorporates the use of artificial skin. Chitosan has been utilized in two ways to address this injury. First, chitosan is used for skin replacement since it has characteristics similar to those of GAGs, which is an essential component in skin. Second, it is used in solution to aid healing and fibroplasia, which is the formation of fibrous tissue in wounds (Kumar, 2000). Overall, there are several studies in which chitosan showed excellent potential as tunable porous scaffolds for tissue engineering purposes. An example of this is shown in Fig. 6.

Silk Silk is a natural fiber of fibroin protein produced by arthropods such as silkworms, spiders, flies, mites, or scorpions. Silkworm silk is used in biomedical applications due to its mechanical properties, biocompatibility, and ease of production. Silk is purified easily, using enzymes or alkali to remove sericin, leaving fibroin behind (see Fig. 7). The molecular weight and dispersity index will vary depending on the source and methods of processing (Zhang et al., 2009). Silk is promising for tissue engineering because of its elasticity, strength, and biocompatibility (Kundu et al., 2013). The silk most commonly used in biomaterials is known as mulberry silk (Kundu et al., 2013). It is produced by the silkworm Bombyx mori. Mulberry silk has been used as surgical sutures for many years, and recently the biomedical applications have greatly increased (Yang et al., 2007). Natural fibers can be twisted into rope, braided, or woven. They can also be fabricated into various forms such as hydrogels, lyophilized powders, porous scaffolds, native silk mats, and silk microparticles (see Fig. 8). They can be dissolved in certain salt solutions and then regenerated into other forms such as films, electrospun fibers, porous scaffolds, or hydrogels (Kundu et al., 2013). Silk fibroin is water soluble, which allows for water-based processing in mild conditions without harsh chemicals (Kundu et al., 2013). Topographic patterns can be added to silk film surfaces using lithography. These patterns encourage alignment and adhesion of fibroblasts and epithelial cells (Lawrence et al., 2009). Natural silk is tough and ductile, with a high strength-to-density ratio. Due to its toughness, strength, and strain-hardening tendency, silk is a promising material for load bearing scaffolds in tissue engineering (Kundu et al., 2013). Silk is found naturally in the form of fibers consisting of a core protein, silk fibroin, coated with the protein sericin. The glue-like sericin causes unwelcome immunological responses in the human body, so natural silk is often purified to remove the sericin.

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Fig. 6 SEM images of chitosan scaffold cross sections, which were fabricated by the freeze-gelation method. This process was done under three different conditions: (A) at a freezing temperature of 80 C, a 0.2 concentration of acetic acid in the chitosan solution, and a 95% concentration of ethanol in the rinse buffer, (B) has nearly identical conditions except it had a 0% concentration of ethanol in the rinse buffer, (C) had similar conditions to (A) as well aside from the difference of a 0.8 concentration of acetic acid in the chitosan solution, and (D) differs from the original conditions shown in (A) with a 20 C freezing temperature. Reproduced from Hsieh, C. Y., et al. (2007). Analysis of freeze-gelation and cross-linking processes for preparing porous chitosan scaffolds. Carbohydrate Polymers 67 (1): 124–132.

Fig. 7

Chemical structure of silk fibroin.

Purified silk fibroin demonstrates mild foreign body responses (Yang et al., 2007). Antigenicity and immunogenicity of silk has been tested in rats, pigs, and humans, resulting in only low levels of inflammation and a minimal immune response. Silk sutures are removed after a period of time, but a long-term immune response has yet to be investigated (Kundu et al., 2013). Silk has a tailorable degradation, which ranges from months to years. The degradation rate depends on the processing during the formation of the material (Zhang et al., 2009). Purified silk fibroin fibers begin degrading 6 months after implantation, losing most tensile strength at the end of 1 year. At end of 2 years, they no longer exist at the site (Yang et al., 2007). Processed silk degrades faster than unpurified silk fibers. The degradation rate depends on the secondary structure of the silk fiber or scaffold, which results during processing. Moreover, macrophage-led biodegradation indicates bioresorbability of silk. During in vitro experiments, osteoblasts and osteoclasts have also been shown to degrade silk (Kundu et al., 2013). Silk has been applied in a variety of disciplines such as vascular grafts, bone, cartilage, ligament, skin, and many others. In vascular grafts, electrospun nanofibers have shown the potential to serve as blood-vessel transplants since they have the ability for cell attachment and the necessary mechanical properties for blood flow (Zhang et al., 2009). In bone tissue engineering, electrospun silk matrices support marrow stromal cell attachment and extracellular matrix formation. When growth factors are added to an electrospun scaffold, bone formation is improved (Zhang et al., 2009). Porous silk fibroin scaffolds (see Fig. 9) have the toughness

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Fig. 8 Biomaterials fabricated from silk fibroin: (A) hydrogels; (B) lyophilized powder; (C) 3D porous scaffolds; (D) native silk mat; (E) silk microparticles. Reproduced from Kundu, B., Raijkhowa, R., Kundu, S. C., Wang, X. (2013). Silk fibroin biomaterials for tissue regenerations. Advanced Drug Delivery Reviews 654, 457–470.

Fig. 9 Applications of silk. Reprinted by permission from Macmillan Publishers Ltd.: Nature Protocols, advance online publication, 22 September 2011. Alam, J., Rahman, W., Mazid, R. A. and Khan, M. R. (2015). Gamma-irradiated gelatin-based films modified by HEMA for medical application. International Journal of Polymer Analysis and Characterization 20(5), 426-434. https://doi.org/10.1038/nprot.2011.379.

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and biocompatibility ideal for bone regeneration. Silk hydrogels also work well, and incubating the scaffolds with cells before implantation increases vascularization (Kundu et al., 2013). When electrospun silk matrices are treated with microwave-induced argon plasma, they show promise as a cartilage repair material. Cartilage tissue, including the menisci of the knees, undergo very slow and frequently inadequate and insufficient natural repair and regeneration due to their avascular nature, often necessitating surgical and suture repair (Beamer et al., 2015; Masoudi et al., 2015; Beamer et al., 2017). They have been shown to improve human neonatal knee articular cartilage chondrocyte adhesion and proliferation (Zhang et al., 2009). Silk-chitosan composite sponges are also known to support the growth of chondrocytes (Kundu et al., 2013). Silk scaffolds can also be used in ligament tissue engineering. Woven silk fibers form a lightweight, strong, elastic material with mechanical properties similar to those of a human’s anterior cruciate ligament (ACL). Modifying braided silk scaffolds with short polypeptides increases collagen production when seeded with mesenchymal stem cells (MSCs). Incorporating microporous silk sponges into a knitted silk mesh mimics a ligament’s extracellular matrix. At 16 weeks postimplantation, the silk partially degraded, allowing for ingrowth of regenerated ligament tissue (Fan et al., 2008). Twisted silk scaffolds or silk-coated poly-lactic-co-glycolic acid (PLGA) nanofibers are both options for ligament regeneration (Kundu et al., 2013). Additionally, knit silk-collagen sponges can be good scaffolds for tendon regeneration when incorporated with stromal cell-derived factor 1 (SDF-1) alpha. It attracts fibroblasts and reduces the accumulation of inflammatory agents (Shen et al., 2010). An additional application for silk is in skin, which is a complex, layered organ. Composite scaffolds best mimic the extracellular matrix of skin by combining silk with collagen I or chitin (Kundu et al., 2013). Silk matrices can be electrospun and functionalized to release molecules such as proteins, DNA, or antibiotics. Functionalized electrospun matrices are used as scaffolds for cell cultures. Scaffolds that release drugs improve biocompatibility and are capable of reducing inflammation and encouraging the migration, proliferation, and differentiation of cells. Silver nanoparticles can be deposited on silk electrospun matrices to be used as wound dressings with antimicrobial properties. When collagen I is added to the silk matrix, it causes keratinocytes to adhere and spread (Zhang et al., 2009).

Cellulose Cellulose is a naturally occurring biomaterial that is commonly used in the biomedical field. It is extremely abundant, easily renewable, and biodegradable. Due to inter- and intramolecular hydrogen bonding between the hydroxyl groups of the neighboring cellulose chains, cellulose is insoluble in water, despite being hydrophilic, and is difficult to dissolve with common organic solvents (Eo et al., 2016). This lack of solubility makes cellulose an ideal biomaterial for tissue engineering purposes since it is sustainable under physiological conditions. Other applications include designing grafts, drug release, aiding in ion exchange, and wound healing. A chemical structure of cellulose can be seen in Fig. 10 (Kang et al., 2015; Torres et al., 2012). Similar to other natural polymers made up of alginates and chitosan, cellulose polymers have sufficient biocompatibility. In addition, cellulose also has adequate mechanical properties, which are necessary to withstand physiological conditions (Torres et al., 2012). Cellulose has also shown to have antiinflammatory and anticancer effects (Eo et al., 2016). Naturally, cellulose does not degrade in humans due to the lack of the enzyme cellulase, which breaks down cellulose. Degradation is one of the bigger problems cellulose faces as a biomaterial. Fortunately, biodegradation can be improved in vitro through periodate oxidation (Torres et al., 2012).

Bacterial Cellulose Bacterial cellulose (BC), also known as microbial cellulose, is a biodegradable, natural cellulose that is synthesized by bacteria. The diameter of BC fibers is 20–100 nm. Bacterial cellulose has high water retention due to being very hydrophilic and having a high surface area to mass ratio. It also has great mechanical strength, exhibits high crystallinity, and is relatively inexpensive to produce (Fu et al., 2013; Mano et al., 2007). The molecular structure is shown in Fig. 11. BC can be synthesized by several types of bacteria, including Cluconacetobacter xylinus. When modified, it can form a material similar to cartilage that is easily molded (Kowalska-Ludwicka et al., 2013). BC can also be produced by the fermentation of Acetobacter xylinum. A limitation is that BC nanofibrils pack tightly to create a dense mesh, which limits opportunities for cell ingrowth. This can be remedied by incorporating porogens during fermentation (Zaborowska et al., 2010). It can also be produced by Gluconacetobacter xylinus and is highly permeable (Fu et al., 2012).

OH HO

3

2

OH 1

4′

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5 OH

Fig. 10

Schematic representation of cellulose structure.

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Chemical structure of bacterial cellulose (Ulery et al., 2011).

Bacterial cellulose was synthesized to make a biocompatible version of naturally occurring cellulose. BC is generally nontoxic, and low cytotoxicity is observed (Fu et al., 2013). When modified, it has chemical and physical properties suitable for reconstructive surgery (Kowalska-Ludwicka et al., 2013). Endothelial cells, chondrocytes, and smooth muscle cells have been shown to adhere to BC scaffolds (Zaborowska et al., 2010). In previous studies, bacterial cellulose did not elicit an immunogenic response aside from temporary inflammation, and thus is biocompatible (Torres et al., 2012). BC is also biodegradable by enzymes found in nature but not in the body. Buffers can be added to encourage bioabsorption (Fu et al., 2013). A common application for bacterial cellulose is found in skin. BC is permeable and absorbent enough to drain fluid from the wound. It is also easily removed after a certain period of recovery, but is still able to conform to the body for extended periods of time (see Fig. 12). Modifying BC with nitrogen plasma improved cell affinity and increased porosity. Additionally, incorporating human epidermal growth factor and collagen increased the proliferation of human fibroblasts. A composite of BC and collagen I makes a superior wound dressing with a high biocompatibility and tensile strength. Silver nanoparticles can also be incorporated into the composite for antimicrobial effects (Fu et al., 2013). Moreover, the transparency of BC is good for wound bandages as it absorbs heat and reduces pain, which is especially helpful for burn patients (Kowalska-Ludwicka et al., 2013).

Fibrin Fibrin is a natural biopolymer formed during the last leg of the coagulation cascade after a division of fibrinogen by the action of thrombin. The coagulation cascade is initiated by an injury or after blood comes in contact with foreign material. Thrombin, produced in the last step of the cascade, divides fibrin peptides A & B to form fibrin monomers (see Fig. 13). The monomers form a fibrin network stabilized through crosslinking by factor XIIIa (see Fig. 14) (Ariëns et al., 2002). Along with platelets and red blood cells, fibrin clots blood to stop bleeding. Fibrin produces a matrix in which cells such as leukocytes, fibroblasts, and endothelial cells attach, move, grow, and organize to generate different functions (see Fig. 15) (Shats et al., 1997).

Fig. 12

(A) pure bacterial cellulose; (B) wet BC-ClAlPc; (C) dry BC-ClAlPc; (D) SEM image of BC-ClAlPc (Clark and Deswarte, 2011).

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Fig. 13 Fibrogen structure. Reprinted with permission from Brown, A. C., and Barker, T. H. (2014). Fibrin-based biomaterials: modulation of macroscopic properties through rational design at the molecular level. Acta Biomaterialia 10 (4): 1502–1514.

Fig. 14 Fibrin polymerization. Reprinted with permission from Brown, A. C., and Barker, T. H. (2014). Fibrin-based biomaterials: modulation of macroscopic properties through rational design at the molecular level. Acta Biomaterialia 10 (4): 1502–1514.

As a naturally produced polymer in the human body, fibrin is highly biocompatible. Fibrin fights off materials that are not biocompatible with the body. Fibrin gel, formed by fibrino-peptides to be insoluble, can be degraded with plasmin-mediated fibrinolysis (Sidelmann et al., 2000). Unlike collagen-based hydrogel, which has fast degradation rate, fibrin has a controllable degradation rate, which is beneficial in tissue engineering. Different types of tissue regeneration require respective degradation rates of the regenerative vehicle, and thus fibrin’s tunable degradation is important (Kjaergard and Weis-Fogh, 1994). Fibrin developed injectable scaffolds are used for the treatment of damaged cardiac and cartilage tissues (Chien et al., 2012). Additionally, fibrin gel-based cell carriers protect cells from environmental forces experienced during the cell delivery process while improving cell viability and tissue growth (Christman et al., 2004).

Gelatin Gelatin consists of a mixture of water and peptide fragments. These peptide fragments are typically derived from animal tissue. This process involves the reduction of crosslinks in collagen as well as the removal of any impurities that may be present. Since gelatin is a derivative of collagen, it can be manipulated in ways that affect the materials crosslink density and chemical reactivity. These variables can be modified for the intended function and application. In addition to its low production cost, gelatin has excellent biocompatibility and is easy to manufacture. Some of the common applications associated with gelatin are drug delivery, hydrogels, and scaffolds. Gelatin has been utilized in the pharmaceutical industry as a medium for drug delivery, most commonly in the form of oral capsules. Although effective, regenerative medicine is now focused on methods that allow for more localized delivery. Microspheres

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Fig. 15 Fibrin PLLA/PLGA composite gels facilitate graft neovascularization and perfusion. (A) Low-power images of fluorescein isothiocyanate (FITC)-Dextran (green) and bright-field images to emphasize the graft area. (B) High-magnification image of the graft area which show the thorough penetration of FITC-Dextran and neovessels 10 days post-implantation. (C) Quantification of the FITC-Dextran and neovessel penetration into the graft area and graft area size (D) showed that the cellularized or noncellularized fibrin and PLLA/PLGA composites provided the most support for graft neovascularization and perfusion (n ¼ 3–6). Reprinted with permission from Brown, A. C., and Barker, T. H. (2014). Fibrin-based biomaterials: modulation of macroscopic properties through rational design at the molecular level. Acta Biomaterialia 10 (4): 1502–1514.

are an example of a matrix that has shown a lot of potential as a targeted delivery system. There are advantages of using microspheres in applications such as drug delivery, bone tissue engineering, and regeneration. Absorption and desorption of substances as well as the kinetic release of the loaded drug components become variables that can be further modified to better suit the intended function (Hossain et al., 2015). Gelatin microspheres containing calcitonin gene-related peptide (CGRP) or substance P promoted bone growth in a rabbit osteoporotic bone defect model. The rate of bone growth was positively correlated with the dose administered (Chen et al., 2016). Aside from microspheres, nanoparticles are also an area of interest. Gelatin nanoparticles were infused with polyethylene glycol and then incubated with breast cancer or BT-20 cells to determine uptake. The results of this study showed that this method may be effective as a long-circulating delivery system in vivo. Furthermore, the nanoparticles have the potential to encapsulate hydrophilic macromolecules and are internalized by tumor cells. Although drug delivery is the most common application for this biomaterial, it is not the only one. A two-component polymer system was created using gelatin films incorporated with 2-hydroxyethyl methacrylate (HEMA) monomer. This material was tested mechanically to determine whether it exhibited properties similar to skin tissue. Once this was confirmed, an additional test was done on burn victims, which resulted in an accelerated rate of healing (Alam et al., 2015).

Synthetic Biomaterials Synthetic biomaterials are not of natural origin and are synthesized for an intended purpose. These biomaterials have several advantages such as high reproducibility, availability, and tunability. Through slight changes during fabrication, it is easy to tune or modify numerous characteristics such as the mechanical properties, degradation rate, and composition of the synthetic biomaterial. A limitation of synthetic biomaterials is that they structurally differ from native tissues and organs. This can affect their biocompatibility and ability to promote tissue remodeling. Synthetic materials also often lack sites for cell adhesion, which may limit their

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regenerative capabilities in vivo. Since they are not of natural origin, synthetic biomaterials are more susceptible to eliciting an immune response and toxicity after implantation. Although this is a valid concern, there are still several synthetic biomaterials that are highly biocompatible and integrate well with the body (Ha et al., 2013; Ige et al., 2012; Bartis and Pongrácz, 2011). Synthetic biomaterials are comprised of metals, ceramics, nonbiodegradable polymers, and biodegradable polymers. They are commercialized and used in clinical settings for treatments such as metal hip implants, plastic intraocular lenses, and many other products utilized in the medical field (Ha et al., 2013). The following sections will detail several synthetic biomaterials and how they are applied in biomedical disciplines.

Polyglycolic Acid Polyglycolic acid (PGA) was one of the initial, degradable polymers researched for biomedical application. Since the 1970s, PGA has been used as the degradable suture DEXON due its material characteristics including a melting point (Tm) greater than 200 C, a glass transition temperature (Tg) between 35 C and 40 C, and a very high tensile strength. The majority of recent studies utilize PGA as a filler material integrated into other degradable polymers. PGA is commonly incorporated into scaffolds for various tissue engineering applications such as bone, tendon, cartilage, tooth, and spinal regeneration. Despite these applications, PGA has limitations as its rapid degradation compromises its mechanical strength, and could potentially cause an undesirable inflammatory response due to the resulting increase of glycolic acid (Ulery et al., 2011). The chemical structure of PGA can be seen in Fig. 16. PGA has proved to be highly biocompatible in most of its applications. Even though there are some reports of potential immunogenic responses when utilizing PGA, most applications have not caused any inflammatory reaction. Additionally, PGA is known for its hydrolytic instability. Hydrolytically unstable polymers are materials that have chemical bonds in their backbone that are susceptible to hydrolysis without an external influence. In the case of PGA, its hydrolytic instability can be attributed to the ester linkage in its backbone. PGA’s random hydrolysis led to the combination with other polymers such as polylactic acid (PLA) to control its degradation rate. In addition to degradation by hydrolysis, PGA also undergoes enzymatic degradation in vivo (Ulery et al., 2011; Clark and Deswarte, 2011). In practice, PGA has been used in an effort to enhance facial nerve regeneration. This was done by placing bone marrow stem cells into a PGA tube, and observing for neural regenerative effects. PGA is suitable for neural regeneration because it is absorbable and has FDA approval for nerve grafting (Anderson et al., 2015). Nerve grafting is a very complex process, and is still in the early stages of research. In the studies conducted thus far, PGA has been shown promising results for producing nerve graft structures (Costa et al., 2013). PGA has also been utilized in wound healing and adhesives. Polyglycolic acid sheets were used in conjunction with fibrin glue spray as an open wound healing material for soft tissues as well as bone surfaces during oral surgery. The PGA adheres to the wound successfully and helps prevent postoperative bleeding as well as inspire epithelialization. These sheets were first used only for soft tissues, and have since been used on hard tissues as well. These PGA sheets can be seen in Fig. 17 (Sakaguchi et al., 2015). As an adhesive, PGA was combined with fibrin sealant to create a very successful tissue adhesive. The combination of the PGA and fibrin created a much stronger sealant than any other biomaterial combination (Shinya et al., 2009).

n Fig. 16

PGA structure.

Fig. 17

(A) Sheets of PGA were cut into small pieces (5–10 mm wide) and (B) approximately 3–10 pieces were used to cover each wound.

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Polylactic Acid Lactic acid is a naturally occurring organic acid that can be produced by chemical synthesis or fermentation. Typically, this process consists of hydrolysis of lacronitrile by strong acids (Figs. 18 and 19). These processes allow for polylactic acid (PLA) to be derived from renewable resources. Currently, PLA is one of the most produced polymers on the market. Primary production of PLA happens through fermentation and is environmentally friendly (Rasal et al., 2010). PLA is a thermoplastic, meaning it can easily cast into many shapes and has potential for 3D printing (Stratton et al., 2017). It is also highly biocompatible, and can be easily absorbed by the body. When lactic acid polymers are manufactured, they can be created in a way that allows for different degradation rates. Hydrolysis kinetics is a multivariable concept that determines how fast implants will degrade. Density, bond strength, H2O concentration, and hydrophilicity all play roles in this degradation process. By changing the degradation rate, the polymer can be designed to release embedded medicine (James et al., 2016). The fastest rates of release are associated with implants that will undergo bulk degradation, while the slower ones only degrade at the surface. Surface degradation is beneficial in applications where sustained release of given compounds are needed (James et al., 2016). The ability to modify the surface of PLA allows for it to be a versatile biomaterial. Chemical modification, plasma treatments, entrapment, and surface coatings can all be used to change the way PLA will function in the body. These methods can be used to modify variables such as hydrophilicity and biocompatibility. However, many have yet to be perfected and can cause some drawbacks due to undesired changes. In biomedical application, PLA’s resorbability makes it an ideal polymer for short-term fixation such as in fracture repair. This type of fixation shows promise for adolescents that are still growing. These devices are ideal for craniofacial surgical operations, an area that is very active as a child grows (Holmes et al., 2004). Conventional biostable devices would not be ideal in this situation, as they would not function properly in the dynamic environment (Fig. 20) (Peltoniemi et al., 2002).

O

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OH OH

H

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CH3

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OH

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Fig. 18 Lactic acid enantiomers. Reprinted with permission from Holmes, R. E., Cohen, S. R., Cornwall, G. B., Thomas, K. A., Kleinhenz, K. K., Beckett, M. Z. (2004). MacroPore resorbable devices in craniofacial surgery. Clinics in Plastic Surgery 313, 393–406.

HO

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O CH3 O O H3C O Lactide

Fig. 19 Hydrolysis of lactic acid to create poly(lactic acid). Reprinted with permission from Holmes, R. E., Cohen, S. R., Cornwall, G. B., Thomas, K. A., Kleinhenz, K. K., Beckett, M. Z. (2004). MacroPore resorbable devices in craniofacial surgery. Clinics in Plastic Surgery 313, 393–406.

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Bone and cartilage regeneration

3D scaffolds

Bone fixation devices

Applications

Ligament reconstruction

Vehicle for long term drug delivery

Injectable micropheres for drug delivery

Fig. 20

Poly(lactic acid) applications for regenerative medicine.

Poly Lactic-co-Glycolic Acid Poly lactic-co-glycolic acid (PLGA) is a copolymer of polylactic acid (PLA) and polyglycolic acid (PGA) (Fig. 21). It is one of the preferred materials for biomedical applications because it can be easily optimized for a variety of uses. PLGA can be modified in terms of its shape, size, degradation rate, and mechanical properties to meet the needs of a given application (Makadia and Siegel, 2011). The copolymer is a biocompatible and biodegradable material with minimal systemic toxicity. As PLGA degrades, it produces metabolic monomers, lactic acid, and glycolic acid, which are metabolized and eliminated from the body (Danhier et al., 2012). Currently, PLGA is widely used in sutures, drug delivery, and tissue engineering scaffolds (Lee et al., 2016). Although PLGA is considered a biocompatible, synthetic polymer, its degradation products may cause inflammatory and foreign body reactions upon implantation. These reactions are initiated by a pH decrease in surrounding tissues, which results from the accumulation of lactic and glycolic acid during degradation (Wei Ji et al., 2012). Composition, crystallinity, average molecular weight, size, and shape are all factors that affect degradation of PLGA. Polymer composition is the most important factor when discussing degradation because the PLA/PGA ratio determines the amount of glycolic acid, which in turn affects its degree of hydrophilicity. The rate of degradation is directly proportional to the glycolic acid concentration (Makadia and Siegel, 2011). Degradation times of 50:50, 75:25, and 85:12 PLGA are 1–2 months, 4–5 months, and 5–6 months, respectively (Ulery et al., 2011). Size and shape have similar effects on the rate of degradation in that larger surface areas cause increased degradation of the matrix. In addition, PLGA’s bulk degradation is reported to be quicker than its surface degradation, which causes drug delivery to be faster in devices with high surface area to volume ratios. Finally, molecular weight is directly related to the polymer chain

O O

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OH HO Glycolic acid

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PLGA structure and hydrolysis.

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Fig. 22 Silk/PLGA-based scaffold for ligament/tendon tissue engineering. (A) Transmission Electron Microscopy (TEM) and (B) SEM images of electrospun bFGF-containing PLGA fibers which show the distribution of proteins as indicated by the black arrows. (C) SEM image of FGF coating with electrospun fibers (eF) on microfibrous knitted silk scaffolds (mF). (D) Bone marrow stromal cell (BMSC)-seeded scaffold prior to rolling into cylindrical ligament/tendon constructs and (E) after 7 days of culture. Reproduced from Sahoo, S., Toh, S. L., Goh, J. C. (2010). A bFGF-releasing silk/PLGA-based biohybrid scaffold for ligament/tendon tissue engineering using mesenchymal progenitor cells. Biomaterials 31(11), 2990–2998.

length, which means that the higher the molecular weight, the lower the degradation rates since the polymer chain is proportionally larger (Makadia and Siegel, 2011). PLGA provides a great variety of options when it comes to drug delivery applications. Its degradation is faster than many polymers, but more importantly, its degradation rate can be controlled. PLGA can be packed in any shape and size such as nanospheres, macrospheres, and millimeter-sized implants (Makadia and Siegel, 2011). It is used in the controlled administration of drugs, peptides, and proteins. PLGA has been employed to deliver chemotherapeutics, proteins, vaccines, antibiotics, analgesics, antiinflammatory drugs, and siRNA (Ulery et al., 2011). PLGA can also be applied in tissue regeneration to engineer scaffolds that promote cell culture, proliferation, and differentiation until the regenerated tissue has stabilized (Danhier et al., 2012; Pan and Ding, 2012). In recent years, studies have shown that mesenchymal stem cells have the ability to differentiate into several tissues including bone, tendon, ligament, and fat. In one study, Sahoo et al. demonstrated the use of a silk/PLGA-based scaffold for ligament/tendon tissue engineering. The scaffold promoted cell proliferation and enhanced the production of collagen that resulted in strong structures that proved to have characteristics to aid ligament/tendon repair (see Fig. 22) (Sahoo et al., 2010). Finally, PLGA scaffolds also provide optimal conditions for bone regeneration. These structures are characterized by interconnected pores and large surface areas that allow tissue ingrowth. Preferred pore sizes range from 100 to 350 mm with more than 90% porosity (Yoshimoto et al., 2003). Several techniques are used in the fabrication of PLGA scaffolds, including gas foaming, microsphere sintering, poregen leaching, electrospinning, and polymer printing (Ulery et al., 2011). For example, PLGA foams were successfully used to culture osteoblasts derived from rat mesenchymal cells; showing proper mineralization and three-dimensional bone formation (Yoshimoto et al., 2003).

Polycaprolactone Polycaprolactone (PCL) is a synthetic, biodegradable polymer. PCL nanofibers can be aligned to mimic an extracellular matrix structure (Manoukian et al., 2017). The average diameter of the nanofibers is 500–900 nm (Engel et al., 2008). PCL was first synthesized in the 1930s, but was not commonly used due to its slow degradation rate and inability to bear heavy loads. It regained popularity after the emergence of tissue engineering in the 1990s (Woodruff and Hutmacher, 2010). The molecular structure is shown in Fig. 23. When combined with collagen type I, it is capable of withstanding a high-pressurized environment for an extended period

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n

Chemical structure of PCL (Costa et al., 2013).

of time. The addition of collagen increases the burst pressure and allows PCL to become stronger as a result (Sell et al., 2009). Natural biomaterials can be added to PCL to enhance polycaprolactone’s mechanical properties and stability while avoiding the need to crosslink and introduce harmful crosslinking agents (Cooper et al., 2011). PCL was determined to be biocompatible after both short-term and long-term studies did not elicit any adverse reactions from the host tissue. Fibrous scaffolds of electrospun PCL blended with gelatin demonstrate exceptional biocompatibility with bone marrow stromal cells. PCL composites also support the growth of adipose-derived stem cells and human coronary artery endothelial cells (Sell et al., 2009). Chitosan–PCL composite scaffolds have also been shown to be compatible with Schwann cells in nerve studies (Cooper et al., 2011). It takes more than a year for polycaprolactone to degrade noticeably, and the total degradation time is up to 4 years. PCL is biodegraded by bacteria and fungi. In the human body, degradation is a two-step process. During the first year, ester groups are hydrolytically cleaved. Next, intracellular degradation occurs (Woodruff and Hutmacher, 2010). Degradation is ultimately dependent on the molecular weight and crystallinity of the polymer or composite (Lam et al., 2008b). The slow degradation of this synthetic polymer is beneficial for neural regeneration as nerves regenerate slowly (Cooper et al., 2011). PCL is a versatile polymer that has been implemented into various biomedical research studies. For example, PCL or PCL/ hydroxyapatite scaffolds have been created using a precision extrusion deposition process. They support the growth and migration of primary fetal bovine osteoblasts and have the correct mechanical properties, pore size, and interconnectivity for bone tissue engineering (Shor et al., 2007). Nonwoven meshes of PCL incorporated with beta-tricalcium phosphate nanoparticles form a composite that mimics the structure of bone tissue at the bone–cartilage interface (Erisken et al., 2008). PCL scaffold porosity can be created and tuned using selective laser sintering to make scaffolds similar to trabecular bone (Eshraghi and Das, 2010). Electropsun scaffolds of PCL/collagen type I composites have also been successfully implanted as an artery in a rabbit model. The composite demonstrated mechanical properties such as tensile strength and elasticity comparable to native arteries. Moreover, the PCL/collagen type 1 composite did not cause platelet adhesion which prevented undesirable clotting (Sell et al., 2009). Nanofiber scaffolds of PCL, collagen type I, and collagen type II support growth and attachment of human coronary artery endothelial cells (Sell et al., 2009) (Fig. 24). In terms of neural regeneration, aligned chitosan–PCL composite fibrous scaffolds resulted in Schwann cells

Fig. 24 Electrospun PCL/collagen composite scaffolds: (A) The gross appearance of scaffold; SEM images of (B) the entire PCL/collagen composite (18x), (C) the scaffold surface (6000x), and (D) the scaffold’s cross-sectional appearance (4000x). Reproduced from Sell, S. A., McClure, M. J., Garg, K., Wolfe, P. S., Bowlin, G. L. (2009). Electrospinning of collagen/biopolymers for regenerative medicine and cardiovascular tissue engineering. Advanced Drug Delivery Reviews 61 (12), 1007–1019.

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aligning in a bipolar fashion along the direction of the aligned fibers, and thus making these scaffolds appropriate for the reconstruction of nervous tissue (Cooper et al., 2011).

Polyetheretherketone Polyetheretherketone (PEEK) is a thermoplastic polymer that was commercialized in the early 1980s and later proposed as a material to be used for medical applications (Panayotov et al., 2016). Typically PEEK composite materials have been used as a replacement for metallic and ceramic implants, which have a higher modulus of elasticity (Kurtz and Devine, 2007). PEEK was first introduced to the medical field as a fracture fixation, and it remains to be used for this purpose. Despite several years and advancements, PEEK is still one of the most popular biomaterials on the market due to its versatile capabilities (Manoukian et al., 2016b). The chemical stability of PEEK has been of interest to medical device companies since its introduction in 1998. At room temperature, the only solvent capable of dissolving PEEK is 98% sulfuric acid (Ferguson et al., 2006). Its modulus of elasticity is nearly identical to cortical bone, which allows it to be an optimal candidate for an interbody device in vertebral fusion applications. Furthermore, PEEK is highly resistant to gamma and electron beam radiation, which allows for easy sterilization. The free radicals, which are produced from these sterilization methods, have a life span of about 20 min. Due to this short lifetime, PEEK will not be considered as a source of secondary electron emission. PEEK is readily visible under MRI because of its natural radiolucency (Johansson et al., 2016). PEEK has been incorporated into numerous practices such as dental reconstruction, spinal surgery, and many others as shown in Fig. 25. A major drawback often observed with titanium and titanium alloy implants is stress shielding. Due to the significant difference in elastic modulus between titanium and bone the strains that they face when stresses are applied are different. This can cause bone loss and ultimately implant failure. Recently PEEK has been proposed as an alternative because of its favorable elastic modulus (Fig. 26). However, PEEK does not have a high resistance to mechanical strength which means it would need to undergo modifications to better suit the desired functionality. PEEK has also been proposed as the implant material for vertebral fusion surgery since its stiffness is similar to bone and will allow the graft to heal uniformly. In addition, PEEK is radiolucent which will allow the doctor to ensure that the implant is properly positioned for optimal healing.

Polyethylene Glycol Polyethylene glycol (PEG) is a linear or branched polyether with a general structure that is terminated with hydroxyl groups (see Fig. 27) (Roberts et al., 2002). Some relevant characteristics for its use in the biomedical field include solubility in water and organic solvents, controlled solubility, nontoxic, and hospitable to biological materials (Harris, 1992). PEG is widely used for biomedical applications because it is a nontoxic polymer that yields nonimmunogenicity and nonantigenicity. It does not harm active cells or proteins, and thus it does not cause immunogenic responses when introduced in the body (Alcantar et al., 2000). Tunable

Cranial reconstruction

3D scaffolds

Hip implants

Applications

Bone reconstruction

Tooth replacement

Vertebral body replacements

Fig. 25

Medical applications of polyetheretherketone.

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Temporary PEEK abutments (A)

(B)

(C)

PEEK dental implants (D)

(E)

(F)

Fig. 26 Examples of PEEK dental implants and temporary abutments (A) PEEK abutment (DENTIN Implants Technologies LTD); (B) PEEK abutment (Nobel Biocare); (C) PEEK abutment (SGC Dental); (D) PEEK Perso A implant (SisoMM); (E) Win! PEEK implant (Champions Implants); (F) Biopik (Biopikimplants). Reprinted with permission from Panayotov, I. V., Orti, V., Cuisinier, F., Yachouh, J. (2016). Polyetheretherketone (PEEK) for medical applications. Journal of Materials Science: Materials in Medicine 277, 1–11.

HO−(CH2−CH2−O)n−H) Fig. 27

Chemical structure of poly(ethylene glycol) (PEG).

biodegradability is one of the most important characteristics in the biomedical applications of PEG. However, PEG is nonreactive and it requires functionalization with cross-linked groups in order to create insoluble networks. Addition of acrylate, thiol, amine, maleimide, or vinyl sulfone reactive groups has been used to functionalize PEG (Zustiak and Leach, 2010). Cross-linked polymerics networks are used to create hydrogels for biomedical applications such as tissue engineering diagnostics, wound dressing, drug delivery, and barrier materials to regulate biological adhesions (Ahmed, 2015). PEG is commonly used in the form of hydrogels. Hydrogels are cross-linked polymer networks characterized by their ability to absorb and retain water inside their structure (Ahmed, 2015). A hydrogel’s water content makes it ideal for tissue engineering applications since their physical properties are similar to soft tissues. In one study, a polyethylene glycol (PEG) hydrogel, composed of PEG vinyl sulfone (PEG-VS) cross-linked with PEG-diester-dithiol, was used to create three-dimensional hydrogel matrices with tunable mechanical properties and degradation. Influential parameters included molecular weight, polymer density, and the distance between thiol and ester group in the cross-linker. Additionally, the hydrogels were modified with the fibronectinderived cell-adhesive peptide, arginylglycylaspartic acid (RGD), to aid cell viability in 3-D culture, and to show cell responses for soft tissue regeneration (Zustiak and Leach, 2010) (see Fig. 28). Mahoney and Anseth et al. developed hydrogels by photopolymerizing methacrylate groups covalently linked to PEG macromers. Scaffolds were created in order to support neural cell growth and function. In their approach, they demonstrated that neural precursor cells tolerate photoencapsulation and culture in PEG hydrogels with only 10% cell death in 24 h (Mahoney and Anseth, 2006). Surface coating of materials is a technique used when developing artificial implants. The goal is to produce biocompatible interfaces between body fluids and materials such as polymers, ceramics, and metals (Alcantar et al., 2000). PEG is widely used for this purpose due to its biocompatibility, water solubility. Alcantar, Aydil, and Israelachvili et al. developed a method that allows grafting of PEG onto activated silica films, which can be deposited in various materials such as metals and plastics. The produced films were smoother, more hydrophilic, and absorbed less proteins (Alcantar et al., 2000).

Polymethyl Methacrylate Polymethyl methacrylate (PMMA) is a lightweight, synthetic polymer that is an economical alternative to polycarbonate when extremely high strength is not necessary. An advantage is that PMMA does not contain potential harmful subunits like bisphenol-A found in polycarbonate. Moreover, the synthetic polymer is easier to handle, process, and less expensive than polycarbonate. In practice, PMMA is often used for craniofacial tissue defects such as skin and dentures (Pielichowski and Njuguna, 2005).

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Fig. 28 Cell viability in degradable 3D PEG hydrogels that contained 100 mM RGD. A live/dead assay was preformed for 24 hours following cell culture, and the findings showed no significant differences in cell viability. Reproduced from Zustiak, S. P., Leach, J. B. (2010). Hydrolytically degradable poly(ethylene glycol) hydrogel scaffolds with tunable degradation and mechanical properties. Biomacromolecules 115, 1348–1357.

PMMA has great mechanical properties and low toxicity. While being popular for hip-joint transplantations because of its inert properties, PMMA displays slow degradation (Ganesh et al., 1997). Therefore, creating a polymer blend of polycaprolactone with PMMA produces a polymer material that is better suited for biomaterial applications. In vitro studies conducted by So-Ra Son et al. 2013 used MTT assays to examine cytotoxicity and proliferation of MG-63 osteoblast cells on blended scaffolds of PCL/ PMMA (Son et al., 2013). The study found the blended polymer material suitable for osteoblast cell proliferation. Further evidence of confocal images and expression of proliferation cell nuclear antigen confirmed proliferation and expression of cells in the 7:3 PCL:PMMA blended polymer environment. PMMA is a nonbiodegradable polymer utilized in applications that require permanent, mechanically stable structures such as bone tissue regeneration (Jessy and Ibrahim, 2014). At high temperatures, the long chain backbone of PMMA separates and reacts with itself to change properties. PMMA can undergo thermal degradation and thermal oxidative degradation in the presence of oxygen, as well as photodegradation, oxidative degradation, and UV degradation. Thermal degradation causes changes in properties of the polymer such as reduced ductility, chalking, color changes, and cracking. PMMA depolymerization at 300–400 C produces volatile monomers of methyl methacrylate (MMA) (Holland and Hay, 2001). PMMA is used in biomaterial applications such as bone cement, lenses, bone substitutes, and drug delivery systems. It is used to remove wrinkles and scars on skin tissue permanently. In dental implants, polymer material PMMA is substituted for missing dental roots. Similar to the physical and mechanical qualities of human dentine, PMMA has a low modulus of elasticity, thermal and electrical passiveness, and ideal porosity (Leigh, 1975).

Conclusion Biomedical engineering is a board field that mainly represents aspects of biomaterials, bioinformatics, biomechanics, bioinstrumentation to aid in medicine and health. In this article, we discuss the use of biomaterials that facilitate the interaction between the body and materials present on the surface level of biomedical devices. This is important in the successful use of biomedical devices because if the immune system rejects the biomaterials present on the device, the device will fail and cause a severe immune response in the human body. With advancements in nanotechnology and materials science, there have been further studies of biomaterials for tissue engineering and drug delivery methods. In this article we discuss both natural and synthetic polymers used in biomaterials and their characteristics. The polymers examined in this article are collagen, chitosan, silk, bacterial cellulose, fibrin, polylactic acid (PLA), polyglycolic acid (PGA), polycaprolactone (PCL), polyetheretherketone (PEEK), polyethylene glycol (PEG), and polymethyl methacrylate (PMMA). Understanding their properties has allowed biomedical engineers to make vast improvements in the field of biomaterials, and the lives of patients.

Acknowledgments Authors acknowledge funding support from the National Institute of Biomedical Imaging and Bioengineering of the National Institutes of Health (award number R01EB020640), Connecticut Regenerative Medicine Research Fund (15-RMB-UCHC-08) and the National Science Foundation Graduate Research Fellowship award in support of this work.

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Further Reading Lam, C. X. F., Hutmacher, D. W., Schantz, J.-T., Woodruff, M. A., & Teoh, S. H. (2008a). Evaluation of Polycaprolactone scaffold degradation for 6 months in vitro and in vivo. Journal of Biomedical Materials Research, 90A(3), 906–919.

Biomimetic Approaches for Regenerative Engineering Nirmalya Tripathy, University of Washington, Seattle, WA, United States Rafiq Ahmad, Jeong Eun Song, and Gilson Khang, Chonbuk National University, Jeonju-si, Jeollabuk-do, Republic of Korea © 2019 Elsevier Inc. All rights reserved.

Introduction Biomimetic Materials Natural Materials Fibrin Collagen Glycosaminoglycans Self-assembling polypeptides Synthetic Materials Mechanisms of hydrogel formation Hybrid Materials Fabrication of Biomimetic Scaffold Prefabricated preparation of scaffolds Porous scaffolds by conventional approaches Fibrous scaffolds by electrospinning Rapid prototyping-based microfabrication Modular hierarchical assemblies Scaffold Functionalization Biomolecules Delivery Importance of Bioreactors Biomimetic Scaffolds for Regenerative Engineering Bone tissue Cardiovascular tissue Conclusions and Future Prospects Acknowledgments Further Reading

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Introduction In general terms, “Biomimetics” aims toward understanding biological fundamentals and implementing for developing biomedically relevant technologies and functional biomaterials mimicking physicochemical, mechanical, and biological characteristics of natural material, for example, for biomedical devices or as scaffolds for tissue regeneration. In regenerative medical therapy, tissue engineering (TE) approaches direct the cells to differentiate at the desired time and place while maintaining the right phenotype. Thus a biomimetic approach to the generation of engineered tissues enables the assembly of functional tissues using biologically derived design requirements. These include synthesis to optimization of certain compositions or properties similar to those of the extracellular matrix (ECM), novel processing technologies to achieve structural features mimicking the ECM on various levels, strategies to emulate cell–ECM interactions, and biologic delivery strategies. This article outlines current biomimetic materials approaches in TE, with improved cellular/tissue functions and regenerative outcomes, showing the biomimetic materials significance for TE and regeneration. Biomimetic scaffolding materials, natural or synthetic, biodegradable or permanent, are tuned into 3D architectures suitable for cell seeding and cultivation. Biomaterial choices are mostly guided by the requirement for enough mechanical strength and restore cell signaling for the tissues. They also act as informational templates for cells by implementing patterning, ligand binding, and sustained release of cytokines. Each tissue consists of its own challenges for designing scaffold. For example, bone tissues consist of dense mineralized matrix that can withstand significant compressive loads, requires scaffolds that are capable for providing mechanically stable framework with interconnected large pores for cell infiltration. Another instance is cardiac muscle, dense syncytium of electromechanically coupled cells that transmit force and deformations in multiple directions, requires soft and elastic scaffolds. As well, custom-designed scaffolds were also finding utility in developing tissue models for drug screening and further disease studies without usage of in vivo models such as 3D printed cardiac tissue, neural constructs for predictive drug screening.

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Biomimetic Materials Natural Materials Fibrin Fibrinogen (Fg) is a monomeric and has the 340 kDa molecular weight and aggregates to produce fibrin polymer subsequently thrombin cleavage. Thrombin acts on fibrinogen to cleave the fibrin peptides on fibrinogen that put off physicochemical polymerization or self-assembly of the molecule. The resulting network is chemically cross-linked by the blood transglutaminase factor XIII a, and its complex fibril structure and cross-linked character depend upon the details of its formation. Fibrin plays a significant responsibility in healing and tissue repair in adults, but less so throughout embryonic development. As a result, moderately few morphogenetic signals concerned in development interrelate particularly with fibrin, although they are most vigorous when immobilized in a 3D matrix. Fibrin has been used as a scaffold protein for the immobilization of adhesion and growth factors. For instance, heparin-binding growth factors have been integrated to fibrin scaffolds, and bind with heparin, then it is immobilized via a heparin-binding peptide enzymatically integrated into the fibrin matrix. Fibrin, not like collagen or fibronectin, is not associated with mature tissue structure, but somewhat as a temporary repair stage ECM component. This matrix fibrous has been used in the form of fibrin glue as a tool for surgeons. And the fibrin glue formulations have important amounts of fibronectin which may unintentionally improve cell migration within the fibrin matrix. Fibrin has been used clinically as a scaffold matrix for the release of keratinocytes, providing mesenchymal stem cells transfected with growth factors, for spinal cord repair and repair of peripheral nerves. Even though fibrin is not an ECM in the common sense, because cells in the local environment do not make it, the material is nonetheless a crucial member of the body’s repertoire of matrices and serves the function of a conditional matrix, being remodeled and replaced with ECM molecules. In difference to fibrillar collagen matrices, where cell migration taking both through mechanisms that are reliant and independent of proteolytic degradation, cell migration in fibrin is almost completely dependent upon cellconnected proteolytic activity like plasmin and MMPs. This difference in cellular behavior from fibrillar collagen perhaps results from the slighter mesh size of the fibrin matrices and the stronger fibril–fibril connections, owing to the nature of network formation and covalent stabilization. A number of proteins are naturally integrated into fibrin matrix throughout coagulation, such as fibronectin and alpha-2-plasmin inhibitor, and these factors covalently cross-linked into the matrix by factor XIIIa. Other biomolecules, such as FGF-2, bind noncovalently to fibrin and are capable of providing growth factor-specific bioactivity, such as potentiating of endothelial cell (EC) proliferation, even when bound. On the other hand, fibrin matrices remain naturally badly active for the majority cell types, leading to their functionalization with ECM peptides, recombinant ECM protein domains, or growth factors. Similar to collagen matrices as described earlier, engineered growth factors or ECM fragments can be included into fibrin noncovalently by linking with a fibrin-binding domain to the recombinant protein. For instance, the native knob: pocket connections throughout fibrin assembly can be demoralized to delay the release of recombinant proteins. However, to have even better incorporation and controlled release, biomolecules can be covalently integrated. In a powerful approach, the factor XIIIa enzymatic substrate sequence of alpha-2-plasmin inhibitor, Asn-Gln-Glu-Gln-Val-Ser-Pro-Leu, can be linked with bioactive peptides or recombinant proteins, allowing these molecules to be covalently included during the fibrin’s natural polymerization process via factor XIIIa-stimulated cross-linking. For instance, VEGF, fibronectin fragments, and parathyroid hormone 1–34 have been engineered and covalently included to be released in a manner that depends on local cell-induced proteolysis. In the case of the PTH1–34 fusion, the hormone is active only when released: modification of its N-terminus with the factor XIIIa substrate fusion inhibits the activity of the hormone and this activity is regained after cell-induced proteolysis between the fusion and the PTH1–34 domain. This engineered peptide in fibrin is currently in human clinical evaluation for local bone repair.

Collagen Collagen is a natural component of ECM existing in many tissues, like skin, bone, tendon, ligament, and other connective tissues. One of the fundamental hypotheses in collagen research, as connected to biomaterials, is that evolutionary bioengineering has formed a material that has perfect properties for biological applications. An essential characteristic of this type of biomaterial is its exceptional assembled structure, widespread happening in nature, and possible to complete humiliate in biological environments; thus, collagen has been extensively used as a practical biomaterial in TE. Collagen is made up 89% of natural matrix and 32% of the volumetric composition of bone. Collagen scaffolds formed according to an identical protocol have exposed exceptional biological performance owing to their high porosity and permeability. Earlier studies noted that scaffolds need a high porosity and surface area combined with a good permeability and pore interconnectivity for cell migration and nutrient perfusion through the cell culturing procedure. The major drawback of collagen as a scaffold substance for bone TE is that it has comparatively poor mechanical properties. Collagen type I, the mainly abundant protein in mammals, has a triple helical structure made of three polypeptide chains containing repeating Gly-X-Y triplets in which the X and Y positions are regularly engaged by proline and 4hydroxyproline, correspondingly, the latter of which is extremely significant for intermolecular hydrogen bonding. Collagen can be readily purified from skin, tendon, and placenta. This ECM matrix protein can be reconstituted to form a fibrillar matrix by rising the pH and temperature of a forerunner solution. Broadly used, collagen type I is regularly treated with proteases to get rid of small nonhelical telopeptides that are located at the ends of the triple-helical domain and that responsible to most of the protein’s cross-species immunogenic nature. However, for clinical application of this material, concerns of immunogenicity and disease transmission stay behind. To evade these risks, methods for the recombinant expression of collagen have been urbanized in several eukaryotic expression systems, and recombinant human collagen types I and III are commercially accessible.

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The in vivo application of collagen gels is imperfect by a short age in mechanical strength; many methods have been explained in order to form collagen matrices that own adequate mechanical proprieties to at least partially oppose cell-induced contraction. For example, chemical glycation can regulate, in convenient manner, the elastic character of collagen gels. Collagen hydrogels crosslinked enzymatically or compressed to comparatively high density. Heat and chemical treatments have also been urbanized to create cross-linked collagen sponges for bone and cartilage repair, even though these materials turn into so thick that they are no longer actually hydrogels. Fascinatingly, the dipole moment of fibrillar collagen gives rise to the possibility to get everlasting microscopic and even macroscopic alignment of fibrillar collagen matrices. For example, collagen fibril alignment under a strong magnetic field has been established to inform special individuality for inducing directed cell migration, such as neurites which cultivate preferentially in the route of fibril alignment. Collagens bind with several cellular receptors that alter the cells’ performance; collagen gel bioactivity can be improved with growth factors that have been engineered to attach the matrix in a noncovalent manner. Collagen binding sequences resultants from collagenase, von Willebrand factor or fibronectin have been recombinantly merged to growth factors in order to holdup their discharged from collagen scaffolds. For instance, collagen injection of vascular endothelial growth factor is capable to recover cardiac function after an acute myocardial infarction.

Glycosaminoglycans Glycosaminoglycans (GAGs), linear polymers containing two repeating disaccharide units, is a significant part of human physiology. As GAGs are a form of biopolymers, few issues needs to be resolved sequentially for enhancing their wide applicability. For instance, GAGs and other biopolymers can have dissimilar molecular weights, purity, structure, and even charge distribution depending on the source. The ECM structural proteins have increased biomechanical and biochemical functions owing to the long unbranched polysaccharides, the GAGs. These polysaccharides, generally, are components of proteoglycans of the ECM, aside from in the case of hyaluronic acid, and it is not covalently linked to a protein core and is entwined inside the extracellular space. These strongly anionic polymers absorb water, which provides compressive strength to the ECM, while the GAGs also directly affect tissue organization via connections with cell-surface receptors. Hyaluronic acid (HA) is a linear polysaccharide that contains alternating units of a repeating disaccharide, b-1,3-N-acetyl-Dglucosamine and b-1,4-D-glucuronic acid. HA is a nonsulfated GAG, and it originates throughout the body, from the vitreous of the eye to the ECM of cartilage tissues. HA can be modified in many ways to alter the properties of the resulting materials, including modifications leading to hydrophobicity and biological activity. HA can be altered in several ways to alter the properties of the resulting materials, including modifications leading to hydrophobicity and biological activity. Hyaluronic acid can be isolated from animal tissue, such as the rooster comb, and can be biotechnologically formed using Streptococcus bacterium. Because the material absorbs enormous amounts of water at equilibrium, it forms a hydrogel that is nonfibrillar, and owing to entanglement connected with its high molecular weight, up to several million Daltons, the gel dissolves only very gradually. Numerous chemical hyaluronic acid derivatives have been prepared and by controlling the functional group (e.g., pendant hydrophobic groups), the type of covalent bond (e.g., stable or hydrolytically sensitive), and the cross-linking density, it is promising to create a wide range of physically various materials. A number of changes of the carboxyl and hydroxyl groups of hyaluronic acid have been created, to crosslink the material into an elastic gel that resists suspension or rendering the polymer controllably more hydrophobic and thus less soluble. Hydrogels and hyaluronic acid have been used for various applications, including keratinocyte transfer for dermal wound healing and chondrocyte transplantation for cartilage repair. Although hyaluronic acid interacts with at least three cellsurface receptors (CD44, RHAMM, and ICAM-1), its biological activity can be considerably increased by the integration of other functional biomolecules. For instance, hyaluronic acid gels have been functionalized with peptides or protein fragments resulting from fibronectin to get better fibroblast proliferation and wound healing. Likewise, cellular infiltration into the hydrogel can be enhanced by using MMP-sensitive cross-linkers, since hyaluronic acid gel degradation is principally due to hyaluronidase while cell migration is ambitious principally by the activities of various MMPs.

Self-assembling polypeptides Inspired by the understanding of protein self-assembly, significant progress has been prepared using supramolecular self-assembly of biomolecules to form nanofibrillar matrices in situ. These approaches use noncovalent intermolecular interactions to fabricate higher order structures by self-assembly of oligomeric peptide, nucleotide, and nonbiological amphiphilic building blocks. Many of these systems necessitate nonphysiologic circumstances for self-assembly, numerous can gel at cell-tolerable conditions. Zhang and coworkers developed a class of nanofibrillar gels with very high water content (> 99%) cross-linked by assembly of selfcomplementary amphiphilic peptides in physiological medium. These ionic self-complementary peptides are characterized by a periodic repetition of alternating ionic hydrophilic and hydrophobic amino acids. Upon exposure to aqueous solutions with neutral pH, they form stable b-strand and b-sheet structures, partitioning the side chains to two sides, one polar and the other nonpolar. They self-assemble to form nanofibers with the nonpolar residues inside and positively and negatively charged residues forming complementary ionic interactions, like a checkerboard. A number of similar scaffolds have been reported and customized to deliver growth factors and cells. For instance, gels of RAD self-assembling peptide have been used to encapsulate several growth factors to accelerate dermal wound healing and myocardial repair. Self-assembled hydrogel peptides are also being used as a tool for three-dimensional cell culture. For example, several types of cells like chondrocytes and neural stem cells have been cultured within these scaffolds. However, these gels biomechanically systematize cells in a three-dimensional fashion; they show no specific cell interaction because they are not equipped with any specific biofunctional ligands. To engineer receptor-mediated biospecificity, active and functional motifs from the ECM have been working

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to significantly enhance their interactions with cells and tissues. Self-assembling oligomeric amphiphiles permit integration of specific biomolecular signals without inhibiting the self-assembling properties and nanofiber formation. The scaffolds used for the laminin-derived peptide Ile-Lys-Val-Ala-Val, encapsulated neural progenitor cells were observed to differentiate into neurons. Following this trend, a range of self-assembling peptides with various functional motifs such as cell adhesion sites or protease sensitive sequences have been produced. Another type of polypeptides that forms hydrogels is derived from the Val-Pro-Gly-X-Gly pentapeptide repeat (where X is a guest amino acid other than proline) found in human tropoelastin. Called elastin-like-polypeptides (ELPs), the chains are soluble in aqueous solution, but as the solution temperature is raised, ELPs become insoluble and aggregate at a critical temperature, termed the inverse transition temperature. By changing the composition and chain length of guest residues, the transition temperature can be accurately tuned between 0 C and 100 C for specific applications such as recombinant protein purification or cell culture. For example, in vitro studies reported that ELPs have been established to support the synthesis and retention of cartilaginous matrix from encapsulated chondrocytes and adult stem cells. Likewise to self-assembling peptides, ELPs have also been customized with ECM ligands derivative from fibronectin, in order to support better cell attachment.

Synthetic Materials The immense opportunities for biomaterials design and functionality enabled by mimicking nature continue to stretch the limits of imagination. As both biological understanding and engineering capabilities develop, more sophisticated biomedical materials can be synthesized that have multifaceted chemical, biological, and physical characteristics designed to achieve specific therapeutic goals. The biological materials defined earlier assist as a point of departure for biomimicry, the use of synthetic materials to summarize salient materials features of natural ECM molecules, such as in situ crosslinking, demonstration of adhesion ligands, binding of growth factors, and susceptibility to cell-derived proteases. Synthetic analogs have several advantages. The methods used for matrix formation and functionalization are becoming progressively straightforward and easy to control. Several reactions can be performed under gentle, repeatedly physiological, circumstances that allow integration of cells or biological molecules with slight loss of function or viability. The use of completely synthetic materials removes the purification problems that can happen with naturally derivative materials as well as decreases the potential for an immune response or pathogen transmission. Lastly, polymeric-constructed hydrogels, which represent a huge class of synthetic materials used for this application, mimic the extremely hydrated viscoelastic properties of the natural ECM, permit for transport by diffusion and interstitial flow, and can present soluble, affinity-bound, or covalently bound biological features. Synthetic cell-responsive hydrogels for use in TE can be designed from a number of hydrophilic synthetic polymers or polysaccharides, with poly(ethylene glycol) (PEG), poly(vinyl alcohol), poly(hydroxyethyl methacrylate), alginate derivatives, and others, and have been reviewed comprehensively in the literature.

Mechanisms of hydrogel formation The mechanical and biological influence of hydrogels can be controlled spatially, so specific signals can be exhibited in a pattern that mimics the morphological gradients found in development and wound healing. The external tools like temperature and light further expand hydrogel manipulation to the temporal scale, so variations can be made to the structure and demonstration of biophysical signals on demand. By using light, biological signals can be committed to or released from a hydrogel, and structural crosslinks can be made or broken to correspondingly harden or deteriorate a material in real time. Light that has entered skin tissue and reached an underlying hydrogel has been used to expose cell-adhesive peptides, which was found to increase the number of inflammatory cells and increase fibrosis in a PEG implant. The vinyl sulfone adapted multi-arm PEG macromers (PEG-VS) to permit a Michael addition reaction among the acrylated PEG and a thiol group, which is utmost often offered as a free cysteine on a peptide or protein. And the PEG-VS are first functionalized with mono-cysteine species, like cell adhesion ligands, growth factor, and their ligands. Polymers can be cross-linked into hydrogel networks covalently by numerous mechanisms, which can be largely grouped into chain-growth polymerizations, such as photopolymerization, and step growth polymerizations, such as Michael addition reactions. An example of photopolymerization contains photo-introduced reactions between PEG diacrylate molecules or among thiolacrylates or thiolenes. The functionalized PEG-VS are then cross-linked into a biodegradable gel network consuming peptides that contain a protease-sensitive substrate sequence lined by cysteine encompassing domains. Michael adding reactions are not restricted to acrylates and also happen between maleimide and thiol groups, as has been revealed for the crosslinking of a p-maleimidophenyl altered dextran. Freshly, sequential copper-free click chemistry has been useful for the creation of hydrogels and their succeeding patterning with biomolecules. Exactly, an azide can react with an alkyne to practice a triazole using a di-fluorinated cyclooctyne moiety throughout hydrogel formation. Orthogonal thiol-ene photocoupling such as with peptides comprising the photoreactive allyl ester Fmoc-Lys-OH can be used to pattern biomolecules. Anseth’s group has used this click chemistry pattern to prepare PEG hydrogels with MMP-cleavable sequences and patterned combination of Arg-Gly-Asp (RGD) peptides to deliver localized cell attachment. In addition to Michael adding reactions and click chemistry, step-wise polymerizations can also be attained through enzyme-catalyzed crosslinking of peptide-functionalized materials. Sperinde and Griffith reported that transglutaminase to crosslink multifunctional glutaminyl PEG with polypeptides comprising alternating lysine and phenylalanine residues. Otherwise, a mixture of multi-arm PEG molecules, each conjugated with one of two different counter-reactive peptide substrates for factor XIIIa, can be cross-linked in the occurrence of this transglutaminase. Comparable to the method used to covalently functionalize fibrin matrices, biofunctional peptides labeled with one of this factor XIIIa. In addition to the covalent crosslinking mechanisms defined earlier, hydrogels can also be designed by physical or ionic connections between molecules. This behavior is detected in

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the self-assembly of peptide amphiphiles into fibrillar b-sheet structures, as defined earlier, and throughout the complexation of polymers or polysaccharides through ions.

Hybrid Materials Two or more materials combined and made composite material by merging often ones that have very dissimilar properties. The two materials work together to provide the composite unique properties. Though, within the composite you can simply tell the different materials separately as they do not dissolve or mix into each other. The major benefit of composite materials is that they are light as well as strong. By choosing a suitable combination of matrix and strengthening material, a new material can be made that accurately meets the requirements of a particular application. Composites also deliver design flexibility because many of them can be molded into complex shapes. Nanocomposites, high performance materials, display uncommon property combinations and unique design potentials. Nanocomposite is a multiphase solid material and it has one of the phases, has one, two, or three dimensions of < 100 nm, or structures having nano-scale recurrence distances between the different phases that make up the material. Nanocomposites are found in nature, for example in the structure of the abalone shell and bone. The use of nanoparticle-rich materials long predates the understanding of the physical and chemical nature of these materials. Jose-Yacaman et al. investigated the origin of the depth of color and the resistance to acids and bio-corrosion of Maya blue paint, attributing it to a nanoparticle mechanism. From the mid-1950s nanoscale organo-clays have been used to control flow of polymer solutions (e.g., as paint viscosifiers) or the constitution of gels (e.g., as a thickening substance in cosmetics, keeping the preparations in homogeneous form) although the resulting product is more efficient, the raw materials are often expensive. Although most materials used in TE scaffolds this far are in their pure form (single component), one polymer often does not meet all the needs for various TE applications. Natural bone matrix is a typical example of organic/inorganic composite material consisting of collagen and mineral (apatites). This natural composite material has an excellent balance between strength and toughness, superior to either of its individual components. As an example of biomimetic scaffolds, polymer/inorganic composite scaffolds will be reviewed below as advantageous scaffolds for bone TE. The term “composite” is usually referred for those materials in which the different phases are separated on a scale larger than the atomic, and properties such as the elastic modulus are significantly changed in comparison with those of a homogeneous material. The engineering of composites and nanocomposites draws on traditional characterization and processing technologies. Research describing structures comprising nanoparticles appears to trust on methods that are being pushed to the limit of resolution. Nanocomposites contrast from conventional composite materials due to the extraordinarily high surface to volume ratio of the reinforcing phase and its exceptionally high aspect ratio. Similar to the major inorganic component of natural bone, the inorganic compound such as hydroxyapatite or a calcium phosphate (CaP) in a composite scaffold delivers good osteoconductivity while the polymer offers the constant structure and design flexibility to attain the high porosity and high surface area, which are essential for anchorage-dependent cells such as osteoblasts to survive and differentiate. By blending and phase separation techniques, polymer/inorganic composite scaffolds have been established with enhanced mechanical properties and osteoconductivity. The HAPcomprising scaffolds progress osteoblastic cell seeding consistency and show significantly improved expression of mature bone marker genes such as osteocalcin (OC) and bone sialoprotein (BSP) above plain polymer scaffolds. Bone tissue formation throughout the scaffold has been established. Some other groups fabricated poly(lactic-co-glycolic acid) (PLGA)/HAP or PLGA/ PCL/HAP composite scaffolds using a technique called salt leaching. Though the salt-leaching technique has restrictions in making highly porous and well-connected pore structures, they also constantly established the enhanced osteoconductivity over the PLGA scaffolds. Another report shows that HAP in the composite scaffolds significantly progresses the protein adsorption capacity, inhibits apoptotic cell death, and offers a more favorable microenvironment for bone tissue regeneration. Considering that protein-scaffold and cell-scaffold contacts occur at the scaffold pore surfaces, a biomimetic method has been established to grow bone-like apatite nanoparticles on prefabricated porous polymer scaffolds in a simulated body fluid to efficiently modify the internal pore wall surfaces with bone-like apatite without changing the bulk structures and properties of the scaffolds. The apatite produced via the biomimetic procedure in an SBF is partially carbonated HAP more similar to the natural bone apatite than the stoichiometric HAP crystals. The moderately carbonated apatites must degrade quicker than the stoichiometric HAP crystals and assist as a better scaffold constituent in terms of new bone tissue remodeling. The growth of apatite crystals is significantly affected by the polymer materials, porous structure, and ionic concentration of the SBF as well as the pH value. To additionally mimic the nanofibrous organic collagen and the incompletely carbonated nanoapatite of the bone matrix, macroporous nanofibrous scaffolds were examined for bone-like apatite deposition. A huge number of nanoapatite particles were designed, and even a uniform and dense layer of nano-apatite was used to shelter the entire internal pore wall surfaces without clogging the macro-pores after a sensibly long time of incubation in an SBF.

Fabrication of Biomimetic Scaffold Previous TE scaffolds encompassing fibrous biodegradable polymer fabrics were formed using textile technologies. There are several polymers like poly(glycolic acid) (PGA), poly(lactic acid) (PLA), and other semicrystalline can be processed into fibers by textile technologies. PGA nonwoven scaffold is widely used in TE research. In a living system, cells collect a various signals and instructions through communication with the adjacent cells within tissues and, more significantly, cellular behaviors are directed by complex biochemical and biophysical information offered by the ECM, as demonstrated by cell fate determination and contact direction

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phenomenon. Originally cells located in a 3D cushioned network are protected against external mechanical stress. Furthermore, the ECM delivers a proper niche for cell adhesion, migration, proliferation, and differentiation due to molecular connections between specific cell membrane receptors and signaling cues from surrounding ECM resources. Consequently, for successful tissue regeneration at a target site, the architecture of scaffolds must emulate the natural design of the ECM and instruct cellular behavior while sufficiently housing the cells. TE scaffolds have been arranged using variety of materials and numerous processing conditions; employment of appropriate fabrication approaches is important for successful regeneration of target tissue.

Prefabricated preparation of scaffolds Polymeric biomaterial plays a significant role in today’s health care technology. Polymer hydrogels were the leading experimentally designed biomaterials for human use. Until recently, conventional methods to develop prefabricated scaffolds have targeted mostly on introduction of open pores and interlinked channels inside biodegradable scaffolds to induce the viability of injected cells. Several processing techniques have been established to fabricate biodegradable 3D foams for scaffolds. And there are two important methods, salt leaching and gas foaming, that have been extensively used because these methods are straightforward, cost-effective, and easy to scale up. Plentiful scaffolds prepared from these techniques have resulted in effective clinical outcomes and are ultimately expected to enter the commercial market. As an alternative, electrospinning permits nanoscaled fibrous design of scaffolds that mimic functional collagen structures. Electrospun nanofibers are characterized by multifaceted fibrous and interconnective porous structures, which make it possible to fabricate extremely structured scaffolds for making cellular growth. Microfabrication techniques are constructed on rapid prototyping (RP) methods including stereolithography, fused deposition modeling, shape deposition manufacturing, selective laser sintering, and 3D bioplotting, and these are measured as effective directions to fabricate custom-made scaffolds with a specified anatomical outline for hard tissue regeneration. Newly, modular assembly approaches have been used to develop biomimicking hybrid tissues with a synthetic architecture loaded with inductive biological signals and specified cells.

Porous scaffolds by conventional approaches Hierarchical computational techniques have been used to design of 3D anatomic scaffolds with porous architecture that maintains the function and mass transport. SFF has allowable fabrication of scaffolds with well-ordered architecture from polymer, ceramic, hydrogel, and even metal biomaterials. Porous scaffolds have considerably better mechanical properties than scaffolds processed using other methods. These enhanced mechanical properties are particularly important for bone-TE, which has much greater stiffness and strength than other tissues. However, even for soft tissue, traditional methods are followed to develop scaffolds that are not sufficient mechanically. The need for enough scaffold mechanical properties, joined with a wide range of scaffold base-material properties, necessitates the capacity to control scaffold mechanical properties through architecture topology design. Even though preliminary steps have been made in linking CTD and SFF, future work should determine how closely designed scaffolds can achieve desired mechanical properties as a function of material and SFF processing method. Biocompatible and biodegradable polymeric foams have been harnessed as temporal structural supports to regenerate various tissues such as bone, cartilage, nerve, ligament, skin, and liver. An open porous geometry with interconnected channels is a requirement for high density cell growth within the scaffold as well as the mass transport of nutrients, oxygen, and metabolic waste; a high cell density and efficient mass transport contribute to cell viability, proliferation, and ultimate treatment into functional tissues. Bioerodible aliphatic polyesters including PLA, PGA, and their copolymer, PLGA, have been extensively used to fabricate biodegradable scaffolds because of their convenient hydrolysis into nontoxic components in vivo. A wide range of biodegradable scaffolds with diverse morphologies have been fabricated by conventional methods such as porogen leaching, gas foaming, emulsion/freeze drying, and expansion in supercritical fluid. In particular, a porogen leaching method was primarily engaged to create a highly porous scaffold with large interconnectivity. In this method, particulate materials at a micrometric level including inorganic salts, carbohydrates, and paraffin spheres are usually embedded into a polymer/solvent paste, followed by elimination through washing after solvent evaporation. The level of porosity and the overall pore size and also the surface properties of the microspheres were delicately tuned. The mean pore size of the microspheres was adjusted from 7.9 to 29.4 mm by varying PLGA concentration in oil phase during double emulsion process. Moreover, the hydrophilic amino groups were introduced on the microsphere surface to facilitate the cellular adhesion and proliferation. When chondrocytes were cultured on the super porous microspheres, large amounts of deoxyribonucleic acids, collagen, and GAGs were produced, implying a high rate of cell growth and expression of a cartilage-specific phenotype. These porous spherical scaffolds could potentially be used to distribute therapeutic cells to damaged sites by injection.

Fibrous scaffolds by electrospinning Electrospinning is a well-known technique for the fabrication of nanoscale fibers. It continues to be studied widely due to its range of advantages such as high surface-to-volume ratio, tunable porosity, and ease of surface functionalization. Certainly, the resulting fibers are enormously useful for applications in TE, drug delivery, and wound dressings. This is an area with vast potential for controlled release research. As electrospun fibers mimic the ECM of tissues in terms of scale and morphology, there is the potential for them to be used as scaffolds. Electrospinning is a flexible tool for developing ultrafine fibers with a wide range of diameters from a few micrometers to several nanometers. The externally applied Coulomb forces are similar to the mechanical forces applied in traditional spinning. The processing flexibility of this technique ensures fiber production from a broad range of precursor materials that consist of synthetic polymers, semiconductors, natural polymers, ceramics, and their combinations. The final product produced

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by electrospinning is often in the form of a 2D nonwoven mesh characterized by arbitrarily oriented nano/micrometer-sized fibers. Electrospun fibrous meshes have attractive characteristics such as a wide surface-to-volume ratio and a highly interconnective porous architecture, and can be assembled into aligned fibers, making these meshes ideal for application in nanoscaled devices. Compared to the conservative methods producing porous materials, through past quite a few years, the electrospinning technique has received more attention since as a simple and powerful means to create nanometer-sized elements. With regard to biomedical applications such as TE scaffolds, drug delivery carriers, and wound care devices, these fibrous materials allow delicate modulation of cellular characteristic and fine control of drug release. Because of their structural and morphological resemblance to the native ECM, electrospun nanofibers have been exploited to attempt to produce an ideal TE scaffold. A multiple range of synthetic polymers such as biodegradable aliphatic polyesters and native biopolymers, for example, collagen, silk fibroin, chitosan, alginate, and hyaluronic acid, have been electrospun as biomimetic and temporal substrates to adapt a variety of cellular activities. For instance, the natural collagen fiber analog generated by electrospinning served as an excellent substrate for EC growth and osteoblast differentiation. Electrospun hyaluronic acid also facilitate chondrocyte growth while the cell retaining typical cartilaginous phenotypes, showing assure in the application of cartilage regeneration. The novel 3D collecting template is based on manipulation of electric fields and forces to direct the ultimate 3D architecture; this approach generated electrospun tubes with diverse tubular configurations.

Rapid prototyping-based microfabrication Rapid prototyping (RP) techniques are defined as automated evidence of each tomographic layer sequence based on automatic 3D images into the ultimate preferred architecture through additive layer-by-layer (LbL) fashion. And this RP technique is one of the most accurate and reproducible avenues for regulating the internal pore size, porosity, pore interconnectivity, mechanical performance, and overall dimensions of TE scaffolds. These features favor the scaffolds to be biomimetic to native tissues or organs with regard to shape and mechano-compatibility. Though, microfabricated products frequently lack the bioactivity to stimulate tissue regeneration. Because product fabrication is frequently implemented at the micrometer scale, integration of nanoscale features into the microstructure is essential to finely tune cellular responses. Photolithography is a technique to replicate various types of patterns through exposure of a photosensitive material via photomasking and vastly implemented for fabrication of microelectronics. However this approach is limited to photosensitive materials and only appropriate for fabrication on planar surfaces. Soft lithography produces an optional set of techniques for microfabrication that does not have these restrictions. Another study reported production of a dual-scale scaffold by combining the processes of RP and electrospinning. In this process, the microfibrous layer of the scaffolds was first built via the RP process and then polymeric nanofibers were directly deposited onto the microfibrous layer by electrospinning. The subsequent microfibrous layers integrated with electrospun nanofibers were constantly laminated onto the previously integrated layers so that a 3D hybrid structure could be fabricated. The hybrid construct consisted of a microfibrous wood pile structure and electrospun nanofibers from different biocompatible materials, namely poly (–caprolactone) (PCL) and collagen. The dually designed scaffold showed enhanced biological function in terms of chondrocyte adhesion and proliferation. In another study, a LbL polyelectrolyte assembly system was applied to a RP-based microstructured scaffold for application in bone TE. Hydroxyapatite and collagen are present in the extracellular components of natural bone tissue and show osteogenic effects on human mesenchymal stem cells. In the same study, they customized hydroxyl apatite nanoparticles with catechol-functionalized hyaluronic acid. The personalized hydroxyapatite nanoparticles with a negative charge and type I collagen with a positive charge were used to generate blocks for LbL assembly, which causes the formation of a nanocomposite multilayer in RP scaffold surface. While adjusting the amounts of hydroxyapatite and collagen on the surface, the degree of hMSC differentiation was observed by measuring alkaline phosphatase (ALP) activity and osteoblast-related genes expression level, including BSP-II, bone morphogenetic protein-2, osteopontin (OP), and OC. ALP activity and the relevant gene expression increased in a pattern consistent with the levels of the two surface components. The achievement of a biomimetic design approach that combines biomimetic structures with surface chemistry opens up new frontiers in the fabrication of multifunctional scaffolds and also offers flexibility in that the integration of other bioactive agents onto the scaffold such as bone growth factors could additionally promote bone tissue regeneration.

Modular hierarchical assemblies Until now, producing large engineered tissues using prefabricated scaffolds has been restricted by the nonhomogeneous allocation of cells within the scaffold and a shortage of vascularization, which leads to cell death in the core of the scaffold. To solve this problem, the approach of modular TE, which targets to assemble biological modules typically consisting of cell masses or cell/polymer constructs, was introduced. A modular scaffold is developed to maintain flexibility and adaptability for tissue reconstruction so that large complex tissues can be rebuilt by the connection of simple modular units in several different ways, and different modules with diverse functionalities can be integrated into the final construct. Multicellular spheroids can be used as modular building units for TE and were formerly self-assembled into a toroidal 3D structure using tissue liquidity mediated by cell–cell and cell–matrix interactions. Furthermore, organ printing is a programmed system that uses bio-ink to assemble modules into 3D functional living structures; for example, multicellular spheroids have been collected with the aid of ECM-mimicking hydrogels. Though, the formation of a stable and strong bioconstruct is a major challenge for successful implantation and functional retention of the engineered tissue. Previous study reported that the nanofilaments mentioned earlier were altered with the RGD peptide moiety and selfassembled with MSCs to formulate composite multicellular spheroids. And these composite spheroids presented improved hMSC viability and increased level of adipogenic differentiation compared to blank spheroids composed only of cells. These

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improved composite periods were used as prefabricated building units for adipose TE. Combinations of a suitable carboxylic acid and n-butylstannoic acid constitute modular gelation systems, in which the creation of a well-defined “tin-drum” nanocluster subsequently underpins the hierarchical assembly of nanostructured fibers, which form self-supporting gel-phase networks in organic solvents. Molecular complementarity is crucial in the development of chemical systems and determinations a significant number of outstanding complications in the emergence of complex systems. All physical and mathematical models of organization within complex systems depend on nonrandom association between components. Molecular complementarity offers a naturally occurring nonrandom linker. Most notably, the formation of hierarchically controlled stable modules massively progresses the probability of attaining self-organization, and molecular complementarity delivers a mechanism by which hierarchically systematized stable modules can develop. Lastly, modularity based on molecular complementarity creates a means for storing and replicating information. Microfibrous templates formed by the RP technique were used to mechanically stabilize the assembled tissue. The microfibrous templates were also changed with fibroblast growth factor (bFGF), an angiogenic growth factor, to stimulate neovascularization. These bioactive templates were consumed to support the 3D assembly of composite spheroids, yielding a biomimetic hierarchical assembly with solid multiscaled structures. Implantation of engineered tissue paradigm in the dorsal subcutaneous pockets of nude mice causes successful formation of active and vascularized adipose tissue, suggesting that retaining the recommended tissue contours and adipose tissue function, the crucial hurdles in regeneration of soft tissue were overcome. This assembly concept and modular design can be extended to reconstruct a variety of multifunctional tissues and organs.

Scaffold Functionalization Advances in biological and medical sciences, materials science, and engineering have triumphed in the past few decades in providing valuable implant materials for human tissue repair for millions of patients all over the world. Currently, several materials were used clinically for implants in orthopedics, dentistry, etc. which include metals, polymers, ceramics, and composites. And these materials served their roles moderately well; some of the currently used implant materials that were not initially developed for medical applications have shortcomings in their biological environment. Though many biodegradable polymers such as PLA, PGA, PLGA, and PCL have been used to create manageable structures for regulating the repair and regeneration of damaged tissues, active control of cell adhesion and downstream cellular events is still challenging due to the absence of biofunctional ligand that can instruct cells. As a result, many efforts have been made to integrate bioactive ECM molecules onto the scaffolds’ surfaces by several modification modalities including noncovalent and covalent binding. ECM proteins can be noncovalently adsorbed onto the surfaces of scaffolds by electrostatic interaction and van der Waals forces. To coat PLLA and PCL scaffolds fibronectin and laminin are used and have been shown to increase the cell adhesion. Furthermore, stem cells appear to favorably differentiate into osteoblasts depending on the type of ECM adsorbed to the scaffolds. Recently ECM microarray study examined the effects of combination of ECM molecules like collagen-I, III, IV, laminin, and fibronectin on selective differentiation of embryonic stem cells, and they also emphasized the importance of ECM proteins as immobilized biological cues on scaffolds. Although, physical adsorption is a comparatively weak force and may not be suitable for TE applications for which continued signaling are required. As a substitute, covalent immobilization of bioactive molecules has been attained by surface etching and plasma/gamma-ray treatment, which involves the cleavage of a degradable polymer backbone to generate carboxyl (–COOH) groups or radicals. These functional groups can undertake additional reactions with ECM proteins and related bioactive peptides (DGEA, RGD, IKVAV, YIGSR) on the surfaces of several scaffolds including PVDF, PLLA, PCL, and PLCL. The covalently conjugated molecules have been shown to be stable under physiological conditions and prolonged implantation in the localized microenvironment maintains the biological activity. Originally, these bioactive ligands enhanced cell adhesion to the surface and were involved in cell-specific signaling. Hydroxyapatite deposition technique on nanofibers using the CaP dipping method by mineralization. Unlike plasma-spraying technique, it does not include high temperatures, thereby overcoming the disadvantages attributed to the high temperatures used during the process, such as the possibility of fracture at the interface between the titanium and the hydroxyapatite due to the residual stress at the interface, and variations in the porosity, composition, crystallinity, and structure of the plasma-sprayed hydroxyapatite. So, new hydroxyapatite coating approaches have attracted great attention in recent years for substituting the plasma spraying high temperature techniques. For increasing the process of osseointegration, it is important that the initial cell capture ratio is high. Another work demonstrated that it is possible to capture MSCs on a substrate such as allograft bone by raising a system where it was able to capture MSCs on allograft bone with an enrichment factor of 3–4X at best. Fibronectin conjugated to PLLA, PCL, and silicone rubber scaffolds significantly improved the survival of osteoblasts, chondrocytes, and myoblasts, while collagenmodified PLLA scaffolds eased the chondrocytes proliferation. RGD peptide-immobilized scaffolds enabled the differentiation of osteoblasts and the myogenic differentiation of muscle precursor cells. Likewise, IKVAV and YIGSR have been described to regulate neural and EC differentiation. The density and distribution of bioactive molecules presented from the surface must be optimized for appropriate control of cell behavior. A high density of bioactive molecules on the surface is connected to greater cell spreading, focal contact formation, and proliferation to a certain extent, although cell migration is relative to protein density on the surface in a biphasic manner. Hence, a nominal amount of signaling peptide can be commonly used for surface immobilization and another study demonstrated that as low as 1 and 10 fmol cm 2 of RGD peptide was sufficient for cell spreading and focal contact formation. Furthermore, the peptide density and the distribution on the surface regulate integrin clustering. The size of integrins is approximately 10 nm and their clustered complex structure is < 100 nm; the distribution of peptides for cell-adhesive mimetics should

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be considered on a nanometer scale. Newly, an adaptable technique to immobilize bioactive molecules based on mussel adhesive proteins was reported. To establish the effectiveness of this technique, dopamine was polymerized onto metals, polymers, and ceramics, and then secondary ligands, for example, ECM bioactive molecules, were covalently immobilized to the polydopamine layer. Likewise, another study employed hybrid mussel adhesive protein attached with short ECM bioactive domains by recombinant DNA technology. These methods are effective in that surface immobilization can be carried out irrespective of the chemical composition of the materials.

Biomolecules Delivery Existing TE approaches focus on restoring impaired tissue architectures using biologically active scaffolds. The perfect scaffold would mimic the ECM of any tissue of interest, stimulating cell proliferation and de novo ECM deposition. A plethora of techniques have been assessed to engineer scaffolds for the organized and targeted release of bioactive molecules to deliver a functional structure for tissue growth and remodeling, in addition to improve recruitment and proliferation of autologous cells within the implant. The controlled supply of growth factors and cells within biomaterial carriers can increase and accelerate functional bone formation. The carrier system can be designed with preprogrammed release kinetics to provide bioactive molecules in a localized, spatiotemporal manner which is similar to the natural wound healing process. The carrier can also act as an ECM-mimicking substrate for stimulating osteoprogenitor cellular infiltration and proliferation for integrative tissue repair. Growth factors are key molecules that stimulate intracellular signal cascades required for tissue regeneration and also integrated into scaffolds where they can deliver positive biochemical signals ranging from promotion of mitogenic activity to stimulation of neovascularization. The growth factors used for creation of engineered tissues are mainly dependent on the target tissue types to be regenerated. Epidermal growth factor, transforming growth factor-b, fibroblast growth factors, and platelet-derived growth factor are frequently used to accelerate wound healing because of their mitogenic induction of epithelial cells and fibroblasts in addition to upregulation of matrix formation, for example, chemical immobilization of recombinant human EGF (rhEGF) on an electrospun scaffold was effectively engaged for wound healing. The rhEGF-adapted nanofibrous scaffold preserved the growth factor activity even under harsh conditions such as enzymatic degradation, subsequently inducing keratinocyte differentiation. Bone morphogenic protein-2 and TGF-b are often essential to stimulate and preserve tissue-specific properties for regeneration of hard tissues such as bone and cartilage. Generally used methods for combining growth factors and other cues involve covalently tethering them to the scaffold material these methods have been used before to stimulate ES cell differentiation and survival. Although such strategies are extremely effective at retaining such cues inside scaffolds, covalent linkages can impede internalization and, in order, affect the activity and function of growth factors. Affinity-based supply systems have been discovered as alternative methods of retaining growth factors inside scaffolds via noncovalent interactions. Growth factor is a soluble signaling molecule that binds with transmembrane receptors on target cells and controls variety of cellular functions. The ultimate biological response elicited from a growth factor depends on the uniqueness of the growth factor and target cell, receptor type, cell number, and other signaling events. So, an important component in designing an organized delivery system is selection of the appropriate single or combination of growth factors for maximized tissue repair. The TE scaffolds can be fixed with or without embedded cells. Previous study reported that cell-loaded scaffolds displayed adverse effects leading to reduced healing processes as a result of immune responses activated by the ECM materials secreted from the preseeded cells. The use of cell-free scaffolds may be more needed in certain cases. When using scaffold alone method, the scaffold has to be particularly adapted with bioactive molecules to induce the infiltration of peripheral progenitor cells or multipotent stem cells. The regenerative conductivity or inductivity directed by the scaffold is mainly due to the composition of the material, which can distribute biochemical cues such as growth factors. For bone lesions repair, growth factor releasing scaffolds have been engaged to ameliorate osteogenesis and initiate the vascularization, both of which can be controlled by the dose and release kinetics of the growth factors. Particularly, collagen sponges combining recombinant human BMP-2 (rhBMP-2), advertised as the INFUSE Bone Graft, are clinically used for treating degenerative disc diseases. In a human clinical trial, the rhBMP-2/collagen sponges showed more reliable osteoinduction than autogenous bone grafts. The lack of autologous cell sources for tissue reconstruction has led to attempts to use pluripotent or multipotent stem cells. Exactly, MSCs are most often used as another cell source for reconstructing connective tissues such as cartilage, bone, adipose, and tendon. Some studies that have used progenitor cells or stem cells for TE have stated that scaffolds loaded with growth factors result in a high degree of cell differentiation. Characteristically, continued release of growth factors and a short diffusion distance between abound growth factor and a target cell are mandatory to maximize the activity of the growth factors because of their short biological half-lives in the body, for example, gelatin microspheres loaded with TGF-b were used to stimulate the assembly of hMSCs into cellular aggregates. Dispersion of TGF-b released from microspheres into the hMSCs of cellular aggregates induces chondrogenesis, as measured by the amount of DNA, GAG, and collagen type II formed. In addition, two types of electrospun nanofibers with dissimilar bFGF release profiles and surface hydrophilicity were matched by measuring the differentiation level of hMSCs. Stem cells grown on electrospun fibers for fast release of bFGF and a more hydrophilic surface presented better collagen production and upregulated gene expression, consistent with fibroblastic differentiation, suggesting that the release kinetics of the growth factor as well as the surface properties of the scaffold could change stem cell fate. Moreover, proper infiltration of blood vessels permits long-term survival of the surrounded tissue construct by stimulating transport of indispensable nutrients, oxygen, and cell-signaling molecules. Angiogenic growth factors including acidic or basic FGF, vascular endothelial growth factor, angiopoietins, and PDGF can be combined into scaffolds to facilitate host tissue in growth and induce angiogenesis.

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PLGA foams fabricated by using solvent casting with salt leaching or supercritical carbon dioxide have been engaged to straight encapsulate bFGF. Protein release was fairly continuous, but an initial burst and comparatively low loading competence were unavoidable. So, heparin and heparin sulfate have been demoralized to immobilize bFGF on the surfaces of various scaffolds on the basis of the strong binding affinity between heparin and bFGF. The heparin facilitated supply of bFGF has separate advantages in terms of protection of early degradation of growth factors and facilitation of cell recognition, eventually resulting in better angiogenesis. Combination of chemically heparinized foams and bFGF or VEGF have showed to be an effective delivery platform for growth factors, with steady and continued release of growth factors over a stretched period and consequently superior angiogenic activity. Heparinized electrospun fibers also displayed a high degree of bFGF surface immobilization and more prolonged release of the growth factor than that of bare electrospun fibers. Likewise, laminin and bFGF were simultaneously adsorbed on heparinized nanofibers. The combination of immobilized biochemical cues, that is, laminin and bFGF, as well as nanofiber alignment, synergistically induced oriented neurite outgrowth and cell migration. Furthermore to localized introduction of a single growth factor, distribution of multiple growth factors in immediate or sequential release mode can enhance treatment efficiency. A polymeric system for dual growth factor delivery was established for therapeutic angiogenesis. Different release kinetics was detected for VEGF-165 and PDGFBB provided from a single PLGA scaffold; implantation of this scaffold induced the rapid formation of a mature vascular network in a rat model.

Importance of Bioreactors In regenerative medical therapy, TE approach has been proposed to fulfill the neotissues demand via in vitro tissue regeneration. In this scenario, bioreactors are the key strategy to translate cells/tissue-based constructs under a controlled environment, into lowcost, large-scale products that are clinically safe and biologically effective. Several experimentation conditions can be optimized including temperature, pH, oxygen tension and perfusion of cells, external stimuli, etc. in order to provide physical and biochemical regulatory signals desired for cellular proliferation and differentiation. For instance, in vivo cells will respond to mechanical stimulation which bioreactors can supply, aiding cells to produce ECM within shorter time span under optimal mechanical stiffness. The mechanical stimulation has also been shown to stimulate stem cells down different lineages, enabling the formation of different cell types. The bioreactors’ designs are specific to desired tissue and application; however some of the common basics should be followed: (i) They should be easily assembled and efficient in tissue formation in a short time span while keeping the products sterile. (ii) The materials must be nontoxic which will be employed in designing bioreactors. This has ruled out most metals as they have the potential of releasing ions into the media which could be highly toxic to cells in most cases. (iii) The bioreactor should come with a sensor for accurate monitoring of culturing conditions. The principle of an in vivo bioreactor is to implant a well-designed TE scaffold with or without cell seeding onto a well-vascularized site of a living organism and harvest when matured. However, future trends of in vivo bioreactors lie with implantable devices which are able to monitor and control the in vivo culturing condition and scaffolds with higher biocompatibility that is suitable for long-term implantation with minimal rejections.

Biomimetic Scaffolds for Regenerative Engineering Bone tissue Bone formation involves three types of bone cells, that is osteoblasts, osteocytes, and osteoclasts that begin with sequential events of proliferation, differentiation, matrix formation, and finally mineralization. Importantly, bones are made of remarkably strong and flexible tissues that are capable to remodel itself perpetually. However, to cure/repair any damage/defects TE has emerged as an extraordinary approach for the treatment with the use of designed nanostructured materials, especially osteoconductive biomaterial scaffolds along with osteogenic cell populations and osteoinductive bioactive factors. Mainly, biomimetic scaffold materials have been used to understand the osteoblastic differentiation of progenitor cells. These ideal biomaterial scaffolds should act as a carrier for cells and growth factors and also should induce formation of bone with surrounding tissues after implantation. In search of perfect biomaterial, various biomaterials including CaP ceramics, bioresorbable polymers, bioactive glasses, functionalized hydrogels, and composites are extensively used to hold cells for bone regeneration. Recently, it has been found that the ECM glycoproteins are expressed during new bone formation and OP, thrombospondin, and BSP as matrix proteins were important in bone cell migration, proliferation, matrix deposition, and mineralization. Hence, biomimetic scaffolds’ integration into bone tissue takes place at the material–tissue interface, which results in initial cell and substrate interactions. Thus, it is important to design bone matrix proteins modified biomimetic scaffolds for excellent cell–matrix interactions to promote bone growth. Initially, incorporation of bioactive peptides protein fragments, that is arginine–glycine–aspartic acid (RGD) is recognized by the cell’s transmembrane integrin receptors, which is the most common adopted strategy to enhance functionality. The composite of RGD and rosette nanotube (RNT) was prepared with a unique surface chemistry with favorable cytocompatibility properties for bone repair. In this study, RNT’s incorporation into a biocompatible hydrogel resulted in a biologically inspired 3-D nanoscale scaffold with RGD-functionalized self-assembled biomimetic features, which was more suitable for 3-D bone regeneration. Moreover, most of the studies were reported utilizing RGD modification on traditional materials on 2-D flat substrates. RGD-containing peptides bound to neonatal calvarial osteoblasts using membrane receptors were reported. The effect of RGD peptides was

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monitored on the specific interaction with bone-derived cells in vitro, where a 15 amino acid sequence containing RGD derived from BSP was immobilized on glass plates. The seeded osteoblasts on the modified surfaces for 2 h showed enhanced cell attachment strength and promoted focal contact formation. In a study, it was reported that osteoblast adhesion for 4 h in serum-free medium was enhanced only on an Arg-Gly-Asp-Ser (RGDS) immobilized surface; however the proliferation rate after 3 days on peptide-modified surfaces was not significantly different than on control surfaces. They also showed the effect of the growth factor OP-1 in conjunction with a cell-binding peptide. OP-1 exhibited synergistic enhancement of mineralization of the osteoblasts cultured on RGD-modified surfaces. Additionally, the translation of these results in vivo remains challenging since a number of ECM proteins are involved in wound healing and bone forming processes. Biomimetic properties of biodegradable polymers were enhanced by chemical modification with RGD-containing peptides. Previously, silk-based materials were used for modification of polymers with RGD-containing peptides to investigate the effect on the long term growth of human osteoblast like cells. The modified silk-based scaffolds showed that immobilized peptides promoted expression of 1-ethyl-3-(dimethylaminopropyl) carbodiimide hydrochloride (EDC) proteins and calcified bone nodule formation compared to nonmodified control surfaces. In another report, biodegradable polymers of PLA and PLGA were coupled with RGD peptides using poly(L-lysine) and fabricated as scaffolds for the growth of osteoblasts. The designed polymers showed enhanced initial cell attachment and also showed significant expression of osteogenic phenotype. RGD peptides modification over biodegradable polymers is advantageous and more preferred over the other model surfaces, that is quartz, glasses, and metals. On the other hand, synthetic biodegradable polymers are getting more attentions due to their specific interactions of incorporated peptides with activated receptors and also useful to study osteoinductive abilities through in vivo implantation. To overcome the challenges faced during the application of biomimetic scaffolds to bone defects with an irregular shape, RGD peptides were further tethered to unsaturated polyester and subsequently crosslinked to form solid constructs, which showed potential as injectable application. Additionally, covalently linked biomimetic hydrogels with RGD peptides synthesized using various macromolecular structures presented good bioactive signals and also showed the capability of modulating marrow stromal osteoblast adhesion. Recently, heparin-binding domain in FN was used as a substitute for RGD peptide sequences, which has ability to be recognized by polysaccharide molecules in the cell membrane and hence can be used to design biomimetic materials for bone TE. In recent report, pectin-based injectable biomaterials were used for bone TE. They used pectin for the first time to modify RGDcontaining oligopeptide and used as an ECM alternative for bone tissue regeneration. The immobilized MC3T3-E1 preosteoblast cells viability and differentiation were monitored through viability, metabolic activity, morphology, and osteogenic differentiation tests. The results showed that preosteoblast cells immobilized in both types of pectin microspheres maintained a constant viability up to 29 days and were able to differentiate. Also, grafting of the RGD peptide on pectin backbone induced improved cell adhesion and proliferation within the microspheres. The cells were grown inside and outside of the microspheres and organized themselves in the 3D structures that produced a mineralized ECM, suggesting the potential of pectin as an injectable cell vehicle for bone tissue regeneration. More recently, functionalized d-form self-assembling peptide hydrogels were utilized for bone regeneration. They demonstrated the utilization of functional motif RGD to modify peptide D-RADA16 for designing and fabricating peptide DRADA16-RGD. Also, they incorporated the angiogenic growth factor bFGF into these peptide hydrogel scaffolds. Finally, they applied the self-assembling peptide D-RADA16-RGD hydrogels (with or without bFGF) to repair the bone defects of SD rat femoral condyle and evaluated the osteogenic ability of D-RADA16-RGD hydrogels. The obtained results showed enhanced extensive bone regeneration.

Cardiovascular tissue The aim of cardiac TE development is the development of bioengineered biomimetic materials to facilitate and provide the physical support and help cardiac regeneration. Mainly, the removal of the damaged cardiac tissue is done by replacing certain functions of the damaged ECM that avoids adverse cardiac remodeling and dysfunction after myocardial infarction. However, the cardiac regeneration is facilitated with the help of biomimetic materials that support cardiomyocyte differentiation and expansion both in vivo and in vitro, support and protect injured cells into the heart and repair them, and guide the formation of regenerative heart tissues through engineered and/or native regeneration. For effective regeneration of cardiac tissues, various critical factors (i.e., replacement of cardiomyocytes, support of electrical conduction, and mechanical force of contraction) must be maintained or established. Presently, nanotechnology has offered tools to design or bioengineer various biomimetic materials that mimic nature or draw inspiration from the nature. These designed biomimetic materials are extensively explored in the field of cardiovascular science and provided opportunities for cellular transplantation using polymeric scaffolds to repair damaged cardiac tissue. Numerous studies were reported to explore the ability of biomimetic materials for creating tissues to support different functions. For treating patients with valvular cardiovascular disease, mostly heart valve augmentation or replacement methods are utilized. However, such valve should be designed using noninflammatory, biocompatible, and nonthrombogenic materials which can grow in pediatric patients. Additionally, the mechanical and structural properties are two important factors needed for heart valve scaffold material replacement. For all these problems, biomimetic materials development is necessary. To replace the damaged native tissues during cardiac or peripheral bypass surgery, synthetic cardiovascular grafts are used when there is limited availability of healthy autologous tissues/ vessels. The lower patency rate of synthetic vascular grafts as compared to natural tissues is mainly due to thrombus formation and intimal hyperplasia following the implantation. However, difference in the mechanical compliance of synthetic grafts compared to the surrounding native tissues caused other complications to patients. Another important limitation is small diameter (< 6 mm) of

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synthetic grafts which causes increased failure rate. Hence, attempts are made to design the vascular grafts with enhanced mechanical stability and biocompatibility. The functional importance of biomimetic cardiovascular grafts relies on the anatomical structure and the biological function of blood vessels. These blood vessels are made of three different tissues layers which play critical role in blood circulation, maintain tensile strength and viscoelasticity of the blood vessel due to collagen and elastin structural components, and provide smooth tissue with long-term durability. Other elements that present an ideal vascular graft need to be designed with having important properties, that is, infection resistance, bio-compatible, suturability, and off-the-shelf availability. Endothelial cells (ECs) seeding is another approach in which vascular grafts are engineered onto synthetic materials to render the luminal surface antithrombogenic. In this approach foreign cells cannot be used as they may elicit an immune response. To avoid this drawback, patient’s own cells are harvested for seeding of the graft lumen before implantation. Typically, cell culture techniques are also introduced in case of limited supply of ECs. However, transplanted ECs have drawback of detachment from the surface during the surface exposure to blood circulation. The biomimetic materials mimicking the fibrillar architecture of the ECM have been used to modify the surface of cardiac specific cells (vascular grafts). The phase separation, electrospinning, and selfassembly methods are used to develop the nano and submicron polymeric fibers. Recently, the fibronectin and RGD-containing peptides were coated on an expanded polytetrafluoroethylene (ePTFE) in order to improve ECs attachment. In this case, the fibronectin was properly adsorbed to the polymer surface and RGD-containing peptides were covalently immobilized. The saphenous vein ECs were seeded on coated ePTFE for 30 min and exposed to the shear stress in an artificial flow circuit, which showed enhanced cell attachment and retention by coating with fibronectin as well as immobilizing RGD-containing peptides. Saphenous vein ECs incubated on the RGD peptides modified ePTFE surface showed excellent cell attachment and retention. In another study, biomimetic polyurethane-based vascular grafts were designed that showed poor ECs adhesion by themselves. Hence, RGDcontaining sequences incorporated were introduced into polyurethanes using chemical and photochemical modification methods. Improved ECs adhesion and proliferation were observed with these RGD containing sequences modified polyurethanes. However, RGD modification might increase platelet adhesion during blood circulation in the human body due to platelets also expresses integrin receptors recognizing RGD sequences. On the other hand, platelet adhesion to RGD-modified surface may also reduce the patency rate of the vascular graft; thereby ultimately it will lead to the failure of the graft. The signaling peptides, that is TyrIle-Gly-Ser-Arg (YIGSR) and Arg-Glu-Asp-Val (REDV) were explored due to their specific interactions with ECs to enhance the endothelialization that is needed for a nonthrombogenic vascular surfaces generation. In a study, YIGSR was immobilized on the surface of poly(ethylene terephthalate) and glass to improve the EC attachment and spreading. The migration of ECs study showed significant increase in the persistence of cell movement that resulted in the random motility coefficient of ECs. Further, REDV peptides derived from III-CS domain of human plasma fibronectin were used to study the interaction of a receptor present in ECs which was due to the peptide surface modification that resulted in good attachment of ECs and spreading. The vascular grafts were modified to modulate ECM protein production during tissue formation and control cellular interactions. The ECM protein production is critical to generate robust vascular grafts, tissue viability, and provide mechanical integration with adjacent tissues. Additionally, the produced ECM must substitute fabricated biomimetic materials for myocardial regeneration and new tissue formation. As it is reported, ECM production by SMCs and ECs is dependent on the density and type of signaling peptides attached onto the glass surfaces. This process can be assessed by the incorporation of 3H-glycine to the expressed proteins. Study showed that the initial cell adhesion on modified surfaces with Val-Ala-Gly-Pro (VAGP), RGDS, Val-Ala-Val-Ala-Gly-Pro (VAVAGP), and Lys-Gln-Ala-Gly-Asp-Val (KQAGDV) was increased. However, the total matrix production was decreased on these surfaces after 7 days as compared to nonmodified surfaces. Further, the ECM production was increased with the addition of TGF-b in the media. This result demonstrated that growth factors and signaling peptides are needed in order to optimize matrix production. Growth factor such as TGF-b1 was also tethered to PEG-based hydrogels through covalent binding and hence increased hydroxyproline production was obtained compared to unmodified hydrogels.

Conclusions and Future Prospects With nanoscience and nanotechnology driving the revolution in the field of materials science and engineering, enables to design and fabricate novel biomaterial scaffolds incorporating various biomimetic features at the nanometer scales, molecular, and genetic level. Owing to the complex and dynamic nature of cell responses toward regulatory signals, imaging of engineered biomimetic constructs during cultivation immensely improves our understanding of cellular responses to micro-environmental parameters with spatial and temporal specificity. Thus, with biology providing fundamental design requirements for regenerative medicine, engineered tissues can acts as models for developmental studies, disease modeling, toxicity screening, and many more.

Acknowledgments This research was partly supported by the Bio & Medical Technology Development Program of the National Research Foundation (NRF) (grant number: 2012M3A9C6050204) funded by the Korean government (MEST) and Technology Commercialization Support Program, Ministry for Food, Agriculture, Forestry and Fisheries (grant number: 814005-03-3-HD020), Republic of Korea, and as well supported by a grant of the Korea Health Technology R&D Project through the Korea Health Industry Development Institute (KHIDI) (grant number: HI15C2996), funded by Ministry of Health&Welfare, Republic of Korea.

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Further Reading Amini, A. R., Laurencin, C. T., & Nukavarapu, S. P. (2012). Bone tissue engineering: recent advances and challenges. Critical Reviews in Biomedical Engineering, 40, 363–408. Cordonnier, T., Sohier, J., Rosset, P., & Layrolle, P. (2011). Biomimetic materials for bone tissue engineering – state of the art and future trends. Advanced Engineering Materials, 13, 135–150. Currie, L. J., Sharpe, J. R., & Martin, R. (2001). The use of fibrin glue in skin grafts and tissue engineered skin replacements: a review. Plastic and Reconstructive Surgery, 108, 1713–1726. Dhivya, S., Ajita, J., & Selvamurugan, N. (2015). Metallic nanomaterials for bone tissue engineering. Journal of Biomedical Nanotechnology, 11, 1675–1700. Dunn, D. A., Hodge, A. J., & Lipke, E. A. (2014). Biomimetic materials design for cardiac tissue regeneration. WIREs Nanomedicine and Nanobiotechnology, 6, 15–39. Gong, T., Xie, J., Liao, J., Zhang, T., Lin, S., & Lin, Y. (2015). Nanomaterials and bone regeneration. Bone Research, 3, 15029. Haghi, A. K., Oluwafemi, O. S., Josmin, P., & Hanna, J. M. (2013). Composites and nanocomposites. Apple Academic Press, 230. Jennifer, P., Mikael, M., & Jeffrey, A. (2010). Biomimetic materials in tissue engineering. Materials Today, 13, 14–22. Liu, Y., Luo, D., & Wang, T. (2016). Hierarchical structures of bone and bioinspired bone tissue engineering. Small, 12, 4611–4632. Meyer, R. A., Sunshine, J. C., & Green, J. J. (2015). Biomimetic particles as therapeutics. Trends in Biotechnology, 33, 514–524. Peter, X. (2008). Ma biomimetic materials for tissue engineering. Advanced Drug Delivery Reviews, 60, 184–198. Prabhakaran, M. P., Venugopal, J., Kai, D., & Ramakrishna, S. (2011). Biomimetic material strategies for cardiac tissue engineering. Materials Science and Engineering: C, 31, 503–513. Pradhan, S., Hassani, I., Clary, J. M., & Lipke, E. A. (2016). Polymeric biomaterials for in vitro cancer tissue engineering and drug testing applications. Tissue Engineering Part B: Reviews, 22, 470–484. Sears, N., Dhavalikar, P., Whitely, M., & Cosgriff-Hernandez, E. (2017). Fabrication of biomimetic bone grafts with multi-material 3D printing. Biofabrication, 9, 025020. Shotorbani, B. B., Alizadeh, E., Salehi, R., & Barzegar, A. (2017). Adhesion of mesenchymal stem cells to biomimetic polymers: a review. Materials Science and Engineering: Materials for Biological Applications, 71, 1192–1200. Shoulders, M. D., & Raines, R. T. (2009). Collagen structure and stability. Annual Review of Biochemistry, 78, 929–958. Stevens, M. M. (2008). Biomaterials for bone tissue engineering. Materials Today, 11, 18–25. Taek, G. K., Heungsoo, S., & Dong, W. L. (2012). Biomimetic scaffolds for tissue engineering. Advanced Functional Materials, 22, 2446–2468. Titorencu, I., Albu, M. G., Nemecz, M., & Jinga, V. V. (2017). Natural polymer-cell bioconstructs for bone tissue engineering. Current Stem Cell Research & Therapy, 12, 165–174.

Bioreactors: System Design and Application for Regenerative Engineering Antonio Valdevit, Stevens Institute of Technology, Hoboken, NJ, United States; and SEA Limited, Columbus, OH, United States © 2019 Elsevier Inc. All rights reserved.

Introduction Bioreactor Considerations Biological Element Selection Biological Environment Mechanical Environment Fluid Flow Gas/Fluid Pressure Substrate Selection Basic Description and Operation for Design of Feedback-Controlled Bioreactor Employing Direct Mechanical Loading Single-Cell Loading Bioreactor Optimization of Loading Compressive Loading and Imaging Further Reading

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Nomenclature a Solenoid radius (m) b Drag Force (N) |BZ | Magnetic field (T) Fmag Magnetic field force (N) I Current (A) L Solenoid length (m) Lw Length of wire (m) N Solenoid turns v Velocity (m/s) m Mass of microsphere (kg) rp Particle radius (m) h Fluid viscosity ¼ 8.9  10 4 Pa s m0 Permittivity of free space ¼ 8.85418782  10 12 m 3 kg 1 s4 A2

Introduction The term bioreactor can be rather confusing due to the varied configurations and widespread applications. Most simply described, a bioreactor is a self-contained entity capable of growing and/or sustaining the biological functions of living organisms. The organisms may be cells, viruses, bacteria, or live tissue. Regardless of organism, the appropriate environment for growth and sustainment of function is required and, more importantly, is specific to the organism if proliferation of the organism and sustainment of biological function are to be achieved. The environment is not solely based on the biofluidic or media conditions. The mechanical conditions also play an important role, especially in the case of biological material comprising the musculoskeletal system. Methods by which mechanical conditions may be generated with a reactor include fluid flow to induce shear, while gas pressure may be used to apply compression. More recently, direct methods of compressive load application with respect to individual cells have been realized using magnetic beads. Another aspect for consideration in bioreactor design is the substrate or layer where the biological element will reside. With the exception of blood cells, most biological material requires a platform upon which to adhere. The adherence permits proliferation and, when conditions are appropriate, normal biological function. The substrate may be rigid, smooth, porous, or even flexible so as to permit application of tensile forces by stretching of the substrate. In short, the design of a bioreactor is an attempt to mimic the normal biomechanical and physiological environment of biological material that may be subsequently perturbed using physical and biochemical signals in an attempt to understand physiological responses and function.

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Bioreactor Considerations In the design of any device involving biological tissue interaction and in the specific case of a bioreactor, it is vital that the design process be parsed into less daunting engineering tasks. When each element has been addressed, one can then initiate the more difficult problem of integrating the various components. It should be noted that assuming that one can simply add the components would be a naïve approach to a bioreactor design. The design process is iterative. One may find that some elements may require modification in order to be integrated into the final design. Furthermore, one must be open to compromises. Bioreactor design and application is a field that is at the forefront of basic science research. As such, components or materials to achieve an optimal design may not be available, may require invention, or may be cost prohibitive. To minimize effort, one should do a thorough literature review to investigate what components, materials, and environmental conditions have been used with success by other investigators employing similar or identical biological material. Conducting such a task will reduce the incidence of selecting a material or environmental condition that has been shown less efficacious or even detrimental in your specific application.

Biological Element Selection The selection of cells and tissue for study can be as varied as the design of bioreactors. Depending on the nature of the study, appropriate cell or tissue selection is required in order to replicate the biomechanical or physiological function in musculoskeletal tissue. One may consider using osteocytes or osteoblasts in the study of bone proliferation and mineralization. In the case of connective tissue, such as cartilage or tendons and ligaments, chondrocytes or fibrocytes are an appropriate biological selection. Alternatively, stem cells may be employed under different mechanical and environmental conditions to study the effects of the cellular environment upon differentiation. Cell use, be it specific or undifferentiated stem cells, may be either human or animal. Prior to selecting a specific cell type, one should conduct a literature review in order to identify if a specific cell type, cell line, or species has been associated with successful culturing and sustainment in the intended application.

Biological Environment The biological environment may be one of the most challenging aspects in any bioreactor design and use. Regardless of cellular type selection, consideration for nutrients, waste disposal, and maintenance of the biological environment are crucial. For example, if one designs a reactor encompassing a scaffold where the media flow is essentially stagnant, most likely many cells residing in the central region of the scaffold will likely be exposed to a reduced oxygen concentration due to physical barriers restricting diffusion and a lack of fluid flow. Under these conditions, cells near the center of the scaffold may experience an elevated level of cellular death as compared to those in the surface of the scaffold where nutrient and oxygen levels are significantly increased as compared to the center of the scaffold.

Mechanical Environment All cells respond to mechanical stimulation. Obviously, cells that constitute bone tissue are the most responsive as evidenced by Wolff’s law. Wolff’s law is not to be taken as a physical law such as Newton’s laws of motion. That is, Wolff’s law is an observation that states that bone will remodel so as to minimize the loading stresses upon the structure. Consequently, increased loading will increase bone remodeling and strength. Reduced loading will reduce the bone volume and hence bone strength. However, cellular response to the mechanical environment is not isolated to osteoblasts, osteoclasts, and osteocytes, which are located in bone. Endothelial cells which constitute lumen as in blood vessels, arteries, and walls of the GI tract are also subjected to a mechanical environment. The difficulty in applying mechanical stimulation in a bioreactor setting lies with orientation, magnitude, and frequency. Different cell types are subjected to multiple combinations of these parameters during different activities and, as such, require a mechanical environment to replicate the physiological state. There are several methods available to establish a mechanical environment. These include use of fluid flow, gas pressure generation, and physical mechanical loading.

Fluid Flow The shear stress (s) upon the cells at the base of the flow channel is given by s ¼ 6mS and/WH2, where m is the fluid viscosity, S is the flow rate, and W and H are the width and height of the channel, respectively (Fig. 1). One must verify that the flow is laminar in order to achieve constant fluid shear stress. This can be somewhat ensured by several considerations. The dimension of the width should be much greater than the area where the cells are located as well as be greater than the height. The length of the channel should be sufficiently long as compared to the region of cellular attachment so as to facilitate uniform flow by avoiding the turbulent regions at the inlet and outlet of the channel. Alternatively, another configuration allows for differential shear stresses to be displayed based on the location from a central inlet. Under this geometric configuration, the shear stress (s) is given by s ¼ 6mS/ 2pDH2, where m is the fluid viscosity, S is the flow rate, D is the distance from the central inlet, and H is the height of the channel

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Fig. 1

Fluid flow reactor mechanism for constant fluid stress due to laminar flow.

Fig. 2

Fluid flow reactor configuration to permit differential fluid stress based on distance from central axis of the inflow.

(Fig. 2). This fluid flow geometry permits examination of the cellular response to shear stress over a range of stress values using a single experiment since the fluid stress varies as the reciprocal of the distance, D, from the inlet.

Gas/Fluid Pressure While fluid flow for shear stress is common, in general they are restricted to small volumes on the order of microscope slides. Cellular volumes may be increased through the use of gas/fluid compression geometries. In these cases, cells to be loaded are placed at the base of a chamber that is filled part way with media. The gas phase above the media is then pressurized by depression of an actuator which decreases the gas volume and hence pressurizes the liquid below (Fig. 3A). The advantages of such an apparatus include ease of use and reduced variability. However, it should be noted that gas exchange due to cellular activity within the fluid media will eventually reach equilibrium and, with continued use, result in a depletion of nutrients and required oxygen for cellular proliferation. These configurations are suitable for short-term and acute experiments. To reduce the complexity and retain reproducible loading, a similar configuration employs a flat compression platen placed upon cells embedded within a matrix or scaffold. Such a static loading condition allows for controlled loading (Fig. 3B). In addition, the simplicity of the mechanism can permit multiple devices to be used with various combinations of matrix/scaffold and platen mass in order to investigate the effects of load and 3-D cell geometry. This particular method can suffer from two disadvantages. The first challenge is that cells located near the surface of the matrix/scaffold will see nutrient media at a significantly higher level as compared to those cells embedded deep within the structure. Such a condition can lead to oxygen and nutrient depletion within the central regions of the matrix/scaffold configuration. A second concern when using this geometry is the effect of Poisson’s ratio. This is especially significant in the cases of soft materials. Examples include experiments involving cartilage, muscle, meniscus, tendons, and ligaments. With these tissue and matrix configurations, compressive loading will lead to lateral expansion of the material. When this occurs, the cells will not only be subjected to compressive loading but also experience shear due to the movement of the matrix/scaffold perpendicular to the direction of loading. Characterization of the shear component of the overall stress will require computational fluid dynamics simulations.

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Fig. 3 (A) Compression of the gas above the fluid phase with a plunger or an actuator serves to compress the gas and in part compressive forces through the fluid upon the cells located at the base. (B) Direct compression of the scaffold or matrix containing the cells can be performed using a plunger or an actuator. One must be aware of the Poisson’s ratio effect due to compression of the scaffold or matrix.

These illustrations display some of the configurations possible under which cells may be subjected to stresses. When designing a novel or traditional loading configuration, it is important to keep in mind that the resulting apparatus should attempt to replicate the conditions in vivo with respect to applied loading, environmental fluid, and substrate. One should also keep in mind that these techniques are generally applied over shorter durations and that long-term experiments will require mechanisms for media replenishment and cellular waste removal.

Substrate Selection Substrate selection is a vital element in bioreactor performance. The surface upon which cells are seeded can not only influence cell migration but also elicit an effect upon physiological function of the cells. Cell types generally display a preferential affinity for certain substrate surface composition and surface morphology. One must consider the hydrophobic/hydrophilic effects of the surface upon the residing cells. Furthermore, surface texture plays an important role. Cells migrate across the surface by elongating or extending the region in response to a preferred interaction with the substrate surface. At some point, the elongated region becomes anchored or adhered to the surface of the substrate. Subsequently, the remaining material of the cell is dragged along the surface through retraction (Fig. 4).

Fig. 4 A graphical representation of cell motion across a substrate. Extension of the cellular membrane is followed by anchoring at a specific distance from the cell and then proceeded by retraction of the cellular body towards the anchor location.

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With the advent of 3-D printing, microcontoured surfaces are possible and play a significant influence in cell migration. In general, cells prefer a substrate for adhesion. Exceptions to this include red blood cells or other cells found in the bloodstream. A more recent element with respect to surface effects has been demonstrated with the implementation of electric fields in various orientations across the surface seeded with cells. Under these conditions, cells tend to align themselves with the field. In general, 2-D considerations as outlined above are less likely to result in cellular apoptosis as waste and nutrient fluids can flow freely and can be replenished with fresh fluids as needed. In the case of 3-D structures or scaffolds, the likelihood of cell death is increased as one approaches the center of the construct. While the cells residing on the surface are exposed to free-flowing and/or fresh media, those cells residing deep within the structure can only obtain nutrients and dispose of waste through diffusion within the structure and, as such, are subjected to depleted nutrients and oxygen, as well as an elevated waste environment. For 3-D structures, the material should provide sufficient porosities for ample diffusion to occur (Fig. 5). In addition, pulsatile flow around the structure could be used to enhance fluid flow within the structure and hence decrease the likelihood of cell death. However, one must keep in mind that the applied loading upon the structure may not be the resultant load upon the cellular material contained within. Consider a 3-D scaffold capable of sustaining several thousand Newtons in compressive load. Imparting 10– 20 N of load as is commonly cited in the literature will result in virtually no loading upon the cellular material within the scaffold. While most scaffolds are confined to basic science research or to nonload-bearing applications, with the advent of 3-D printing, the realization of surgically sized, physiologically load-sustaining scaffolds is possible. In the case of the examples presented for spinal fusion and segmental bone replacement applications, the scaffolds are approximately 30 mm in diameter and 10 mm in height. The spinal scaffold consists of interwoven hyperbolic struts in concentric patterns adjusted for decreased modulus transition as one traverses from the periphery to the central region of the scaffold (Fig. 6A). The segmental bone replacement scaffold consists of intertwined conduits arranged to mimic the plywood configuration of bone. This dual modulus design facilitates the low modulus cancellous bone in the center and the increased high-strength modulus of cortical bone at the periphery (Fig. 6B). Both designs are 3-D printed from biodegradable polylactic acid (PLA). These unique geometries manifest themselves with failure loads of 4900 and 9600 N for the spinal and bone replacement scaffold, respectively. While these static properties are essential, both scaffolds also sustain runout loads of 800 and 1100 N to 5 million cycles, respectively, without cracking. With respect to biological response, in the case of the spinal device, alkaline phosphatase activity (indicative of proliferating osteoblast cells) displayed a significant increase after only 14 days when stimulated as compared to unstimulated controls (Fig. 7A). In the case of the segmental replacement scaffold, the calcium colorimetric assay evaluating osteoblast mineralization was manifested significantly greater than stimulation frequencies of 2 Hz as compared to unstimulated controls with 0.5 and 5 Hz frequencies. This 2 Hz frequency continued to manifest elevated mineralization throughout 21 days of stimulation as compared to the other frequencies examined (Fig. 7B). These two scaffolds represent the use of 3-D printing in the case of surgical sized implants. As such, in order to maintain cellular viability at the center of these devices, many of the concepts presented will be applicable. When considering 3-D materials evaluations to characterize the mechanical response of the material under static and dynamic loading are required to ensure appropriate loading of the biological material when conducting in vitro experiments. Often, loading a single structure is unfeasible, and as such one would like to take advantage of loading multiple 3-D structure simultaneously. To achieve this requires fabrication of appropriate fixtures that can adjust to specified strains or stresses prior to initialization of

Fig. 5 Regardless of scaffold design or geometry, fluid flow throughout the scaffold is important if one is to avoid cellular apoptosis in the center of the scaffold due to buildup of waste materials and lack of nutrient and oxygen flow.

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pulsatile loading. This is most easily accomplished using individual plungers of known mass to reside upon each 3-D structure (Fig. 8). A plate is fitted over all the plungers and secured to each plunger using fasteners. Such a procedure will ensure that each 3-D structure is initialized with the plunger mass as initializing stress or strain. Subsequent loading will then ensure that each 3-D structure is imparted with the same loading profile. Using commercial materials, testing frames in conjunction with bioreactor vessels such as the one described here can facilitate accurate dynamic and sinusoidal loading at controlled levels. While the bioreactor vessels may themselves be 3-D printer or fabricated by traditional means, the use of commercial testing frames becomes not only impractical, but also a significant expense. Furthermore, to employ these vessels in these testing frames requires transport from sterile incubator conditions to testing facilities which may lead to contamination and variability during the course of a long-term experiment. The ultimate solution to such a challenge lies in the development of a bioreactor that can employ feedback control in order to reproducibly and accurately apply sinusoidal loading conditions. The subsequent section of this guide will provide a description for operation and design for the elements required for such a device.

Basic Description and Operation for Design of Feedback-Controlled Bioreactor Employing Direct Mechanical Loading At the heart of any loading system is a transducer. Transducers can come in a variety of geometries and technologies in order for one to evaluate physical quantities. In general, when transducers are coupled to signal conditioning circuits, an electrical output based

Fig. 6 (A) A surgical sized spinal fusion scaffold capable of sustaining physiological loading can be 3-D printed from polylactic acid (PLA). The varying strut geometry permits a gradual modulus change from a high-strength periphery to a reduced modulus core. (B) In the segmental replacement of long bone, a high-modulus periphery is required to replicate cortical bone. Centrally, a reduced modulus core permits fluid flow perpendicular to a radial within the scaffold.

Fig. 6

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(continued).

ALP activity in stimulated hFOB cells P 50 years. In comparison with the arthroscopic repair, mini-open rotator cuff repair resulted in superior repair integrity and shoulder function. However, the correlation between the rotator cuff integrity and patient satisfaction and function remained controversial after repair. Moreover, arthroscopic surgery was twice as expensive as open surgery initially, but no significant difference in overall cost-effectiveness was found after 2 years. Despite the advances in arthroscopy, miniopen rotator cuff repair remains a useful technique. Furthermore, a systematic review aimed to determine the effectiveness of early or conservative rehabilitation for patients after surgical repair of rotator cuff tears. The study showed no difference in function, pain, range of motion, or retears ratio between early and conservative rehabilitation. However, early mobilization may be beneficial for small and medium tears. Moreover, a metaanalysis of RCTs and non-RCTs compared the clinical outcomes of autograft with allograft tendons in patients with posterior cruciate ligament (PCL) regeneration. PCL plays an important role in maintaining the stability of the knee joint. PCL rupture can lead to meniscal and articular cartilage damage, which can accelerate degeneration of the knee joint. The results obtained in the meta-analysis data suggest that there were equivalent effects on the Lysholm knee function score in patients who underwent autograft versus allograft PCL reconstruction. In addition, the patients with autograft tendons had a higher Tegner activity scale than those with allograft tendons. Altogether, the most important finding of this study is that the application of an autograft does not bring about more successful outcomes than an allograft. Achilles tendon rupture is another tendinopathy associated with physical exercise and age-related degeneration that affect both athletic and general population. Multiple RCTs have shown conflicting results while comparing surgical with nonsurgical treatment for acute Achilles tendon rupture. Based on a systematic review of overlapping metaanalyses, the current best evidence shows that surgical treatment may be preferred at centers that do not have functional rehabilitation, and nonsurgical intervention may be the best choice at centers offering functional rehabilitation. Moreover, four RCTs involving 169 patients were conducted to determine whether augmented repair provides better clinical benefits compared with nonaugmented repair. This metaanalysis suggested that no statistical difference was found between augmented and nonaugmented repair groups in terms of infection rate, patient satisfaction, and rerupture rate. Further, prospective randomized studies need to be performed to assess clinical outcome of these two techniques for acute Achilles tendon rupture.

Skeletal Muscle Skeletal muscle is one of three types of muscle that comprises approximately 45% of human body mass and is essential for generating forces that are involved in movement and locomotion. It is a highly organized structure encompassing nerves, blood vessels, myofibers, and extracellular connective tissue. In a minor injury situation such as muscle strain, the skeletal muscle can self-repair through the activation of muscle satellite cells (MuSCs, muscle stem cells). The formation of myofibers and their integration into muscle tissue results because of the proliferation and differentiation of satellite cells into myoblast, in response to a muscle injury (Fig. 1). The stem cell niche (interaction between the satellite cells and their environment) plays a key role in the muscle regeneration process. The current treatments include applying ice and resting the strained tissue for a few days. Also, nonsteroidal anti-inflammatory drugs (NSAIDs) and acetaminophen help in reducing pain and swelling. A systematic review and metaanalysis of 41 studies performed between 1985 and 2015 reported that the use of NSAIDs is more effective in the treatment of lower body muscle injuries when compared with upper body injuries. Additionally, after a short-term acute muscle injury, NSAIDs reduced pain, blood creatine kinase level, and strength loss. Severe damage to skeletal muscle such as traumatic injuries, myopathies, and tumors can lead to an irreversible loss of muscle mass. So, the formation of fibrotic scar tissue in the affected tissue after injury prevents the muscle regeneration. When severe trauma occurs, the current treatment includes the implantation of healthy, vascularized, and innervated autologous tissue (well known as muscle flaps, isolated from the vicinity of the injury) in the damaged area. However, the limitations of this treatment are the donorsite morbidity and long periods of rehabilitation, leading in most of cases to muscle regeneration with poor functionality and strength.

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Fig. 1 (A) Immediate response to muscle injury. Upon injury, MuSCs activate, self-renew, proliferate, and differentiate into myoblasts. Fibroadipogenic progenitors (FAPs) are also activated and secrete transient extracellular matrix. Debris is removed by the infiltration of neutrophils in the injury tissue. (B) Regeneration in young niche. Recruitment of neutrophils and FAP apoptosis is induced by the action of pro-inflammatory M1 macrophages. Also, angiogenic and anti-inflammatory factors are secreted by anti-inflammatory M2 macrophages. The process of differentiation and fusion of myoblasts is essential in the new muscle fibers formation. (C) Regeneration in aged/pathologic niche. Stiffening is the result of changes in the MuSC niche composition. The inflammation process persists due to the dysregularization of the inflammatory response. Fibrosis is produced by the FAPs’ oversecretion of ECM. Reproduced from Han, W. M., Jang, Y. C. and García, A. J. (2017). Engineered matrices for skeletal muscle satellite cell engraftment and function. Matrix Biology 60–61, 96–109.

Ligaments Ligaments are connective tissue comprising fibrous bands that connect the bones together at the joints and support some internal organs. Dense bundles of collagenous fibers and fibrocytes (bone marrow-derived mesenchymal progenitor cells that are involved in the vascularization and healing processes) are the main components of ligaments. Ligaments can be classified in two types: (1) white ligament composed by collagenous fibers that are not stretchy, very strong, and allow subjective freedom movement and (2) yellow ligaments rich in elastin fibers that are very tough and elastic. The anterior cruciate ligament (ACL) is one of the most important ligaments that play an important role in the knee stability. The ACL does not self-repair due to its avascular nature, that’s why some treatments like surgical reconstruction and rehabilitation are needed for ACL restoration. In the United States alone, > 200,000 people suffer ACL rupture each year, so conventional treatments are being practiced using surgical reconstruction with/without allograft or autograft. Two clinical studies were done to compare outcome between patients treated with rehabilitation alone with early ACL reconstruction and patients with rehabilitation and optional delay ACL reconstruction for 5 years. One hundred twenty-one cases were studied, 62 patients with early ACL reconstruction and 59 with delay ACL reconstruction. In all cases, patient received similar rehabilitation. The studies showed that patients who opted for delayed ACL reconstruction had greater instability in the knee joint compared with the early ACL reconstruction group. However, the study also concluded that both groups (early and delay ACL reconstruction) did not show any radiographic differences and thus did not provide better results after 5 years. Another study of 3–5 years clinical trial compared autograft with allograft ACL reconstruction, including 64 patients. Symptoms like activity level, physical examination, and knee testing were recorded. Overall, the International Knee Documentation Committee ratings and Cincinnati knee score showed no clinical statistical difference between groups. This study provided evidences for using allograft as an alternative reconstruction treatment. Recently, a 10 years clinical trial compare allograft with autograft for primary ACL reconstruction, involving 99 patients. There were 4 autograft (8.3%) and 13 allograft (26.5%) failures that required further revision and

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reconstruction. The results concluded that there was no significant difference in the scoring (Tegner or International Knee Documentation Committee ratings), but patients with allograft implant showed three times higher failure rate compared with those who received the autograft implants. In order to overcome the limitation of allograft and autograft, a different approach using synthetic and natural materials was explored for ACL reconstruction patients. Clinical trial using the artificial ligament, GORE-TEX (made by polytetrafluoroethylene (PTFE)) showed only one prosthetic breakage in 30 patients after 2 years. Further, in another study, the implantation of PTFE ligaments resulted in improvements of objective and parameters after 18 months. However, after a certain period, GORE-TEX graft was withdrawn from the market because the implantation of these materials resulted in complication in patients such as synovial reaction and inguinal lymphadenopathy. On the other hand, polyester like polyethylene terephthalate (PET) was employed as synthetic ligament and upon implantation in 130 sportsman showed good results in terms of joint stability, and complete integration in the host tissue was observed after 2 years. Besides, in another study, the implantation of PET in 47 patients improved subjective parameters, Tegner activity level, and higher stability.

Regenerative Engineering Approaches for Soft Tissue Regeneration Adipose Tissue Regenerative Engineering The regenerative engineering of adipose tissue is the field that addresses the pathologies of adipose tissue and the clinical needs. The use of new mechanisms of tissue regeneration based on specific cell precursors and new technologies for biomaterials production may address the limitations of conventional treatments (Fig. 2). Some studies have described the development of different cell-laden artificial extracellular matrices (ECMs) for adipose tissue regenerative engineering applications with promising results. In this sense, synthetic and natural polymers have been proposed to manufacture 3-D biodegradable structures. Key factors such as pore size, total porosity, and pore interconnectivity are essential properties necessary to produce 3-D structures that determine the transport of nutrients and metabolites that in the end modulate cell colonization and extracellular matrix production. Among synthetic polymers, poly(lactic acid) (PLA), poly(glycolic acid) (PGA), and poly(lactic-co-glycolic acid) (PLGA) have been used as cell delivery platforms. Specifically, insulin, insulin-like growth factor-1 (IGF-1), basic fibroblast growth factor (bFGF)-loaded PLGA-polyethylene glycol (PLGA-PEG) microparticles, and insulin and IGF1-laden PLGA-PEG microparticles/PLGA scaffold constructs were implanted in a subcutaneous rat model, enhancing the free fat graft volume and weight. In another case, after 35 days of in vitro culture, 3T3-L1 cells (adipogenic cell line) seeded within PGA scaffolds showed a lipid accumulation with mature and unilocular morphology. Finally, the formation of mature fat pads was observed after 35 days of subcutaneous implantation in nude mice. Moreover, the natural polymers offer a good alternative for tissue regeneration due to their similarity with some biomacromolecules present in the in vivo environment and with the ECM. In addition, natural polymers possess good biocompatibility, and they may prevent foreign body reactions, frequently detected with synthetic polymers. Collagen, hyaluronic acid (HA),

Fig. 2

Strategy for adipose tissue regenerative engineering combining adipose-derived cells with engineered scaffolds.

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gelatin, and fibrin have been widely used. For instance, subcutaneous implantation of preadipocytes/collagen porous constructs in immunodeficient mice showed new vessels and adipose tissue formation. Decellularized adipose tissue (DAT) is another natural substrate widely used for adipose regenerative engineering. DAT matrices are obtained through a decellularization process of adipose tissue into acellular 3-D porous substrates. Recently, with the development of new technologies, engineered adipose tissue constructs have been produced using 3-D bioprinting technique. Well-defined 3-D printed tissue constructs have been obtained by a combination of DAT (as bioink) and human adipose tissue-derived mesenchymal stem cells (hASCs). Good cell penetration, tissue remodeling, and adipose tissue formation were observed after 2 weeks of in vivo subcutaneous implantation of the DAT/hASC constructs in nude mice.

Skeletal Muscle Regenerative Engineering New muscle regeneration strategies have emerged to overcome the limitation of the current treatments. The design and development engineered muscle constructs based on cell seeding into 3-D scaffolds until the tissue becomes functional and their subsequent transplantation into the affected tissue in patients is the main concept for muscle regenerative strategies. Recently, some studies have shown promising results. For instance, mesenchymal stem cell (MSC)-laden alginate cryogel (with porous size between 70 and 150 mm) has been investigated as multifunctional substrate. The engineered material was used as MSC artificial niche and, at the same time, as IGF-1 and vascular endothelial growth factor (VEGF) release platform to enhance the MSC paracrine effect on muscle progenitor’s cells. 7, 28, and 56 days after the implantation of the artificial niches in a muscle trauma model in rats, the muscle strength improved, reducing considerably the fibrosis, and muscle fiber density increased. Another study reported the fabrication of mouse mesoangioblast-laden PEG-fibrinogen hydrogels. After the implantation of the constructs under the skin on the surface of the anterior tibial muscle, the attraction of host vessels and nerves induced by the mesoangioblasts was observed, as well as the formation and maturation of aligned myofibers. Finally, an artificial muscle like normal tibialis anterior resulted after replacing the ablated tibialis anterior with a mouse mesoangioblast-laden PEG-fibrinogen construct. Keratin-based hydrogels have been widely used as cell/growth factor delivery platform for functional muscle regeneration. Skeletal muscle progenitor cells, IGF-1, and/or bFGF-laden keratin hydrogels were implanted in a murine model of volumetric muscle loss injury to the latissimus dorsi (LD) muscle. The influence of topographical cues like groove on the surface of different materials including HA and poly(L-lysine) polyelectrolyte complex, poly(dimethylsiloxane) (PDMS), engineered gelatin methacrylate, and polystyrene substrates on muscle derived cells has been investigated. For instance, myoblast alignment and myotube formation were observed on grooved micropatterned PDMS substrates. Interestingly, additional cells added on top of the previous myotube layer followed the same pattern, so the cells attached, aligned, and fused formed a 3-D tissue multilayer patch.

Tendon Regenerative Engineering The use of cell niche matrices is the new regenerative engineering approach to speed the healing process of rotator cuff tendon injuries. The main concept is to use them as tissue bridge between the tendon and the bone facilitating cell growth and collagen deposition. Oriented electrospun nanostructure scaffolds have been produced to mimic: (1) the tendon extracellular matrix (scaffolds composed by fibers with diameter in a range of 50–500 nm, like collagen structure) and (2) the biomechanical properties of the rotator cuff. Scaffolds of poly(lactic-co-glycolic acid) (PLGA) and polycaprolactone functionalized with poly [(ethyl alanato)1 (p-methyl phenoxy)1] phosphazene (PNEA-nPh) fabricated by electrospinning technique are promising nanofiber matrices for tendon repair. Specifically, when the polycaprolactone/PNEA-mPh hybrid scaffolds were implanted in a rat model of rotator cuff augmentation, the morphology of regenerated tendon was like the suture repair group (see Fig. 3). However, the implantation of rat MSC/polycaprolactone/PNEA-mPh constructs resulted in a tissue morphology like the intact tendon indicating a faster tendon

Fig. 3 (A) Randomly oriented PLGA scaffolds (nanofibers diameter between 400 and 800 nm). (B) Repair augmentation using PLGA scaffolds. Reproduced from Taylor, E. D., Nair, L. S., Nukavarapu, S. P., McLaughlin, S. and Laurencin, C. T. (2010). Novel nanostructured scaffolds as therapeutic replacement options for rotator cuff disease. Journal of Bone and Joint Surgery 92, 170–179.

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remodeling process. Only traces of donor rat MSC were observed after 6 weeks suggesting that the observed therapeutic effect is due to the delivery of growth factors, immunomodulators, etc. MSC-laden PLGA scaffold implanted in rabbit model of an infraspinatus tear resulted in the formation of Sharpey fibers and fibrocartilage. In groups with cell-material constructs, the resultant enthesis had better tensile strength and greater collagen fiber formation compared with controls (cell-free scaffolds). On the other hand, several growth factors can augment rotator cuff healing process. bFGF-loaded PLGA electrospun scaffolds were employed in a rat model of a chronic rotator cuff tear. During the initial stage of the healing process, local delivery of bFGF induced high fibrocartilage formation and better collagen fiber organization compared with controls (bFGF-free scaffolds).

Ligament Regenerative Engineering The strategy of ligament regenerative engineering involves the use of biocompatible and biodegradable biomaterials with the necessary mechanical properties and the cell source, preferably primary cells isolated from autologous healthy ligament tissue. Ideally, the development of scaffolds as cell support, cell maturation in vitro followed by the implantation of the engineered constructs in patients, is the principal aim of this concept. From the clinical point of view, the principal advantages are the simple surgical procedures, minimal patient morbidity, good mechanical stability, and minimal risk of disease transmissions or infections. Ligament– bone interface is another important aspect to consider. This is a multilayered transition zone involving tissues with different mechanical properties that’s why the fabrication of constructs, able to mimic the ligaments in vivo, is a challenge for regenerative engineering field. Natural and synthetic materials have been extensively used for ligament reconstruction. For example, collagen and collagenplatelet sponges were used to regenerate the anterior cruciate ligament (ACL). The scaffolds were implanted in a minipig model of lateral ACL transection (ACLT). Normal suture repair procedure was used as control (see Fig. 4). Thirteen weeks after the implantation of collagen sponges, mature fibroblast and vascular tissue were observed in the healing area, and no inflammatory cells and

Fig. 4 Gross appearance of control group (suture repair procedure) (A) and scaffold groups (B). In both cases, the repair tissue happens from the anterior slope of the tibial spines to the intercondylar lateral wall. Change in cartilage structure was not observed. Individual synovialized band was observed in the repair tissue. Reproduced from Fleming, B. C., Magarian, E. M., Harrison, S. L., Paller, D. J. and Murray, M. M. (2010). Collagen scaffold supplementation does not improve the functional properties of the repaired anterior cruciate ligament. Journal of Orthopaedic Research 28, 703–709.

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residues of collagen sponges were detected. Collagen sponges alone were not sufficient to enhance the healing process; however, the combination of collagen and platelet improve the tissue repair. Silk is another interesting material that has widely been used as ligament replacement substitute. Tensile strength and toughness are its main advantages compared with other biomaterials from natural sources. MSC-loaded braided silk scaffold constructs were used to repair the ACL using a porcine model. Twenty-four weeks after implantation, the new regenerated ligament was visually like the ACL of the control group. Implanted materials were almost degraded, and a layer of fibrous tissue was formed due to cell infiltration and ECM production. Besides, positive type I collagen and almost negative type III collagen and tenascin-C markers were found in the newly regenerated tissue. Furthermore, poly(hydroxyesters) like PLLA and PGA have been employed to reconstruct the ACL due to their biocompatibility, degradation rates, and good mechanical properties. In an interesting study, a million of primary rabbit ACL cells were seeded on PLLA braided scaffolds and implanted in a rabbit ACL model. Twelve weeks after implantation, the constructs showed a dense connective tissue and vascularization throughout the replacement, production, and deposition of collagen fibers surrounding the PLLA fibers. The tensile mechanical properties of rabbit ACL and ligament replacements were measured. Significant differences between the strength retention of construct and cell-free scaffolds were observed at 12-week time point. In conclusion, the combination of cells and biomaterials yielded better results than cell-free materials because of better cell infiltration along the scaffold and more mature collagen remodeling during the transfer of load to the neoligament. The challenging regeneration of soft tissues can thus be taken to a next level by making advancement in tissue engineering by integrating it with areas of advanced materials sciences, physics, stem cell science, and developmental biology. The emergence of new regenerative approaches would bring a paradigm shift in the soft tissue regeneration field.

Further Reading Benjamin, T., Corona, S., & Greising, M. (2016). Challenges to acellular biological scaffold mediated skeletal muscle tissue regeneration. Biomaterials, 104, 238–246. Choi, J. H., Gimble, J. M., Lee, K., Marra, K. G., Rubin, J. P., Yoo, J. J., Vunjak-Novakovic, G., & Kaplan, D. L. (2010). Adipose tissue engineering for soft tissue regeneration. Tissue Engineering. Part B, Reviews, 16, 413–426. Escobar Ivirico, J. L., Bhattacharjee, M., Kuyinu, E., Nair, L. S., & Laurencin, C. T. (2017). Regenerative engineering for knee osteoarthritis treatment: Biomaterials and cell-based technologies. Engineering, 3, 16–27. Han, W. M., Jang, Y. C., & García, A. J. (2017). Engineered matrices for skeletal muscle satellite cell engraftment and function. Matrix Biology, 60–61, 96–109. Harris, K., Driban, J. B., Sitler, M. R., Cattano, N. M., & Hootman, J. M. (2015). Five-year clinical outcomes of a randomized trial of anterior cruciate ligament treatment strategies: An evidence-based practice paper. Journal of Athletic Training, 50, 110–112. James, R., & Laurencin, C. T. (2014). Musculoskeletal regenerative engineering: Biomaterials, structures and small molecules. Advances in Biomaterials, 123070, 12 p. Kwee, B. J., & Mooney, D. J. (2017). Biomaterials for skeletal muscle tissue engineering. Current Opinion in Biotechnology, 47, 16–22. Peach, M. S., Ramos, D. M., James, R., Morozowich, N. L., Mazzocca, A. D., Doty, S. B., Allcock, H. R., Kumbar, S. G., & Laurencin, C. T. (2017). Engineered stem cell niche matrices for rotator cuff tendon regenerative engineering. PLoS ONE, 12, e0174789. Pei, B., Wang, W., Fan, Y., Wang, X., Watari, F., & Li, X. (2017). Fiber-reinforced scaffolds in soft tissue engineering. Regenerative Biomaterials, 4, 257–268. Taylor, E. D., Nair, L. S., Nukavarapu, S. P., McLaughlin, S., & Laurencin, C. T. (2010). Novel nanostructured scaffolds as therapeutic replacement options for rotator cuff disease. Journal of Bone and Joint Surgery, 92, 170–179. Yang, G., Rothrauff, B. B., & Tuan, R. S. (2013). Tendon and ligament regeneration and repair: Clinical relevance and developmental paradigm. Birth Defects Research Part C: Embryo Today, 99, 203–222.

Characterizing the Properties of Tissue Constructs for Regenerative Engineering Yusuf Khan, University of Connecticut Health Center, Farmington, CT, United States © 2019 Elsevier Inc. All rights reserved.

Introduction Evaluating the Mechanical Properties of Tissues and Tissue Constructs Mechanics of Materials Anisotropy Viscoelasticity Methods for Testing Mechanical Properties of Biological Tissues and Synthetic Materials Compression and tensile testing 3-Point and 4-point bending Torsion testing Indentation testing Evaluating Tissue Construct Biocompatibility and Cell-Construct Interactions Toxicity Cell viability and proliferation Differentiation In Vivo Models for Tissue Construct Evaluation Hard Tissue Defect Models Soft Tissue Defect Models Conclusion References Further Reading

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Introduction At the intersection of tissue engineering and regenerative medicine lies regenerative engineering. Defined as “the integration of tissue engineering with advanced material science, stem cell science, biophysical stimulation, and areas of developmental biology,” regenerative engineering describes the newest strategies for replacing biological tissues. It leans heavily on the latest developments in materials science, stem cell biology, and clinical translation and as places importance on understanding how these materials, cellular processes, and their clinical translation potential can be measured in terms of mechanical properties, cellular biocompatibility, cell-material interactions, and preclinical in vivo testing. Later, we discuss methods of evaluating hard and soft tissues, materials, cell behavior and cell-material interactions, and preclinical animal models that can help pave the way from benchtop to bedside.

Evaluating the Mechanical Properties of Tissues and Tissue Constructs Properly matching the mechanical properties of the tissue to be regenerated has long been a fundamental design criterion for tissueengineered constructs. Whether replacing hard tissues like bone or soft tissues like ligament, tendon, muscle, cartilage, or blood vessel, properly matching the mechanical properties is essential. In hard tissues like bone, a failure to match the compressive modulus of the construct to that of the surrounding bone can lead to failure of either the construct or the tissue around it. Engineering a construct with inferior mechanical properties can jeopardize the mechanical integrity of the injury site by chancing catastrophic failure or simply eliminate that construct from consideration of such load-bearing defect sites. Over-engineering a construct such that the mechanical properties are considerably higher than those of the surrounding tissue can lead to a phenomenon called stress shielding in which the mechanical stresses applied to the defect area are absorbed and solely supported by the construct that is mechanically overmatched to the surrounding bone tissue, shielding the surrounding tissue from physical loading and stimulating the body to resorb the bone tissue from those unloaded areas. Since the body responds to physical loading by stimulating the production of more bone, it contrarily responds to a lack of physical loading, as would be experienced in stress shielding, by resorbing the unloaded bone. So regardless of whether the mismatch under- or over-compensates for the mechanical strength of the surrounding bone, it leads to hindrances in healing. The same mismatch of mechanical properties can be seen in soft tissues like ligament, which has a very specific s-shaped stress–strain curve that has to be mimicked in an engineered construct, and blood vessel, where the compliance of the synthetic vessel should best approximate that of a native vessel to minimize the disruption of blood flow as it passes from native vessel to synthetic conduit and back to native vessel. Disruptions in this flow regime can lead to changes in the forces applied to the vessel wall stimulating overproliferation of cells near the vessel-construct anastomosis,

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a phenomenon called neointimal hyperplasia. This can cause the lumen of the vessel and/or construct to become partially or completely obscured. Below is a summary of the mechanical properties of some hard and soft tissues and the mechanical properties of some of the more common materials used to repair or restore these tissues. Of note is the dramatic mismatch in these properties of many of these materials, clarifying the importance of novel approaches to designing constructs. It is important to note that for some materials there is a very wide range of mechanical properties. This is the case for several reasons. For the human tissue, the mechanical properties of a particular tissue, skin for instance, can vary considerably depending on where it was harvested prior to testing. The skin on the plantar surface of the foot, for instance, is considerably different than that of the face, for instance. The wide range of properties seen in the synthetic materials may be less intuitively obvious, but these properties can vary based on material testing protocols, processing techniques, synthesis techniques, modifications to materials, and the inherent variability that exists within materials of similar type. While this variability in material properties may initially appear as a challenge to the regenerative engineer when making choices for the design of a construct, it is actually an advantage that a material with very similar characteristics can be formed to have such a wide spectrum of mechanical strength (Table 1).

Mechanics of Materials Anisotropy Beyond compression and tension are a number of other mechanical characterizations that are relevant to tissues and should guide the development of tissue replacement constructs. Many tissues, for instance, exhibit anisotropic mechanical properties. Anisotropy describes the property of a material that exhibits different properties in the x, y, and/or z direction. Mechanical anisotropy suggests that a material compressed or placed in tension in one axis would not necessarily display the same mechanical response in a different axis. For example, a natural ligament that is tested in tension along its longitudinal axis would prove to have dramatically stronger mechanical properties than that same ligament tested in tension across its transverse axis. Some biological materials exhibit anisotropic mechanical properties like ligament, tendon, and cortical bone, while others exhibit more of an isotropic mechanical profile like cartilage (to some degree) and trabecular bone. A microscopic examination of the hierarchical structure of these tissues often reveals a structural basis for either isotropy or anisotropy.

Viscoelasticity Traditionally materials under mechanical loading express some amount of deformation in response to applied load, or strain in response to stress. The material is loaded until failure and can be represented by a simple stress–strain curve (Fig. 1). Many soft tissues within the human body, however, respond to mechanical loading in a way that resembles both a fluid and a solid, or as having a combination of viscous and elastic mechanical properties, respectively. These viscoelastic tissues respond differently in traditional tensile testing depending on how parameters like strain rate are varied in the testing. Strain rate, the rate at which the crosshead moves in a tensile testing machine, can result in different mechanical properties for viscoelastic tissues and therefore should be chosen carefully depending on the tissue, or engineered construct, being tested. Other testing method variations can reveal tissue behaviors under mechanical loading such as stress relaxation, creep, and hysteresis. Stress relaxation occurs when Table 1

Mechanical properties of biological tissues and synthetic materials used to regenerate those tissues

Tissue/material type

Ultimate tensile strength (MPa)

Elastic modulus in tension (MPa)

Ultimate compressive strength (MPa)

Elastic modulus in compression (MPa)

Cortical bone

124–174 (Pal, 2014)

Trabecular bone Ligament Tendon Cartilage

2.5 (Athanasiou et al., 2000) 50–100 (Holzapfel, 2001) 50–100 (Holzapfel, 2001) 9–40 (Holzapfel, 2001)

17,000–19,000 (Pal, 2014) 483 (Athanasiou et al., 2000) 345 800–2000 1–10 (Pal, 2014)

90–167 (Athanasiou et al., 2000) 0.15–14 (Athanasiou et al., 2000) – – –

Aorta

0.3–0.8 (Holzapfel, 2001)

–

Skin Hyaluronic acid Chitosan

1–20 (Holzapfel, 2001) – 0.5–45 (Foster et al., 2015; Jana et al., 2012) 37

0.05–0.1 (Isnard et al., 1989) 0.02–0.1 (Liang, 2010) – 23–68 (Jana et al., 2012)

5000–15000 (Athanasiou et al., 2000) 12–900 (Athanasiou et al., 2000) – – 1–20 (Barker and Seedhom, 2001) –

– – 0.3–1.7 (Jana et al., 2012)

– 0.02–0.15 (Levett et al., 2014) 5.5–18 (Jana et al., 2012)

–

–

35–350 (Eshraghi and Das, 2010)

0.6–38 (Eshraghi and Das, 2010)

600–7000 (Bergström and Hayman, 2016) 12–317 (Eshraghi and Das, 2010)

Poly(lactide) Poly(caprolactone)

1.1–16 (Eshraghi and Das, 2010)

It is important to note that the ranges of properties given for tissues and materials reflect the effect of testing protocols, processing techniques, material synthesis techniques, modifications to materials, and the inherent variability that exists within tissues and materials of similar type.

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Fig. 1 Stress–strain curve. Stress (y-axis) is a measure of the load placed on an object and strain (x-axis) is the material deformation in response to the stress. The linear portion of the (solid) curve indicates the elastic region, in which the material will revert to its original form if the load was removed. The curved portion of the (solid) curve is the plastic region, beyond which the material is permanently disfigured. Once the material breaks (point of material failure) the stress drops to 0.

a tissue is placed under tension and the crossheads are moved a fixed distance apart and held in that position such that the tissue is held in tension with a constant load (see Fig. 2). Over time as the tissue is held in tension at a fixed distance the stress exerted on the load cell by the tissue relaxes or decreases, indicating that the tissue is no longer under the same tension as it was at the onset of testing. This is a phenomenon of viscoelastic tissues, as a purely elastic tissue would not undergo this stress relaxation. Creep occurs when a soft tissue or material is placed in tension with a constant strain rate and the tissue or material constantly deforms as the crosshead continues to move, rather than resist the crosshead movement and ultimately fail (see Fig. 3). When the crosshead stops moving and returns to its original position, removing the load from the tissue, the stress–strain curve follows a different path than when initially loaded, unlike a purely elastic material that would follow the same path. This discrepancy between the initial and recovery path of the stress–strain curve of viscoelastic tissues or materials is termed hysteresis (see Fig. 4) and gives insight into how a tissue, or a material, will recover after the loading is removed. The tissue or material may or may not fully regain its original shape but the rate at which it recovers may be different than a purely elastic material.

Methods for Testing Mechanical Properties of Biological Tissues and Synthetic Materials Evaluating the mechanical properties of tissue or materials is most effective when following specific protocols such that the results of testing from one laboratory can be accurately compared to another laboratory. This is critical given the varied results that can be obtained depending on the specific testing protocol as described earlier. Consistency in mechanical testing can be achieved by observing published protocols or using testing standards like those published by the American Society for Testing and Materials (now simply known as ASTM) or the International Organization for Standards (ISO). These standard organizations publish detailed protocols for a full range of materials evaluation, including mechanical testing. For instance, ASTM standard D695 provides a protocol for testing the compressive strength of solid plastic materials, specifying the dimensions of the sample (aspect ratio), the crosshead speed of the mechanical testing machine, and the collection of data. Typically force and displacement data are recorded by the mechanical testing machine and are readily converted into stress and strain data, either by the testing machine’s

Fig. 2 Stress relaxation occurs when a material is loaded initially and held under constant strain for a period of time, during which the load experienced by the tissue decreases without a change in strain.

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Creep occurs when a material is loaded with a constant fixed strain rate but the applied stress does not increase.

Fig. 4 The hysteresis curve shows the stress–strain relationship as a material is loaded (blue line) and unloaded (red line). After the material is loaded it recovers its initial shape differently than a perfectly elastic material.

accompanying software or by the user from the raw force/displacement data. Plotting the stress versus strain curve yields information such as the elastic modulus, yield stress, fracture stress, and toughness.

Compression and tensile testing As described earlier, compression testing of a material involves a sample of the material of dimensions indicated by the appropriate standard by bringing two compression platens together to compress the sample. Specific protocols call for the compression of a cylinder of material but the specific shape of the material can vary by standard protocol. Applied force and crosshead displacement data are collected in real time and the test continues until the sample fails (as indicated by the shape of the force-displacement curve) or the tester chooses to end the test. As reported earlier, compression testing can reveal the ultimate compressive strength, yield strength, elastic modulus, and toughness of a sample. Tensile testing is done in a similar manner but rather than being compressed between two platens the sample is connected on either end to tensile testing grips. Typically when samples are tested in tension, the sample is formed into a dogbone shape such that the ends of the sample are wider than the middle which serves two purposes; (1) it increases the surface area available for the tensile grips to hold onto the sample to provide more surface area for gripping and (2) it biases the sample to more likely fail along the mid-region rather than at the tensile grip since the mid-region is thinner and therefore mechanically less robust. This strategy allows for stress concentrations that may exist at the tensile grips to have less of an effect on the failure of the material, allowing for a more accurate test. Tensile testing can reveal the ultimate tensile strength, young’s modulus, and yield strength. Hard tissues like bone are often tested in compression, as are implant materials. One may question the validity of pure compression in testing tissues like bone since it is rarely ever in pure compression, but having a uniform testing strategy to compare bone between anatomical sites or between subjects can serve to eliminate other testing variables and permit proper comparison.

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3-Point and 4-point bending Bending tests are conducted by placing a length of material across a span and pushing down along the span to bend the material until failure. Bending tests reveal the elastic modulus of bending, flexural stress, and flexural strain of a material. 3-Point bending involves placing the material across a span supported on either ends of the material and bringing down a point source to the center of the span and bending the material until failure while recording applied force and crosshead displacement (see Fig. 5). The 3-point description comes from the two points of support at the ends of the material and the one point of deflection brought down to the middle of the material. 4-point bending tests are conducted similarly to 3-point bending tests except that rather than one point source being brought down to the center of the span of material two points slightly separated from the center of the material are brought down in contact with the material. This separation of the two point sources spreads the region of bending out from the center such that a larger portion of the material is tested than with only one point of deflection. Hard materials like bone and implants can be tested using bending tests as a measure of the tissue or material in both tension (the bottom of the sample as it is tested) and compression (the top of the material as it is tested).

Torsion testing Torsion testing involves the twisting of a sample along an axis and is a useful test for acquiring information like torsional shear stress, maximum torque, shear modulus, and breaking angle of a material or the interface between two materials. Typically a longitudinal sample is placed in a torsion tester and one end of the sample is twisted around the long axis until failure, during which the force, or in the case of rotation the torque, and the displacement, or in the case of rotation the angular displacement, are recorded. Torsion testing is appropriate for materials that may experience a torsional load like a metallic bone screw, an intramedullary rod, rubber tubing that may become twisted, or to measure the shear strength of a bond between an implant and native tissue like bone.

Indentation testing In contrast to traditional mechanical testing that provides one mechanical assessment of an entire material, indentation methods can mechanically characterize a material at the macro-, micro-, and nano-scale level. This becomes a valuable tool when looking at tissues and complex tissue constructs in which several tissue types or materials may be integrated together. Doing traditional tensile, compressive, or torsion testing of heterogeneous tissues or tissue constructs would provide one numerical value for each test performed, but that one value may not accurately characterize the various tissue or material types. Indentation studies at both the micro and nano level allow for the precise mechanical testing of either small volumes of material or across very small, specific regions of a tissue or construct. Whether nanoindentation or microindentation is performed depends on the specific questions being asked but is ultimately dictated by the physical size of the indenter, which are typically either spherical or pyramidal in shape. The choice of tip shape depends on the information to be collected and the brittleness or softness of the material itself. Nanoindentation, or depth-sensing indentation, requires a small load from a point source to be applied to a tissue or material surface to yield a small, localized deformation. The loading and resulting displacement of the probe are monitored and recorded and from this data the material hardness and modulus can be calculated. Loads used in nanoindentation can range from micronewtons to millinewtons and displacements can range from nanometers to micrometers. Advantages of this mechanical testing technique include the ability to evaluate complex tissues and constructs with precision and accuracy with relatively little destruction of the sample. Tissues like cartilage that can have regional differences in mechanical properties within the same tissue can benefit from nanoindentation. Mechanical properties like elastic modulus, hardness, storage and loss modulus, and compliance can be calculated with indentation tests.

Fig. 5 3-Point (left) and 4-point (right) bending test apparatus. 3-point bending provides three points of contact; two supports and one center point where the loading is applied. 4-point bending provides four points of contact; two supports and two points where loading is applied. The two points spread the loading region along the specimen so a larger portion of the material is tested in bending.

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Evaluating Tissue Construct Biocompatibility and Cell-Construct Interactions While characterizing the mechanical properties of a tissue construct is critical to its overall success, it is equally critical to characterize the construct in terms of its interactions with the building blocks of regenerated tissues: the cells. Cellular evaluation is a vast area of study and fills textbooks with theory, technique, and outcome measurements and an exhaustive evaluation is beyond the scope of this article. There are, however, some fundamental considerations and techniques that provide an important foundation to future studies. Here, we discuss these fundamental analyses, the theory behind their utility, and some suggested methodologies that help determine the overall biocompatibility of a tissue construct. The term biocompatibility refers to the degree of response a material may invoke after being implanted. Generally speaking, any foreign material that is implanted into the human body will invoke a response but that response can vary widely from toxic resulting in serious injury or death to benign that raises no concerns but should be understood, to a response that is predicted and desired because it may lead to enhanced healing.

Toxicity The toxicity of a material or construct refers to how toxic it is, or to how much damage it can do to an organism in general. When considering regenerative engineering constructs under development, one may choose to focus on the toxicity of a construct specifically to cells, or its cytotoxicity. A cytotoxic material would be one that caused considerable harm or death to the cells on, in, or near the construct, but it is reasonable to assume that a material or construct that is toxic to cells would be toxic to larger organs or tissues made up of similar cells. While the toxicity of a material is less of a concern for biomaterials that have been evaluated previously and in use currently new materials would certainly need to undergo toxicity testing. These materials can be evaluated by placing cells directly on or in proximity to the material to determine whether any compounds or eluents that may leach from the material are toxic. This latter case is especially important in the development of resorbable materials since an intact material may be cytocompatible but its degradation products may not be. For instance, polyesters such as polylactide, polyglycolide, and poly(lactide-coglycolide) are well-studied and widely used degradable biomaterials and are inherently nontoxic to cells that attach, migrate, and proliferate while in contact with the material. These polymers, however, degrade when in contact with an aqueous medium and their degradation products are lactic and glycolic acid, both highly acidic and potentially damaging to cells on or near the material. Despite these acidic degradation products, these polymers are widely and successfully used but serve as a good example of why the degradation products are as important as the material itself when it comes to cytotoxicity studies. Testing for cytotoxicity generally involves determining whether the cell membrane is intact. An intact cell membrane indicates a live cell and a ruptured cell membrane indicates a dead cell. Assays have been developed to evaluate the integrity of the cell membrane in which solutions are added to a cellularized environment, be it a tissue construct or cells in culture, that either stain the intracellular components of a dead cell or react with extracellular components that have been released from a cell because of the damaged cell membrane. For instance, trypan blue is used to visually determine a live or dead cell. Since trypan blue does not pass through the cell membrane, a live cell with an intact cell membrane will not stain blue while a dead one with a damaged and therefore permeable membrane will. Another common method is a live/dead assay, often coupled with fluorescent markers. The fluorescent markers are tagged to either a molecule that passes through the cell membrane or one that does not, the former fluorescing at a different wavelength than the latter and therefore distinguishing between live and dead cells. Other assays interact with compounds normally maintained within the cell membrane in live cells but are released into the extracellular milieu if the membrane is broken, indicating a dead cell. Some cytotoxicity assays provide an overall level of cell death based on the entire population of cells without the specificity of which cells are live or dead, while other methods allow the determination of which cells are alive and which are dead. The choice of assay depends on the information needed; an overall picture of cytotoxicity of a material or construct or a regional measure of cell survival within a construct or material.

Cell viability and proliferation Cell viability and proliferation are important metrics of a material or construct for tissue repair. They can provide important feedback on the suitability of surface modifications, three-dimensional architecture, oxygen transport, degradation product compatibility, and many other aspects of the material or structure. While often used interchangeably cell viability and proliferation are similar to each other in that both are measures of live cells but distinct in that viability is a measure of cellular activity and overall health, while proliferation is a measure of the rate of growth and production of daughter cells of a cell population. Either may be suitable as a tool to measure the health of a cell population and the suitability of a construct or material for cell survival but the assays and techniques used to assess each can vary and care should be taken to ensure that the method of analysis or choice of assay is appropriate. Generally speaking, proliferation is required to have a construct fully populated by cells which in turn mandates a fully interconnected pore structure or other means of adequate nutrient and waste exchange to allow the cells to remain viable within the depths of the construct. If the construct has insufficient nutrient transport cells will not proliferate to the center of the construct and new tissue will not be deposited. Measuring proliferation is one measure of how well a construct can support the residence of cells throughout its structure. Common assays used to test cellular viability include the MTT and MTS assays, which measure the viability of a cell through its metabolic activity, DNA synthesis assays which can translate to actual cell numbers, and manual cell counts using a dye such as trypan blue, which, as discussed earlier, will permeate a damaged cell membrane as an indicator of a nonviable, or dead, cell.

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Differentiation For a tissue construct to facilitate the regeneration of tissues, it must be able to sustain not only the proliferation of cells but also their differentiation as well. Cell differentiation involves the maturation and subsequent specialization of a cell. As cells mature in their life cycle they become more specialized, or more focused in their function. For instance a mesenchymal stem cell can contribute to the regeneration of multiple tissue types depending on its gene expression and subsequent maturation, or differentiation. Once it differentiates its overall potential to become multiple tissue types is reduced to one specific tissue type. Differentiation is not limited to stem cell populations though. An osteoblast, for instance, will differentiate into an osteocyte and its function will become more specialized. Understanding how cells differentiate, what causes them to differentiate, and how to control their differentiation are powerful tools available to the regenerative engineer through the design of tissue constructs. Materials scientists and engineers have developed methods to influence or control cellular differentiation by modulating material aspects like composition, topography, architecture, delivery of payload, and mechanical properties to name a few. Evaluating the differentiation of a cell is one way of measuring cellular function on a construct but also its overall health. As cells differentiate they undergo a number of changes that can be measured or tracked to assess the nature and extent of differentiation. Surface receptors, protein expression, and genetic expression are a few of the ways the regenerative engineer can assess the extent of differentiation of a cell and gather a picture of the functionality of those cells on the construct. Whether as a passive measure of cellular function on a cell or as a measure of the intended influence on a particular cell, differentiation is an important tool in assessing the efficacy of a tissue construct.

In Vivo Models for Tissue Construct Evaluation In vitro cellular analysis provides important insight into the potential utility of a tissue construct as a tool for regenerating tissue and offers the advantage of an environment that is easy to control and while it can act as an important first approximation of efficacy it is only an estimate of the in vivo environment. To truly evaluate a tissue construct an in vivo model is necessary. Also unlike human studies where tissue is evaluated after injury or onset of disease animal models can chronicle the progression of injuries. Many different in vivo models exist to test toxicity, biocompatibility, mechanical stability, and tissue repair potential. An exhaustive list of in vivo models is beyond the scope of this article, but here we provide some defect models as examples of the different approaches that can be taken. Within musculoskeletal regenerative engineering there are several models of hard and soft tissue injuries and defects. Below are descriptions of each, the type of injury they model, why they are chosen, and how they inform the clinical implementation of the tissue construct.

Hard Tissue Defect Models The in vivo models used in hard tissue construct evaluation can focus on critical size or noncritical size defects. A critical size defect, typically made in long bones, is one that will not heal on its own. Unlike a noncritical size defect that, if left alone, would eventually heal, a critical size defect is of sufficient size to prevent spontaneous repair. In regenerative engineering we mostly focus on those tissues that cannot heal themselves and therefore use critical size defects. The presumption is if the engineered construct can heal a critical size defect in an animal model it can support large-scale defect healing in humans. To this end several long bone models have been used in mice, rats, and rabbits, where different animals have different physical sizes and healing times. Mice tend to heal faster than rats and rabbits while the larger size of the rat and rabbit make for easier defect formation and, depending on the construct, more realistic preclinical models. Long bone defects allow for the evaluation of cortical bone repair while calvarial defects, circular defects in the skulls of animals, allow for the evaluation of trabecular bone. While long bone defects in these animals would provide a good assessment of long bone repair in humans, a calvarial defect in mice may provide a good assessment tool for trabecular bone found in the facial bones while providing a bony area large enough to create a defect and assess healing using a construct. Trying to create a jaw or facial bone defect in a mouse may prove to be too challenging and certain types of constructs may not be able to be manufactured on such a small scale. Some of the long bones that have been evaluated in these animals are rat tibial segmental defects, rat femoral segmental defects, rabbit ulnar segmental defects, and rabbit radial segmental defects.

Soft Tissue Defect Models The in vivo models used for soft tissue construct evaluation focus on the repair and/or regeneration of tissues like muscle, ligament, tendon, and cartilage. Soft tissue defect models can assess the construct’s ability to provide mechanical stability to the defect site, to encourage cellular migration and infiltration, and over time to evaluate the overall healing of the defect. As with hard tissue models, many soft tissue models focus on defects that do not fully heal on their own such as full ligament and tendon tears, but can also evaluate methods of increasing the rate and extent of tissue repair for injuries that may heal on their own. The specific animals chosen for soft tissue defects depend on the animal anatomy, physical size, and the biomechanics of how the tissue in question is loaded. For instance, the rat is physically small and can be challenging to work with when considering the repair or regeneration of tissues like the anterior cruciate ligament (ACL) or rotator cuff, but studies have demonstrated that due to anatomic similarities between the rat and human rotator cuff the rat is uniquely suited as a model for human rotator cuff tears and subsequent surgical

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repairs. In this instance, the challenge of the small size is outweighed by the benefits of having a model that more accurately matches the biomechanical loading and anatomical structure of a human rotator cuff. Rotator cuff defect models typically involve the transection of one or more tendons in the shoulder or chronic overuse of the tendon leading to injury or a combination of both. Repair typically involves the resuturing of the transected tendon back to the bone of origin, but depending on the construct that has been developed to facilitate repair the strategy may vary. For anterior cruciate ligament repair small animals are used as well but larger animals offer fewer challenges surgically, which may result in less surgical error or variability between animals. Small animals have been used for cell-based ligament repair where a surgical implantation of a construct is not required, but larger animals like rabbits have found favor when a construct is being surgically implanted and sutured to existing tissues. In these instances the knee cavity is opened and the intact ACL is transected. Tunnels are drilled through the tibia and femur (proximal and distal to the transected ACL) and the construct is fed through each of the tunnels such that the two bones are connected by the construct, much like the native ACL. Once the construct is adequately secured within each bone (which can be done by a variety of methods), the soft tissue around the knee is sutured closed and the animal is free to ambulate. This model allows the construct to be mechanically loaded just as the native ACL was prior to transection and is exposed to the same extracellular milieu as the native ACL, providing an excellent recapitulation of the native ACL environment for testing. For cartilage repair, there are a few different animal models. Some models involve the creation of a focal defect, which is a defect in the cartilage or the cartilage and underlying trabecular bone, with well-defined borders. Other models involve the creation of osteoarthritis, a degenerative disease of the cartilage that results in more diffuse, less-localized, cartilage damage across an articulating surface. Each of these models of injury is created in different ways and is designed to measure different treatment strategies. The focal injuries can be created by impact or surgical removal of cartilage with or without the underlying bone while the less-localized injuries like osteoarthritis can be created surgically by creating grooves in the cartilage, surgically manipulating the meniscus to induce osteoarthritis, or surgically transecting the ACL to induce osteoarthritis. The choice of model depends on the treatment strategy to be tested. The choice of animal may depend on the questions being asked or the budget available. Small animal cartilage defects are used quite commonly where one of the above methods is applied to mice, rats, rabbits, and guinea pigs, but the anatomy of small animal cartilage and underlying bone is somewhat different from humans so larger animals have been used as well. The anatomical similarity of the cartilage and underlying bone in animals like dogs, goats, sheeps, and pigs to humans is better but the cost associated with these studies is considerably higher. Each approach, small or large animal, has its benefits and drawbacks so the selection of a model is often dictated by the questions being asked or the stage of development of the construct.

Conclusion The field of regenerative engineering presents an exciting new paradigm for tissue construct design and implementation that seeks to take advantage of the newest discoveries in materials development, stem cell biology, developmental biology, and clinical translation. Here, we have discussed some of the fundamental aspects of tissue construct evaluation from mechanical testing to cellular evaluation and preclinical testing. While not an exhaustive analysis of the area it presents a starting point for the budding regenerative engineer and provides a scope from which to build a thorough testing programs for novel tissue constructs. Each distinct area is constantly under development and therefore constantly changing and being updated, but a foundational knowledge of these areas will serve the regenerative engineer well as they move forward with their novel approaches to complex tissue repair.

References Athanasiou, K. A., Zhu, C. F., Lanctot, D. R., Agrawal, C. M., & Wang, X. (2000). Fundamentals of biomechanics in tissue engineering of bone. Tissue Engineering, 6(4), 361–381. Barker, M. K., & Seedhom, B. B. (2001). The relationship of the compressive modulus of articular cartilage with its deformation response to cyclic loading: Does cartilage optimize its modulus so as to minimize the strains arising in it due to the prevalent loading regime? Rheumatology (Oxford), 40(3), 274–284. Bergström, J. S., & Hayman, D. (2016). An overview of mechanical properties and material modeling of polylactide (PLA) for medical applications. Annals of Biomedical Engineering, 44(2), 330–340. Eshraghi, S., & Das, S. (2010). Mechanical and microstructural properties of polycaprolactone scaffolds with one-dimensional, two-dimensional, and three-dimensional orthogonally oriented porous architectures produced by selective laser sintering. Acta Biomaterialia, 6, 2467–2476. Foster, L. J., Ho, S., Hook, J., Basuki, M., & Marçal, H. (2015). Chitosan as a biomaterial: Influence of degree of deacetylation on its physiochemical, material and biological properties. PLoS One, 25, 10(8). Holzapfel, G. A. (2001). Biomechanics of soft tissue. In J. Lemaitre (Ed.), Handbook of material behavior: Nonlinear models and properties (pp. 1057–1075). Cachan: LMT-Cachan. Isnard, R. N., Pannier, B. M., Laurent, S., London, G. M., Diebold, B., & Safar, M. E. (1989). Pulsatile diameter and elastic modulus of the aortic arch in essential hypertension: A noninvasive study. Journal of the American College of Cardiology, 13(2), 399–405. Jana, S., Florczyk, S. J., Leung, M., & Zhang, M. (2012). High-strength pristine porous chitosan scaffolds for tissue engineering. Journal of Materials Chemistry, 22, 6291–6299. Levett, P. A., Hutmacher, D. W., Malda, J., & Klein, T. J. (2014). Hyaluronic acid enhances the mechanical properties of tissue-engineered cartilage constructs. PLoS One, 9(12). Liang, X. (2010). Boppart SA. Biomechanical properties of in vivo human skin from dynamic optical coherence elastography. IEEE Transactions on Biomedical Engineering, 57(4), 953–959. Pal, S. (2014). Mechanical properties of biological materials. In Design of artificial human joints and organs (pp. 23–40). New York, NY: Springer.

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Further Reading Athanasiou, K. A., Zhu, C. F., Lanctot, D. R., Agrawal, C. M., & Wang, X. (2000). Fundamentals of biomechanics in tissue engineering of bone. Tissue Engineering, 6(4), 361–381. Buffinton, C. M., Tong, K. J., Blaho, R. A., Buffinton, E. M., & Ebenstein, D. M. (2015). Comparison of mechanical testing methods for biomaterials: Pipette aspiration, nanoindentation, and macroscale testing. Journal of the Mechanical Behavior of Biomedical Materials, 51, 367–379. Sah, R. L., & Ratcliffe, A. (2010). Translational models for musculoskeletal tissue engineering and regenerative medicine. Tissue Engineering. Part B, Reviews, 16(1), 1–3.

Clinical and Laboratory Aspects of Hematopoietic Stem Cell Transplantation ST Avecilla and MM Cushing, New York Presbyterian Hospital, Weill Cornell Medical College, New York, NY, USA © 2019 Elsevier Inc. All rights reserved.

Hematopoiesis: A Brief Introduction to How Blood Cells Are Created and Formed The Utility of Hematopoietic Stem Cell Transplantation HPC Collection HPC-M (Bone Marrow) HPC-A (Peripheral Blood Stem Cells) HPC-C (Umbilical Cord) HPC Characterization HPC Manipulation Cryopreservation HPC Infusion Adverse Reactions to HPC Infusions Complications after HSCT Advances in HSCT Further Reading Relevant Websites

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Glossary Allogeneic transplant Stem cells utilized are from a donor. Apheresis Medical procedure during which a patient or donor has their blood separated and a specific fraction collected via a continuous flow centrifuge. Autologous transplant Stem cells utilized are from the patient’s own cells (rather than from a donor). Conditioning regimen Chemotherapy or irradiation given immediately prior to a transplant to eradicate the patient’s disease and suppress immune reactions, which could lead to rejection. Cryoprotectant Liquid substance capable of inhibiting intracellular ice formation in the process of freezing cells for long-term storage. Cytokine Signaling protein, which stimulates the proliferation of cells via specific receptors. Differentiation The process whereby an unspecialized stem cell acquires the features of a specialized cell such as a mature red blood cell. Differentiation is controlled by the interaction of a cell’s genes with the physical and chemical conditions outside the cell, usually through signaling pathways involving proteins in the cell surface. Engraftment A state where a hematopoietic progenitor or stem cell has homed to the bone marrow and reconstituted hematopoiesis for the recipient of the transplant. Flow cytometer A device that measures various optical parameters of single cells in suspension. Graft-vs-host disease (GvHD) Common complication of allogeneic stem cell transplant where immunocompetent T cells in the donor graft recognize and attack host tissue as a foreign target. Hematologic malignancy Cancer of the blood or bone marrow. Hematopoiesis Process of production, multiplication, and specialization of blood cells in the bone marrow beginning with a hematopoeitc stem cell. Hematopoietic progenitor cell (HPC) A cell with the capability to proliferate and generate differentiated blood cells, but which is already at a further stage of differentiation than a stem cell and may only differentiate into its designated target hematopoietic cells. In addition, HPCs only can divide a limited number of times, unlike stem cells. For the purpose of this article, the term HPC will include HPCs and hematopoietic stem cells. Hemostasis Arrest of bleeding by vasoconstriction and coagulation. HPC mobilization Process by which drugs are used to cause release of stem cells from the bone marrow into the peripheral blood. Peripheral blood stem cells (PBSCs) HPCs circulating in peripheral blood that are collected by apheresis from an autologous or allogeneic donor. Stem cell An undifferentiated, omnipotent cell with an unlimited capability of self-renewal.

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Hematopoiesis: A Brief Introduction to How Blood Cells Are Created and Formed Blood is a liquid tissue that has many important functions in the human body. In addition to serving a critical role in transporting oxygen and nutrients, blood and its constituents are essential in host defenses, hemostasis, and many other functions. Hematopoiesis is a process whereby stem and progenitor cells, through a process of proliferation and differentiation, result in the generation of the cellular elements of blood. This differentiation process is influenced by lineage-specific cytokine signaling such as granulocytecolony stimulating factor (G-CSF), erythropoietin, thrombopoietin, in addition to microenvironmental stimuli such as interaction with bone marrow stroma and blood vessels. Due to the ephemeral nature of the components of blood, hematopoiesis is crucial in the continuous replenishment of functional cells. There are several instances where clinical treatment of disease results in the irreversible destruction of the bone marrow, which can be fatal if a compatible hematopoietic progenitor cell transplant is not performed.

The Utility of Hematopoietic Stem Cell Transplantation Hematopoietic stem cell transplantation (HSCT) refers to the intravenous infusion of autologous or allogeneic hematopoietic progenitor cells (HPCs), collected from bone marrow, peripheral blood, or umbilical cord blood (CB), to replace damaged or aberrant stem cells in a patient. A list of common indications for HSCT is given in Table 1. In the currently practiced methods of treating hematologic malignancies, HSCT is a frequently chosen modality in conjunction with chemotherapy and/or total body irradiation. Nonmalignant conditions that are treated with HSCT include severe sickle cell disease, severe combined immunodeficiency, in addition to marrow failure states such as aplastic anemia. Most chemotherapy regimens employ cytotoxic agents which are intended to kill the neoplastic cells. The nonneoplastic bone marrow, which has a high proliferative baseline, is also severely and adversely affected and may be irrevocably damaged. In order to rescue the patient from a bone marrow depleted state, an HSCT must be done. Due to the fact that the immune cells are primarily derived from the bone marrow and associated tissues, it is highly important that the donor HPCs come from a human leukocyte antigen (HLA)-matched party. The requirement for HLA matching even outweighs the need for A/B/O blood group matching. Sources for HLA-compatible HPCs can vary from autologous HPCs to those harvested from related or unrelated donors. Transplantrelated mortality and morbidity may be affected by numerous aspects of the transplant (Table 2). Prior to collecting an allogeneic donor candidate, several health criteria must be assessed. The donor must pass a physical examination in addition to answering an HPC donation questionnaire, which assesses infectious disease risk. Infectious disease marker assays are performed on the donor’s blood to ensure that no infectious agents are transmitted unknowingly.

Table 1

Common indications for HSCT

Autologous

Allogeneic

Multiple myeloma/amyloidosis Non-Hodgkin lymphoma Hodgkin lymphoma Acute myeloid leukemia Neuroblastoma Germ cell tumors

Acute myeloid leukemia Acute lymphoblastic leukemia Myeloproliferative disorders Multiple myeloma Non-Hodgkin lymphoma Hemoglobinopathies (thalassemia/sickle cell) Inborn errors of metabolisms Autoimmune disorders Severe congenital immune deficiencies

Table 2

Aspects of HSCT, which may affect morbidity/ mortality

Toxicity and effectiveness of conditioning regimen Degree of HLA matching Infection prevention and treatment Autologous vs allogeneic transplant Source of transplant (bone marrow, peripheral blood, or umbilical cord) Time from transplant until engraftment Incidence and severity of graft-vs-host disease

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HPC Collection HPC-M (Bone Marrow) The first recognized source of HPCs was bone marrow. It was discovered that a small population of the harvested cells had the property of giving rise to the full repertoire of hematopoietic cell lineages and most importantly they could reconstitute a depleted bone marrow for the lifetime of the recipient. While no longer the most popular method of extracting HPCs for transplant, hematopoietic progenitor cell-bone marrow (HPC-M) has some advantages to the methods developed in later years to harvest HPCs for transplant (Table 3). First and foremost, the prospective donor need not take any medication or cytokine which may have both short- and long-term negative health effects. The extraction process is a surgical procedure during which the donor is placed in a prone position under local or general anesthesia for pain management. Once the donor is properly prepared, an 11–14 gauge needle is inserted into the posterior iliac crest until it is in the marrow space. Marrow is then aspirated from the donor until multiple locations accessible from the single insertion site have been exhausted. Typically, the marrow is collected with heparin as the anticoagulant and a biocompatible solution for cell suspension such as Plasma-Lyte A (Baxter Healthcare, Deerfield, IL). Additional anticoagulant such as acid citrate dextrose (ACD) may be added to further prevent sample clotting. Due to the relative rarity of the HPCs within the marrow space, a large volume of material is usually extracted which often includes a significant contamination by peripheral blood. The harvest is monitored by total nucleated cell (TNC) yield with a usual target of >3.0  108 nucleated cells/kg of recipient weight. Drawbacks of harvesting HPC-M include the donor’s exposure to anesthesia, blood loss, postcollection pain, bruising, and other less common symptoms. Due to the blood loss, a great majority of donors require at least one red blood cell transfusion (both autologous and less commonly allogeneic), which is also a risk to the donor. Because T-lymphocytes are not produced in the marrow, the product tends to have much less CD3 T-cell content which can result in a lower incidence of graft-vs-host disease (GvHD).

HPC-A (Peripheral Blood Stem Cells) In the initial experience with HSCT and patients with hematologic malignancies, it was found that in some chemotherapy regimens, patients demonstrated a low but substantially increased number of HPCs circulating in the peripheral blood. With this observation, HPC mobilization regimens were developed which utilized chemotherapy as a mobilization agent. This technique could ethically only be used on patients with malignancies and not with allogeneic healthy donors since the risk of chemotherapy is considered unacceptable. Once cytokines such as G-CSF were discovered and developed for clinical use, mobilization protocols utilizing G-CSF became available. Typically allogeneic and/or autologous donors will receive a subcutaneous injection of G-CSF daily for 5 days with HPC collection via apheresis (Figure 1) occurring on the fifth day. Of key importance is the donor vascular access due to the fact that peripheral vein access is the preferred and least invasive route to obtaining peripheral blood. Donors with poor peripheral venous access can still undergo HPC mobilization and apheresis harvest of HPCs, however, a central line venous catheter or comparable vascular access will first have to be placed. While central vascular access allows for higher apheresis flow rates and typically shorter collection cycles, it comes with the risks of having a central line placed which include infection, air embolus, and bleeding. In addition to cytokine and chemotherapy mobilization regimens, newer agents have been developed to further enhance HPC mobilization. An agent that is FDA approved for HPC mobilization in specific patient populations (lymphoma and multiple myeloma) is plerixafor. Plerixafor (formally known as AMD3100) is a small antagonist molecule of the chemokine receptor CXCR4, which is found on HPCs in addition to other cells. It is hypothesized that disruption of CXCR-4 chemotactic signaling of HPCs while in the bone marrow niche allows for the peripheral mobilization of HPCs into the bloodstream. The apheresis machine is anticoagulated with ACD-A and therefore products collected via this modality have ACD-A as the primary anticoagulant. Collection is monitored by enumerating CD34/CD45 positive cells and collection is done until a goal dose is reached e.g., >2–5  106 CD34þ cells/kg. Unlike HPC-M, apheresis collected HPCs have a substantial number of T cells which has implications in both graft-vs-host and graft-vs-leukemia (GvL) effects of the transplant. An additional benefit of HPC-A is that engraftment times are usually shorter in comparison with HPC-M and HPCs from human umbilical vein CB. See Figure 2 for an example of a HPC-A.

HPC-C (Umbilical Cord) The most recently described, readily available source of hematopoietic progenitor cells for clinical use is human umbilical vein CB, HPC-C. An insightful observation was made upon the microscopic examination of CB smear preparations that there was an Table 3 HPC source HPC-M HPC-A HPC-C

Hematopoietic progenitor comparison Cell dose (per kg recipient weight) 8

2–4  10 nucleated cells 5  106 CD34þ cells >1.5  107 nucleated cells

Speed of engraftment

GvHD risk

HLA match stringency

Moderate Fast Slow/Delayed

Moderate Highest Lowest

High High Moderate-high

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Figure 1

Apheresis collection on a cobe spectra.

Figure 2

HPC-A.

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immature appearing white cell population that resembled blast cells. Further research into this cell population revealed that this blastlike population was in fact a benign hematopoietic progenitor cell population, which could be used successfully in HSCTs. Immediately prior to giving birth, prospective donor mothers are screened with a donor questionnaire and upon giving birth, maternal and CB samples are obtained for appropriate infectious disease marker testing. Once the child is born, the umbilical cord is cross-clamped and the attached placenta is placed in a sterile container and transferred to a workroom where the CB is harvested under as sterile conditions as possible. The harvested CB is then processed to isolate the mononuclear cell fraction and analyzed for transplant appropriateness, which include TNC count, CD34þ/CD45þ cell count, microbiological cultures, and HLA typing. Once the testing samples have been obtained, the product is cryopreserved and banked for future use in HSCT treatments. Due to the immature, immune state of the product, HPC-C is associated with lower rates of GvHD in addition to successful transplants having an increased tolerance for HLA mismatches. Because HPC-C are collected, characterized, and stored before a specific recipient is identified, they represent a rapidly obtainable source of HPCs that can be used without a long lead-time. Conversely, HPC-A and HPC-M both require identification of a suitably HLA-matched donor with genotyping in addition to the performance of a medical procedure to harvest the HPCs (apheresis or surgical marrow collection), which may be accompanied

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with possible HPC mobilization in the case of HPC-A. Due to the small nature of the umbilical cord, there is an intrinsic limit of the quantity of HPCs that can be harvested per cord. Because HPC dose is one of the most important predictors of HSCT outcome (the higher the dose, the more likely engraftment will occur), HPC-C as the only source of HPCs from single donors for HSCT has limited success in larger patients (mostly adults).

HPC Characterization In order to properly prepare an HPC product, several parameters are measured. First a sample of the HPC product is taken and analyzed with a hematology analyzer in order to obtain a TNC count. In the case of HPC-M, the TNC count is used to measure harvest adequacy with the assumption that the typical CD34þ cell content is approximately 1–3% of the total TNC count. For a more accurate assessment of CD34þ cell content, the sample is subjected to flow cytometric analysis. Antibodies to CD34 (HPC marker) and CD45 (pan-leukocyte marker) antigens are used in conjunction with a viability stain to accurately quantify the number of live CD34þ cells that have the proper cellular characteristics to qualify as HPCs (small cells with low granularity, similar to lymphocytes). Viability can also be assessed using a vital stain such as trypan blue in which viable cells exclude the dye; however, trypan blue is considered to be less sensitive than flow cytometric viability using 7-AAD. To test the product for contamination risk, which may occur at any stage between collection and processing, a sample of product is inoculated in the automated microbial culture media at several points in processing. Finally, some laboratories perform HPC potency testing. The most common example of potency testing is the colony forming unit (CFU) growth assay. Briefly, a sample of HPC product is plated in a defined, cytokine-rich medium and allowed to incubate for 14 days, after which the cell culture is examined for the growth of colonies characteristic for different classes of progenitors (see an example in Figure 3). While CFU growth assays are helpful, at this time they cannot assess the presence of hematopoietic stem cells but rather the committed progenitor content (HPC).

HPC Manipulation ABO mismatched grafts are acceptable for HSCT, but there may be related complications, such as acute or delayed hemolysis, or delayed or lack of red cell engraftment. The cells may be red blood cell depleted, but this is often at the expense of the loss of some HPCs.

Cryopreservation While minimally manipulated HPC products are preferred due to higher levels of viability and lower chances that processing adversely affects the ‘stemness’ and viability of the HPC products, there are clinical situations where cryopreservation is an essential component of HPC processing. First and foremost, when autologous HSCT therapy is indicated such as in multiple myeloma and some forms of lymphoma, chemotherapy induction and conditioning regimens cannot occur concurrently with HPC collection due to the cytotoxic effects of the regimens. In order to harvest viable HPCs, the patient is first mobilized and his or her HPCs are collected, characterized, processed, and cryopreserved for later use after ablative chemotherapy and radiation have been administered. Another example where cryopreservation is widely used is in HPC-C. Situations may also occur where potential donors are

Figure 3

CFU assay.

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Figure 4

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Cryofacility.

unavailable at the time the patient needs an HSCT and, if available beforehand, the donor may preemptively donate HPC-A with subsequent cryopreservation for use in the HSCT when appropriate. Due to the fact that water expands when it transitions from liquid to solid phase, and the subsequent expansion results in cellular damage, cells cannot be stored in a frozen state without proper cryoprotection. Industry-wide, dimethyl sulfoxide (DMSO) is used in various concentrations as an agent that crosses the plasma membrane of cells and displaces water such that the cells become relatively dehydrated. DMSO is also thought to inhibit cell-damaging ice formation and promote water solidification in an amorphous, vitreous state that results in reduced amounts of cellular damage upon thawing. Once infused, DMSO is converted to dimethyl sulfone and dimethyl sulfide and eliminated by urinary excretion. Renal excretion accounts for almost half of the administered DMSO dose. Serum half-life of DMSO is about 20 h; the half-life of dimethyl sulfone, a renally excreted metabolite, is about 72 h. A small amount of DMSO is converted to dimethyl sulfide, which is excreted through the lungs for about 24 h and is responsible for the characteristic odor of the patient’s breath following a stem cell infusion. Cryopreserved HPCs are stored in liquid nitrogen tanks, most often at vapor phase (150  C) until they are needed for transplant (Figure 4 for an example of a cryofacility).

HPC Infusion The infusion of an HPC product (i.e., the transplant) is usually an anticlimactic event that resembles the intravenous infusion of any blood product. The product may be infused fresh, as is most often the case for allogeneic transplants, or thawed if it was cryopreserved after collection, as is most often the case for autologous transplants or HPC-C infusions. Usually the bags are thawed near the patient’s room with a 37  C water bath to shorten the exposure of the HPCs to DMSO. Cryopreserved HPCs should always be administered through a central venous line. A 10% DMSO solution has a high molality that can cause pain when infused through a peripheral line. The thawed HPC product may have a volume of 2–25 ml kg1 of the recipient. The product can be administered via a syringe or directly from the bag. Direct infusion from the bag can avoid the risk of contamination and cell loss during transfer from bag to syringe. At the end of the infusion, the bag and line should be flushed with normal saline to minimize cell loss. The line can also be flushed during the infusion to increase the infusion rate. The stem cell product should never be irradiated or transfused through a leukoreduction filter. The dose of DMSO given during the infusion of thawed stem cells varies, but most centers set an upper limit of 1 g of DMSO per kg of patient weight over a 24-h period. To minimize the time of contact between the thawed stem cells and DMSO, HPC infusions should occur as rapidly as possible. The infusion speed of DMSO cryopreserved products reported in the literature varies between 5 and 20 ml/min. A slower infusion rate may avoid some of the adverse reactions associated with DMSO infusion and volume overload. Patients should have their symptoms and vital signs carefully monitored during the infusion and for several hours afterward. Noncardiogenic side effects usually occur during the infusion and often resolve after the infusion is stopped. Bradycardia, however, has been reported to occur several hours after the infusion.

Adverse Reactions to HPC Infusions The incidence of adverse reactions to HPC infusion varies depending on the type of product infused (bone marrow, peripheral blood stem cells, or umbilical cord and autologous vs allogeneic), whether the product is fresh or thawed, and whether the product

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was washed after thawing vs bedside thawing. Reported rates of reactions generally range from 20 to 30% of infusions. Little data exist on the incidence of each individual type of reaction. As with all blood components, there is a risk of microbial contamination with HPCs. Contamination of HPCs can occur during any step in the transplant process, including collection, processing, and thawing. Many transplantation centers perform HPC product cultures before and after processing, but unlike other products, HPCs are not readily replaceable making the management of positive culture results far more complex. Many factors pertaining to the infectious organism, the product, and the patient must be taken into account when evaluating the culture result. The virulence of the identified organism is of great importance, with gramnegative rods raising particular concern. The possibility of replacing the contaminated product should be investigated. If, for instance, only a fraction of the vials for a transplant are culture positive and an adequate number of CD34 positive cells are present in the remaining vials, then the contaminated vials can simply be discarded. Recollection is another option to be evaluated. The stage of the patient’s treatment is critical because it affects how long the transplant can be delayed, particularly if the patient has already undergone a conditioning regimen. If proceeding with the infusion of a culture positive HPC product, prophylactic antibiotic therapy may be administered. When time allows pretransplant, the identity and antibiotic susceptibilty of the contaminant organism can be determined and targeted preemptive antibiotic therapy can be done. Allergic reactions are common in a stem cell recipient and manifest in a similar way to allergic transfusion reactions, with common features including pruritus, urticaria, and orbital swelling. Upper and lower airway obstruction and anaphylaxis may occur during more severe reactions. The allergen may be a protein in an allogeneic donor’s plasma or a product added during processing, such as dextran or DNAse. Allergic reactions can often be prevented by premedication with antihistamines, steroids, or a combination of both. Febrile reactions most commonly occur due to the accumulation of platelet- or leukocyte-derived cytokines in the product or to leukocyte antibodies in the recipient that react with the leukocytes in the product. These reactions may also occur as a result of bacterial contamination of the product or hemolysis due to plasma or red cell antigen incompatibility in allogeneic transplants. Thawed products have additional, associated adverse events. The thawed HPC product contains many components besides the HPCs themselves, including DMSO, plasma, white blood cells, granulocyte debris, and red blood cell debris (including free hemoglobin). Most reactions associated with thawed products are related to DMSO and include nausea, vomiting, abdominal cramping, hyper- or hypotension, bradycardia, chest tightness, fever, chills, and headache. DMSO gives off an odor that has been compared to creamed corn or garlic. The odor may last for 1–2 days and may cause nausea; sucking on hard candy may help eliminate the nausea. HPC products that are washed to reduce the DMSO content have less infusion-related toxicity.

Complications after HSCT There is a high mortality associated with hematopoietic stem cell transplant. Complications may occur that are related to the patient’s underlying disease, previous treatments, conditioning regimen, transfusions, and medications. In addition, the patient’s immune system may become dysregulated. Early complications are often related to the time to engraftment. An engraftment delay leaves the patient susceptible to infections and bleeding. HPC donor source, chemotherapy regimen, and glucocorticoid treatment can influence the type and propensity for infections, which include bacterial, fungal, and viral etiologies. Mucositis and gastrointestinal toxicity are frequent early treatment complications, especially in the first 14 days after transplant. Hepatic veno-occlusive disease (sinusoidal obstruction syndrome) is a serious complication that is more common in allogeneic transplant vs autologous with a mortality of 20–50%. Late complications of stem cell transplant include avascular necrosis of bone, cataract formation, bronchiolitis obliterans, and gonadal failure. The development of a secondary malignancy after transplant, often myelodysplastic syndrome or leukemia, infrequently occurs. Graft failure results when the recipient’s immune system rejects the infused donor cells and engraftment does not occur. It is less likely to occur after fully myeloablative conditions, but it is more common after reduced intensity conditioning or after the infusion of products with lower numbers of stem cells (such as umbilical CB). Relapse of the underlying disease is the most frequent cause of graft failure. Risk factors for relapse include transplant for advanced disease, autologous transplant, T-cell depleted graft, and reduced intensity conditioning regimens. Other risk factors for graft failure include HLA mismatch between donor and recipient and damage to the stem cell product during processing (rare). Donor lymphocyte infusion, collected from donors with or without G-CSF mobilization, can also be helpful in the treatment of patients with progressive disease after transplant. Immunologic events after allogeneic transplant can be both beneficial and deleterious. GvHD is a cause of morbidity, mortality, and decreased quality of life after allogeneic stem cell transplant. GvHD is caused by donor T cells that recognize antigenic disparities between the donor and recipient and cause a reaction. GvHD causes selective epithelial damage of target organs including the skin, liver, and the gastrointestinal tract. Acute GvHD occurs within the first 100 days after transplant. GvHD is a clinical diagnosis with laboratory (biopsy) confirmation. Risks for GvHD include HLA and gender disparity, increased age of the donor or recipient, and higher quantity of T cells in the product. T cells are the main mediators of GvHD; however, although T-cell depleted grafts decrease the risk of GvHD, they increase the risk of engraftment failure. Glucocorticoids are the main treatment for GvHD; antithymocyte globulin, mesenchymal stem cell, and photophoresis have also been used to treat steroid-refractory GvHD with variable responses. The severity of GvHD is inversely related to the risk of relapse. Patients with GvHD have a lower risk of relapse (due to GvL effect).

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GvL is a term that refers to the adoptive immunotherapeutic effect of transplanted allogeneic hematopoietic cells against leukemia cells. This condition parallels GvHD and measures that decrease GvHD also decrease GvL. Monitoring for engraftment and GvHD is important for a transplant program and is required by accrediting bodies such as Foundation for the Accreditation of Cellular Therapy and AABB (formerly American Association of Blood Banks). It is crucial that programs perform outcome analysis (i.e., time to neutrophil and platelet engraftment following product administration) and relate this to product efficacy. Overall and treatment-related morbidity and mortality must be assessed at 100 days and 1 year after transplantation.

Advances in HSCT Although HSCT has made remarkable progress since its inception in the 1950s, there is still room for significant advancement. Efforts to use reduced intensity conditioning regimens, improve stem cell mobilization and homing after transplant, and develop procedures to enhance ex vivo expansion of stem cells are all under way. In addition, manipulation of the stem cell product to include the sorting of cells into subsets (natural killer cells, T cells, regulatory T cells, etc.) may allow the selective infusion of only the desired components, thus avoiding any unwanted complications.

Further Reading Copelan, E. A. (2006). Hematopoietic stem-cell transplantation. N. Engl. J. Med., 354, 1813–1826. Delaney, M., & Haspel, R. L. (2009). Thawing and infusing cellular therapy products. In E. M. Areman, & K. Loper (Eds.), Cellular Therapy: Principles, Methods, and Regulations (pp. 375–382). Bethesda, MD: AABB. Donmez, A., Aydemir, S., Arik, B., et al. (2012). Risk factors for microbial contamination of peripheral blood stem cell products. Transfusion, 52, 777–781. Gratwohl, A., Baldomero, H., Aljurf, M., et al. (2010). Hematopoietic stem cell transplantation: a global perspective. JAMA, 363, 2091–2101. Kumar, S., DeLeve, L. D., Kamath, P. S., & Tefferi, A. (2003). Hepatic veno-occlusive disease (sinusoidal obstruction syndrome) after hematopoietic stem cell transplantation. Mayo Clin. Proc., 78, 589–598. Linenberger, M. L. (2009). Collection of cellular therapy products by apheresis. In E. M. Areman, & K. Loper (Eds.), Cellular Therapy: Principles, Methods, and Regulations (pp. 251–260). Bethesda, MD: AABB. Oran, B., & Shpall, E. (2012). Umbilical cord blood transplantation: a maturing technology. Hematol. Am. Soc. Hematol. Educ. Program, 2012, 215–222. Sauer-Heilborn, A., Kadidlo, D., & McCullough, J. (2004). Patient care during infusion of hematopoietic progenitor cells. Transfusion, 44, 907–916. Spitzer, T. R. (2009). Bone marrow collection. In E. M. Areman, & K. Loper (Eds.), Cellular Therapy: Principles, Methods, and Regulations (pp. 236–250). Bethesda, MD: AABB.

Relevant Websites http://www.aabb.org. http://www.accessdata.fda.gov/scripts/cdrh/cfdocs/cfcfr/cfrsearch.cfm. http://www.celltherapysociety.org/. http://www.factwebsite.org/. http://marrow.org.

Dental Stem Cells M Nakashima and Y Hayashi, Department of Dental Regenerative Medicine, Center of Advanced Medicine for Dental and Oral Diseases, National Center for Geriatrics and Gerontology, Research Institute, Obu, Japan © 2019 Elsevier Inc. All rights reserved.

Introduction: Mesenchymal Stem Cells Dental Stem Cells Dental Pulp Stem Cells Stem Cells from Human Exfoliated Deciduous Teeth Stem Cells from Apical Papilla Dental Follicle Progenitor Cells Periodontal Ligament Stem Cells Isolation of Dental Stem Cells Characterization of Dental Stem Cells Surface Markers Dentin/Pulp Periodontal Ligament Bone Neuron Blood Vessel Trophic Factors Clinical Application Conclusions References

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Glossary Adult stem cells Undifferentiated cells, found throughout the body after embryonic development, that multiply by cell division to replenish dying cells and regenerate damaged tissues. Also known as somatic stem cells, they can be found in juvenile as well as adult animals and humans. Mesenchyme A type of undifferentiated loose connective tissue that is derived mostly from mesoderm, although some are derived from other germ layers, e.g., some mesenchymes are derived from neural crest cells and thus originate from the ectoderm. Neural crest cells A transient, multipotent, migratory cell population unique to vertebrates that gives rise to a diverse cell lineage including melanocytes, craniofacial cartilage and bone, smooth muscle, peripheral and enteric neurons and glia. Regeneration Growth anew of lost tissue or destroyed parts or organs. Repair Improve or restore damaged or injured parts or organs.

Introduction: Mesenchymal Stem Cells Since the term stromal stem cells and mesenchymal stem cells (MSCs) have been introduced in the 1980s, MSCs have been isolated from almost all adult tissues (e.g., bone marrow, adipose tissue, synovium, dermis, periosteum, peripheral blood, solid organs (e.g., liver, spleen, and lung), and teeth). MSCs are rare and quiescent populations in their niche within fully specialized tissues, having self-renewal properties and multilineage differentiation potential in vitro (Dominici et al., 2006). MSCs possess potential immunomodulatory, antiinflammatory, and trophic effects. MSCs are heterogeneous, comprising a subset of stem cells (or different subsets of stem cells) and more committed progenitor cells. According to the International Society for Cellular Therapy, the term ‘multipotent mesenchymal stromal cells’ is used for the plastic adherent cells independent of their origins, and the MSCs should be termed only for the subset. Due to lack of standardization in isolation and culture techniques and definitive and specific markers, the three minimal criteria are proposed for MSCs to be defined: (1) plastic adherence, (2) expression of CD73, CD90, and CD105, and lack of CD11b, CD14, CD19, CD79a, CD45, and HLA-DR expression, and (3) their trilineage (ecto-, meso-, endodermal lineages) differentiation potential (Dominici et al., 2006). Regenerative medicine has especially been paying attention to clinical applications of MSCs-based therapy to repair or regenerate damaged tissues. Advantages of harnessing MSCs are their easy availability, high proliferative activity, fewer of the ethical concerns associated with embryonic stem cells and low immunogenicity. Several clinical studies have already demonstrated therapeutic utility of MSCs in diseases such as graft-versus-host disease (GVHD), myocardial

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infarcts, diabetes, and different types of neurological disorders (Sensebé et al., 2010). The therapeutic effects of MSCs are not restricted to their differentiation ability but depend on their potency to modulate local environment, activate endogenous stem/ progenitor cells, and secrete various trophic factors (Phinney and Prockop, 2007). The regenerative mechanisms, however, are not completely understood. This section will address the current knowledge on MSCs from dental tissues, especially their characterization and regenerative potential.

Dental Stem Cells All stem cells involved in odontogenesis originate in mesenchyme except ameloblast progenitor cells. The mesenchyme of the oral and facial region originates almost exclusively from the paraxial mesoderm up to the third week of development. Neural crest cells of the ectoderm migrate into the hyoid arches in the fourth week, and most of the mesenchyme is ultimately of neuroectodermal origin. These cells are known to have an ectomesenchymal origin (Ulmer et al., 2010). The various MSC populations are isolated from teeth, and include cells from the pulp of both adult teeth (dental pulp stem cells (DPSCs)) and exfoliated deciduous teeth (stem cells from human exfoliated deciduous teeth (SHEDs)) and, from the root apical part of the papilla (stem cells from apical papilla (SCAPs)), from the tissue (dental follicle) that surrounds the unerupted tooth (dental follicle progenitor cells (DFPCs)), and from the periodontal ligament (PDL) that links the tooth root with the bone (periodontal ligament stem cells (PDLSCs)). These distinct populations differ in terms of self-renewal capability, proliferation rate, stem cell marker expression and differentiation potential, etc. Their biological properties are described in more detail in the following.

Dental Pulp Stem Cells The postnatal pulp contains several niches of potential progenitor/stem cells, which is important for mediating reparative dentin formation. There is the ‘true’ or ‘mother’ adult stem cell central to the niche. This subset of undifferentiated cells displays an infrequent, yet unlimited self-renewal, and gives rise to a renewed mother stem cell and a daughter transit amplifying progenitor cell at mitosis. These daughter progenitor cells have more limited self-renewal capacity, but are highly proliferative. Cell–cell and cell– matrix communication is critical to maintain the stem cell status within the niche (Sloan and Waddington, 2009). The niche usually maintains a quiescent state and specific signals derived from precise area of niche permit stem cells to stay alive, and change their number and fate (Scadden, 2006; Chen et al., 2011). The niche resides predominantly in the perivascular and perineural sheath regions of the pulpal cavity (Shi and Gronthos, 2003), from where they migrate to the site of injury (Téclès et al., 2005). Additional niche in the odontoblast–subodontoblast layers and the pulpal stroma, however, is also suggested by elevated expression of Notch signals (Løvschall et al., 2005). DPSCs can be isolated by a colony isolation method from the adult dental pulp (Gronthos et al., 2000), which can be acquired from third molars, pulpectomized teeth or unneeded teeth for the purpose of orthodontic treatment. DPSCs exhibit a self-renewal property, a high proliferative capacity, and a multilineage differentiation potential to give rise to a variety of cell types, such as osteoblasts, chondrocytes, adipocytes, myoblasts, endotheliocytes, and melanocytes, as well as neurons and glia, being of neural crest origin (Ulmer et al., 2010; Kawashima, 2012). In vivo, DPSCs can differentiate into odontoblasts and induce host cells to participate in regeneration by generating dentin–pulp-like complex after subcutaneous transplantation in conjunction with hydroxyapatite/tricalcium phosphate (HA/TCP) into immunocompromised mice (Gronthos et al., 2000, 2002; Batouli et al., 2003). A dentin–pulplike complex with well-established vascularization can be regenerated de novo in emptied root canal by DPSCs ectopically (Cordeiro et al., 2008; Prescott et al., 2008; Huang, 2009a, 2010). DPSCs can also differentiate into osteoblasts and endothelial cells to produce adult bonelike tissue with an integral blood supply (d’Aquino et al., 2007). Differentiation of DPSCs into mature oligodendrocytes or functionally active neurons is demonstrated after transplantation into the embryonic mesencephalon (Arthur et al., 2008). Preincubated DPSCs toward odontogenic, adipogenic, and myogenic lineages differentiate along distinct pathways after transplantation into immunocompromised mice (Kerkis et al., 2006; Zhang et al., 2008). The multilineage differentiation potential of DPSCs strongly suggests their possible applications in regenerative medicine. The colony-derived populations of DPSC are heterogenous and contain more than one stem cell population. Therefore, various strategies have been tried to isolate a more defined clonal subset of stem/progenitor cells using immunoselection of some cell surface antigen markers such as STRO-1 (Shi and Gronthos, 2003; Laino et al., 2006; Yang et al., 2007), and CD105 (Nakashima et al., 2009; Iohara et al., 2011) by flow cytometry or magnetic-activated cell sorting. Isolation of the ‘true’ or ‘mother’ adult stem cells, which exhibited lower level of the DNA-binding fluorescent dye, Hoechst 33342, than that of the rest of the pulp cells (Iohara et al., 2006), is an alternative method. So far, at least two different stem cell populations are identified. One is originating from the neural crest (ectomesenchyme) and the other is of mesenchymal origin. However, the isolation results of different stem populations are still inconclusive due to lack of specific cell surface markers. Thus, a novel isolation method is necessary to be developed before subsequent characterization of distinct DPSC populations.

Stem Cells from Human Exfoliated Deciduous Teeth SHEDs are isolated from dental pulp derived from exfoliated deciduous teeth. Compared to DPSCs, SHEDs exhibit higher growth rates, and differentiate into a variety of cell types to a greater extent, including neural cells, chondrocytes, adipocytes, osteoblasts, and

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myocytes (Miura et al., 2003; Kerkis et al., 2006; Wang et al., 2010, 2012a; Nourbakhsh et al., 2011). Higher expression of pluripotent markers such as Oct4, Sox2, Nanog, and Rex1, lower neurosphere formation, and lower expression of neuronal markers after neuronal induction are found in SHEDs compared with DPSCs (Govindasamy et al., 2010). The use of embryonic stem cells is associated with ethical concerns, and autologous postnatal stem cells with multipotency have the limitations of readily available sources. On the other hand, deciduous teeth are disposable and readily accessible and SHED is an attractive alternative. When SHEDs are injected into tooth slice with scaffold and transplanted subcutaneously into immunocompromised mice, SHEDs are able to differentiate into odontoblasts and endothelial cells to generate dentin–pulp-like tissue (Sakai et al., 2010). SHEDs also induce recipient murine cells to differentiate into bone-forming cells to repair critical-size bone defects (Miura et al., 2003; Seo et al., 2008). SHEDs are capable of differentiating into blood vessels that anastomosed with the host vasculature (Cordeiro et al., 2008), neuronal developmental potential is demonstrated by injecting SHEDs into the dentate gyrus of the hippocampus of immunocompromised mice (Miura et al., 2003). Transplantation of SHED spheres into the rat striatum of Parkinson disease partially improves the apomorphineevoked rotation of behavioral disorders (Wang et al., 2010). These features and accessibility of tissue resource make SHEDs a potential source of stem cells to repair and regenerate vasculogenic and neurogenic diseases and damaged tooth structures.

Stem Cells from Apical Papilla A unique population of dental stem cells known as stem cells from the root apical papilla is residing in the tips of growing tooth roots of the permanent immature teeth (Sonoyama et al., 2008). The apical papilla tissue is only present during root development before or during the tooth erupts into the oral cavity. In developing teeth, root formation starts as the epithelial cells from the cervical loop proliferate apically, and influence the differentiation of odontoblasts from undifferentiated mesenchymal cells and cementoblasts from follicle mesenchymal cells (Linde and Goldberg, 1993). There is an apical cell-rich zone lying between the dental pulp and the apical papilla. The distinction between the dental pulp and the apical papilla is that the apical papilla represents a precursor tissue for the radicular pulp. SCAPs proliferate faster with greater population doubling than DPSCs (Sonoyama et al., 2006). SCAP have osteo/dentinogenic, neurogenic, and adipogenic differentiation potential, but not myogenic and chondrogenic differentiation potential (Sonoyama et al., 2006; Abe et al., 2011). SCAPs, similar to DPSCs, are more committed to osteo/dentinogenicity. SCAPs express osteo/dentinogenic markers and growth factor receptors similar to DPSCs, but express lower level of dentin sialophosphoprotein, matrix extracellular phosphoglycoprotein, transforming growth factor (TGF) bRII, fibroblast growth factor receptor (FGFR) 3, and Flt-1 (VEGFR-1) than do DPSCs (Sonoyama et al., 2008). In clinical case of apexogenesis, SCAPs give rise to odontoblasts to induce root formation since SCAP residing in the apical papilla survive in an infected immature tooth with periapical periodontitis (Chueh and Huang, 2006). When SCAPs are transplanted into immunocompromised mice with HA/TCP as a scaffold, they differentiate into odontoblasts to form the dentin structure (Sonoyama et al., 2006). Most human tissues from early in their development are not clinically available for stem cell isolation. On the other hand, the root apical papilla is accessible postnatally from extracted wisdom teeth. The embryonic-like properties (i.e., in the process of development) and the higher proliferative potential of SCAP make this population suitable for cell-based regeneration and preferentially for forming root.

Dental Follicle Progenitor Cells The dental follicle is a loose ectomesencyme-derived connective tissue sac surrounding the enamel organ and the dental papilla of the developing tooth germ before eruption. The dental follicle plays a crucial role in tooth development and contains progenitors for cementoblasts, PDL fibroblasts, and osteoblasts. DFPCs are isolated from the dental follicle of impacted third molars or during their tooth eruption usually between 15 and 28 years of age (Morsczeck et al., 2005; Yao et al., 2008). DFPCs are fibroblastlike, colony-forming, and plastic-adherent cells, and have the ability to form compact mineralized nodules in vitro. Putative stem cell markers, Notch-1, nestin, STRO-1, and FGFRI-IIIC are expressed (Morsczeck et al., 2005). DFPCs have the ability to differentiate toward cementoblasts, PDL fibroblasts, osteoblasts, adipocytes, and neurons, showing greater plasticity than other dental stem cells (Morsczeck et al., 2005, 2008, 2009; Kémoun et al., 2007; Yao et al., 2008; Coura et al., 2008). DFPCs differentiate into PDL fibroblasts that secrete collagen and interact with fibers on the surfaces of adjacent bone and cementum to form PDL. DFPCs can form PDL after transplantation into severe combined immune deficiency (SCID) mice (Yokoi et al., 2007), but hard tissues such as dentin, cementum, or bone are not identified (Yagyuu et al., 2010). Dentin noncollagenous proteins extracted from dentin can stimulate DFPCs to differentiate into cementoblast lineages. Col-I facilitates the differentiation of DFPCs along the mineralization process. Enamel matrix derivatives or bone morphogenetic protein (BMP)-2/-7 activate DFPCs toward the cementoblastic phenotype (Kémoun et al., 2007). DFPCs are heterogenous populations containing highly undifferentiated state of PDL-type lineage and cementoblastic or alveolar bone osteoblastic lineage, suggesting they play a role in tissue regeneration as much as the individual lineages might do. DFPCs transplanted into the alveolar fossa in conjunction with a dentin matrix-treated scaffold contribute to the formation of rootlike tissues with a pulp–dentin complex and a PDL connecting a cementum-like layer to host alveolar bone (Guo et al., 2012b). Further research is necessary for potential uses of DFPCs.

Periodontal Ligament Stem Cells The PDL is a fibrous connective tissue derived from the dental follicle and originates from neural crest cells. PDL connects the cementum to alveolar bone and maintains the root of the tooth in the alveolar socket by acting as a ‘shock absorber’ during

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mastication. PDL contains progenitor cells that migrate and differentiate into either cementoblasts or osteoblasts in response to lesions (McCulloch, 1985). A population of stem cells from PDL (PDLSCs) has been identified (Seo et al., 2004). PDLSCs can differentiate into cementoblasts, adipocytes, odontoblasts, chondrocytes, myotubes, neurons, astrocytes, and oligodendrocytes under defined culture conditions (Seo et al., 2004; Sonoyama et al., 2006; Gronthos et al., 2006). In vivo, they generate cementum/PDL-like structures similar to native PDL as a thin layer of cementum after transplantation with HA/TCP into immunocompromised mice (Seo et al., 2004). After transplantation into the sockets of the mandible in minipig models, autologous PDLSCs with HA/TCP and gelform scaffolds form a bioroot encircled with PDL connecting to the surrounding bone (Sonoyama et al., 2006). After transplantation into apical involvement defects of a canine advanced periodontitis model, PDLSCs show better regenerating capacity of PDL, alveolar bone, and cementum as well as peripheral nerves and blood vessels in the regenerated PDL space compared with DPSCs (Park et al., 2011). These findings suggest that PDLSCs have a potential function of repair and regeneration of PDL.

Isolation of Dental Stem Cells Dental stem cells can be isolated by using the explant outgrowth technique or the enzymatic digestion technique. The former technique is that the tissues cut into 2-mm3 pieces are directly incubated in culture dishes and stem cells migrating out from the tissues are collected after 2–3 weeks (Stanislawski et al., 1997; Spath et al., 2010). The latter technique is that the cell suspensions after digestion with collagenase or a combination of collagenase and dispase are seeded in culture dishes to be colonized (Gronthos et al., 2000). These populations are highly proliferative but heterogenous in cellular differentiation and multipotentiality. Therefore, a more defined clonal subset of stem/progenitor cells are isolated using immunoselection of some cell surface antigen markers by flow cytometry or magnetic-activated cell sorting. Those cell surface markers include STRO-1 (the perivascular cell marker), CD34 (the hematopoietic/endothelial marker), c-kit/CD117 (the putative stem/progenitor cell protooncogene marker), b1 integrin (preferentially expressed in primitive cells), low-affinity nerve growth factor receptor (LANGFR/p75/CD271) (the embryonic neural crest cell marker), CD105/endoglin, and stage-specific embryonic antigen (SSEA)-4. Furthermore, the ‘true’ or ‘mother’ adult stem cells, side population cells, are isolated from dental pulp and PDL by flowcytometry using the DNAbinding fluorescent dye Hoechst 33342 exclusion (Nakashima et al., 2009; Hynes et al., 2012; Kawanabe et al., 2012). Those fractionated subpopulations show higher proliferation and migration activities, a longer proliferative life span, and greater multidifferentiation potential including angiogenesis and neurogenesis compared with unfractionated colony-derived populations (Nakashima et al., 2009; Iohara et al., 2011). There are two different origins in dental stem cells: mesenchymal stem/progenitor cells and more embryonic neural crest stem cells (Coura et al., 2008; Degistirici et al., 2008; Waddington et al., 2009; d’Aquino et al., 2011; Dangaria et al., 2011; Janebodin et al., 2011; Abe et al., 2012; Achilleos and Trainor, 2012; Kaku et al., 2012). SSEA-4þ cells isolated from dental follicle have all the characteristics of neural crest cells and share some features with embryonic stem (ES) cells: high level of embryonic stem markers (TRA-1-60, TRA-1-81, OCT-4, SSEA-4), mRNA expression of Nanog and Rex1, pluripotency in vitro and in vivo, integration into the inner cell mass of blastocytes, high level of telomerase activity, and ability to form embryoid bodies (d’Aquino et al., 2011). Fractionated stem cells might be more suitable for some tissue regeneration since some regenerated tissues are much higher in volume compared with unfractionated colony-derived populations (Iohara et al., 2011). A novel simple method should be developed prior to clinical trials using cost-effective and potential disposable apparatus, by which dental stem cells can be isolated from a small amount of tissue with efficiency and in safety. Potential utility of differences in migratory activity among the different cell populations in place of cell surface antigen marker expressions is recently demonstrated (Murakami et al., 2013).

Characterization of Dental Stem Cells Surface Markers (Table 1) Dental stem cells are composed of more than 95% of CD29þ, CD44þ, CD73þ, CD90þ, and CD105þ cells. Less than 2% of dental stem cells express the panleukocyte marker CD45, the hemotopoietic/endothelial cell marker CD34, the monocyte and macrophage markers CD11 and CD14, the B-cell marker CD79 and CD19, or HLA class II. Different types of dental stem cells share common stem cell antigen marker expressions (Huang et al., 2009b; Kerkis and Caplan, 2012). Nevertheless, the data currently available on other markers such as CD34, CD117, and CD146, are inconsistent, and the clear definition of an exclusive population of dental stem cells is difficult. This is based on the following facts: (1) Each type of dental stem cells contains several stem cell subpopulations. (2) The cells are related to the plastic adherent and cultured population, which changes the phenotype easily during cell culture. (3) The heterogeneity also depends on isolation and culture methods, which are varied in different laboratories. There are many variables that affect the final outcome in the composition of subpopulations: initial plating density, coating of culture dishes, composition of cell culture media (glucose content, calcium concentration, supplementation with antioxidants), supplement (bovine serum, human serum, platelet lysate, or growth factors), oxygen supply (hypoxia), and method of subculturing and cryopreservation. Therefore, standardization of the isolation and culture procedure is needed for good reproducibility of results from different laboratories and studies.

558 Table 1

CD(þ)

CD()

Regenerative Engineering j Dental Stem Cells Surface markers of dental stem cells

CD10 CD13 CD24 CD29 CD44 CD59 CD71 CD73 CD90 CD105 CD106 CD117 CD146a CD166 CD271 STRO1 CD3 CD14 CD18 CD31 CD34 CD45 CD150

DPSC ( Lindroos et al., 2008; Nakashima et al., 2009; Rodrı´guez-Lozano et al., 2011)

SHED ( Govindasamy et al., 2010; Rodrı´guezLozano et al., 2011; Sucha´nek et al., 2010)

B

B

B B B

B B

B B (5–10%)

B B B B B B B B B B (9%)

B

B B

B B B

B B B

B B B B

SCAP ( Huang et al., 2009b; Rodrı´guezLozano et al., 2011) B B B B B B B B

DFPC ( Guo et al., 2012a; Rodrı´guezLozano et al., 2011)

PDLSC ( Huang et al., 2009b; Rodrı´guez-Lozano et al., 2011)

B B

B B

B B B

B B B

B B B

B B B

B

B

B

B

B B

B B

B B (>18%)

B B B B

a CD146 is reported in Nakashima, M., Iohara, K., Sugiyama, M., 2009. Human dental pulp stem cells with highly angiogenic and neurogenic potential for possible use in pulp regeneration. Cytokine Growth Factor Rev. 20, 435–440.

Regenerative Potential (Table 2) Dental stem cells offer potential for regeneration of damaged tooth tissues such as dentin, pulp, and PDL. Bone, nerve, and blood vessel are also repaired or regenerated by dental stem cells.

Dentin/Pulp DPSCs, SHEDs, and SCAPs are derived from pulp tissue or the precursor of pulp, and have dentin/pulp repair/regeneration potential (Huang et al., 2009b; Volponi et al., 2010; Estrela et al., 2011). When DPSCs are transplanted alone or together with BMP2 on the exposed pulp in the cavity, the dentin–pulp complex is induced (Nakashima and Akamine, 2005; Iohara et al., 2006, 2009). DPSCs and SCAPs form the dentin–pulp complex after subcutaneous transplantation into immunocompromised mice, whereas SHEDs form mineralized tissue without the distinct pulp–dentin complex (Huang, et al., 2009b). In ectopic tooth transplantation models, dentin–pulp complex with well-established vascularity can be regenerated de novo in emptied root canal space of tooth slice or tooth root by either DPSCs or SHEDs (Huang, 2011). In orthotropic pulpectomized tooth model as a regenerative endodontic therapy, pulp tissue with good vasculature and innervation is completely regenerated after transplantation of DPSCs with cellderived factor-1 (SDF-1) (Iohara et al., 2011). Dentin formation is also induced in the coronal and the apical parts of the regenerated pulp tissue to prevent microleakage as well as along the dentinal wall. Transplantation of DPSCs with SDF-1 yields a significantly larger amount of the regenerated pulp tissue compared with transplantations of bone marrow or adipose-derived stem cells (Ishizaka et al., 2012).

Periodontal Ligament (Hynes et al., 2012) Transplantation of PDLSCs into immunocompromised mice validates the ability of PDLSCs to form functional cementoblast-like cells and cementum/PDL-like tissue, including Sharpey’s fibers. The regenerative capacity of PDLSCs is presented in a range of dental defects in various animal models with scaffold including HA/TCP and collagen, demonstrating the regeneration of normal periodontal tissues, containing organized cementum, alveolar bone, and the PDL attachment apparatus. Multilayered cell sheets of PDLSCs supported by scaffold including polyglycolic acid (PLGA), hyaluronic acid, HA/TCP, and collagen enhance periodontal regeneration with the formation of new cementum and well-oriented PDL fibers. The autologous DFPCs generate new cementum, alveolar bone, and Sharpey’s fibers of PDL into the apical involvement defect. However, PDLSCs have the best regenerative capacity compared with DFPCs and DPSCs, showing good innervation and vascularization in addition to regeneration of PDL, alveolar

Table 2

Biological properties of dental stem cells

Properties

DPSC ( Bakopoulou et al., 2011; Demarco et al., 2011; Gronthos et al., 2000; Iohara et al., 2011, 2009, 2008; Nakashima et al., 2009; Sugiyama et al., 2011; Waddington et al., 2009)

SHED ( Demarco et al., 2011; Morsczeck et al., 2010; Nishino et al., 2011; Nourbakhsh et al., 2011; Wang et al., 2012a; Seo et al., 2008)

SCAP ( Bakopoulou et al., 2011; Demarco et al., 2011; Ding et al., 2010; Sonoyama et al., 2008)

DFPC ( Demarco et al., 2011; Morsczeck et al., 2010; Tomic et al., 2011)

PDLSC ( Feng et al., 2010; Ke´moun et al., 2011; Liu et al., 2008)

Location

Permanent tooth pulp

Exfoliated deciduous tooth

Apical papilla of developing root

Periodontal ligament

Proliferation rate Heterogeneity Multipotency

þ þ Odontoblast, osteoblast, chondrocyte, myocyte, neurocyte, adipocyte, choroneal epithelial cell, melanoma cell, iPS þ Bone regeneration, neural regeneration, myogenic regeneration, dentin–pulp regeneration

þþ þ Odontoblast, osteoblast, chondrocyte, myocyte, neurocyte, adipocyte, iPS

þþ þ Odontoblast, osteoblast, neurocyte, adipocyte, iPS

Dental follicle of developing tooth þþ þ Odontoblast, osteoblast, neurocyte, adipocyte,

þ Bone regeneration, neural regeneration, tubular dentin, wound healing

þ Bone regeneration, neural regeneration, dentin–pulp regeneration, root formation

Migration Regenerative potential

þ Bone regeneration, PDL regeneration

þ Bone regeneration, PDL regeneration, root formation

B

B

B

B

B

B

B B

B

B B

B B

B B

Regenerative Engineering j Dental Stem Cells

Trophic effect Antiapoptosis (Iohara et al., 2008; Jewett et al., 2010; Sugiyama et al., 2011; Wanachottrakul et al., 2011; Zhao et al., 2012) Immunosuppressive (Ding et al., 2010; Ishizaka et al., 2013; Jewett et al., 2010; Morsczeck et al., 2010; Tomic et al., 2011; Wanachottrakul et al., 2011; Zhao et al., 2012) Proliferation Migration

þþ þ Odontoblast, osteoblast, chodrocyte, neurocyte, cementoblast

559

560

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bone, and cementum. DPSCs transplanted into various periodontal defects result in inconsistencies among different research groups. A root/periodontal complex with significantly better compression strength is formed when a root-shaped HA/TCP block containing SCAP coated with PDLSC-seeded gel foam is implanted in tooth socket.

Bone (Ulmer et al., 2010; Hynes et al., 2012; Kawashima, 2012; Kerkis and Caplan, 2012) DPSCs transplanted with HA/TCP, PLGA, collagen, nanofiber hydrogel, or HA nanohydroxyapatite/collagen/poly (L-lactide) exhibit bonelike structure rather than dentin. Pretreatment of DPSCs with BMP2 promotes osteogenesis. SHEDs with a proper scaffold such as platelet-rich plasma have also the ability to form mature bone. PDLSCs show better osteogenic properties than SHEDs when treated with retinoic acid in combination with insulin. Critical size bone defect is repaired by DPSCs and SHEDs.

Neuron Dental stem cells differentiate into neural lineage in vitro. SHEDs injected into the hippocampus of immunocompromised mice after 1 week cultivation in the neuronal differentiation medium express neural markers, indicating neuronal differentiation in vivo (Miura et al., 2003). DPSCs transplanted into the mesencephalon of day-2 chicken embryo acquire a neuronal morphology and function, suggesting DPSCs exposed to the appropriate environmental cues differentiate into active neuron (Arthur et al., 2008). DPSCs coordinate axon guidance via CXCR-4 and stromal SDF-1/CXCL12 axis to induce neuroplasticity within a receptive host nervous system (Arthur et al., 2009). Transplantation of DPSCs into the hippocampus promotes proliferation, cell recruitment, and maturation of endogenous neural stem/progenitor cells by secreting neurotrophic factors, suggesting their therapeutic potential as a stimulator and modulator of the local repair response in the central nervous system (Huang et al., 2008). Significant recovery from neurological dysfunction in cerebral ischemia and spinal cord injury models are reported after transplantation of DPSCs or SHEDs, suggesting their neural regenerative potential in the central nervous system (Yang et al., 2009; de Almeida et al., 2011; Sugiyama et al., 2011, Sakai et al., 2012). In a peripheral nerve injury model, artificial nerve conduits containing DPSCs promote nerve regeneration with myelinated fibers (Sasaki et al., 2008, 2011).

Blood Vessel DPSCs and SHEDs accelerate blood flow and vasculogenesis/angiogenesis in an ischemic hind-limb model and a peripheral nerve injury model (Iohara et al., 2008; Nakashima et al., 2009; Sasaki et al., 2008). DPSCs differentiate into osteoblasts and endotheliocytes synergically after transplantation into immunocompromised rats and lead to generation of adult bone structure with an integral blood supply, suggesting that angiogenesis may be regulated by distinct mechanisms (d’Aquino, 2007). When SHEDs are seeded in biodegradable scaffolds within tooth slices and are transplanted into immunodeficient mice, they also differentiate

Table 3

Soluble factors secreted by dental stem cells Soluble trophic factors and cytokines

BDNF (Iohara et al., 2011; Ishizaka et al., 2012, 2013; Sakai et al., 2012; Sugiyama et al., 2011) EGF (Kim et al., 2012) FGF2 (Kim et al., 2010, 2012; Nakamura et al., 2009) GDNF (Gale et al., 2011; Howard et al., 2010; Ishizaka et al., 2012, 2013; Sakai et al., 2012; Sugiyama et al., 2011) GM-CSF (Iohara et al., 2008, 2009, 2011; Ishizaka et al., 2012, 2013) HGF (Iohara et al., 2008; Su et al., 2012) IGF1 (Kim et al., 2012; Ochiai et al., 2012) MMP3 (Iohara et al., 2008, 2009, 2011; Ishizaka et al., 2012, 2013) NPY (Iohara et al., 2008; Iohara et al., 2011; Ishizaka et al., 2012, 2013) NGF (Iohara et al., 2011; Ishizaka et al., 2012, 2013; Kim et al., 2010, 2012; Nakamura et al., 2009; Sakai et al., 2012; Sugiyama et al., 2011) PDGF (Iohara et al., 2008; Kim et al., 2010, 2012) TGF-beta (Derringer and Linden, 2007; Gale et al., 2011; Howard et al., 2010; Kim et al., 2012; Nakamura et al., 2009; Ochiai et al., 2012) VEGF (Derringer and Linden, 2007; Dissanayaka et al., 2012; Iohara et al., 2008, 2009, 2011; Ishizaka et al., 2012, 2013; Kim et al., 2010, 2012; Zhou et al., 2004)

DPSC

SHED

SCAP

DFPC

PDLSC

Brain-derived neurotrophic factor

þ

þ

þ

þ

þ

Epidermal growth factor Fibroblast growth factor 2 Glial cell-line derived neurotrophic factor

þ þ þ

þ þ

þ

Granulocyte macrophage colony-stimulating factor Hepatocyte growth factor Insulin-like growth factor 1 Matrix metalloproteinase-3

þ þ þ þ

þ

þ

þ þ þ

þ

Neuropeptide Y

þ

Nerve growth factor

þ

þ

þ

þ

þ

Platelet-derived growth factor beta Transforming growth factor beta

þ þ

þ þ

þ

þ

þ

Vascular endothelial growth factor

þ

þ

þ

þ

þ

þ þ þ

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into endothelial cells in addition to odontoblast-like cells in the regenerated pulp tissue within the slices (Cordeiro et al., 2008; Sakai et al., 2010).

Trophic Factors (Table 3) Among the various types of cell-to-cell signaling, paracrine signaling transmits over short distances between different cell types. Dental stem cells secrete a broad panel of growth factors and cytokines and provide instructive cues via paracrine signaling. The factors secreted by dental stem cells suppress the local immune system, inhibit apoptosis, enhance mobilization of endogenous stem/progenitor cells, and promote regeneration including angiogenesis and neurogenesis. These immunomodulatory, antiapoptotic, migratory, angiogenic, and neuroprotective/neuroregenerative effects, which are referred to as trophic effects (Nakashima et al., 2009; Nakashima and Iohara, 2011, 2014; Nakashima and Huang, 2013), are distinct from the direct differentiation of dental stem cells into repair/regenerated tissue. Studies in vivo in animal models of pulpectomy, ischemic hind limb, ischemic cerebrum, and spinal cord injury indicate these trophic effects of dental stem cells (Iohara et al., 2011; Sugiyama, 2011; Ishizaka et al., 2012, 2013; Inoue et al., 2012). This understanding may lead to the rational design of new therapies utilizing dental stem cells for damaged tissue and degenerative disorders.

Immunomodulatory Roles (Table 3) (Kerkis and Caplan, 2012; Wang et al., 2012b) The two outstanding hallmarks of dental stem cells (DSCs) and MSCs are the homing to sites of injury and ischemia and immunomodulation. The immunomodulatory role of DSCs and MSCs is due to (1) activated T-cell apoptosis via the FAS ligand/FASmediated death pathway, (2) upregulation of T regulatory cells (Tregs), which results in immune tolerance, and (3) suppression of natural killer cells, as these are effector cells of innate immunity and play a key role in cytotoxic potential and secretion of preinflammatory cytokines such as tumor necrosis factor a. While the suppressive effects of DSCs and MSCs on T lymphocytes are known, their effect on B-lymphocytes is not clear. In some studies, MSCs inhibit B-cell proliferation, whereas in other investigations they have a stimulatory role on proliferation of B lymphocytes. Allogeneic MSCs have been used in immunotherapy in systemic lupus erythematosus, rheumatoid arthritis, and multiple sclerosis. It is noteworthy that allogeneic MSCs have been used to treat GVHD. Although the immunosuppressive and immunomodulatory roles of DSCs and MSCs are known, it is important to point out that major histocompatibility complex (MHC) class I is low and MHC class II antigens are absent in MSCs. The candidate immunomodulators secreted by DSCs and MSCs include TGF-b, hepatic growth factors, interleukins 6 and 10. In conclusion, the immunomodulatory roles of DSCs have a profound effect on the clinical effectiveness by inhibition of T-lymphocyte function, and upregulation of Tregs stimulating immune tolerance.

Clinical Application Clinical examination has demonstrated the potential efficacy and safety of transplanting autologous PDL progenitor cells with HA/ TCP in the treatment of human periodontitis with deep intrabony defect. All three patients show no inflammation in the treatment area and no systemic disorder associated with transplantation (Feng et al., 2010). Another clinical trial has demonstrated that implantation of a biocomplex composed of DPSCs and a collagen sponge scaffold resulted in optimal bone repair and complete bone regeneration in oromaxillofacial bone defect (d’Aquino et al., 2009).

Conclusions Apart from all the basic scientific evidence on dental stem cells accumulated during the last decade, we are at the heralding of a new era of stem cell therapy. Dental stem cells are one of the most potential adult stem cells. Preclinical studies have started autologous or allogenic cell therapies harnessing dental stem cells. Further studies using in vivo models are needed to augment the understanding of the regulation of the migration, growth, and differentiation of dental stem cells or trophic effects of dental stem cells by interactions with resident cells and immune cells, growth factors, and cytokines during regeneration/repair. The lack of standardized isolation methods and culture protocols needs to be also overcome to guarantee high cell quality. A safe and efficacious method to isolate good manufacturing practice-grade dental stem cells should be developed prior to clinical trials. The cost consuming manufacturing needs to be evaluated and improved before regenerative dentistry/medicine can replace the conventional treatment. One possibility might be off-the-shelf allogeneic stem/progenitor cells and banking for every day usage. Dental stem cells are on the direct path to their clinical usage for the regenerative treatment of a multitude of diseases and injuries.

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Drug and Gene Delivery for Regenerative Engineering Morgan A Urello, Tianzhi Luo, Bing Fang, Kristi L Kiick, and Millicent O Sullivan, University of Delaware, Newark, DE, United States © 2019 Elsevier Inc. All rights reserved.

Introduction Drug- and Gene-Delivery Mechanisms Diffusion-, Degradation-, and Swelling-Controlled Drug Release Microenvironment-Stimulated Drug Release Hydration and swelling pH-Responsive Thermoresponsive Externally Stimulated Drug Release Thermoresponsive systems Light-responsive Magnetic or electric field-responsive Cell-Triggered Drug Release Cellular interactions Cellular by-product-responsive Drug- and Gene-Delivery Systems Ceramics Hydrogels Electrospun Fibrous Scaffolds Micro and Nanocarriers Types of Therapeutic Molecules in Drug and Gene Delivery Small-Molecule Drugs Protein-Based Therapeutics Growth factors Monoclonal antibodies Nucleic Acids Genes RNAi Delivery System Selection/Comparison Therapeutic Delivery in Chronic Wound Care Delivery System Small-Molecule Drug Delivery Protein-Based Therapeutic Delivery Gene-Based Therapeutic Delivery Current Obstacles and Future Applications Further Reading

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Glossary Controlled therapeutic delivery Spatiotemporally controlled administration of therapeutics to avoid off-target effects and achieve localized drug delivery within the desired ranges of concentration and duration. Gene Primary sequence of nucleic acids that encodes a protein. Growth factor A class of signaling proteins that stimulate cell proliferation, differentiation, survival, inflammation, and/or tissue repair. Monoclonal antibody Antibodies produced from a single cell clone with monovalent affinity to a single epitope on a targeted antigen. Regenerative engineering Restoration or regeneration of living tissue through integration of materials science and biochemical signaling. “Smart” polymers Polymers purposefully engineered to change chemically and/or physically in response to changes in temperature, pH, light, or other physicochemical stimuli (e.g., “stimuli-responsive polymers”). siRNA Small RNA with the capacity to bind and catalyze degradation of target mRNA sequences through assembly of the RNAinduced silencing complex (RISC).

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Substrate-mediated delivery The administration of a drug or biologic within a scaffold to promote localized, controlled delivery. Systemic delivery The administration of medications, drugs, biologics, or drug-loaded nanocarriers into the circulatory system.

Introduction The delivery of drugs, such as small molecules and growth factors, represents a vital part of regenerative engineering approaches aimed at restoring or regenerating living tissue through integration of materials science and biochemical signaling. Efficient drug administration aids in the activation of specific reparative pathways that promote cellular recruitment, proliferation, and differentiation within the injured or diseased site. Moreover, in more recent years, gene delivery has garnered increasing interest in regenerative engineering by providing a means to reprogram cells and propagate/promote desired cellular lineages and behaviors. Systemic administration of therapeutic agents often results in rapid clearance from circulation and/or degradation before activity is manifested. To overcome these obstacles, therapeutics are generally administered using extraphysiological and repetitive dosing regimens; however, such routines may cause serious side effects including immune activation and carcinogenic responses. For instance, the only FDA-approved recombinant platelet-derived growth factor (PDGF) therapy for diabetic neuropathic ulcers has been demonstrated to reduce ulcer healing time by 32%, yet PDGF application has also been linked to a fivefold increase in cancer-related deaths in diabetic patients. Advances in biomaterials science have provided more elegant solutions for combating the aforementioned issues by providing platforms for targeted, sustained delivery of therapeutics with tight spatial and temporal control. Traditional regenerative engineering approaches employ a wide variety of biomaterials, ranging from ceramics to polymers, that form 3-D structures to guide repair. Ideally, implanted biomaterials provide a scaffold for cellular attachment, proliferation, and differentiation while also serving as a platform for controlled therapeutic delivery. Over the past 25 years, advances in biomaterials science have greatly expanded the choice of biomaterials, such that polymeric and inorganic materials can be engineered with a range of mechanical properties, degradation rates, and chemical functionalities. Increased biomaterial diversity has allowed for selective control of cell behaviors and also has permitted therapeutic delivery with greater spatial and temporal control. Biomaterials can regulate delivery kinetics via a multitude of mechanisms including diffusion-based, environmentally responsive, and celltriggered release. Within therapeutic delivery systems, bioactive molecules may be physically entrapped or immobilized though a variety of mechanisms, depending upon both the material design and the therapeutic agent’s properties. Nonspecific electrostatic, van der Waals, and hydrophobic interactions have been shown to facilitate the immobilization of various proteins and DNA vectors onto different biomaterials and to successfully facilitate sustained therapeutic delivery. Alternatively, biomaterials and therapeutics have been engineered to contain complementary functional groups to mediate binding, including biotin–avidin pairs, antibody–epitope pairs, and extracellular matrix (ECM)-specific targeting. Such approaches may enable substrate specific binding, reduce off-target effects, and improve control over release kinetics. The goal is to design delivery approaches able to provide the required duration and dosage of therapeutic delivery while also maintaining and/or enhancing the specific biological activities of the therapeutic. For example, the biological activity of growth factors is dependent not only on local concentration but also upon the context of growth factor presentation within the cellular microenvironment. While some growth factors are active only when tethered to a biomaterial, others are active only after release and cellular internalization. These differences must be accommodated when choosing or designing an optimized delivery system. Additionally, drug- and gene-delivery systems have employed a wide range of micro and nanocarriers. The administration of smaller devices is typically less invasive, and nanocarriers additionally benefit from unique pharmacokinetic properties that can prolong circulation. Delivery via nanocarriers can also be used to improve hydrophobic drug solubility and bioactivity, which increases the quantity and efficacy of deliverable drug. Furthermore, polymeric and metal-based micro and nanocarriers have been utilized for site-specific therapeutic delivery and/or controlled release from scaffold-mediated delivery systems. The inclusion of cell- and tissue-targeting moieties on these carriers has enabled targeted delivery to multiple organs, diseased tissues, and tumors, while their inclusion in bulk delivery systems has enabled enhanced control over therapeutic delivery kinetics, multitherapeutic delivery, improved therapeutic stability, and cellular uptake. For instance, polyplex-based DNA nanocarriers, formed via electrostatic interactions between DNA and cationic polymers, have been demonstrated in numerous studies to enhance nuclease resistance and improve cellular uptake, resulting in improved gene-transfer efficiency. This review will provide an overview of the diverse array of drug- and gene-delivery systems in regenerative engineering, with a focus on material-mediated delivery systems. Micro/nanocarriers will be discussed in the context of scaffold-mediated delivery. For more information on their micro/nanocarrier design following systemic administration, the suggested reading, including a review of stimuli-responsive nanocarriers by Ganta, S. et al., can be referenced. Different delivery mechanisms and smart material designs will be discussed broadly, followed by an overview of delivery system selection based on the objective of the application and the properties of the therapeutic. The article will conclude with an application-focused section detailing delivery system design and application in chronic wound repair, an area of great interest in regenerative engineering.

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Drug- and Gene-Delivery Mechanisms The development of an appropriate drug delivery system plays a vital role in determining the rate and efficacy of tissue regeneration. The release kinetics and bioactivity of the therapeutics, as well as how the delivery systems interact and integrate into the surrounding host environment need to be considered. In most cases, release is controlled at least in part by diffusion, where net molecular transport is driven by a concentration gradient, and delivery kinetics are largely controlled by the fluid properties and physical characteristics within the scaffold. Scaffold degradation and swelling properties largely influence release. The advent of “smart” or stimuli-responsive polymers, which display significant physiochemical changes in response to changes in their environment, has provided a means for triggering release using either external stimuli, such as heat, light, or a magnetic field, or microenvironmental stimuli, such as pH changes, presence/absence of enzymes, or temperature changes. Furthermore, some delivery systems have been designed to respond to cell-mediated stimuli, including protease release and generation of reactive oxygen species (ROS). The types of stimuli-responsive systems are summarized and compared in Fig. 1.

Diffusion-, Degradation-, and Swelling-Controlled Drug Release Diffusion regulates drug release from most delivery systems. Therapeutics must pass through the water-insoluble material (e.g., a ceramic or polymer-based matrix) that forms the delivery device for release to occur. Diffusion can be controlled at a macromolecular level, such as through pores in a matrix, or on a molecular scale, by passing between polymeric chains. Diffusion-controlled delivery systems are typically categorized as either matrix-based or reservoir diffusion systems. Within matrix-based systems, the therapeutic is distributed throughout the delivery scaffold. Water permeation promotes uniform swelling, leading to a volume expansion in the bulk polymeric matrix and a subsequent increase in matrix pore size. Alternatively, within reservoir diffusion systems, the therapeutic is encapsulated within a permeable polymer membrane. Swelling is observed as a nonuniform volume expansion, wherein only the membrane region allows water permeation and swelling. Within both systems, effective diffusion only occurs when the pore size of the swelled matrix is considerably larger than the dimensions of the therapeutic molecule. In many diffusion-based systems, release is characterized by an initial burst release phase followed by a steady, slower release phase. Within hydrophilic or porous matrices, this two-phase release is the result of an initial period of rapid water uptake that triggers the burst release, followed by uniform degradation throughout the matrix structure leading to steady drug release. Alternatively, within hydrophobic and inert matrices, the materials behave as reservoir systems in which zero-order release kinetics are obtained without an initial burst release phase due to the surface-only wetting phenomenon. However, bioactive, biodegradable materials are typically employed in regenerative engineering. To improve control over therapeutic release kinetics while avoiding an initial burst release phase, various approaches have been developed. One approach is to modify the microstructure of the delivery scaffold. A wide range of cross-linking techniques, including UV photopolymerization and various physical and chemical cross-linking approaches, have been used to improve both scaffold stability and therapeutic retention. These cross-linked structures are better able to achieve sustained, tailorable delivery through physical entrapment of drug within tailorable scaffolds. For example, Y. Tabata and coworkers regulated hepatocyte growth factor (HGF) release from chemically cross-linked gelatin hydrogels by varying the amount of the cross-linking reagent glutaraldehyde and thereby altering the resulting hydrogel degradability. Moreover, when implanted subcutaneously, cross-linked gelatin-mediated delivery of HGF induced significantly enhanced angiogenic activity as compared with free HGF codelivered with a gelatin scaffold lacking any growth factors. Polymer–polymer and polymer–peptide cross-linking strategies are also commonly employed. For instance, polyethylene glycol diacrylate (PEGDA) is

Fig. 1 A summary of the stimuli-responsive drug and gene delivery release mechanisms and the responsive components incorporated into the delivery system.

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commonly cross-linked with thiol-modified natural polymers including gelatin, hyaluronic acid, and chondroitin sulfate. This approach has been implemented by R. A. Peattie and coworkers to mediate the codelivery of bioactive vascular endothelial growth factor (VEGF) and basic fibroblast growth factor (bFGF) to promote enhanced angiogenic activity within murine models. Alternatively, both physical and chemical methods have been employed to modulate the affinity between the therapeutic molecule and the matrix and have achieved better controlled delivery from polymeric networks and porous scaffolds. Nonspecific electrostatic interactions between ionic polymers and charged drugs are commonly used to improve drug retention and delay release. Phosphate- and amino-functionalized polymers have achieved sustained, tailorable release of cationic and anionic drugs, respectively. For instance, A. Concheiro and coworkers demonstrated that the copolymerization of 4-vinylpyridine and poly(hydroxyethyl methacrylate) increased the loading and retention of nonsteroidal anti-inflammatory drugs (NSAIDS) by up to 20-fold and expanded retention from hours to a week without altering the mechanical properties of the hydrogel. In another study, W. Jiskoot and coworkers showed that model protein release from methacrylated-gelatin hydrogels was directly controlled by varying protein– gelatin electrostatic interactions. Alternatively, natural growth-factor binding sites within ECM proteins such as fibronectin, fibrinogen, and collagen have been incorporated into delivery systems to improve therapeutic retention/release profiles. For instance, the growth factor-binding capacity of fibrinogen has been utilized to efficiently deliver low dosages of fibroblast growth factor-2 (FGF2) and placenta growth factor-1 (PIGF-1) from fibrin matrices to facilitate wound healing in diabetic mice (db/db). Furthermore, J. A. Hubbell and coworkers prepared PEG-based matrices functionalized with growth-factor binding sequences isolated from fibrinogen and discovered that these synthetic matrices could sequester multiple growth factors and fully recapitulate the effect of fibrin within a diabetic mouse impaired healing model. In another study, R. J. Levy and coworkers prepared collagen implants functionalized with antiadenoviral antibodies, and they showed that these implants had the capacity to achieve highly localized and sustained gene delivery when applied to stents in pig coronary arteries. Others have achieved sustained delivery through incorporation of particulate systems within polymeric matrices. Within these composite systems, the matrix contains therapeutic molecules that are preloaded in secondary controlled release carriers such as micelles, liposomes, or other particle-based delivery vehicles capable of sustained release. Secondary delivery systems have been fabricated using both synthetic (e.g., polyvinyl alcohol (PVA), poly(lactide-co-glycolide) (PLGA), and polycaprolactone (PCL)) and natural polymers (e.g., gelatin, alginate, and chitosan). These nanocarrier constructs have been successfully used to preserve the bioactivity of the therapeutic, sustain its release, and even enable tailorable, multitherapeutic release. For example, K. T. Nguyen and coworkers demonstrated the codelivery of vascular endothelial growth factor (VEGF) and PDGF-BB from electrospun chitosan/ poly(ethylene oxide) (PEO) fibers, and they showed that codelivery accelerated wound healing in full-thickness rat skin wounds. Specifically, the encapsulation of VEGF and PLGA nanoparticles preloaded with PDGF-BB within the nanofibers facilitated a fast delivery of VEGF and a sustained delivery of PDGF-BB that resulted in significantly enhanced angiogenesis and reepithelization. Additionally, secondary carriers have been functionalized to enable specific, tailorable affinities for both natural and synthetic delivery substrates and thereby enhance control over delivery. For instance, polyethylenimine-DNA (PEI-DNA) polyplexes have been incorporated into collagen by functionalizing the PEI with either biotin or collagen mimetic peptide (CMP), such that collagen-based substrates can be modified with polyplexes through biotin–avidin binding or CMP–collagen hybridization, respectively. The application of functionalized micro and nanocarriers has many additional benefits in regard to preserving therapeutic bioactivity and targeted delivery as discussed later in this article.

Microenvironment-Stimulated Drug Release Advances in materials science over the past three decades have enabled the development of environmentally responsive polymers and, in turn, better integration of therapeutic delivery with delivery site processes. These biomaterials are preengineered to release therapeutic molecules in response to physiologically relevant stimuli, including aqueous solutions and changes in pH and/or temperature. Stimuli may induce simple one-time release or on–off delivery profiles. The response within the material typically occurs through an induction of swelling or degradation or by stimulation of a reversible phase transition that facilitates rapid diffusion or degradation-mediated release of an immobilized therapeutic.

Hydration and swelling Delivery systems are often preengineered to swell upon application in aqueous biological fluids or media and thereby trigger drug delivery. The application of hydrophilic or porous materials facilitates rapid water uptake, and system parameters such as crosslinking density and charge density determine the degree of equilibrium swelling and the therapeutic diffusion rates. For example, increasing the number of ionic groups within a hydrogel is known to increase its swelling capacity. The increase in swelling is caused by an increase in counterions within the gel, which produces an increase in osmotic pressure. To utilize this delivery mechanism, delivery systems are often composed of polymers like hydroxyethylcellulose, whose ionic pendant groups can release loaded therapeutics upon swelling in aqueous biological fluids or media. Other delivery systems have been engineered with hydrolyzable chemical linkages that can be used to regulate scaffold degradation rates or directly immobilize therapeutics via a biodegradable linkage. In hydrolytic degradation, polymer bonds react with water molecules and break apart. Commonly employed hydrolytic polymer types include polyanhydrides, polyamides, and polyesters. The primary factor affecting the rate of hydrolysis is the partial charge of the reactive carbon atom; therefore, different chemical groups have inherently different reactivities and can be chosen based on the desired degradation rate for a drug-delivery application. Furthermore, the hydrolysis rate of ester and amide groups in hydrophilic polymer networks can be altered through

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the incorporation of adjacent charged molecules. For instance, M. P. Lutolf and coworkers reported that varying the charge of hydrolytic cross-links within a polymeric network could promote as much as a 12-fold difference in degradation, leading to significant changes in model protein retention periods ranging from 6 to over 70 days. Additionally, E. Alsberg and coworkers demonstrated the capacity to achieve tailorable and sustained delivery of siRNA by covalent incorporation of cationic linear PEI into photocrosslinked dextran hydrogels through hydrolyzable ester linkages.

pH-Responsive Swelling and collapsing behavior within polymer networks may also be regulated by pH. A local decrease in pH is associated with multiple conditions including inflammation, wound healing, and myocardial ischemia due to the overproduction of lactic acid, increased concentration of acidic by-products from bacterial metabolism, and/or glycolytic activity of infiltrated neutrophils. For this reason, pH is one of the most well-studied stimuli in drug delivery, and a plethora of pH-responsive materials have been documented. Ionic pH-responsive polymers have the capacity to accept or release protons due to changes in pH because their structures consist of either acidic groups, like carboxylic or sulfonic acids, or basic groups, like amines. Thus, the electric charge of the polymer changes in response to pH, and these changes in charge can alter solubility and/or morphology. For instance, polyanions (e.g., albumin, poly(methacrylic) acid (PMAA), and polyacrylic acid (PAA)) contain a large number of ionizable acid groups, which accept protons at low pH and release protons at high pH. Thus, when pH increases, the polymer swells due to electrostatic repulsion of the negatively charged groups, and these pH-driven changes in swelling increase the diffusivity of the therapeutic. A similar effect occurs in cationic polymers (e.g., polylysine, chitosan, and PEI) when the pH decreases. The charge of a polymer at a given pH is based upon its pKa, which in turn is determined by the polymer composition and molecular weight. pH-responsive polymers have been incorporated into bulk delivery systems like hydrogels and electrospun scaffolds, and they have also been incorporated into micro and nanocarriers. Additionally, polycationic materials, such as PEI, polyamidoamine (PAMAM), and other dendrimers, are commonly utilized in nonviral gene delivery. Polycationic polymers are commonly used to complex and protect nucleic acids through electrostatic interactions. The pH-sensitive nature of these polymers can also enable endosomal escape through the “endosomal buffering” mechanism. This mechanism occurs in response to decreases in pH following cellular uptake, when the polymer becomes increasingly charged, and correspondingly, the increased counterion concentration within the endosome increases the osmotic pressure and results in endosomal lysis. Moreover, pH-triggered release has been achieved using acid-sensitive bonds like hydrazones or acetal groups or derivatives of N-ethoxybenzylimidazole that undergo accelerated hydrolysis under mildly acidic conditions. These bonds are used within scaffold cross-linkers or as direct tethers between therapeutics and scaffolds.

Thermoresponsive Thermoresponsive materials are another useful tool in therapeutic delivery. The human body has a narrow range of temperatures, and significant deviations are often indicative of medical problems including infection. Moreover, external sources may be used to heat or cool tissues, and therefore, external heating or cooling can be used to trigger release as discussed in the next section. Thermoresponsive systems are composed of polymers that undergo significant changes in solubility due to the inclusion and interactions of hydrophobic and hydrophilic moieties in the polymer chain. The balance point temperature is referred to as the lower critical solution temperature (LCST), at which point the polymer favors neither hydrogen bonding with the polymer nor hydrogen bonding with water. The most commonly employed temperature-responsive materials include those with LCSTs close to body temperature such as poly(N-isopropylacrylamide) (PNIPAAm), poly(methyl vinyl ether) (PMVE), and elastin-like polypeptides. These materials exhibit sol-to-gel transformationsdfor example, transitions from liquid to solid phasednear body temperature. In turn, these materials have been utilized to achieve noninvasive therapeutic delivery of drugs, proteins, and cells, which can be mixed with the polymer in its soluble state at temperatures below the LCST and subsequently delivered by topical application or injection. Exposure to physiological temperatures induces the phase separation of the polymer and the concomitant formation of a therapeutic gel reservoir, enabling improved control over delivery. This approach is often coupled with additional stimuliresponsive delivery systems to achieve multitriggered therapeutic release as discussed later in the article. Thermoresponsive materials also have been used to achieve oscillatory drug release from hydrogels, electrospun porous scaffolds, gel-based microparticles, and nanoparticles.

Externally Stimulated Drug Release Advances in material design have enabled the release of therapeutic in response to external stimuli such as changes in temperature, light, and electric fields. These responsive systems provide additional spatiotemporal control over delivery and, in some cases, an “on–off” switch for delivery that enables marked improvements in delivery localization and safety.

Thermoresponsive systems Thermoresponsive systems triggered by external stimuli typically respond to a wider range of temperatures than those engineered for responses to microenvironmental stimuli (e.g., body temperature). As previously described, these systems employ polymers such as poly N-isopropylacrylamide, poly(ethylene glycol), and elastin-like polypeptides, engineered to undergo temperature-dependent and reversible sol-gel transitions through the inclusion of hydrophilic and lipophilic moieties within the polymer chain. Polymers with an LCST near but below body temperature are ideal for therapeutic delivery. Therapeutic agents can be mixed with the polymer

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solution in liquid form below the LCST, and the polymer will rapidly form a gel-like therapeutic reservoir following application in response to body temperatures. Moreover, external heat may be applied to achieve oscillatory therapeutic delivery. For instance, S. K. Sia and coworkers demonstrated on-demand remote release within implantable N-isopropylacrylamide-co-acrylamide pellets via ultrasound-induced temperature changes. Other groups have similarly demonstrated remote release from thermoresponsive polymers via topical application of heat or hyperthermia or through inducing localized temperature changes using light or magnetic and electric fields. These noninvasive approaches will be discussed in greater depth in the next sections.

Light-responsive Alternatively, light-sensitive systems have also been developed for on-demand release in response to illumination of a specific wavelength in the ultraviolet, visible, or near-infrared regions. A common strategy is to incorporate chromophores such as o-nitrobenzyl, trisodium salt of copper chlorophyllin, coumarin, or metals. Within these systems, the chromophore absorbs light of the appropriate wavelength. In some cases, the absorbed light dissipates as heat, increasing the local temperature and inducing cargo release from a secondary thermoresponsive component in the materials (e.g., PNIPAAm, PMVE, and elastin-like polypeptides). Alternatively, some light-responsive systems rely upon photosensitive polymers that demonstrate reversible photoisomerization behaviors including open-ring transitions and cis–trans conversions including spiropyran, coumarin, and azobenzene derivative-based systems. These materials can exhibit reversible swelling behaviors and/or “on–off” affinities for drugs in response to light stimulation.

Magnetic or electric field-responsive “On–off” therapeutic release has also been achieved through the reversible application of magnetic fields by incorporating magnetic nanoparticles within polymeric networks. Magnetic field application causes the network to vibrate, which can result in increased diffusion rate of encapsulated therapeutics. Network vibrations can also trigger sol-gel transitions when the magnetic nanoparticles are incorporated into thermoresponsive systems composed of polymers such as PNIPAAm, as the thermoresponsive component will respond secondarily to magnetic field-induced heating. On the other hand, the application of external electric fields has similarly been used to trigger release through incorporation of pH-sensitive polymers with high concentrations of ionizable groups. Upon application of an electric field, the subsequent change in pH disrupts hydrogen bonding between the polymer chains, which in turn can cause polymer degradation or deformation (i.e., swelling or deswelling) and subsequent drug release.

Cell-Triggered Drug Release Within the field of regenerative engineering, many biologically inspired materials utilize native regenerative processes to mediate delivery. Synthetic biomaterials have been designed to contain biological cues, including domains of ECM molecules, growth factors, and protease substrates, to promote cell behaviors, therapeutic retention, or cell-triggered therapeutic release.

Cellular interactions To promote cellular interactions, natural biomaterials like collagen, laminin, and fibronectin are commonly incorporated into the delivery systems. In some applications, cell-adhesive peptide motifs are incorporated into delivery systems and have several advantages over the use of whole proteins. These advantages include enhanced stability against conformational change, control over ligand density and location, minimized immune response, and cost-effectiveness. Common cell-adhesion motifs include RGD (fibronectin, laminin, vitronectin, and various collagen), YIGSR and IKLLI (laminin), GFOGER and DGEA (type-I collagen), and REDV and RGD/PHSRN (fibronectin). In addition to cell adhesion, the incorporation of ligands can affect canonical cell behaviors including proliferation, viability, migration, and therapeutic responsiveness. For instance, synergistic signaling between certain integrins and growth-factor receptors has been linked to significant alterations in cellular adhesion and/or growth-factor signaling. Furthermore, in cases where a therapeutic is stably immobilized within a scaffold-based delivery system, cellular invasion is the key factor in determining delivery kinetics.

Cellular by-product-responsive Cell-triggered therapeutic delivery can also be achieved through the application of materials that are responsive to cellular byproducts, which include proteases, ROS, and reductants. In contrast to materials that respond through chemically induced mechanisms such as hydrolysis, which often produce bulk degradation, proteolytic materials enable localized, cell-triggered degradation and drug delivery. Proteolytic delivery systems often are fabricated from natural materials such as collagen, laminin, or fibronectin that innately contain specific proteolytic domains. Alternatively, such delivery systems can be formed from synthetic materials in which protease-labile peptide sequences (i.e., collagenase- and plasmin-sensitive peptides) are incorporated as cross-linkers. For example, cell-mediated release of VEGF has been reported from PEG-based hydrogels conjugated to the cell-adhesive peptide RGD through the incorporation of matrix metalloproteinase 2 (MMP 2)-labile peptidic cross-linkers. Additionally, altering the density of proteolytic sequences has been shown to control cell-mediated degradation and subsequent therapeutic delivery; as the density of these sites increases, degradation and release also increase. Because many pathological conditions are associated with elevated, cell-produced ROS, delivery systems composed of hydrogels, polymeric nanoparticles, and inorganic nanoparticles have utilized a series of ROS-labile linkers and materials to achieve spatiotemporally controlled release in both the extra- and intracellular environment. For example, C. L. Duvall and coworkers developed

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a thermoresponsive hydrogel with the capacity to release a model drug in an H2O2 dose-dependent manner. Moreover, the potential for producing materials for cell-mediated, multidrug release was achieved by K. Kiick and coworkers using a liposome-PEG hybrid hydrogel whose thioether succinimide cross-links degrade upon exposure to thiols such as glutathione, which is found at elevated concentrations in intracellular compartments.

Drug- and Gene-Delivery Systems The scaffolds employed in regenerative engineering have been synthesized from a diverse array of organic and inorganic polymers, ceramics, and composites. Ideal scaffolds mimic the 3-D structure of the native tissue and have the ability to perform as a delivery platform for therapeutic drugs, proteins, and genes. In this section, recent developments in scaffold-mediated delivery and carrier selection will be discussed.

Ceramics Inorganic 3-D scaffolds and substrates, particularly calcium-based substrates for bone repair, constitute a major thrust in regenerative engineering. Hydroxyapatite (HAp) is the calcium apatite species in bone, making it an especially important naturally occurring mineral in the inorganic scaffold/substrate group. HAp also is an excellent biomaterial in bone tissue engineering due to a series of unique and appealing properties such as biocompatibility, bioresorbability, and osteogenicity. With rational design approaches in scaffolds, controllable delivery and directed differentiation of stem cells can be achieved. HA-based materials can be applied either as implants or implant coatings that enhance local osteogenesis. HA also is an excellent candidate material for the controlled delivery of specific drugs and/or genes at localized sites, and it effectively promotes the differentiation of nearby stem cells into targeted lineages (e.g., chondrogenic and/or osteoblastic cells). HAp colloids also have been developed, as these materials can be used to form coatings or construct the inorganic phase of hybrid materials/composites. Tricalcium phosphate (TCP) is another important member of the calcium-based inorganic scaffold/substrate family. TCP has been used clinically as a bone substitute in grafting operations. It is also considered advantageous as a drug/gene-delivery material due to its excellent biocompatibility and lack of significant inflammatory responses. TCP can transform into HAp through hydrolysis, such that many of the materials in this class are actually a mixture of TCP and/or HAp and other organic materials. Several investigations have shown that incorporating TCP into an organic gel scaffold demonstrates a better wound-healing effect. In many cases, polymer coatings are used in conjunction with TCP-based materials to reduce the brittleness of the pure inorganic phase and to enhance the controllable drug release profile, and the applied polymers also can provide additional drug carrier functions. For example, M. Zhang and coworkers have shown that TCP scaffolds with PCL coatings enable sustained and steady bone morphogenetic protein-2 (BMP-2) delivery over a 2-week period that promoted enhanced osteogenic activity within rat models. Furthermore, to enable controlled deformation of the material to accommodate the geometry of different defect sites, efforts have been made to develop injectable composites containing TCP beads. Calcium ions and inorganic phosphate together have been shown to have a boosting effect on osteogenesis. Reports have also documented the functions of TCP scaffolds when directly used as growth-factor carriers including BMP-2 and VEGF. Biphasic calcium phosphate (BCP) scaffolds contain both hydroxyapatite and beta-tricalcium phosphate. As a mixed material, BCP exhibits many superior properties that combine the benefits of its two components, such as highly hierarchical porosities, biocompatibility, bioresorbability, osteoconductivity, and mechanical strength appropriate for bone grafts. However, in order to enable BCP-based materials to conform to the geometries necessary in various applications, polymer/BCP ceramic composites are often prepared for scaffold use in bone tissue engineering. Furthermore, the FDA-approved bioactive protein, bone morphogenetic protein-2 (BMP-2), has been loaded into BCP-based scaffolds to form a biomaterial with improved osteoinductivity. Micro/ macroporous BCP (MBCP) scaffolds also have been used as carriers for mRNA encoding for hBMP-2, enabling sustained release, enhanced osteogenic stem-cell differentiation, and mineralization. As a bioactive coating, BCP is also functional in drug delivery, triggering early osseointegration in titanium implants. Amorphous calcium phosphate (ACP) is the first solid phase precipitated after the rapid mixing of aqueous solutions containing Ca2 þ and PO3  ions, and ACP also is an essential mineral phase formed in mineralized tissues. It does not have a periodic order of crystalline structures, making it distinct from other forms of calcium phosphates. ACP has high drug adsorption efficiency, and the amorphous surface facilitates release of the bound drug molecules. A variety of interesting and promising studies have been performed with ACP-based scaffolds or composites for tissue engineering. For example, in situ precipitation of ACP and ciprofloxacin crystals in chitosan hydrogels produces a composite scaffold suitable for drug delivery and controlled release of therapeutics like BMP-2. ACP porous microspheres have a high capacity for drug loading and can provide pH-responsive drug release, demonstrating promise for applications in drug delivery. Strontium has been doped with amorphous calcium phosphate to form porous microspheres (SrAPMs). SrAPMs were incorporated into collagen, and the SrAPM–collagen scaffolds could effectively stimulate osteogenesis and promote bone regeneration. In all cases, therapeutics (either drugs or genes) can be incorporated into ceramics/composites in multiple different ways, including physical adsorption, chemical conjugation, and encapsulation. The most straightforward loading procedure is physical adsorption, in which drugs/genes or drug/gene-containing nanocarriers are loaded and retained on ceramic/composite substrates by electrostatic or van der Waals interactions. For example, in aqueous media, hydroxyapatite crystal surfaces gain a positive charge

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due to release of OH– ions. As the majority of proteins present negative charges at physiological pH range, electrostatic interactions between hydroxyapatite and proteins can be readily established in aqueous condition. The roughness of the hydroxyapatite helps protein adsorption. Moreover, with appropriate modification/functionalization, the inorganic carriers also are able to chemically bond to the cargoes. For instance, L. Bachas and coworkers demonstrated biphosphonate linkers enable oriented immobilization of proteins on to HAp through hydrazine bonding. In other examples, the HAp surface is modified with amino acid, which can chemically connect to growth factors through carboxyl–amine chemistry.

Hydrogels Hydrogels are three-dimensional hydrophilic networks capable of absorbing a large amount of water. These networks are composed of polymeric chains that are cross-linked via chemical conjugation or alternatively by physical interactions such as entanglements, crystallites, van der Waals interactions, and/or hydrogen bonding. Because of their excellent biocompatibility, biodegradability, nontoxicity, and critical role in the ECM, native macromolecules such as sugars and proteins are widely used as the building blocks for hydrogels. Polysaccharides including alginate, chitosan, cellulose, dextran, and hyaluronic acid have been widely studied, as have proteins such as collagen and gelatin. However, although natural materials exhibit a variety of well-defined hierarchical structures and important biological functions, such as cell recognition and adhesion, their use in therapeutic applications can be limited by immune responses and susceptibility to enzymatic degradation. To overcome these limitations, synthetic polymers also have been used to fabricate hydrogels for drug and gene delivery. Among the various types of polymeric hydrogels, hydrogels synthesized from PEG, an FDA-approved polymer, have been extensively studied with promising preclinical and clinical results. Biomimetic peptides also have been widely studied in hydrogen applications, where they serve either as biological stimuli or as cross-linking sites. For example, CMP, synthetic peptides that mimic the triple-helical conformation of native collagens, have been conjugated to four-arm PEG. When equipped with a CMP domain, the four-arm PEGs form a hydrogel with these peptides serving as physical cross-linkers. Moreover, at temperatures above the melting temperature (TM) of the CMP, the unfolding of the triple helices induces hydrogel disassembly, allowing a convenient mechanism to thermally trigger the release of encapsulated drugs. Due to their biocompatibility, high swelling in aqueous media, and responsiveness to pH, temperature, and other stimuli, hydrogels have been extensively investigated as drug-delivery systems for molecules ranging from nonsteroidal antiinflammatory drugs (NSAIDs) to proteins. Compared with other drug-delivery systems, hydrogels closely resemble living tissues, with high water content, well-designed mechanical properties, and minimal tendency to adsorb proteins from bodily fluids. Additionally, the pore size of hydrogels can be easily manipulated via changing the chemical composition and cross-linking ratio of the polymeric network, which can in turn be utilized as a method to control the loading and releasing of encapsulated drugs. Due to their high water content, drug release from hydrogels is typically relatively fast and occurs over a period of hours to days. For applications in which a slower release rate is desirable, a wide range of strategies have been employed, including physical entrapment and covalent conjugation to the hydrogel network. One of the most widely used methods to reduce the release rate is to introduce electrostatic interactions between ionic polymer networks and oppositely charged drugs. To improve the strength of the pairwise charge–charge interactions, multivalently charged polymers, such as phosphate-functionalized polymers, are often used to form the hydrogel. For example, functionalizing a PNIPAAm-based hydrogel with polyoxyethyl phosphate-containing comonomer drastically improved the encapsulation of cationic lysozyme into the hydrogel. In another example, adding positively charged N-(3-aminopropyl)methacrylamide or 4-vinylpyridine monomer into poly(hydroxyethyl methacrylate) networks increased the amount of encapsulated NSAIDs by more than one order of magnitude and extended the period of sustained release up to approximately 1 week. Chemical conjugation is widely used to load drugs into hydrogel matrices, and these methods can enable stimuli-responsive release depending on the nature of the covalent bonds linking the drug to the gel. The details on environmentally triggered release applications are discussed later in the article. Thermoresponsive hydrogels are commonly developed using the principles previously discussed employing a wide range of polymers that exhibit temperature-responsive phase transitions. The most common characteristic of these polymers is the presence of hydrophobic groups, such as methyl, ethyl, and/or propyl groups. Hydrogels with desired thermoresponsiveness can be injected into the body in a liquid state with encapsulated drug, followed by gelation in the body at physiological temperature to form a crosslinked hydrogel. Additionally, the transition temperature of the hydrogel can be easily tuned by making copolymers of hydrophobic (e.g., NIPAAm) and hydrophilic (e.g., acrylic acid) monomers and adjusting the ratio of the hydrophilic and hydrophobic segments of the polymer to tailor therapeutic release. Generally, a higher content of hydrophobic polymers in the hydrogels, a lower transition temperature is obtained. The pH-sensitive hydrogels are another commonly studied class of stimuli-responsive hydrogels. Hydrogels with pH-sensitivity are primarily used in delivering therapeutics to tumor cells due to the existence of an acidic pH within the tumor stroma. The pHsensitive hydrogels are usually constructed using polymers with acid-sensitive bonds that can be easily cleaved in acidic conditions. For example, hydrogels cross-linked via Schiff’s base reactions are generally stable at physiological pH yet can be degraded under mildly acidic conditions due to the cleavage of the imine bond. The pH responsiveness also can be obtained by using polymers with ionizable chemical groups whose chemical properties such as swelling ratio and water solubility are altered based upon the charge state of such groups. For instance, S. A. Hegazy and coworkers designed pH-responsive hydrogels that were copolymerized from PEG/acrylic acid. The diffusion coefficient of the encapsulated model drug ketoprofen was highly dependent upon both the pH and the ionic strength of the medium, as the degree of ionization of the acrylic acid increases with increases in pH resulting in greater

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numbers of fixed charges, chain repulsion, and thus hydrogel swelling, while increases in solution ionic strength result in increases in the number of counterions and thus a decrease in repulsive forces and swelling. In tissues and cells, enzymes are secreted to change the structure and properties of ECM, facilitating cell proliferation, differentiation, and tissue regeneration. The incorporation of enzyme substrates as a biologically responsive component in hydrogel networks is a useful strategy to mimic the proteolytic sensitivity of ECM and enable the triggered release of encapsulated drugs. For example, by introducing the matrix metalloproteinase (MMP)-responsive peptide thymosin b4 (Tb4), Langer and coworkers designed a PEG-based hydrogel that was conducive to human umbilical vein endothelial cell (HUVEC) adhesion, survival, migration, and organization. Incorporation of Tb4 significantly increased the secretion of MMP-2 and MMP-9 from encapsulated HUVECs, which subsequently triggered the degradation of the hydrogel itself and the release of Tb4. The enzymatically controlled release of the peptide induced vascular-like network formation within the PEG hydrogels. These results indicated that the Tb4encapsulating, enzymatically degradable hydrogels may be useful as scaffolds for in situ regeneration of ischemic tissues. In addition to temperature-, pH-, and enzyme-responsive hydrogels, smart hydrogels that respond to other stimuli also have shown promise in regenerative medicine. For example, electric current has been applied as an external trigger to induce cargo release. Polyelectrolyte hydrogels are used for this purpose due to their capacity to undergo swelling or shrinkage when an electric field is applied, allowing the delivery of encapsulated therapeutics such as edrophonium chloride and hydrocortisone in an “on–off” manner when the applied electric field is switched. Light also has been used as a trigger to induce responses within hydrogels. Light-sensitive hydrogels are advantageous in terms of their control mechanism since light stimuli can be applied instantly with precisely controlled location and intensity. For example, the addition of bis(4-di-methylamino)phenylmethyl leucocyanide into a PNIPAAm-based hydrogel produced a network that swelled in response to UV irradiation but shrank when the light was shut off. This property was conferred by the bis(4-di-methylamino)phenylmethyl leucocyanide, which is usually a neutral molecule but can dissociate into ion pairs under UV irradiation. Such a property may have potential for light-sensitive drug-delivery applications. In addition to the widely used environmental triggers discussed earlier, other types of stimuli responses also have been introduced into hydrogels for drug delivery purposes, including sensitivity to magnetic fields, pressure, specific ions, thrombin, and various antigens. Readers interested in details of these types of hydrogels are redirected to previous published reviews in the suggested reading.

Electrospun Fibrous Scaffolds Fibrous scaffolds produced by electrospinning have become increasingly popular in regenerative engineering due to their high surface area-to-volume ratio and porosity, which together simulate the structure of protein fibers within the ECM. Moreover, the versatility of electrospinning in terms of polymer content, fiber structure, and functionalization has made it possible to fabricate scaffolds with a broad range of mechanical properties, bioactivities, and mass transport properties. For instance, drug diffusion rates within these porous scaffolds can be tailored by alteration of polymer degradation kinetics, matrix crystallinity, porosity, geometry, and other chemical phenomena. The diversity and biomimicry of electrospun fibrous scaffolds have led to their use in numerous regenerative medicine applications including wound healing and nerve and spinal muscle regeneration. Several methods have been developed for incorporation of therapeutic molecules into electrospun scaffolds, including surface modulation, blending, and coaxial processing. In most cases, therapeutic delivery from electrospun scaffolds is either achieved by binding the therapeutic onto the fiber surface or by encapsulating it within the fiber. While localized delivery has been achieved using both methods, encapsulation usually offers greater control over therapeutic diffusion rates and overall release kinetics. For example, blending approaches require the therapeutic to be codissolved or dispersed with the polymer in the organic solvent that evaporates during the electrospinning process. This one-step fabrication technique has been successfully used to achieve sustained release of hydrophobic drugs such as rifampicin and paclitaxel from hydrophobic polyester polymers and hydrophilic drugs such as doxorubicin from hydrophilic polymers including gelatin, PEG, and PVA. One challenge is that insufficient solubility and/or heterogenous distribution of drugs within the polymer solution can limit drug and polymer choices and often produces materials that exhibit an inconsistent burst release of encapsulated therapeutic. Core–shell electrospinning techniques, such as coaxial and emulsion electrospinning, overcome solubility complications and have a proven capacity to enhance control over release while maintaining the bioavailability of unstable biologics. Unlike blended electrospun fibers, which are produced from a single phase, core–shell fibers are fabricated using dual-solvent approaches that minimize contact between the biologic compound and organic solvents. Furthermore, the core–shell structures generated have a polymeric shell that deters direct contact between biomolecules and the external environment. Core–shell fibers are commonly fabricated either through coaxial electrospinning, which employs an inner jet and outer jet to generate fibers with a drug core and a protective polymeric shell, or through emulsion electrospinning, in which two immiscible solutions (typically an apolar polymer solution and an aqueous therapeutic solution) are spun simultaneously to generate the core-sheath fiber morphology. Both approaches have been used to obtain the sustained release of active therapeutics. For instance, C. H. Wang and coworkers constructed PLGA/hydroxyapatite (HA) composite scaffolds for the delivery of BMP-2 plasmid DNA. These investigators demonstrated improved therapeutic loading efficiency and extended release periods when chitosan BMP-2 plasmid DNA complexes were encapsulated through emulsion electrospinning versus scaffold dipping following fiber fabrication. The same group also demonstrated BMP-2 protein delivery via PLGA/HAp composite fibrous scaffolds that were generated using emulsion electrospinning. These structures maintained native conformations of the BMP-2, and the BMP-2 release profile could be controlled via varying HAp content.

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Furthermore, the versatile nature of electrospinning has provided an ideal platform for tailored, multitherapeutic delivery. For instance, T. W. Wang and coworkers demonstrated the capacity to tailor the release of four angiogenic growth factors by incorporating the therapeutics either directly or preloaded into gelatin nanoparticles into electrospun collagen/hyaluronic acid nanofibers. These multicomponent structures produced accelerated wound closure and elevated collagen deposition and enhanced bloodvessel maturation when applied in a diabetic rat wound model. Electrospun scaffolds also have been developed using stimuli-responsive polymers and biologically inspired strategies. For instance, the thermoresponsive polymer PNIPAAm was incorporated into an electrospun fiber mat and shown to retain its LCST behavior after electrospinning. Fibers fabricated from PNIPAAm copolymers or alternatively cospun with PNIPAAm and polymers such as polystyrene and/or PEO have shown reversible deswelling/swelling behaviors that are conducive to thermally triggered drug delivery. Additionally, H. Q. Liu and coworkers described the preparation of pH-responsive, cross-linked poly(styrene-co-(maleic sodium anhydride)) (SMA) and SMA-cellulose acetate composite nanofibers that exhibited improved mechanical strength relative to classically cast hydrogels and also exhibited pH-dependent swelling in aqueous solution. Alternatively, J. Weng and coworkers demonstrated pH-triggered release of the drug paracetamol from electrospun scaffolds though the incorporation of acid-labile acetal groups into the backbone of poly(D,L-lactide)-poly(ethylene glycol). Fibrous scaffolds responsive to additional types of external stimuli, including light, electric fields, and magnetic fields, also have been demonstrated. For instance, S. Ramakrishna and coworkers prepared skin grafts in which the photosensitive polymer poly(3-hexylthiophene) (P3HT) and epidermal growth factor were encapsulated into core–shell-structured gelatin/poly(L-lactic acid)-co-poly(ε-caprolactone) nanofibers (Gel/PLLCL/ P3GF(cs)) by coaxial spinning. Light stimulation in these scaffolds enhanced healing within in vitro wound-healing models. Furthermore, E. T. Kang and coworkers achieved externally controlled, UV-triggered release of the prodrug a-cyclodextrin-5fluorouracil through the incorporation of photosensitive azo groups into electrospun poly(vinylbenzyl chloride-glycidyl methacrylate) nanofibers. UV light application triggered the release of bound drug via UV-induced isomerization of the azo groups from trans to cis confirmation. Similarly, biologically inspired strategies have been employed to achieve controlled therapeutic delivery. For instance, one of the most common bioinspired approaches is to coat electrospun scaffolds with ECM components like collagen to promote enhanced cellular invasion and growth-factor binding and release. Alternatively, biomacromolecules with growth-factor affinities have been directly incorporated or covalently tethered to the scaffold surface to prolong release. B. S. Kim and coworkers improved BMP-2 retention on PLG scaffolds through covalently modifying its surface with heparin. The functionalized scaffold promoted enhanced bone formation relative to scaffolds directly loaded with BMP-2 through sustained BMP-2 release. In other cases, growth factor has been covalently bound to electrospun scaffolds, mimicking native ECM GF sequestering and release. Furthermore, proteasedegradable electrospun scaffolds have been developed through the incorporation of protease-labile sequences. These proteolytically sensitive scaffolds offer a promising platform for cell-triggered delivery in a variety of regenerative engineering applications.

Micro and Nanocarriers A large variety of micro and nanocarriers have been utilized in the delivery of therapeutic small-molecule drugs, proteins, and DNA. Particles have been prepared from a multitude of organic and inorganic materials including nondegradable and biodegradable polymers, lipids, self-assembling amphiphilic molecules, dendrimers, and metals. Material selection is largely dependent on the type and administration route of the drug, as well as the therapeutic objective. For instance, liposomes are lipid-based carriers that consist of an outer lipid bilayer and an inner aqueous space that is ideal for the delivery of soluble factors. Alternatively, micelles are composed of a self-assembling lipid monolayer with a hydrophobic core that is ideal for the delivery of hydrophobic drugs. Additionally, a number of carriers have been specifically engineered for nucleic acid delivery through electrostatic interaction with the negatively charged nucleic acid phosphate backbone. These carriers are typically composed of cationic polymers or cationic lipids that form nucleic acid complexes known as polyplexes and lipoplexes, respectively. Alternatively, lipopolyplexes, composed of polymer–nucleic acid cores and a lipid shell, also are employed, benefiting from the unique properties of both materials. Nucleic acid carriers must protect DNA from enzymatic degradation, promote cellular uptake, and stimulate intracellular unpackaging. Parameters such as polymer hydrophobicity/hydrophilicity, charge density, biodegradability, and the molecular weight can be adjusted to optimize the carrier. In most applications, micro and nanocarriers are delivered via bolus injection, and controlled delivery is achieved by either passive or active targeting. Passive targeting is commonly reliant on carrier size. For instance, microparticles are unable to cross the majority of biological barriers, and therefore, microparticles are most effective when applied at the delivery site where they avoid clearance and can remain present for weeks. In contrast, nanoparticles have the capacity to pass through biological barriers such as the cellular membrane but are typically cleared from the body within days and commonly accumulate in the liver, spleen, and/or kidney. Prolonged circulation and evasion of the mononuclear phagocyte system have been achieved through the incorporation of hydrophilic polymers or glycolipids with flexible chains, such as PEG or GM1, which occupy the immediate area adjacent to the carrier and sterically block interactions with serum proteins and cell-surface receptors. Multiple active targeting strategies also have been employed in regenerative engineering to target specific tissue components and/ or cell types. Active targeting is typically achieved through biologically inspired strategies including the display of cell-specific ligands, antibodies, and/or ECM-binding peptides. For instance, multiple peptides with the capacity to bind exclusively to intact or remodeled collagen, a primary ECM component whose state is indicative of many regenerative processes, have been identified. Sequences have been derived from sources such as collagen-I platelet receptors, collagenase, and decorin or found through phage

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display. Alternatively, CMPs have been engineered to hybridize with remodeled collagen through integration into the natural collagen triple helix. CMP tethers have been utilized to create self-assembling nanocarriers with potential collagen-binding affinities, and these tethers also can retain cytoactive factors and fluorophores within areas of excessive collagen remodeling including joints and wounds. The control over cellular uptake also has been achieved through the inclusion of receptor-specific ligands and cationic cell penetrating peptides. In the case of DNA delivery, where nuclear delivery is essential, additional components such as nuclear localization signal sequences, nuclear proteins such as histones, and histone-mimetic peptides have been incorporated into carriers. Once retained in the delivery site, micro and nanocarriers may provide an additional mechanism to modulate release via diffusion or stimuli-responsive methods. Moreover, scaffold-mediated delivery systems composed of micro or nanocarriers encapsulated increased capacity to regulate release and overcome pharmacological limitations while preserving the bioactivity or stability of the therapeutic. For instance, therapeutic often is loaded into microgels whose diffusion-based delivery may be modulated by cross-link density. Growth factor-loaded gelatin microcarriers have been utilized to modulate the delivery of bioactive GF including insulinlike growth factor (IGF), BMP-2, and transforming growth factor-beta 1 (TGF-beta 1) from PEG-based hydrogels. These structures have been shown in several examples to support cartilage repair. “Smart” or stimuli-responsive materials, coupled with microenvironmental differences in pH and/or temperature that are typical in diseased or healing tissues, can be utilized to trigger therapeutic delivery via increased drug diffusivity or carrier destabilization. Therapeutic molecules have been loaded into biocompatible micro and nanocarriers constructed from stimuli-responsive materials as previously discussed. For instance, T. G. Park and coworkers prepared thermoresponsive, super-expandable pluronic/PEI polyplex nanocarriers with the capacity to efficiently deliver small interfering RNA (siRNA) into the cytosol and subsequently silence the targeted mRNA. To trigger release following cellular uptake, the temperature was lowered to 20 C, and the decreased temperature caused the system to undergo a phase transition with dramatic swelling that led to an 800-fold size increase and a subsequent increase in therapeutic diffusivity. In most cases, thermoreponsive micro and nanocarriers are liposomes, polymeric micelles, or nanoparticles typically composed of components that exhibit LCST behavior such as PNIPAAm; however, metal-based carriers may be secondarily incorporated to provide a mechanism to raise the system temperature in response to light or magnetic and electric fields. Additionally, pH-sensitive nanocarrier delivery systems often have advantageous features such as enhanced cellular uptake, endosomal/lysosomal escape induced by the proton sponge effect (osmotic swelling), and the capacity for surface charge reversion. For instance, the natural pH-responsiveness and biocompatibility of chitosan and chitosan-based nanocarriers make them ideal candidates for controlled release in naturally acidic environments such as wounds or intracellular environments. H. Q. Zhao and coworkers demonstrated that the grafting of chitosan onto PEI polyplexes preserved the buffering capacity of the PEI while achieving transfection efficiencies comparable with the standard transfection agent Lipofectamine and improvements of over 40% in cell viabilities in chondrocyte and synoviocyte cell lines. Other stimuli responses also have been demonstrated. M. O. Sullivan, T. H. Epps, III and coworkers have fabricated siRNA mPEG-b-poly(5-(3-(amino)propoxy)-2-nitrobenzyl methacrylate) (mPEG-b-P(APNBMA)) polyplexes with the capacity to silence targeted mRNA sequences in response to UV-stimulated siRNA release. Alternatively, M. F. Mieler and coworkers reported lightresponsive release of a model drug from gold’silver nanorods coated with DNA-cross-linked polymeric shells. Within these systems, the nanorods acted as photothermal convertors absorbing light near infrared that was then converted to heat and thus elevated temperatures and membrane permeability. In other examples, cell-triggered release of polyplex has been achieved through integration into protease-labile scaffolds or the use of proteolytic linkages. For instance, M. O Sullivan, K. L. Kiick, and coworkers fabricated DNA–PEI polyplex collagens in which release/retention of polyplex was tailored through variation of CMP display on the polyplex. The transfection levels in the murine NIH/3T3 fibroblast cell line were approximately an order of magnitude greater when protease expression was stimulated via TNF-alpha treatments.

Types of Therapeutic Molecules in Drug and Gene Delivery Within the field of regenerative engineering, therapeutics can be categorized as small molecular drugs, proteins, or nucleic acids (Fig. 2). The primary objective is to augment or enable the complex reparative process. Each therapeutic type has its own distinct benefits and delivery obstacles.

Small-Molecule Drugs The majority of small-molecule drugs have been identified based on their capacity to modulate key signaling pathways involved in tissue repair. For instance, pyrvinium is an FDA-approved drug used to treat infection by inhibiting the Wnt pathway, which drives the expression of several inflammatory molecules that are upregulated during bacterial infections. P. P. Young and coworkers demonstrated that daily administration of pyrvinium into subcutaneous, PVA sponges in mice generated better-organized and better-vascularized granulation tissue and an increased tissue proliferative index. The same group also evaluated the therapeutic value of pyrvinium in murine myocardial repair and demonstrated that the administration of a single dose via coronary artery ligation reduced adverse cardiac remodeling based on the decrease in the left ventricular internal diameter in diastole relative to the control. Alternatively, deferoxamine is an FDA-approved iron chelator that has been in clinical use for decades. Specifically, it is

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A summary of the therapeutic types in regenerative engineering and potential obstacles to efficient delivery.

known to aid in wound healing by inhibiting HIF-1alpha degradation, and in so doing, it decreases oxidative stress in the wound bed and stimulates subsequent decreases in necrotic tissue and enhanced wound healing in murine wounds. In addition to antiinflammatory drugs such as pyrvinium and NSAIDs and chelators for reducing oxidative stress, another class of drugs commonly used in regenerative engineering is statins, which are broadly characterized as lipid-lowering agents. The application of statins demonstrated promising results in multiple regenerative engineering applications including wound healing, cardiovascular repair, and bone growth. For instance, the anabolic effect of both simvastatin and lovastatin increased the expression of BMP-2 mRNA. In general, small-molecule drugs have lower manufacturing costs and higher stability relative to biologics, which enables longer shelf life, increased half-lives within the body, and more flexibility in regards to chemical modification/conjugation. Furthermore, the treatments are less complicated and already well understood by the FDA and the pharmaceutical industry. However, smallmolecule drugs are typically nonspecific and have been documented in many cases to have adverse systemic effects and rapid clearance.

Protein-Based Therapeutics The application of protein-based therapeutics has been enabled by the advent of cost-effective recombinant DNA technology, which is used to clone, express, and purify proteins of interest such as growth factors and antibodies. Growth factors play vital roles in the pathways underlying healing within all tissues, and the delivery of growth factors can significantly augment the reparative process. On the other hand, numerous monoclonal antibodies with the capacity to target specific proteins and cells have been identified. This targeting capacity has been utilized for controlled therapeutic delivery and pathway inhibition.

Growth factors Growth factors are signaling proteins capable of triggering specific cellular behaviors including migration, proliferation, and differentiation, through interaction with transmembrane receptors. Most preclinical trials to date have been conducted with angiopoietins, which promote blood-vessel maturation and stability; BMP-2, which stimulates osteogenic differentiation and cell migration; fibroblast growth factor (FGF), which encourages migration, proliferation, and survival of endothelial cells; PDGF, which regulates endothelial cell proliferation and migration; and VEGF, which promotes angiogenesis. Due to the innate role of growth factors in the healing cascade, growth-factor administration has been demonstrated to reestablish endogenous healing responses through coordination of multiple regenerative cascades. Despite their promise, growth factors typically have incredibly short half-lives within the body, generally on the order of hours (i.e., for FGF-2 7.6 h, for VEGF165  1.5 h, and for PDGF-BB < 4 h). The instability of these protein-based therapeutics is exacerbated in injured areas where protease activity is elevated, and in the case of chronic wounds, cells in the wound bed often exhibit alterations in behavior leading to limited growth-factor responses. To overcome growth-factor instability in regenerative medicine applications, extraphysiological and repetitive dosage regimens often are applied; however, these types of dosing approaches increase the danger of growth-factor toxicity, off-target responses, and oncogenicity. Protein engineering can be used to enhance growth-factor stability, but preserving engineered growth-factor bioactivity is not trivial. Moreover, the delivery of multiple growth factors at different times and concentrations is required to fully recapitulate the regenerative process and account for the synergistic nature of growth factors.

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Monoclonal antibodies Therapeutic monoclonal antibodies have the capacity to specifically bind to cells or proteins. Accordingly, this targeting capacity has been used in a variety of contexts in drug delivery. In regenerative engineering applications, a common application of monoclonal antibodies is to bind to growth factors involved in pathological pathways and thereby inhibit their activity. For instance, the application of the FDA-approved anti-VEGF agents bevacizumab, ranibizumab, pegaptanib, and aflibercept have revolutionized the treatment of neovascularization retinal disorders by inhibiting angiogenesis. The application of anti-TGF-beta 1, 2, and 3 in wound healing and prevention of hypertrophic scar formation has been widely explored, and findings in a rabbit ear impaired healing model strongly suggested that TGF-beta 1, 2, and 3 concentrations are important during distinct periods during both early and late stages of repair. Additionally, the balance of TGF-beta 1, 2, and 3 plays a distinct role in hypertrophic scar formation. Monoclonal antibodies generally have longer half-lives than growth factors, with typical half-lives on the order of days as opposed to hours; however, the large size of the antibodies typically causes poor perfusion and complications such as cytokine release syndrome and immunogenicity. Many of these obstacles can be overcome though delivery using smart delivery systems, as previously discussed.

Nucleic Acids Nucleic acid-based strategies in regenerative engineering can generally be categorized as gene or RNA interference (RNAi) therapies. Both types of therapies can be used to reprogram cells in vivo or ex vivo, and nucleic acid delivery approaches can also be used to increase or reduce expression of proteins controlling wound-healing cascades. Nucleic acid-based strategies can be especially beneficial in regenerative engineering because they can induce cell-mediated expression/delivery of signaling molecules such as growth factors, and the cellular production/secretion processes mimic those involved during endogenous delivery of these molecules.

Genes Gene delivery can be used to overcome some of the shortcomings of protein-mediated delivery. The successful delivery of therapeutic plasmid DNA or viral vectors into target cells facilitates the expression of fresh, bioactive proteins with appropriate posttranslational modifications, and these delivery methods also microlocalize delivery. Accordingly, the delivery of growth-factor genes, as opposed to the direct delivery of growth-factor proteins, has been shown to have significant therapeutic benefits within the wound environment. Viral vectors are the most common tool for delivering genetic material into a cell, owing to the innate ability of viruses to transfect cells. Numerous viruses have been used as vectors, and viral vectors can be purposefully chosen to facilitate permanent or temporal transfection and to a certain degree enable targeting of a specific cell host. To date, most gene-therapy clinical trials utilize retroviruses to achieve long-lasting transgene expression and adenoviruses to obtain transient transfection. While viral vectors obtain high-efficiency transfection, off-target delivery and immunogenic and oncogenic risk factors have greatly limited their clinical success. Plasmid DNAs are attractive tools in gene delivery due to their ease of modification and the ability of plasmids to selfreplicate in bacterial hosts. Moreover, tissue-specific promoters may be incorporated to achieve localized delivery; however, plasmid DNA requires a carrier to protect it from nucleases and enable efficient transfection.

RNAi Alternatively, nucleic acids with the capacity to silence protein expression, such as siRNA or microRNA, are often explored in regenerative engineering. These approaches exploit the ability of noncoding small RNAs to knock down gene expression through binding to specific mRNAs and tagging them for nuclease destruction (siRNA) or to reduce gene expression through physically blocking translation into proteins (microRNA). Like plasmid DNA, small RNAs require a delivery vehicle to protect them from degradation and induce cellular uptake. Unlike plasmid DNA, the active compartment for small RNAs is the cytosol; therefore, nuclear transport is not required. In multiple regenerative engineering applications, excessive proteolytic activity makes MMPs an attractive target for small-RNA delivery. For example, H. S. Yoo and coworkers recently demonstrated that the MMP-responsive delivery of MMP-2 siRNA in diabetic ulcer models has the capacity to reduce wound closure time, decrease MMP-2 expression, and increase cytokeratin levels as compared with nontreated controls. Similar to anti-VEGF monoclonal antibodies, VEGF and its receptor are also intuitive targets for knockdown, and small-RNA approaches also have used to reduce excessive angiogenesis in ocular disorders to prevent macular degeneration. Additionally, microRNAs have been demonstrated to play a prominent role in the survival of cardiac progenitor cells, and thus, their delivery shows potential in cardiac regeneration. Major obstacles for small-RNA delivery include its high cost, potential off-target effects, inefficient uptake into cells, and lack of effective release within the delivery site.

Delivery System Selection/Comparison Within this review, several types of delivery systems used in regenerative engineering have been discussed. To determine the correct type of delivery system for a given application, several factors must be simultaneously considered. These include the physicochemical properties of the therapeutic, the pathology of the delivery site, and the potential systemic versus localized effects of the delivery system itself. Typically, a therapeutic for use in regenerative engineering is chosen based on its impact on healing cascades. For instance, NSAIDs such as ibuprofen often are administered due to their well-recognized ability to reduce pain and inflammation

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through blocking cyclooxygenase (COX) enzymes, which trigger the formation of prostaglandin-inflammatory signaling molecules. In many applications, growth factors or the genes that encode them are delivered due to their roles in cellular differentiation, proliferation, migration, and other behaviors essential for tissue repair. Alternatively, siRNA can be delivered to silence or reduce the expression of proteins inhibiting regeneration such as overexpressed proteases in chronic wounds. Once a therapeutic is selected, a delivery system must be chosen based on drug properties such as molecular weight, aqueous solubility, charge, serum stability, and the characteristics of the targeted delivery site. For instance, hydrophobic drugs may need to be loaded into liposomes, micelles, or the hydrophobic cores of electrospun fibers to enhance their solubility within aqueous biological systems. Similarly, the half-life of the therapeutic should be considered, and precautions should be taken, as appropriate, to preserve bioactivity. Precautions may include direct alteration of the therapeutic to enhance its stability, immobilization to a substrate to reduce aggregation, or loading into polymeric- or lipid-based particles (or the cores of electrospun scaffolds) to reduce drug contact with biological fluids. On the other hand, charged drugs may be adsorbed or immobilized onto polymeric or ceramic scaffolds to promote localized delivery, and in the specific case of nucleic acid delivery, charge can be used to form a complex that preserves integrity and promotes cellular uptake. Considerations of the delivery site and disease pathology are also of great value in therapeutic delivery. Extracellular delivery is commonly achieved utilizing bulk or microcarrier delivery systems, whereas intracellular delivery is commonly reliant on nanocarriers or viral vectors. However, in tissue engineering, the application of implantable or injectable 3-D scaffolds may provide structure for cellular adhesion and ECM deposition and a platform for controlled bioactive cue delivery. On the other hand, systemic applications of targeted micro or nanocarriers can provide a less invasive approach for reaching less accessible tissues or cell types. Additionally, specific pathologies often are associated with characteristic microenvironmental changes that can be used to modulate release. For instance, acute wounds progress from an alkaline state to an acidic state when healing begins; however, chronic wounds beds continue to have an elevated pH over a prolonged duration. The pH of stage 1 pressure ulcers has a pH similar to that of intact skin (pH 5.4–5.6), while the pH values of stage 2 and 3 ulcers have been reported as 6.9 and 7.6, respectively. The microenvironmental differences in pH can be used as a logical trigger for therapeutic release. Moreover, the enhanced proteolytic activity, ECM degradation, and ROS production during the early stages of acute healing and within chronic wounds can be harnessed to trigger release in responsive systems. In applications in which “on–off” delivery is of value, the accessibility of the delivery site to external stimuli should also be considered. For instance, the application of UV light, electric current, and heat/cold induction are sensible means for modulating responsive delivery for topical applications; however, for certain applications, exogenous stimuli with enhanced penetration depth are required, such as near infrared radiation (NIR), radiofrequency electromagnetic fields (EMFs), and ultrasound. For instance, high-power NIR was reported to penetrate at least 3 cm into the brain when wavelengths of 810 and 980 nm with 10–15 W power were applied, which demonstrates the potential to pass through skin and bone and remotely trigger therapeutic release in traumatic brain injury. Furthermore, at 400 kHz, 99% of EMF radiation penetrates into 15 cm of tissue, and background heating of the tissue is insignificant. The localized and systemic impacts of the delivery system itself must also be taken into account. For instance, biocompatibility is vital for avoiding immune responses. Moreover, delivery systems may be engineered to have direct therapeutic value on their own. In many applications, delivery systems are purposefully designed to mimic the ECM and subsequently perform its native roles in facilitating cellular attachment, proliferation, and phenotypic commitment. Biomimicry is typically achieved through incorporation of native ECM components, such as collagen, or addition of mimetic peptides that contain either ECM-derived cell binding sites or ECM structures such as the collagen triple helix. For instance, collagen-based materials are commonly employed as both gels and sponges in regenerative medicine, where they can enhance healing by serving as both sacrificial substrates for proteolytic activity and scaffolds for cellular adhesion. Other studies have highlighted important physical characteristics that determine the ability of a delivery system to trigger cellular responses. For instance, the diameter and orientation of electrospun fibers within tubular nerve guides and wound dressings were determined to impact nerve regrowth and wound closure, respectively.

Therapeutic Delivery in Chronic Wound Care Delivery System Impaired healing is the result of alterations in normal physiological processes typically caused by aging or diabetes. Acute wound repair proceeds through four distinct but overlapping phasesdhemostasis, inflammation, proliferation, and remodeling; however, chronic wounds become stalled in prolonged, exaggerated inflammatory phases. As a result, the wound bed is characterized by excessive oxidative stress, proteolytic activity, and bacterial infections. Given the localized nature of this application coupled with the inherent lack of blood flow within chronic wounds, bulk, localized therapeutic strategies are highly beneficial. The application of a 3-D scaffold such as a hydrogel or electrospun fiber mat has been used to achieve spatiotemporal control over delivery while also providing additional therapeutic benefit by serving as an analog for the ECM though facilitating cellular adhesion and phenotypic commitment. For instance, the ECM component collagen is a primary component in many wound dressing approved by the FDA for diabetic foot ulcer care, including Allograft, Dermagraft, and Promogran. These collagen-based products have been demonstrated to increase fibroblast proliferation and decrease protease activity. For instance, wound closure was documented in 91% of DFU patients treated with Dermagraft within 12 weeks, whereas only 78% of patients exhibited the same outcome in a control group treated with the conventional wet-to-dry dressing approach.

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Small-Molecule Drug Delivery The primary goal of any therapeutic approach in chronic wound healing is to replicate or enhance normal healing processes. A plethora of small-molecule drugs and biologics have shown potential in restoring coordination between the intracellular, intercellular, and extracellular pathways vital for repair. For instance, as previously discussed, FDA-approved small-molecule drugs such as pyrvinium are used to treat wound infection through inhibiting the expression of several inflammatory molecules, and deferoxamine (DFO) is used to reduce oxidative stress in the wound bed through inhibiting hypoxia-inducible factor 1alpha degradation. However, systemic delivery of these agents is not a viable option for treating chronic wounds due to the toxicity of these agents and their short plasma half-life exhibited in in vitro and in vivo preclinical studies. Accounting for the delivery site and the properties of the drug, G. C. Gurtner and coworkers developed a local transdermal drug-delivery system of potential clinical value as pictured in Fig. 3. To enable the relatively large and hydrophobic DFO to penetrate the stratum corneum, the lipophilic outermost layer of skin, DFO/polyvinylpyrrolidone (PVP) complex-loaded reverse micelles were fabricated and encapsulated within a slow-releasing ethyl cellulose matrix. Upon topical application of the DFO patch to the skin, the reverse micelles were released from the degradable polymeric matrix, and they subsequently penetrated through the stratum corneum into the hydrophilic, aqueous environment of the dermis; this process caused the reverse micelle to disintegrate, enabled the DFO/PVP complex to disassociate, and ultimately freed DFO for transdermal delivery. This system was successfully used in diabetic murine models to prevent ulcer formation and promote diabetic wound healing through reduction of hyperglycemiainduced oxidative stress. As our understanding of the fundamental and intricate pathways underlying repair improves, additional drugs of value in chronic wound repair continue to be uncovered.

Fig. 3 (A) Development of a transdermal drug delivery system for DFO. DFO aggregates with PVP and surfactants to form reverse micelles (RMs). RMs are dispersed in the polymer ethyl cellulose. After release from the polymer matrix the RMs enter the stratum corneum and disintegrate. PVP dissolves and DFO is delivered to the dermis. DFO TDDS improves healing of diabetic ulcers. (B) Full-thickness ulcer wounds of diabetic mice treated with a transdermal DFO TDDS formulation or vehicle control (n ¼ 10). TDDS were replaced every 48 h. Duscher, D., Neofytou, E., Wong, V. W., Maan, Z. N., Rennert, R. C., Inayathullah, M., Januszyk, M., Rodrigues, M., Malkovskiy, A. V., Whitmore, A. J., Walmsley, G. G., Galvez, M. G., Whittam, A. J., Brownlee, M., Rajadas, J. and Gurtner, G. C. (2015). Transdermal deferoxamine prevents pressure-induced diabetic ulcers. Proceedings of the National Academy of Sciences of the United States of America 112, 94–99.

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Protein-Based Therapeutic Delivery Protein-based therapeutics such as growth factors and monoclonal antibodies have been extensively studied in chronic wound clinical trials. The delivery of various growth factors from polymeric hydrogels and electrospun scaffolds has been demonstrated to promote faster wound closure, angiogenesis, and collagen deposition within various impaired healing animal models. The objective of the therapy is to replace the degraded growth factors in the hostile wound bed in order to recapitulate normal repair; therefore, delivery systems must promote sustained growth-factor delivery and bioactivity by overcoming the inherently short half-lives of most of these factors in the protease-rich wound environment. In some cases, delivery systems are engineered to codeliver multiple growth factors, typically through encapsulating one therapeutic directly into a polymeric scaffold and preloading the second therapeutic into micro or nanocarriers that can act as a secondary barriers to regulate diffusion or facilitate stimuli-triggered release. Dual growth-factor delivery often enhances healing in animal wound models as compared with delivery of only one growth factor; however, in clinical delivery, synergistic interactions between different growth factors and the ECM often are not accounted for. As shown in Fig. 4, J. A. Hubbell and coworkers presented a multifaceted approach for engineering the cellular microenvironment to greatly enhance VEGF-A and PDGF-BB levels in wound healing or alternatively BMP-2 and PDGF-BB levels in bone repair. Specifically, a multifunctional recombinant fibronectin fragment engineered to contain (i) a factor XIIIa substrate fibrin-binding sequence; (ii) the 9th–10th type III FN repeat (FN III9–10) containing a major integrin-binding domain; and (iii) the 12th– 14th type III FN repeat (FN III12–14), which binds growth factors promiscuously, including VEGF-A165, PDGF-BB, and BMP-2 was covalently cross-linked into a fibrin matrix. The modified fibrin was loaded with VEGF-A and PDGF-BB, and this loaded scaffold was shown to significantly enhance angiogenesis leading to healing in a diabetic murine wound model as compared with fibrin alone, modified fibrin without growth factors, or growth factors that were individually applied without fibrin. These findings highlight the potent synergistic signaling between certain integrins and growth-factor receptors and show the importance of harnessing these biomimetic interactions when designing biomaterials and delivery systems. Dual regulation through integrin signaling and growth-factor signaling also has been shown to present a viable strategy to attenuate growth-factor potency in other regenerative engineering applications such as osteodifferentiation. Alternatively, therapeutic monoclonal antibodies are utilized to neutralize proteins that are elevated in chronic wound repair versus acute wound repair, including TGF-beta, flightless, or interleukin-6. For instance, N. H. Voelcker and coworkers demonstrated significantly enhanced wound healing when flightless I neutralizing antibodies were delivered from silicon nanoparticles, as compared with a no-treatment control or wounds in which the antibody alone was administered.

Gene-Based Therapeutic Delivery The capacity of nucleic acid delivery to reprogram cells is of particular value in chronic wound repair. Dr. David Margolis reported the results of the first clinical trial in humans for gene therapy in wound healing in 2000. Chronic wound gene therapies have become increasingly clinically relevant over the past few decades due to their potential ability to enhance or inhibit protein expression and restore wound-bed homeostasis. For instance, bioactive scaffold-mediated delivery of genes encoding for PDGF, VEGF, FGF, and/or EGF has been used to upregulate protein expression and facilitate enhanced angiogenesis, epithelization, and overall faster wound closure times within impaired animal wound models; however, vector escape and immune responses have greatly inhibited the clinical translation of these therapies. RNA interference therapies have encountered similar obstacles. Biologically inspired approaches that not only account for the wound environment but also harness pathways known to occur in excess in chronic versus acute wounds have great potential in restoring wound-bed homeostasis. For instance, MMPs are expressed in excess in chronic wounds. Under normal wound-healing conditions, these proteases facilitate ECM remodeling and promote wound repair, yet in chronic wounds, the overexpression of these factors contributes to the pathology of the wound bed. M. O. Sullivan, K. L. Kiick, and coworkers harnessed ECM remodeling to achieve efficient gene delivery by stably integrating DNA polyplexes into collagens using adjustable CMP tethers, such that gene release and expression were dependent upon MMP protease expression and subsequent collagen turnover. Furthermore, mechanistic studies from this group suggested that CMP–collagen hybridization persisted after scaffold release, such that natural endocytic clearance mechanisms for collagen facilitated higher levels of CMP-polyplex uptake and expression in cells. Alternatively, H. S. Yoo and coworkers developed an MMP-inspired treatment (Fig. 5) in which antiMMP siRNA polyplexes were displayed on fibrous electrospun scaffolds using MMP-labile peptide linkers. These structures were shown to silence MMP expression and encourage wound recovery in a diabetic murine model. The continued development of controlled, biologically inspired nucleic acid therapies has great promise in chronic wound repair.

Current Obstacles and Future Applications Within the field of regenerative engineering, an ever-growing understanding of the processes underlying repair is rapidly expanding the library of possible therapeutics; however, delivery remains a major concern. While beneficial under the right conditions, the therapeutics required to promote regeneration commonly cause serious adverse side effects ranging from blindness to cancer when delivered to off-target locations. To overcome these obstacles, better control over therapeutic delivery is a necessity. The continued development of “smart” polymers and inorganic particles has great potential in overcoming many of these dangers. Moreover, there is a great need for new delivery strategies that better replicate the complex, multicomponent presentation of healing factors during the normal course of tissue healing. Strategies for mediating multitherapeutic delivery and better recapitulating

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Fig. 4 (A) A multifunctional recombinant FN fragment is engineered to display the integrin-binding domain (FN III9–10) linked to the GF-binding domain (FN III12–14) and to comprise the substrate sequence a2PI1–8 for factor XIIIa. The fragment is covalently cross-linked into a fibrin matrix during the natural polymerization process of fibrin via the transglutaminase activity of factor XIIIa. Delivering VEGF-A165 and PDGF-BB within functionalized fibrin matrices enhances skin wound healing in diabetic mice compared to treatment with fibrin only, fibrin functionalized with FN III9–10/12–14 only, fibrin containing GFs only, and fibrin functionalized with FN III9–10/12–14 containing GFs. (B) After 7, 10, and 15 days, wound closure and granulation tissue area were evaluated by histology. (C) Wound histology (hematoxylin and eosin staining) at 10 days. Black arrows indicate wound edges; red arrows indicate tips of epithelium tongue. The granulation tissue (pink-violet) is characterized by a large number of granulocytes with nuclei that stain in dark-violet or black. Muscle under the wounds is stained in red. Fat tissue appears as transparent bubbles. Scale bar, 1 mm. Higher magnification (5) of the granulation tissue is shown on the right. Martino, M. M., Tortelli, F., Mochizuki, M., Traub, S., Ben-David, D., Kuhn, G., Muller, R., Livne, E., Eming, S. and Hubbell, J. A. (2011). Engineering the growth factor microenvironment with fibronectin domains to promote wound and bone tissue healing. Science Translational Medicine 3, 100RA89.

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Fig. 5 (A) Schematic showing the preparation of the MMP-responsive nanofibrous mesh and the proposed mechanism for delivering siRNA into diabetic ulcers with the modified nanofibers. (B) Wound recovery of diabetic ulcers in C57BL/6 mice where wound recovery rate was calculated by comparing the closed wound area. (C) In vivo expression levels of MMP-2, keratin 5, keratin 14 and GAPDH in re-epithelialized tissues at day 3 and day 7. (B) Normalized expression levels of MMP-2, keratin 5 and keratin 14 with respect to GAPDH expression based on the electrophoresis results in a. * Indicates a statistical significance (P > .05). Kim, H. S. and Yoo, H. S. (2013). Matrix metalloproteinase-inspired suicidal treatments of diabetic ulcers with siRNA-decorated nanofibrous meshes. Gene Therapy 20, 378–385.

biologically inspired cellular interactions with the delivery system have the potential to enhance therapeutic potency and in turn reduce therapeutic dosage requirements. Pursuit of better, holistic delivery strategies may overcome the clinical barriers that have largely inhibited the commercial use of growth factors and prevented the approval of gene-based therapies.

Further Reading Andreadis, S. T., & Geer, D. J. (2006). Biomimetic approaches to protein and gene delivery for tissue regeneration. Trends in Biotechnology, 24, 331–337. Briquez, P. S., Hubbell, J. A., & Martino, M. M. (2015). Extracellular matrix-inspired growth factor delivery systems for skin wound healing. Advances in Wound Care, 4, 479–489. Cao, S. G., Hu, B. H., & Liu, H. Q. (2009). Synthesis of pH-responsive crosslinked poly styrene-co-(maleic sodium anhydride) and cellulose composite hydrogel nanofibers by electrospinning. Polymer International, 58, 545–551. Cui, W. G., Qi, M. B., Li, X. H., Huang, S. Z., Zhou, S. B., & Weng, J. (2008). Electrospun fibers of acid-labile biodegradable polymers with acetal groups as potential drug carriers. International Journal of Pharmaceutics, 361, 47–55.

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Davis, H. E., & Leach, J. K. (2011). Designing bioactive delivery systems for tissue regeneration. Annals of Biomedical Engineering, 39, 1–13. Foster, A. A., Greco, C. T., Green, M. D., Epps, T. H., & Sullivan, M. O. (2015). Light-mediated activation of siRNA release in diblock copolymer assemblies for controlled gene silencing. Advanced Healthcare Materials, 4, 760–770. Fu, G. D., Xu, L. Q., Yao, F., Li, G. L., & Kang, E. T. (2009). Smart nanofibers with a photoresponsive surface for controlled release. ACS Applied Materials & Interfaces, 1, 2424–2427. Ganta, S., Devalapally, H., Shahiwala, A., & Amiji, M. (2008). A review of stimuli-responsive nanocarriers for drug and gene delivery. Journal of Controlled Release, 126, 187–204. Hastings, C. L., Roche, E. T., Ruiz-Hernandez, E., Schenke-Layland, K., Walsh, C. J., & Duffy, G. P. (2015). Drug and cell delivery for cardiac regeneration. Advanced Drug Delivery Reviews, 84, 85–106. Hoare, T. R., & Kohane, D. S. (2008). Hydrogels in drug delivery: Progress and challenges. Polymer, 49, 1993–2007. Jeon, O., Song, S. J., Kang, S. W., Putnam, A. J., & Kim, B. S. (2007). Enhancement of ectopic bone formation by bone morphogenetic protein-2 released from a heparinconjugated poly(L-lactic-co-glycolic acid) scaffold. Biomaterials, 28, 2763–2771. Jin, G. R., Prabhakaran, M. P., & Ramakrishna, S. (2014). Photosensitive and biomimetic core–shell nanofibrous scaffolds as wound dressing. Photochemistry and Photobiology, 90, 673–681. Kim, H. S., & Yoo, H. S. (2013). Matrix metalloproteinase-inspired suicidal treatments of diabetic ulcers with siRNA-decorated nanofibrous meshes. Gene Therapy, 20, 378–385. Lai, H. J., Kuan, C. H., Wu, H. C., Tsai, J. C., Chen, T. M., Hsieh, D. J., & Wang, T. W. (2014). Tailored design of electrospun composite nanofibers with staged release of multiple angiogenic growth factors for chronic wound healing. Acta Biomaterialia, 10, 4156–4166. Lee, K., Bae, K. H., Lee, Y., Lee, S. H., Ahn, C. H., & Park, T. G. (2010). Pluronic/polyethylenimine shell cross linked nanocapsules with embedded magnetite nanocrystals for magnetically triggered delivery of siRNA. Macromolecular Bioscience, 10, 239–245. Lu, H. D., Dai, Y. H., Lv, L. L., & Zhao, H. Q. (2014). Chitosan-graft-polyethylenimine/DNA nanoparticles as novel non-viral gene delivery vectors targeting osteoarthritis. PLoS One, 9, e84703. Martino, M. M., Tortelli, F., Mochizuki, M., Traub, S., Ben-David, D., Kuhn, G., Muller, R., Livne, E., Eming, S., & Hubbell, J. A. (2011). Engineering the growth factor microenvironment with fibronectin domains to promote wound and bone tissue healing. Science Translational Medicine, 3, 100RA89. Munsell, E. V., Ross, N. L., & Sullivan, M. O. (2016). Journey to the center of the cell: Current nanocarrier design strategies targeting biopharmaceuticals to the cytoplasm and nucleus. Current Pharmaceutical Design, 22, 1227–1244. Nie, H. M., & Wang, C. H. (2007). Fabrication and characterization of PLGA/HAp scaffolds for delivery of BMP-2 plasmid composite DNA. Journal of Controlled Release, 120, 111–121. Nie, H., Soh, B. W., Fu, Y. C., & Wang, C. H. (2008). Three-dimensional fibrous PLGA/HAp composite scaffold for BMP-2 delivery. Biotechnology and Bioengineering, 99, 223–234. Ordeig, O., Chin, S., Kim, S., Chitnis, P. V., & Sia, S. K. (2016). An implantable compound-releasing capsule triggered on demand by ultrasound. Scientific Reports, 6, 22803. Saravanakumar, G., Kim, J., & Kim, W. J. (2017). Reactive-oxygen-species-responsive drug delivery systems: Promises and challenges. Advanced Science, 4, 1600124. Schmaljohann, D. (2006). Thermo- and pH-responsive polymers in drug delivery. Advanced Drug Delivery Reviews, 58, 1655–1670. Turner, C. T., McInnes, S. J. P., Melville, E., Cowin, A. J., & Voelcker, N. H. (2017). Delivery of flightless I neutralizing antibody from porous silicon nanoparticles improves wound healing in diabetic mice. Advanced Healthcare Materials, 6. Urello, M. A., Kiick, K. L., & Sullivan, K. L. (2014). A CMP-based method for tunable, cell-mediated gene delivery from collagen scaffolds. Journal of Materials Chemistry B, 2, 8174–8185.

Ethics of Issues and Stem Cell Research: the Unresolved Issues Z Master, Albany Medical College, Albany, NY, USA; and University of Alberta, Edmonton, AB, Canada © 2019 Elsevier Inc. All rights reserved.

Introduction In the Beginning: The Discovery of hESCs Hype Surrounding Stem Cell Research Personhood and the Moral Status of the Human Embryo Harms to Women Alternative Technologies to the Derivation of Pluripotent Stem Cells Nondestructive Techniques Altered Nuclear Transfer Cytoplasmic Hybrids Parthenotes hESC-Derived Ova Nonegg and Nonembryo Techniques to Derive Pluripotent Stem Cells The Reality of SCR: IPSCs Are Not Replacing hESCs Stem Cell Translation and Commercialization Challenges Stem Cell Markets Stem Cell Tourism Patient and Public Perceptions Strategies to Mitigate Stem Cell Tourism Summary Acknowledgments References

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Introduction There has been considerable hype surrounding stem cell research (SCR). This hype, presented as exaggerated positive and negative portrayals of SCR, can skew public perception and sway ethics discourse toward a particular point of view (Master and Resnik, 2013). Upon the isolation of human embryonic stem cells (hESCs) in 1998, ethics discussions focused primarily on the moral status of human embryos and to a lesser degree, the physical and social harms to women egg providers. This moral discourse had received so much attention and publicity that it influenced, at least in part, the conduct of science to identify new ways of creating stem cells. In 2007, the discovery of induced pluripotent stem cells (iPSCs) was heralded in the popular press as absolving the ethical issues surrounding the moral status of human embryos and harms to women (Caulfield and Rachul, 2011) and thus, moral conversations began shifting toward the translation and commercialization of SCR, and importantly, the premature translation of stem cell interventions delivered to the public – a phenomenon called ‘stem cell tourism.’ But to date, no satisfactory or ethically sustainable solution to the initial moral issues has been resolved. As the science of SCR continues to advance and scientists discover new ways of creating embryos, the unresolved ethical issues surrounding moral status and harms to women resurface. In this article, I will discuss how the hype surrounding stem cells has influenced ethics discourse and how both scientific discoveries and ethics debates influence one another. I will cover topics such as the moral status of human embryos and the harms to women debates, the challenges of having a translation and commercialization ethos, and the ethics of stem cell tourism.

In the Beginning: The Discovery of hESCs Stem cells can be derived from a host of sources including embryonic, fetal, and adult tissue. These cells are different than most differentiated adult cells because they are capable of dividing into other cell types through a process called differentiation and can self-renew creating more undifferentiated stem cells. Different cell types vary in their ability to differentiate (denoted as potentiality). Adult stem cells have limited potentiality and can differentiate into a few cell types (multipotent) or a single cell type (unipotent). The hematopoietic stem cell is an example of a multipotent stem cell capable of differentiating into all of the cells of hematopoietic origin including T and B lymphocytes, erythrocytes (red blood cells), megakaryocytes (platelets), eosinophils, neutrophils, monocytes, and mast cells. hESCs are pluripotent and can divide into each and every cell type of the human body.

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Totipotency is when a cell(s) (e.g., a single cell embryo) can divide to become a new organism. Both adult and hESCs are desirable candidates for regenerative medicine to repair or replace diseased, destroyed, or degenerated cells and tissues. The derivation of hESCs was first accomplished by isolating cells from the inner cell mass (ICM) of day 3–5 embryos called blastocysts, which leads to its destruction (Shamblott et al., 1998; Thomson et al., 1998). Because of their potentiality and higher division rate, hESCs were sought after more than adult stem cells and hence ethical debates honed in on the destruction of human embryos. Yet another layer of scientific and ethical complexity was added to the mix – how would transplanted stem cells bypass immune rejection? The transplantation of allogeneic hESCs would not likely survive the host immune response. Scientists explained that cloning hESCs to genetically match the host through a technique known as somatic cell nuclear transfer (SCNT) may bypass immune rejection. SCNT is a technique where oocytes are enucleated of their genetic material and the nucleus of a diploid cell (e.g., from a patient’s skin cell) can be transferred into the enucleated egg. After nuclear transfer, the egg can be parthenogenically activated either through chemical or electrical means permitting it to undergo development. At the blastocyst stage, hESCs can be derived from the ICM and this time, the cells are a nuclear genetic match to the patient and when they are transplanted into the patient, they should be recognized as self and will not be immune rejected. This process was commonly referred to as therapeutic or research cloning. Instead of deriving hESCs, if the cloned embryo was transferred into a woman, then the resulting child would be a living clone of the patient – a process referred to as reproductive cloning, the technique used to clone Dolly (Campbell et al., 1996). Yet both reproductive cloned animals and cloned animal ESCs by SCNT exhibit a series of problems (reviewed in Master et al., 2007a). Performing SCNT to derive hESCs has recently been accomplished and the technique has been made more efficient so fewer ova are required (Tachibana et al., 2013).

Hype Surrounding Stem Cell Research With the advantage of hindsight, it is safe to say that many players engaged in the debate (e.g., scientists, reporters, governmental spokespersons, and politicians) made outlandish exaggerations of stem cell and cloning research merely as a means to swing moral discourse and scientific practice in a particular direction, or simply halt the science (Nisbet, 2004; Nisbet and Lewenstein, 2002). Hype can be positive or negative exaggerated representations predicting the future of SCR (Master and Resnik, 2013). There are several ethical issues with hyping SCR: (1) as science is communicated to the public primarily through the media (Nelkin, 1995), hyping SCR can distort the public’s understanding of science; (2) positive portrayals of cures can promote premature translation of research (Knowles, 2009; Ryan et al., 2010); and (3) unmet promises can lead to a loss of public trust in SCR (Ogbogu, 2006; Doerflinger, 2008; Downey and Geransar, 2008; Bubela et al., 2009; Wilson, 2009). Several popular media sources have demonstrated positive and negative hype as it relates to SCR (Kitzinger and Williams, 2005; Zarzeczny et al., 2010; Murdoch et al., 2011). The positive hype surrounding SCR has focused on (1) promising cures for many diseases; (2) creating gametes that may be used in assisted reproduction (Master, 2006); (3) developing new diagnostic and therapeutic means to treat cancer based on cancer SCR; and (4) being a major economic engine (Caulfield, 2010). The negative hype surrounding SCR and cloning has focused on cloning human armies, creating grotesque animal–human chimeras, the commodification of life including human ova, and killing unborn life (tantamount to killing babies). Despite the hyped promises and perils of SCR, several cogent moral arguments can be made to permit or prohibit the use of embryos for SCR. (Ethical issues related to reproductive cloning, the banking of stem cells, and research ethics are beyond the scope of this article.)

Personhood and the Moral Status of the Human Embryo From 1998 till 2007, the moral status of the human embryo and defining personhood dominated moral discussions on SCR. Not surprisingly, questions about when life begins and what is the moral status of human embryos stemmed from earlier arguments on abortion, which created the appropriate conceptual basis for a similar set of arguments delivered by the same set of actors (Jasanoff, 2005). Persons are ascribed inalienable rights such as the right to life, liberty, and security; interference of privacy; freedom of thought, belief, opinion, and expression; and freedom from persecution, among others (UN General Assembly, 1948). If persons are given the right to life, it was believed that if embryos were labeled as persons, they deserve protection from being destroyed. The attention given to defining personhood and ascribing moral characteristics of human embryos took foothold and became instrumental as it was believed to be the key to winning the debate. One logical flaw in this line of argument is that being a person does not automatically guarantee moral protection and not being a person does not automatically mean one can treat it as a means for research. One’s view of the moral status of human embryos, whether it is religious or secular, would at minimum, influence, if not fully guide, their views on the ethical permissibility of SCR. This is likely to include whether embryos can be purposely created for SCR, whether only excess embryos can be used, or whether embryos cannot be used for SCR. Human embryos can be created for different purposes. They may be created specifically for reproductive purposes (assisted reproduction) or they may be created specifically for research. Because assisted reproductive procedures require the creation of several embryos of which one or more may be transferred into a woman in order to achieve pregnancy, excess or surplus embryos may remain. Surplus embryos originally created for reproductive purposes that are no longer needed can be donated for research. Some observers of the SCR debate maintain that it is ethical to use only surplus embryos, but not create embryos specifically for SCR

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because these excess embryos would otherwise be destroyed or remain in a state of cryopreservation. Others, however, claim that excess embryos should only be used for reproductive purposes and cannot be used for research. These positions, in part, depend on an individual’s views of moral status. Perhaps a central and powerful voice in US debates on SCR stems from the notion that a human embryo has full moral status equivalent to persons. In general, upholders of this view are unlikely to accept the destruction of human embryos for any purpose irrespective of how laudable the research goals are or for what purpose the embryos were created, although some believe that even if embryos had full person status they can be used for hESC research (Rizzieri, 2012). Many such critics of SCR have strong moral convictions that destroying human embryos is an illicit act because it suppresses innocent human lives that ought to be protected (Paul, 2003; Vatican, 2008). In other words, the destruction of human embryos treats them only as a means for research and not as ends. According to this view, human embryos at the moment of conception have full moral status. One logical problem with this viewpoint is that it conflates species identity with membership in a moral community permitting anything genetically human to be considered moral (Marquis, 2002). Analogously, a plate of human epithelial cells would deserve the same moral status as embryos because they are merely human. A second problem with this line of argument is when, developmentally speaking, does an embryo become a person? Many supporters of this view would argue personhood is attained at the moment of conception, which may be biologically defined as the moment a sperm penetrates the outer membrane of the egg and deposits its nuclear genetic material. But at this stage, the two nuclear genomes from the male and female gametes exist separately as pronuclei and have not yet combined, which occurs a bit later in development during syngamy. (Syngamy is the point in development where the nuclear membranes of the two pronuclei – one from the sperm and another from the egg – breakdown and the nuclear genetic material combine and form a single nucleus.) So perhaps personhood is attained during syngamy. But even after syngamy, an embryo can split naturally to create twins until embryonic day 14 during gastrulation and the formation of the primitive streak when natural twinning cannot occur afterward. Identical twins in society are considered separate persons so perhaps embryonic day 14 is the pinnacle moment when personhood is attained. Yet perhaps when neurogenesis or other identifiable events occur throughout development are morally important time points. Claiming personhood at discrete developmental events in a continuous developmental process is to some degree arbitrary, and perhaps based on one’s own moral value and importance placed on particular biological points (Green, 2002). Instead of claiming full moral status at discrete developmental time points, another view generally used to justify full prohibition of using embryos for SCR is the potentiality view. The potentiality view states that if an embryo is allowed to develop fully, it will inherit the moral characteristics to the status of persons because if allowed to develop and be born, it will come to have a future like ours (FLO) and thus, should be protected from unjust killing (Marquis, 2002). As moral persons, we have values, interests, and cherish our lives and hence it is wrong to kill potential persons as it is to kill actual persons because it deprives potential persons with an FLO. Devolder claims that individuals who assign a moral difference permitting the use of surplus embryos, but prohibit the creation of embryos for SCR have professed beliefs on the potentiality theory of moral status, which she claims cannot justify this distinction (Devolder, 2005). Yet to many others, an embryo is not a person because it does not have any intrinsic value or interests, and irrespective of whether it will eventually be a person, at its current stage of development it simply does not have full moral status. It logically flows that if the embryo does not have moral status it does not deserve moral protection. An argument generally made by proponents of SCR is that persons must have cognitive capacities. Because persons are conscious and have the capacity to feel joy, pain, and other emotions is why persons make plans, have values, and take an interest in their future lives (Feinberg, 1992). Mary Anne Warren outlines five cognitive criteria to warrant membership in the moral community: “(1) consciousness (of objects and events external and/or internal to the being), and in particular the capacity to feel pain; (2) reasoning (the developed capacity to solve new and relatively complex problems); (3) self-motivated activity (activity which is relatively independent of genetic or direct external control); (4) the capacity to communicate, by whatever means, messages of an indefinite variety of types, that is, not just with an indefinite number of possible contents, but on indefinitely many possible topics; (5) the presence of self-concepts, and self-awareness, either individual or racial, or both” (Warren, 1973). Others believe that sentience is a requirement for moral community membership as sentient beings are rational, have an interest in avoiding pain, and are selfaware (Bortolotti and Harris, 2005). Although none of these criteria are independently necessary, an entity lacking all of them cannot be a moral person. Yet the problem in this line of argument is that individuals with severe developmental delay, unconscious individuals, newborn infants, and perhaps the elderly senile may be excluded from the moral community, but society grants these individuals the rights of persons. Moreover, sentience can also be ascribed to nonhuman animals making it morally unpermissive to use certain animals in research or for other uses. One issue with all of these moral theories that ascribe personhood characteristics is that they equate personhood with moral protection and nonpersonhood with using embryos for SCR. Albeit this argument has an intuitive appeal, it is not so straightforward as we afford many nonperson entities moral protections. Second, many in the public will not clearly articulate their professed beliefs of a certain moral theory and neatly align it to their choice on the conduct of human embryo research. Human embryos can be respected whether they are persons or nonpersons (Steinbock, 1992, 2001). Embryos are nonsentient beings incapable of feeling pain, pleasure, or having interests in their own welfare, but does this mean we can do whatever we want with them? We respect art, the environment, monuments, flags, burial grounds, and other nonmoral entities. Neither a van Gogh painting nor the deceased have moral status, but we still refrain from defacing art, or tramping around in cemeteries or digging up the dead. Similarly, we respect human embryos because they are developing human life and have moral value even if they do not have full moral status (Steinbock, 1992). Perhaps respect should be given because a human embryo is a salient symbol of human life and how we treat

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human embryos as a society demonstrates how we value human life more generally (Robertson, 1995). In this sense, we respect human embryos more than other human cells and tissues and perhaps this is why we are reluctant to use them in instrumental ways. The issue with the respect for embryo argument is that it is unclear how much respect should be paid to human embryos and whether this means they can or cannot be used for SCR. Some would grant respect to mean that human embryos can be sacrificed for socially valuable ends including SCR or for improving assisted reproduction procedures, but they are not to be wasted; they certainly cannot be used for trivial purposes such as making jewelry or for cosmetic testing (Steinbock, 2001). Others might grant permission to use surplus embryos for SCR because of their symbolic representation of life, but not to create embryos specifically for research (Ryan, 2001). Similar to respecting human embryos for their moral value and/or symbolic nature, several theorists contend that human embryos are sacred and as such should not be violated. The belief in the sacredness of human embryos can be based on secular beliefs or religious doctrines of life and the living. The Oxford Dictionary of English first defines sacred in a religious context as something “connected with God or a god or dedicated to a religious purpose and so deserving veneration”; it also defines sacred in secular ways as something “regarded with great respect and reverence by a particular religion, group or individual” and “regarded as too valuable to be interfered with” (The Oxford Dictionary of English, 2005). Religious-based sacredness is generally well understood, but a secular meaning of the term requires further explanation. Dworkin claims that human life is intrinsically valuable despite whether someone places value in it (Dworkin, 1993). For example, most believe that death of someone young or even an abortion is bad because it prematurely ends life even if it is not bad for a particular individual. This would mean that embryos are intrinsically valuable irrespective of whether some people have an interest in the life of embryos. Yet, many also believe that premature death is only bad if it fails to satisfy someone else’s desires or interests and as such, life is extrinsically sacred; these individuals believe that for something to have intrinsic value, it must have inherent value meaning value must derive from its internal features (Uniacke, 2004). Thus, being biologically alive does not give the entity intrinsic value and for life to have intrinsic value it must be conscious and have value in itself. In other words, many would believe that ending embryonic life is bad, not because an embryo has interests in its own existence, but because it leads to a negative experience to others. Most people who believe embryos are sacred would not want to have them deliberately destroyed for SCR, while a few may believe that despite embryos being ‘sacred,’ sacred life can be sacrificed for certain purposes like SCR. Although many commentators engaged in the embryo and SCR debate use terms like sacredness, dignity, and respect (including terms like ‘commodification’ or ‘instrumentalization’), many have argued that their amorphous and undefined nature makes these terms less action guiding and more about a visceral or emotional response using it as a way to stop or prohibit an action like using embryos for SCR (Macklin, 2003; Harmon, 2009; Caulfield and Ogbogu, 2011). Yet despite the limitation of these concepts being difficult to articulate and understand what they mean in terms of policy choices on SCR, many in the general public, politicians, and scholars readily use these terms toward establishing policies on SCR (Master and Crozier, 2011). These terms mean different things to different people and are highly personal. They have been used in ethics and policy debates not only in SCR, but in genetics and biotechnology time and time again. Albeit not cleanly defined and well-theorized, they should not be so easily dismissed as they seem to play a significant role in public ethics discourse and in shaping policy.

Harms to Women A major concern in the embryo–SCR debate is the possible physical and social harms to women who provide ova for research. Whether we want to create an international bank of hESCs or develop individual clonal lines for each patient’s needs, the number of eggs needed for such efforts would be unimaginable and possibly create a social dependency on women for their eggs (Dickenson, 2004; Testa and Harris, 2005; Beeson and Lippman, 2006; George, 2007; Baylis, 2008). Yet despite this significant moral concern, moral reflection and dialog on the harms to women have received far less play in comparison to defining moral status (Dickenson, 2006). Although human ova can be obtained from a variety of sources including oophorectomies, ovariectomies, cadavers, and fetal ovaries, most oocytes used for research are obtained from women egg providers. Obtaining eggs from women requires providing ovulatory drugs to stimulate follicular growth and development in order to retrieve several mature ova. Generally, oocytes are retrieved surgically through transvaginal follicular aspiration using ultrasound guidance under sedation. In cases where ovaries are inaccessible from the vagina, ova can be retrieved laproscopically under light anesthesia. The physical risks associated with ovarian stimulation is ovarian hyperstimulation syndrome (OHSS), which occurs through the loss of controlled ovarian stimulation. OHSS symptoms ranges in severity from minor abdominal distension and possible diarrhea to increased abdominal pain, low urine production, blood clots, respiratory distress, and thromboembolic episodes (George, 2007). Although uncertain, there remains a possible risk of ovarian and uterine cancer from long-term ovarian stimulation (Rossing et al., 1994; Wakeley and Grendys, 2000; Ness et al., 2002; Althuis et al., 2005). Surgical risks include bleeding, vaginal hemorrhage, pelvic infections and injuries, and risk from low-dose anesthetics. Beyond the physical risks, there are social risks in the possible solicitation and exploitation of women for their eggs. There are several reports of advertisements targeted to women egg providers offering substantial payments for eggs from young fecund women (Padawer, 2002; Steinbock, 2004). Although much of these substantial payments are to recruit egg donors for reproductive purposes, payments can also be offered to women to provide eggs strictly for research purposes (Dickenson, 2004). Giving eggs strictly for research alters the risk-benefit ratio as women assume all the risks and receive no direct benefit, whereas benefits are

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afforded to science and society and other individuals such as researchers (Dickenson, 2006; George, 2007). The ethical concern is that substantial financial incentives for obtaining eggs for research may influence a woman, particularly one in financial need, to assume the physical risks associated with ova stimulation and retrieval that they would normally otherwise not consider (Steinbock, 2004). As such, substantive financial inducements may compromise the woman’s ability to make an autonomous decision and exercise informed consent about donating ova for research. Countries with populations facing higher poverty rates with limited research ethics oversight may be more easily exploited (Waldby and Cooper, 2008). A demand for ova for research may also establish some form of reproductive tourism for eggs (Crozier and Martin, 2012). Yet is it prima facie unethical to pay women for their eggs? If a woman understands her choice and agrees on an arrangement where she will receive financial benefit it would not necessarily be considered exploitive or unfair since both parties obtain their desires (Gruen, 2007; Nuffield Council on Bioethics, 2011). It seems ethically acceptable for models to be paid for their looks and talents, athletes for their physical endurance and abilities, so why is it unethical for a woman to sell her ova? Perhaps this can be explained by drawing analogy that selling ‘sacred’ things like organs, blood, and gametes for profane things like cash seems morally offensive where it is permissible to exchange like things, i.e., cash for a television (Hirschman, 1991). In this sense, some believe that egg donation should be done on a strictly altruistic basis.

Alternative Technologies to the Derivation of Pluripotent Stem Cells Partly as a reaction to the moral status and other ethical issues, scientists began developing alternative technologies to derive pluripotent stem cells in order to circumvent ethical issues.

Nondestructive Techniques The first is a set of two nondestructive techniques – namely, blastomere biopsy and blastocyst transfer method (BTM) – aimed to derive pluripotent stem cells from part of the embryo and without destroying it, transfer the rest of the embryo into a woman to create a child. Blastomere biopsy is a relatively standardized procedure used in preimplantation genetic diagnosis (PGD) where one or two blastomeres from six- to eight-cell embryos are extracted and used for genetic analysis. After genetic analysis, the desired embryo is transferred into a woman. Several children have been born from PGD. In our case, blastomere biopsy would be used to derive pluripotent stem cells from blastomeres as demonstrated in mice (Chung et al., 2006) and humans (Klimanskaya et al., 2006) while the rest of the embryo is transferred to create a child. The BTM is a theoretical procedure that isolates a substantive portion of the ICM of a blastocyst to derive pluripotent stem cells while the remainder of the embryo is transferred into a woman to create a child (Liao, 2005). Isolation of ICM cells would be done using a needle similar to the one used in intracytoplasmic sperm injection (ICSI) to create embryos for male infertility. There are several issues with these nondestructive techniques. First, both the procedures attempt to absolve the moral status issue by treating the embryo not only as a means for research, but also as an end by permitting it to be born. Blastomeres in mice have been shown to be totipotent and can give rise to a new organism suggesting the same is probably true for human blastomeres (Kelly, 1977). Similarly, it has been shown that ICM and hESCs when introduced into human embryos serve to create an entirely new organism (Nagy et al., 1990, 1993; Ueda et al., 1995). If blastomeres, and potentially ICM cells, are totipotent and can be coaxed to create an organism anew, then these two techniques are not absolving the moral status issue, but merely preventing the creation of a second individual – a genetic twin (Devolder and Ward, 2007; Master et al., 2007a). A second problem with these nondestructive techniques is that they involve micromanipulation, more extensively in the case of BTM, which purposely subjects the embryo to potential physical harm that could affect is survivability and developmental potential in vitro and in the worst situation, present higher risk for congenital malformations in children (Master, 2005, 2006; Pearson, 2006). Two meta-analyses on preimplantation genetic screening (PGS) have shown that PGS has a lower pregnancy and birth rate than in vitro fertilization (IVF) or IVF with ICSI alone (Checa et al., 2009; Mastenbroek et al., 2011). The third issue is that the proposal, although both ethically and scientifically interesting, is completely impractical. No individuals undergoing assisted reproduction would opt to place their embryos at risk and potentially lower their chances for a successful pregnancy all in the name of science, and perhaps receiving a vial of stem cells genetically matched to their offspring (Master, 2005). Fourth, the nondestructive techniques still require ova from women and thus the potential harms to women remains an issue.

Altered Nuclear Transfer William Hurlbut proposed a technique called altered nuclear transfer (ANT), which uses SCNT to derive genetically matched hESCs, but where the somatic cell has a mutation that prematurely arrests the embryo from further development, generally preventing implantation, but the embryo can develop to a blastocyst stage where hESCs can be derived (Hurlbut, 2005). Proof of concept of ANT has been shown in mice (Meissner and Jaenisch, 2006). Hurlbut claims that because these embryos cannot implant, they are ‘nonembryos’ and as such they can be used to derive hESCs; this technique is a morally superior alternative to SCNT or IVF embryos because these latter methods create and destroy human embryos that undergo full development (Hurlbut et al., 2006). However, many scholars have questioned the classification of ANT-created embryos as ‘nonembryos’ (Melton et al., 2004; Guenin, 2005; Devolder, 2006). Hurlbut’s proposal focuses on the capacity of an embryo to undergo normal development, but it remains ambiguous at which point in development we would classify an entity that develops abnormally or ceases

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development prematurely a ‘nonembryo’ (Gruen and Grabel, 2006). Certainly, embryos gone awry and developing into a teratoma would not be classified as an embryo, but it is unclear whether an embryo that ceases development before implantation would be classified differently. In addition, just because an embryo does not develop normally and classifying it as a ‘nonembryo’ does not preclude it from moral protection (Guenin, 2005). Why cannot this entity be classified as an embryo with limited potential? We must also not ignore the fact that we are purposely manipulating cells and creating an embryo-like entity in order to harvest hESCs, and as such what does this say about how we regard early human life (Cheshire and Jones, 2005; Jones and Cheshire, 2005). It is thus unlikely that opponents of embryo research will permit the creation of entities by ANT as a source of hESCs. In addition, ANT, like the nondestructive techniques, still requires ova from women.

Cytoplasmic Hybrids An alternative to using human eggs that has received some attention is to create pluripotent stem cells from hybrid embryos commonly known as cytoplasmic hybrids or interspecific cytoplasmic hybrids. In some cases, this technique has been advertised as a way to sidestep the shortage of human eggs for SCR, not as a means to circumvent the harms to women (Tecirlioglu et al., 2006; Minger, 2007; Bahadur et al., 2008; Camporesi and Boniolo, 2008; Hammond and Holm, 2008). The technique uses ova obtained from nonhuman animals such as rabbits (Chen et al., 2003) or cows (Chang et al., 2003) for nuclear transfer experiments using the nucleus of a human cell. Similar to SCNT, the hybrid human–animal ovum would be parthenogenically activated and pluripotent stem cells could be derived. The use of animal eggs for nuclear transfer experiments may be advantageous for many basic scientific research endeavors such as to study diseases (i.e., mitochondrial disorders), nuclear reprogramming, how cells differentiate, and how embryos develop (Crawford et al., 2008; Swerdlow, 2007), but it is unlikely to be used for human transplantation and regenerative medicine. Scientifically, there may be reasons to doubt the efficacy of cytoplasmic hybrid research to resolve social justice concerns arising from the demand for human ova because (1) the cells derived may display differential gene expression patterns than their human embryo counterparts (Chung et al., 2009; Ledford, 2009); (2) the resulting pluripotent stem cells may display differences in energy levels than normal hESCs due to dissimilar mitochondria combinations from the animal ovum (St John and Lovell-Badge, 2007); (3) the generation of pluripotent stem cells from cytoplasmic hybrid embryos may be at risk of carrying cross-species contamination (HFEA, 2007); and (4) cytoplasmic hybrids may cause moral confusion by crossing species boundaries (Baylis, 2008).

Parthenotes Diploid oocytes can be activated by parthenogenesis (denoted as parthenotes) and can undergo embryonic development until implantation. Because parthenotes lack imprinted paternal genes, they are unable to implant and arrest development prematurely (Lyle, 1997), but they can be used to derive ESCs (Cibelli et al., 2002; Vrana et al., 2003; Kim et al., 2007). As parthenotes cannot develop similar to embryos, several commentators have argued that they can be used to derive hESCs and sidestep ethical objections related to moral status (Fangerau, 2005; Kiessling, 2005; Marchant, 2006; Rodriguez et al., 2011). However, many people may still find it morally offensive to create these embryo-like entities and it is unclear how useable hESCs derived from parthenotes would be as a source for cellular replacement therapies. Lastly, the ethical issue surrounding the harms to women remains with the use of parthenotes as a source of hESCs.

hESC-Derived Ova A second alternative to lessen the harms to women would be to create ova from the differentiation of hESCs instead of obtaining them from women (Newson and Smajdor, 2005; Testa and Harris, 2005; Master, 2006; Hammond and Holm, 2008). Both mouse and human oocytes have been shown to be created from ESC differentiation (Hubner et al., 2003; Clark et al., 2004; Lacham-Kaplan et al., 2006). There may be several scientific and safety issues of using hESC-derived ova for nuclear transfer experiments: (1) hESCderived gametes may not be effectively reprogrammed to initiate de novo embryonic development to the blastocyst stage in order to derive pluripotent stem cells (Testa and Harris, 2005); (2) embryos created from hESC-derived ova may display altered gene expression and the resulting pluripotent stem cells may not function properly; and (3) hESC-derived ova will still be used to create embryos that would be considered by some to have significant moral weight and would not resolve the moral status debate. Although the use of this technology is beneficial for basic research, much more research is required before they can be used for clinical ends (Marques-Mari et al., 2009).

Nonegg and Nonembryo Techniques to Derive Pluripotent Stem Cells The debates surrounding the moral status are intractable and unlikely to be resolved upon further moral reflection or from a greater understanding of human development. The alternative technologies to derive pluripotent stem cells fail to resolve one or more ethical challenges and in some cases create new ones. One way to absolve many of these ethical issues is to derive pluripotent stem cells from nonovum and nonembryo sources (Master and Crozier, 2012). Here, I review two strategies that aim to derive pluripotent stem cells from nonegg and nonembryonic sources by inducing pluripotency of adult differentiated cells in some manner.

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The first technique involves inducing pluripotency of somatic cells through fusion with stem cells (Do et al., 2006). In particular, Cowan and colleagues demonstrated that the fusion of hESCs with human fibroblasts resulted in a tetraploid hybrid cell that had similar morphology, growth rates, DNA methylation patterns, antigen expression profiles, and differentiation patterns to hESCs (Cowan et al., 2005). However, several moral concerns remain: (1) there is continued dependence on using established hESCs for fusion experiments since they were the products of human embryos that were intentionally destroyed in creating them; (2) it remains unclear whether the somatic genomes of hybrid cells have been sufficiently reprogrammed such that the fused cells can differentiate into a diverse population of cell types required for cell restorative therapies (Yamanaka, 2007); (3) most of the reagents used to fuse cells are cytotoxic, immunogenic, and nonspecific allowing many cell types to fuse in vitro (Sullivan and Eggan, 2006); (4) using tetraploid cells could pose a tumorigenic risk (Sullivan and Eggan, 2006) and may also be immune rejected (Ambrosi and Rasmussen, 2005); and (5) the resulting cells may have different properties and potentially posttransplantation rendering them clinically unusable (Phimister, 2005). Despite the scientific unknowns and feasibility in utilizing cell fusion to create pluripotent stem cells, the technique does seem to liberate the need for ova from women and reduce the moral risks surrounding the destruction of human embryos, but has received relatively little interest as a way to move the SCR field forward, both ethically and scientifically. A second nonegg and nonembryo technique that has truly revolutionized the field of SCR is iPSC technology. The discovery of iPSCs was a remarkable accomplishment and has swept how SCR is practiced today and for this discovery, Professors Shinya Yamanaka and John Gurdon were jointly awarded the Nobel prize for reprogramming differentiated cells into a pluripotent state – Yamanaka for iPSCs and Gurdon for SCNT (Abbott, 2012). Pluripotency can be induced in adult differentiated cells through the overexpression of key genes that reprogram the cell into a pluripotent state (Takahashi and Yamanaka, 2006). The original technique used viruses to infect cells in order to overexpress genes of interest, but since then, nonviral vectors, exogenous factors, and modified culture medium have been used to generate iPSCs (Huangfu et al., 2008; Yu et al., 2009; Zhao et al., 2010). The discovery of iPSCs has been a major accomplishment in understanding cellular reprogramming, but much work still needs to be done if iPSCs are to proceed to the clinic. Several reports comparing iPSCs with hESCs have shown lower efficiencies of iPSCs to differentiate into certain cell types (Hu et al., 2010); differences in gene expression profiles (Chin et al., 2009; Ghosh et al., 2010); iPSCs carrying epigenetic memory (Kim et al., 2010; Polo et al., 2010; Lister et al., 2011); iPSCs carrying more genetic variations and mutations or chromosomal aberrations (Mayshar et al., 2010; Gore et al., 2011; Laurent et al., 2011; Hussein et al., 2011); an increase in tumorigenic markers found in iPSCs (Malchenko et al., 2010); and potential immune recognition of iPSCs (Zhao et al., 2011). These studies demonstrate the need for further research to understand the molecular and cellular mechanisms of differentiation and development.

The Reality of SCR: IPSCs Are Not Replacing hESCs The hype surrounding iPSC research in the media has been astronomical and the greatest benefit portrayed is that iPSCs are free of moral concern (Caulfield and Rachul, 2011). Even George W. Bush in his Eight State of the Union address echoed that the iPSC breakthrough can extend the frontiers of medicine without destroying human life (Bush, 2001; Gottweis and Minger, 2008). Despite iPSCs obviating many ethical concerns (Meyer, 2008), they are certainly not replacing hESC research and instead, both are used as research tools (Scott et al., 2011). Moreover, there are ethical challenges to iPSC research, a significant one surrounding moral complicity where supporters of iPSC research are morally complicit in embryo destruction because iPSC research continues to need and depend on hESC research (Brown, 2009, 2013; Hyun, 2011). There are also a host of other ethical issues regarding banking of iPSC or other cell types, i.e., informed consent, genetic privacy, and withdrawal (Sugarman, 2008; Zarzeczny et al., 2009). Yet given that iPSC research requires and depends on hESC research and thus the continued use of human embryos and need for ova from women donors, can there be a moral and political compromise? Although many believe a loss of personal integrity in a moral compromise (Devolder, 2005; Devolder and Harris, 2005; Tannsjo, 2007), there is value in compromising ethically and politically in SCR and a policy framework could be designed to permit SCR using embryos and eggs for a certain period and then prohibiting their use when nonegg and nonembryo techniques are sufficiently well developed and can replace hESC research (Master and Crozier, 2012). Yet despite whether a compromise can or cannot be settled, it seems that bioethical exchange surrounding moral status and harms to women has been significantly stunted after the discovery of iPSCs and instead, new ethical and policy issues focus toward the translation and commercialization of SCR and stem cell tourism (Zarzeczny et al., 2009; Caulfield et al., 2012a; Kato et al., 2012). Although understanding the ethical and scientific challenges to stem cell translation and commercialization is an understandable and required next step, as basic science continues to explore the molecular mechanisms of differentiation, reprogramming, and development, new types of embryos are created and the same moral issues surrounding personhood, cloning, and harms to women are revisited (Annas, 2011; Caplan, 2011; Callaway, 2011; Darnovsky et al., 2011; Kaiser, 2011, 2013; Young, 2011). Moral reflection and discourse by bioethicists on these ‘older’ debates should still continue strongly.

Stem Cell Translation and Commercialization Challenges Beginning in the early 1980s, there has been a change in culture in academic research toward a movement to translate and commercialize research. In the United States, several factors may account for the cultural orientation from basic academic research toward

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translation and commercialization including the passing of the Bayh-Dole Act (1980), public perception of the importance of biomedical research, intellectual property decisions to commercialize research (i.e., Diamond vs Chakrabarty decision), and John Moore vs Regents of University of California court decision that property rights are awarded to scientists for their intellectual labor as opposed to the biovalue of human tissue (Annas, 1988; Mowery et al., 2001; Lemmens, 2004). This shift is seen through increases in patents and licenses (Mowery et al., 2001; Shane, 2004), funding policy (Woolf, 2008; Wadman, 2010; Nature Cell Biology, 2012; Rohn, 2012; Reed et al., 2012; Trounson and Dewitt, 2012), and the potential pressure to translate and commercialize results (Caulfield et al., 2008; Murdoch and Caulfield, 2009; Bubela and Caulfield, 2010; Caulfield et al., 2012). Academic success has always been measured by the quality and quantity of publications and funding for research, but the new metrics of success for academic researchers include patents, licenses, and the creation of start-up companies (Lemmens, 2004). The commercialization crusade has been deemed to raise a specter of ethical and integrity issues (Maienschein et al., 2008) including (1) minimizing data sharing and increase in secrecy (Louis et al., 2001; Caulfield et al., 2008; Hong and Walsh, 2009; Joly et al., 2010); (2) increase in bias of research during data collection, interpretation, and reporting (Sismondo, 2008); (3) ghost authorship (Sismondo, 2007; Wislar et al., 2011); and (4) premature translation and harms to research participants and the public (Knowles, 2009; Wilson, 2009; Ryan et al., 2010). Yet despite the push to translate research into products and services that directly profit society, only modest clinical success in SCR has been seen to date (Geffner et al., 2008; Baker, 2009; Davis, 2012; Gupta et al., 2012; Hare et al., 2012; Holding, 2012; Makkar et al., 2012; Schwartz et al., 2012; Underwood, 2012). An analysis of clinical trials shows that current clinical research focuses on hematopoietic and mesenchymal stem cells compared to what is reported in the news, which emphasizes hESCs, and neurological and cardiovascular conditions (Bubela et al., 2012). There are several challenges to the clinical translation of SCR. The first challenge would be to identify and utilize the appropriate animal model for human translation (Lindvall, 2012) as not all animal data may be good predictors in determining effectiveness in humans (Knight, 2008; Regenberg et al., 2009b). As the push to translate and commercialize may promote bias, scientists may overestimate effectiveness and underestimate contrary results (Weaver, 2010). Second, stem cells offer novel challenges as a therapeutic when compared to drug-based interventions. There is a risk of tumor formation, undesired motility of stem cells into other areas, genomic instability, and potential risks from viral and other vectors that may be used to modify cells in order to control their behavior posttransplantation (Glass et al., 1999; Master et al., 2007b; Fink, 2009; Goldring et al., 2011). Third, stem cell translation into humans requires microbiological, viral, and other contamination testing, assessing purity of populations, and scaling up to manufacture sufficient quantities of cells in a controlled manner (Halme and Kessler, 2006; ISSCR, 2008; Rayment and Williams, 2010; Goldring et al., 2011). Lastly, stem cells and stem cell products are captured within national regulatory frameworks; their classification and evaluation is different between different countries making it difficult for sponsors to seek regulatory approval (von Tigerstrom, 2008; Goldring et al., 2011; von Tigerstrom, 2011). First-in-human trials are also likely to be difficult because of the unknown outcomes and probabilities in assessing risks, which in turn can undermine a systematic analysis of risk during ethics review (Fung and Kerridge, 2013; Kimmelman, 2012).

Stem Cell Markets The positive hype surrounding SCR has spawned several markets that play on the theme that stem cells have rejuvenating and regenerative properties. The new fountain of youth in the beauty industry centers on products containing the revitalizing power of stem cells. These stem cell–based products are sold directly to consumers and include antiaging hand and facial creams, soaps, shampoos, moisturizers, mascara, eye liner, and other products. A second stem cell–based industry that also hypes the regenerative properties of stem cells is stem cell–based vitamins. Stem cell vitamins aim to increase the endogenous production of adult stem cells and ‘may aid in the repair of damaged tissue’ and ‘in the regeneration of new tissue’ (Vita-Stim Concentrate, 2012). Although the use of stem cell–based lotions and vitamins are relatively benign, the same is unfortunately not true for stem cell tourism (Master and Resnik, 2011a).

Stem Cell Tourism The term stem cell tourism is commonly used, perhaps less than ideally, to describe the phenomena of patients seeking stem cell treatments/interventions that are understudied and untested (Zarzeczny et al., 2012). Stem cell tourism originally received its name as a form of medical tourism where patients traveled to destinations to receive stem cell interventions: patients from countries like the United States, Canada, Australia, and the United Kingdom traveled to China, India, and Mexico where possible lax regulations and enforcement permitted the existence of such clinics (Ryan et al., 2010; Zarzeczny et al., 2010; Sipp, 2011). Yet more clinics are also popping up in highly regulated countries. For example, a few clinics in the United States are offering autologous adult stem cell transplantation therapies for back pain and minor injuries (Cyranoski, 2010a; Lysaght and Campbell, 2011; von Tigerstrom, 2011). Stem cell tourism is a direct-to-consumer, Internet-based market. It is unclear how many people have actually sought stem cell interventions, but it can be estimated to be at least in the tens of thousands and likely a lot more (Master and Ogbogu, 2012). We know a fair bit about stem cell tourism from patient and public web blogs and interviews, media reports, and the analysis of provider Web sites: (1) providers offer to treat a range of diseases varying in severity from backaches and sport injuries, hair loss, erectile dysfunction, to very severe debilitating diseases and injuries such as blindness, cerebral palsy, Parkinson’s disease, and spinal cord injury; (2) providers use all sorts of cell types including adult, fetal, and ESCs, which are administered surgically,

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topically, intravenously, and orally; (3) providers charge hefty sums and require repeated treatments (Lau et al., 2008; Regenberg et al., 2009a; Ryan et al., 2010); (4) clinics and the media offer mainly positive portrayals of the interventions and underemphasize risks (Lau et al., 2008; Zarzeczny et al., 2010); (5) clinics rely on patient testimonials as evidence of efficacy; and (6) there are several risks including tumors, lesions, meningitis, tremors, and death (Dobkin et al., 2006; Amariglio et al., 2009; Barclay, 2009; Miller, 2009; Cyranoski, 2010b; Mendick and Palmer, 2010; Thirabanjasak et al., 2010; Mendick and Hall, 2011; Vogel, 2011).

Patient and Public Perceptions Patients and the public have a diverse array of perspectives on SCR, therapies, and tourism. Many patients who seek stem cell interventions have dealt with conventional medical treatments and feel they have exhausted all their options and that a stem cell treatment is their only viable opportunity (Rachul, 2011). Patients believe they are informed and understand the experimental nature of the therapy, that stem cells can cure them, they are contributing to the advancement of science, and patients seem to distrust the health care and research systems in their home countries (Ryan et al., 2010; Chen and Gottweis, 2013; Rachul, 2011; Einsiedel and Adamson, 2012; Master et al., 2013). There are also some patients who received an intervention with no improvement or relief of symptoms who have become skeptical of the exaggerated claims of stem cell treatments. Interestingly, most healthy volunteers see positive aspects about a decision to go and pursue stem cell treatments (Einsiedel and Adamson, 2012).

Strategies to Mitigate Stem Cell Tourism There are three types of proposals that aim to reduce or ameliorate stem cell tourism. The first involves developing and further strengthening both domestic and international regulations and enforcement to curtail the market. Such strategies have been performed with varying success in China (Cyranoksi, 2009; Nature, 2010), Germany (Stafford, 2009; Vogel, 2011), and several other nations. Regulatory authorities and medical licensure boards are pursing and disciplining providers who have gone off the reservation (Ward, 2010; FBI, 2011a, 2011b). Legal action is a powerful means to deter illegal stem cell activity, yet the nature of the international market makes it challenging to enforce and regulate stem cell tourism. Moreover, international laws, regulations, or policies may take time to develop, and those who wish to evade the rules can move to a more permissive regulatory environment or simply break the rules and accept the penalties, especially when they are not severe (Master and Resnik, 2011b). A second means to quell stem cell tourism is to educate the public and patients. There are several amazing sources of patient information, but it seems that patients who are severely ill and desperate for a possible treatment are not likely to heed the warnings of untested stem cell treatments (Master and Resnik, 2011b; Einsiedel and Adamson, 2012). If patients are distrusting of their national health research and regulatory governance systems as some perception data indicate, it is also unlikely that educating them about the safety and efficacy of stem cell treatments will help in promoting trust or dissuade future stem cell tourists from seeking treatment; however, it is still a considerably worthwhile effort to provide robust information to potential tourists even if it does not deter them from undertaking the intervention (Master et al., 2013). A third response to stem cell tourism is for frontline physicians to inform their patients about the dangers of stem cell tourism (Caulfield et al., 2012b; Zarzeczny and Caulfield, 2010). A similar frontline approach involves scientists sharing stem cell reagents and materials for specific ‘research’ purposes and not to provide such materials to potential providers who wish to start up their stem cell tourism business; this would require scientists to evaluate the curriculum vitae and protocols from the requesters of materials and to stipulate specifically how the materials are to be used in material transfer agreements (Master and Resnik, 2011b). Although much is known about stem cell tourism, further qualitative research understanding the perceptions of a range of stakeholders is needed. For example, many celebrities and public figures have recently sought stem cell interventions, i.e., athletes for sports-related injuries. How is this shaping the public’s view on stem cell treatments? If physicians are to stop stem cell tourism, how much do they know about SCR, how many patients talk with their physicians, and what advice do physicians provide? Similarly, if education is going to stop stem cell tourism, how much education is out there, is it being targeted to the right audiences, and how effective will it be as a deterrent to future stem cell tourists? These and many other research questions still need to be addressed (Master and Ogbogu, 2012).

Summary The ethics and hype in science and technology have been especially rampant in SCR to the point where it has influenced the conduct of science (to make ethical stem cells) and where the hype may have contributed to prematurely market regenerative products and interventions – including the stem cell tourism market. The hype and promises surrounding SCR have also likely contributed to developing a translational and commercialization ethos within the field. While many of the problematic issues surrounding moral status and harms to women who provide eggs for research are difficult to resolve, deriving stem cells from nonegg and nonembryonic sources (e.g., through iPSC research) does seem to be a step in the right direction as they skirt around these ethical issues. As iPSC and SCR show tremendous promise as a possible curative for many diseases and injuries, much more scientific research is still needed. Accurate portrayals of SCR need to be emphasized and hype needs to be tempered so that the public has realistic expectations of where the science is, where it is heading, and when the fruits of labor will be delivered to them.

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Acknowledgments Part of this work was supported by the Cancer Stem Cell Consortium (CSCC) with funding from Genome Canada and the Ontario Genomics Institute (OGI-047); the Canadian Stem Cell Network; and the Interdisciplinary Chronic Disease Collaboration funded by Alberta Innovates. I would like to thank Professor Timothy Caulfield and Dr Lisa Campo-Engelstein for providing helpful feedback to the article and project support by Mr Daniel G. Frederick, Ms Robyn Hyde-Lay, and Ms Faria Grant.

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Eye Diseases and Stem Cells H Ouyang, DH Nguyen, and K Zhang, Shiley Eye Center and Institute for Genomic Medicine, University of California at San Diego, La Jolla, CA, USA © 2019 Elsevier Inc. All rights reserved.

Introduction Anatomy of the Eye Schematic Overview of the Retina Diseases of the Retina Age-Related Macular Degeneration Pathophysiology of AMD Diagnosis of AMD Treatment of AMD Glaucoma Pathophysiology of POAG Diagnosis of POAG Treatment of POAG Retinitis Pigmentosa Diabetic Retinopathy Sources of Stem Cells Embryonic Stem Cells Induced Pluripotent Stem Cells Mesenchymal Stem Cells Retinal Progenitor Cells Müller Glial Cells Stem Cell Therapy in Retinal Diseases RPE Transplantation Photoreceptors Cells Transplantation Ganglion Cells Transplantation Cornea Injury and Stem Cell Therapy Potentials and Limitations of Stem Cell-Based Therapy References

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Glossary AMD Age-related macular degeneration CLET Cultivated corneal limbal epithelial transplantation CNS Central nervous system DP Diabetic retinopathy ESCs Embryonic stem cells IOP Intraocular pressure iPSCs Induced pluripotent stem cells LESC Limbal epithelial stem cells MGCs Müller glial cells MSCs Mesenchymal stem cells POAG Primary open-angle glaucoma RD Retina degeneration RGCs Retinal ganglion cells RP Retinitis pigmentosa RPE Retinal pigment epithelium RPCs Retinal progenitor cells SC Stem cell TACs Transit amplifying cells VEGF Vascular endothelial growth factors

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Introduction Visual impairment or blindness can dramatically reduce the quality of life for an individual. It can be devastating to suddenly or gradually lose the ability to read, drive, or recognize one’s family and friends. Vision loss can result from common eye diseases which lead to damage of the retina and optic nerve, such as age-related macular degeneration (AMD), diabetic retinopathy (DR), retinitis pigmentosa (RP), and glaucoma. Other injuries to the cornea can also cause scarring and subsequent blindness. Vision loss is a considerable public health burden. According to recent statistics reported by the Prevent Blindness America organization, the annual cost of adult vision problems in the U.S. is around $51.4 billion, and the cost will continue to rise with 50 000 or more people becoming blind each year. Transplantation of stem cells in animal models of eye diseases or clinical trials demonstrated that stem cell-based therapy is a promising treatment to restore vision in patients with retinal diseases or ocular surface injury.

Anatomy of the Eye The eyeball contains three main layers: the outer fibrous layer, the middle uveal layer, and the inner neural layer (Figure 1). The surface of the eye is protected by the conjunctiva, which is composed of stratified columnar epithelium. Underneath the conjunctiva is the sclera, commonly known as the white part of the eye. The sclera is made of a tough tissue that connects with the cornea. The border between the sclera and the cornea is called the limbus. The pupil is located in the center of the iris and can change size to accommodate the amount of incoming light. The middle layer of the eyeball houses the aqueous humor, lens, blood vessels, lymphatic vessels, ciliary body and its extension, the iris. The lens can change shape and in doing so change how much light rays are ‘bent’ before they go through a viscous and jellylike substance called the vitreous, before finally arriving at the retina. The retina is a crucial light sensing structure by which visual information is translated into electrical impulses and transmitted to the visual cortex of the brain by the optic nerve. The pathway light travels consist of first traveling through the cornea and pupil. An amount of refraction happens at the cornea, after which light is further focused by the lens. It continues through the vitreous and onto the retina. The retina transfers light signals to the brain which processes the information into images that we see.

Figure 1

Schematic diagram of the anatomy of the eye and layers of the retina.

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Schematic Overview of the Retina There are nine layers of the retina (Figure 1). Starting most internally and immediately behind the vitreous, the layers are the internal limiting membrane, the stratum opticum, the ganglion cell layer, the inner plexiform layer, the inner nuclear layer, the outer plexiform layer, the outer nuclear layer, the external limiting membrane, and the layer of photoreceptors. There are millions of photoreceptors called cones and rods, which sense color or light intensity, respectively. Rods are distributed more heavily around the peripheral retina while cones are mainly concentrated in the macula that is at the back of the retina. The central part of the macula is called the fovea, containing the highest amounts of cones and thus providing the sharpest sense of sight. Rods are totally absent at the fovea. Rods and cones transmit visual signals to neurons called bipolar cells, which in turn pass the signals to retinal ganglion cells (RGCs) whose nerve fibers merge to form the optic nerve. Support cells, called horizontal cells, facilitate communications between photoreceptors. Amacrine cells are support cells that control communications between bipolar cells and RGCs. Müller glia and astrocytes are another two support cells among the neural retina.

Diseases of the Retina Age-Related Macular Degeneration AMD is a progressive disease of the macula, which is an important component in the retina that is responsible for visual acuity and color vision. AMD is among the top three causes of irreversible blindness among elderly people worldwide, while in the United States it is the leading cause of visual impairment (Pascolini et al., 2004). It is estimated that about 1.75 million people in the United States are affected by AMD, and this figure is expected to increase to 3 million by 2020 due to the rapid aging of the American population. Of note, about 7 million Americans currently have evidence of early changes in the retina that are at risk for progression to AMD (Friedman et al., 2004). Visual impairment is a feared disability, and is related to a significant reduction in the quality of life and overall health status.

Pathophysiology of AMD AMD is a multifactorial disease with a complex mechanism of development that is not completely understood. It has been well documented that AMD results from genetic and environmental factors. Risk factors for developing AMD include age, cigarette smoking, obesity, cardiovascular disease, sunlight exposure, and certain genetic variants in the physiologic pathways that regulate inflammation, angiogenesis, and innate immunity (Lim et al., 2012). Evidence of early AMD development is drusen, which are accumulations of extracellular debris under the retinal pigment epithelium (RPE), a crucial component of the retina. Drusen are categorized by location, size, and texture. They are observable on clinical exams and photographs of the retina, and their presence without evidence of vision loss defines early AMD. Many studies have found that substantial progression in size and number, as well as extensive involvement of drusen in the area of the macula, is associated with the risk of developing the two advanced forms of AMD with vision loss, including geographic atrophy (GA) and choroidal neovascularization. Compared to the normal eye (Figure 2), GA, or ‘dry’ AMD, is defined as the presence of a well-demarcated area of at least 175 mm in diameter in the retina where there is marked atrophy of the RPE, which manifests as focal depigmentation without evidence of abnormal blood vessel growth (Bird et al., 1995) (Figure 3). Choroidal neovascularization, or ‘wet’ AMD, is diagnosed when there is

Figure 2

Normal eye. Fundus photograph.

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‘Dry’ age-related macular degeneration, or geographic atrophy. Fundus photograph.

Figure 4

‘Wet’ age-related macular degeneration, or choroidal neovascularization. Optical coherence tomography image.

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presence of growth of new blood vessels under the retina that can cause bleeding, fibrosis, and detachment of the retina (Figure 4). The two advanced forms of AMD can lead to significant deterioration of vision.

Diagnosis of AMD Patients with AMD typically present with a progressive decrease in central vision, which manifests as difficulty seeing straight lines, dark patches in the visual field, and decrease of color vision. In some cases of ‘wet’ AMD in which bleeding from the abnormal blood vessels cause a more acute insult to the retina, sudden and more severe decreases in vision may occur due to detachment of the retina. Diagnosis of AMD requires a complete ophthalmologic examination, which typically includes visual acuity measurements, intraocular pressure (IOP) reading, slit-lamp examination, indirect ophthalmoscopy, and ocular imaging studies. Fundus photography, optical coherence tomography, and fundus fluorescein angiography are several imaging modalities that are very helpful in the diagnosis of AMD and categorization of the stage and type of the disease. They are instrumental in detecting the presence of fluid under the retina and disruptions of the layers of the retina that reflect advanced AMD, as well as provide objective means to evaluate effectiveness of treatment in the ‘wet’ type of AMD.

Treatment of AMD Currently, there is no treatment for ‘dry’ AMD. In the case of ‘wet’ AMD, intraocular injection of antiangiogenic medications is available to halt and/or minimize the growth of abnormal blood vessels. These medications target vascular endothelial growth factors (VEGF), which are key components in the generation of new blood vessels. Current anti-VEGF therapy includes ranibizumab (‘Lucentis’), bevacizumab (‘Avastin’), and aflibercept (‘Eylea’). Several short-term studies have indicated that intravitreal bevacizumab results in improvement in visual acuity that is similar to the improvement seen with ranibizuma (Spaide et al., 2006; Algvere et al., 2008). In 2011, an NIH sponsored randomized clinical trial provided further evidence that ranibizumab and bevacizumab provided similar effects in neovascular AMD when administered on the same schedule (Martin et al., 2011). Other treatment modalities include laser photocoagulation, in which direct laser beams target choroidal vessels to stop their growth.

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Glaucoma Glaucoma is a chronic and progressive disease of the optic nerve, which causes the loss of peripheral vision that may progress to total loss of vision if untreated. Glaucoma is categorized based on the angle of the iridocorneal angle, or the angle between the iris and the cornea. Primary open-angle glaucoma (POAG) is the most common type of glaucoma in the Western world, and is characterized by a normal and open iridocorneal angle, in contrast to angle-closure glaucoma in which the iridocorneal angle creates obstruction of the outflow of the aqueous humor, the circulation that nourishes the cornea and lens, and causes an acute increase in IOP. POAG can be, but is not always, associated with elevated IOP. The molecular pathogenesis of this disease is largely unknown, and lowering IOP is the only available treatment. However, POAG patients eventually progress to blindness despite aggressive IOP treatment. Risk factors for POAG include age, elevated IOP, family history of POAG, myopia, and being of African descent. Some research studies have also found that having a low diastolic pressure is another risk factor (Kwon et al., 2009). Glaucoma poses a heavy public health burden, and is currently the second most common cause of blindness worldwide. In the United States, adult-onset POAG is defined as occurring after 40 years of age and is estimated to affect two out of every 10 adults aged 40 and older. In people with African ancestry, glaucoma is the leading cause of irreversible blindness, and about one-third of the blindness is due to POAG (Leske, 2007).

Pathophysiology of POAG POAG is characterized by slow and progressive optic nerve degeneration due to the death of RGCs, which are neurons with critical functions in the visual pathway. Loss of RGCs and their axons lead to characteristic appearances of the optic nerve that are observable on clinical examination and imaging studies. Typically, compared to the normal eye (Figure 2), the glaucomatous eye has evidence of thinning of the neuroretinal rim which enlarges the optic nerve cup, and increases the optic cup to optic disk ratio (Figure 5). At the same time, there is also loss of surrounding support cells and vasculature. The loss of RGCs and their axons are the cause of visual impairment in POAG (Quigley, 2011).

Diagnosis of POAG As POAG is a chronic and progressive disease that causes painless loss of vision, many patients with POAG may not seek medical attention until the damage is advanced with a substantial loss of visual field. Besides a complete ophthalmologic examination which includes measurement of the IOP, imaging studies to assess the severity of damage to the optic nerve are crucial. Automated visual field testing provides an objective assessment of the degree of loss of visual field, while direct photography of the optic nerve provides a direct evaluation of optic nerve cupping.

Treatment of POAG Controlling IOP is important in the management of POAG. Surgical and medical alternatives exist for lowering IOP. Typically, a patient with POAG is started on eye drops that lower IOP and maintained on the medication or combination of medications as long as the pressure is kept within a normal range. When medical treatment fails, more invasive options are needed such as laser therapy to the drainage network of the eye to enhance outflow of the aqueous humor, or surgical procedures where drainage devices are placed inside the eye to facilitate the drainage of intraocular fluid and therefore lower the IOP.

Figure 5

Primary open-angle glaucoma. Image of the optic nerve.

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Retinitis Pigmentosa A degenerative disease of photoreceptor cells in the retina, RP (Kempen et al., 2004) is a group of diffuse retinal dystrophies that is genetically and clinically diverse. RP is an inheritable disease, with many cases due to mutations of the rhodopsin gene, which is responsible for the production of the pigment of photoreceptors. In some cases, RP is a sporadic mutation without any previous family history. RP can present as a solitary ocular defect or in combination with other systemic pathologies such as hearing loss (Usher syndrome) and kidney disease (Bardet–Biedl syndrome). Patients typically present with loss of peripheral and night vision, usually in the second decade of life but may manifest earlier depending on the pattern of inheritance. RP is characterized by pigmentary changes in the retina which give the disease its name, as well as changes in the retinal blood vessels and degeneration of the photoreceptors. Optical imaging studies to detect pigmentary changes and disruption of retinal layers, as well as eletroretinography, or the study of the function of various retinal cells including photoreceptors, are extremely helpful in the diagnosis of RP. Currently there is no cure for RP.

Diabetic Retinopathy Diabetes is one of the most common diseases worldwide. It is estimated that the number of patients with diabetes in the world will reach 429 million by 2030. Approximately 4.1 million people in the United States have DR, of which almost 1 million has sightthreatening retinopathy (Kempen et al., 2004). It is most prevalent in Latino Americans. DR (Jiang et al., 2002) is a disease of the retinal blood vessel networks that can lead to irreversible blindness. It is a severe and the most common complication of diabetes. The two most important risk factors for DR are the duration of diabetes and the adequacy of glucose control. DR can be classified into two forms, nonproliferative and proliferative, depending respectively on the absence or presence of abnormal retinal blood vessels and bleeding. The major mechanisms by which proliferative DR results in vision loss are (1) retinal vascular leakage and exudation resulting in macular edema (Thompson et al., 1996), and (2) retinal capillary closure leading to retinal ischemia and secondary neovascular proliferation with its attendant complications of vitreous hemorrhage and tractional retinal detachment. On ocular imaging studies, disruption of the retinal layers due to accumulation of fluid and blood from abnormal vessels leakage can be observed. Treatment of proliferative DR includes laser therapy to target abnormal fluid accumulation in the eye, injection of antiangiogenic medications to limit the growth of aberrant blood vessels, and in severe cases surgeries to prevent further injuries to the retina. DR is a chronic disease, and timely diagnosis and close ophthalmologic monitoring are crucial to prevent the onset or progression of vision loss.

Sources of Stem Cells Embryonic Stem Cells Embryonic stem cells (ESCs) are derived from the inner cell mass of the blastocyst-stage embryos. They possess unlimited selfrenewal capabilities and the ability to differentiate into any adult cell types within human tissue (Evans and Kaufman, 1981). ESCs hold great therapeutic promises in the regeneration of functional cell types including neurons (Reubinoff et al., 2001; Zhang et al., 2001), hepatocytes (Shirahashi et al., 2004), and cardiomyocytes (Laflamme et al., 2007), among others. Successful differentiation to eye field cells from human ESCs has been done since 2009. Data have shown that ESCs can differentiate into photoreceptor progenitors and RPE in a sequence and time course that mimic in vivo human retinal development (Meyer et al., 2009). The Sasai group has demonstrated that an optic cup can form by self-organization both in human and mouse ESC culture (Nakano et al., 2012; Eiraku et al., 2011). This ESC-derived retina structure grows into multilayered tissue containing two types of photoreceptors, rods and cones. Whether this optic cup can be induced to form an entire eye remains to be seen, but ESCs do hold potential applications to the treatment of retinal degenerative disorders.

Induced Pluripotent Stem Cells Induced pluripotent stem cells (iPSCs) are adult cells that have been made to resemble ESCs. iPSCs are induced in vitro to express certain genes and factors such as Oct3/4, Sox2, Klf4, and c-Myc that are used in ESCs. iPSCs have been shown to have regenerative properties and have also been shown to be similar in characteristics and abilities to ESCs (Takahashi and Yamanaka, 2006). It has been proven that human iPSCs have a similar potential as ESCs to form RPE and photoreceptors (Meyer et al., 2009). iPSCs transplantation does not require immunosuppression and has no ethical issues. However, iPSCs and ESCs differ in clinical outcomes. Moreover, potential risks exist during integration because of the oncogene expression for iPSCs generation (Puzio-Kuter and Levine, 2009). Transplantation of iPSCs was observed to form tumors in the three germ layers in mice.

Mesenchymal Stem Cells Mesenchymal stem cells (MSCs) are cells that are collected from bone marrow biopsies and umbilical cord blood, which are used for autologous cell transplantation (Bianco and Robey, 2001). They can also be harvested from adipose tissue, placenta (In ’t Anker

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et al., 2004), cord blood, and liver (Campagnoli et al., 2001). Haematopoetic stem cells are mobilized from bone marrow to the bloodstream which makes it less invasive to obtain these cells, though methods of mobilizing mesenchymal cells have proven to be challenging. MSCs differentiate into mesodermal cells (bone, cartilage, and muscle), but have also been shown to have the ability to dedifferentiate into nonmesoderm cells in vitro (Jiang et al., 2002). An in vivo animal model has shown that when MSCs was injected into the subretinal space, the stem cells were integrated and differentiated into photoreceptors, and also slowed the degradation of retinal cells in Royal College of Surgeons rats (Inoue et al., 2007).

Retinal Progenitor Cells Retinal progenitor cells (RPCs) are found from fetal or neonatal retinas. These cells are used to create all of the retinal cells that are formed during embryonic development. In the ectoderm, cells are pushed by transcription factors into a multistep cascade to develop retinal tissue. RPCs can be used to grow neuronal and glial cells (Müller cells), retinal support cells, and can enter all retinal layers to adopt characteristics of several retinal cells (Reh, 2006). This ability suggests that RPCs can be useful in treating retinal degenerative diseases.

Müller Glial Cells It is more accepted that the central nervous system could be regenerated from radial glial cells in recent years. Müller glia is a kind of radial glia located in the retina. Studies have demonstrated that species like zebra fish could regenerate retina from Müller glia by reentering the cell cycle and producing neuronal progenitor cells (Fimbel et al., 2007). A population of Müller glia with stem cell characteristics has also been found in the human adult retina. In vitro markers of neural progenitors are expressed and some of these cells even express markers of mature retinal neurons under different culture conditions (Lawrence et al., 2007). Experiments on the transplantation of Müller glial cells (MGCs) to rodent ganglion cell injury model showed that predifferentiated cells can integrate into the host RGC layer and restore partial scotopic threshold response in electroretinography studies (Singhal et al., 2010).

Stem Cell Therapy in Retinal Diseases The eye is thought to be an ideal organ for cell transplantation due to its special structural characteristics. The optical media of the eye, composed of cornea, lens, and vitreous, are extremely clear when healthy, which facilitate intravitreal and subretinal injection of cells. After surgery, transplanted cells are assessed by imaging techniques such as scanning laser ophthalmoscope and optical coherence tomography. Furthermore, cells could be ablated by laser treatment if overproliferation happened.

RPE Transplantation The RPE is a monolayer of polarized pigmented cells that lies between the photoreceptor outer segments and the choriocapillaris. It plays a crucial role in maintaining the outer blood retinal barrier. The apical surface of the RPE interdigitates with the adjacent photoreceptor outer segments and participates in phagocytosis, transportation, and production of essential components. RPE dysfunction usually results in photoreceptor degeneration and leads to several retinal disorders, such as AMD and RP (Kempen et al., 2004). Despite attempts to transplant autologous RPE to the macular region in patients with AMD, there was no postoperative improvement in visual acuity at 1 year (Del Priore et al., 2001). Other disadvantages for autologous RPE grafts still remain, including the difficulty and impracticality of obtaining sufficient quantities of autologous RPE cells, and the disadvantage that autologous cells always carry the same genetic defect that led to the host disease. Fresh fetal and adult RPE cells have been employed to replace lost RPE in patients with GA (Weisz et al., 1999), but considerable variability in the quality of the obtained RPE is one of its limitations. These issues highlight the urgent need of exploring stem cell-based therapy to create functional RPE for transplantation. Recently, the US Food and Drug Administration has approved the use of human ESC (hESC)-RPE in two human clinical trials in patients with Stargardt’s dystrophy and AMD. A preliminary report in 2012 showed the description of hESC-derived cells transplanted into human patients with Stargardt’s macular dystrophy and dry AMD. In this report, hESCs are differentiated into greater than 99% pure RPE under defined condition. Submacular injection of 5  104 hESC-RPE cells was delivered into a preselected macular site with native albeit compromised RPE and photoreceptors. After 4 months, there were no identified signs of hyperproliferation, tumorigenicity, or immune-mediated transplant rejection in either patient. And it is encouraging that during the observation period neither patient lost vision. Best corrected visual acuity improved from hand motions, and vision also seemed to improve in the patient with dry AMD (Schwartz et al., 2012). Successful RPE transplantation is not only determined by the quality and integration of cells, but also how these cells could survive the rejection environment in the subretinal space. Moreover, the use of hESCs is still ethically controversial. In this case, RPE-like cells generated from patients’ own iPSCs may be considered a better source. These iPSC-RPE cells are similar to RPE cells according to the morphology, function of photoreceptor outer segment phagocytosis and shared markers of developing and mature RPE cells.

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Photoreceptors Cells Transplantation Rods and cones are two types of photoreceptors in the retina. Rods are responsible for sensitivity to light and dark changes, movement and shape of objects, while cones provide color vision (green, red and blue) and sharp images. When photoreceptors are completely lost in retinal diseases, such as RP, vision restoration is impossible even with therapy of RPE transplantation. Therefore, a replacement of photoreceptors is necessary in such cases. Preliminary data showed that transplantation of fetal retinal cells into the midbrain was capable of making pretectal connections and enabled the response to light (Klassen and Lund, 1990). Fetal retinal sheets as a cell mixture has also been used in patients with AMD and RP and showed variable effects (Seiler and Aramant, 2012). Some improvement in visual function is more likely contributed by the positive factors secretion from transplanted tissue. The feasibility of photoreceptor transplantation as a therapeutic strategy was shown by transplanted photoreceptor precursors into adult Gnat1/ mice, which lack rod function and thus provide a model of congenital stationary night blindness. These precursors could form synaptic connections with other retinal cells in the recipient retina and are light-responsive with similar function to normal photoreceptors in adults (Pearson et al., 2012). RPCs transplantation in animal studies also demonstrated the ability to migrate and differentiate into photoreceptors (Bartsch et al., 2008). Unlike ESCs and iPSCs, RPCs do not have unlimited self-renewal ability and have to be reobtained from fetus. Alternative strategies including generating photoreceptors from ESCs and iPSCs could be used as inexhaustible sources. Human ESC-derived RPCs have been transplanted into Crx-deficient mice, an animal model for Lebers Congenital Amaurosis, with the improvement in light response (Lamba et al., 2009). An optimized protocol for differentiating photoreceptors from ESCs has been established recently that offers a direct transplantation of photoreceptors to patients (Meyer et al., 2009). However, besides the ethical restrictions and immune rejection following implantation, other disadvantages of poor cell quality, difficulty in purification and viability of photoreceptors remain to be solved.

Ganglion Cells Transplantation RGCs receive visual signal from photoreceptors. Their axons bundle together to form the optic nerve and extend into the brain. The death of RGCs is the cause of injury to the optic nerve in the common neurodegenerative disease of glaucoma (Quigley, 2011). ESCs (Aoki et al., 2008) and MSCs (Johnson et al., 2010) could serve as a potential source of replacement for damaged RGCs. Transplanted cells have been shown to migrate and integrate into the retina. However, to date, there has been no evidence of functional improvement. There are protection effects in transplantation to rat model of glaucoma, but the reason is thought to be due to the secretion of neurotrophic factors (Cho et al., 2005). MGCs have been reported to retain the ability to divide indefinitely in vitro (Limb et al., 2002). Recently, it has been reported that after an intravitreal injection of MGCs into RGC-depleted retina, vision is partly restored and the injected cells can be differentiated into RGC precursors with the integration into the retina (Singhal et al., 2010). Cell replacement by MGCs transplantation may therefore offer a more promising therapy after irreversible damage or RGC loss in glaucoma. Of note, a successful transplants/replacement means the transplanted cells must not only migrate and integrate into the ganglion cell layer, but also have the ability to differentiate into functional RGCs and generate axons which could extend to the optic nerve and beyond, which is a big additional challenge when compared to RPE and photoreceptor cell transplantation.

Cornea Injury and Stem Cell Therapy The cornea is composed of five layers, including the outermost stratified corneal epithelium, Bowman’s layer, corneal stroma, Descemet’s membrane, and innermost corneal endothelium. Highly resistant tight junctions formed between neighboring corneal epithelium cells protect the eye from injury and foreign body (Klyce, 1972). The transparency of the cornea depends on the epithelial integrity and stromal avascularity. The cornea also provides the majority (two thirds) of the refractive power of the eye (Meek et al., 2003). So the clarity of the cornea is essential for visual acuity. There is a population called limbal epithelial stem cells (LESCs), which are located at the junction of cornea and sclera in an area known as the limbus. These LESCs are presumed to be primitive. They can symmetrically divide to self renew and asymmetrically produce daughter transit amplifying cells (TACs) which migrate centripetally to populate the basal layer of the corneal epithelium (Tseng, 1989). The TACs divide and migrate superficially, becoming more differentiated and finally turning into postmitotic terminally differentiated cells. LESCs deficiency can be the result of primary or acquired disease which affects corneal wound healing and surface integrity, such as chemical (alkali/acid) or thermal burns, aniridia and Stevens-Johnson syndrome. Conjunctivalization, neovascularization, chronic inflammation, recurrent erosions, ulceration, and scarring are possible sequelae, among pain, opacity, and vision loss (Holland and Schwartz, 1996). Allogeneic corneal transplantations, a surgical procedure of transplanting the cornea from a donor, have been successful in restoring patients’ vision to some extent. Still there are several issues to be faced. First, the number of available donors is not sufficient to meet demand and second, the increasingly popular laser eye surgery for refractive error often makes the cornea unsuitable for transplantation.

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Recently, cultivated corneal limbal epithelial transplantation has been described as a promising treatment for limbal stem cell deficiency patients. Pellegrini et al. reported that cultured human limbal stem cells from biopsies of limbal tissue could be a viable source of cells for transplantation to treat burned human corneas. They obtained limbal stem cells from the healthy eyes of 112 patients with ocular burns, and transplanted the cultured cells onto the patients’ damaged eyes. After a 10-year follow-up, permanent restoration of a transparent and self-renewing corneal epithelium was found in three-quarters of the study patients (Rama et al., 2010). Since then, various substrates have been used for limbal stem cell expansion, such as human amniotic membrane (Mariappan et al., 2010), fibrin gels, and temperature-responsive culture inserts. Other sources of epithelial stem cells have also been tried as an alternative option. Transplantation of autologous oral mucosal epithelial stem cells showed a well-reconstructed cornea for patients who have limbal stem cell deficiency, which presented benefits for the treatment of bilateral severe disorders of ocular surface without immunosuppression requirement.

Potentials and Limitations of Stem Cell-Based Therapy With the characteristics of unlimited proliferation and differentiation ability for multiple cell types, stem cells represent a promising approach for vision restoration. A successful replacement of dysfunctional cells by transplantation holds great hope for patients. However, although ESCs have been considered to be an ideal candidate for stem cell-based therapy, ethical objections exist because of its source of human embryo donation. For iPSCs, the use of viruses in generating the cells raises a serious safety issue that transgene integration may lead to mutation within the host genome. In addition, along with their high pluripotency, ESCs and iPSCs also showed the ability of teratoma formation. Tumor growth has been found in mice after the transplantation of a neutrally selected ESC (Arnhold et al., 2004). Indeed, controlling the direction of stem cell differentiation remains a major challenge. To date, the protocol used for differentiation of specific cells requires the use of complex factors and takes many steps. In addition, immune rejection of grafted cells is also a considerable cause of transplantation failure. Despite many challenges that will need to be solved before stem cells become effective clinical therapy, the unique characteristics of stem cells and several successful clinical trials make stem cell-based therapy applicable and promising for vision restoration. Keep in mind that if an optic cup could be generated from a single stem cell, miracle is just about time.

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Kempen, J. H., O’colmam, B. J., Leske, C., Haffner, S. M., Klein, R., Moss, S. E., Taylor, H. R., Hamman, R. F., West, S. K., Wang, J. J., Congdon, N. G., Friedman, D. S., & GRP, E. D. P. R. (2004). The prevalence of diabetic retinopathy among adults in the United States. Arch. Ophthalmol., 122, 552–563. Klassen, H., & Lund, R. D. (1990). Retinal graft-mediated pupillary responses in rats: restoration of a reflex function in the mature mammalian brain. J. Neurosci., 10, 578–587. Klyce, S. D. (1972). Electrical profiles in the corneal epithelium. J. Physiol., 226, 407–429. Kwon, Y. H., Fingert, J. H., Kuehn, M. H., & Alward, W. L. (2009). Primary open-angle glaucoma. N. Engl. J. Med., 360, 1113–1124. Laflamme, M. A., Chen, K. Y., Naumova, A. V., Muskheli, V., Fugate, J. A., Dupras, S. K., Reinecke, H., Xu, C., Hassanipour, M., Police, S., O’sullivan, C., Collins, L., Chen, Y., Minami, E., Gill, E. A., Ueno, S., Yuan, C., Gold, J., & Murry, C. E. (2007). 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In vitro characterization of a spontaneously immortalized human Muller cell line (MIO-M1). Invest. Ophthalmol. Vis. Sci., 43, 864–869. Mariappan, I., Maddileti, S., Savy, S., Tiwari, S., Gaddipati, S., Fatima, A., Sangwan, V. S., Balasubramanian, D., & Vemuganti, G. K. (2010). In vitro culture and expansion of human limbal epithelial cells. Nat. Protoc., 5, 1470–1479. Martin, D. F., Maguire, M. G., Ying, G. S., Grunwald, J. E., Fine, S. L., & Jaffe, G. J. (2011). Ranibizumab and bevacizumab for neovascular age-related macular degeneration. N. Engl. J. Med., 364, 1897–1908. Meek, K. M., Dennis, S., & Khan, S. (2003). Changes in the refractive index of the stroma and its extrafibrillar matrix when the cornea swells. Biophys. J., 85, 2205–2212. Meyer, J. S., Shearer, R. L., Capowski, E. E., Wright, L. S., Wallace, K. A., Mcmillan, E. L., Zhang, S. C., & Gamm, D. M. (2009). Modeling early retinal development with human embryonic and induced pluripotent stem cells. Proc. Natl. Acad. Sci. U.S.A., 106, 16698–16703. Nakano, T., Ando, S., Takata, N., Kawada, M., Muguruma, K., Sekiguchi, K., Saito, K., Yonemura, S., Eiraku, M., & Sasai, Y. (2012). Self-formation of optic cups and storable stratified neural retina from human ESCs. Cell Stem Cell, 10, 771–785. Pascolini, D., Mariotti, S. P., Pokharel, G. P., Pararajasegaram, R., Etya’ale, D., Negrel, A. D., & Resnikoff, S. (2004). 2002 global update of available data on visual impairment: a compilation of population-based prevalence studies. Ophthalmic Epidemiol., 11, 67–115. Pearson, R. A., Barber, A. C., Rizzi, M., Hippert, C., Xue, T., West, E. L., Duran, Y., Smith, A. J., Chuang, J. Z., Azam, S. A., Luhmann, U. F., Benucci, A., Sung, C. H., Bainbridge, J. W., Carandini, M., Yau, K. W., Sowden, J. C., & Ali, R. R. (2012). Restoration of vision after transplantation of photoreceptors. Nature, 485, 99–103. Puzio-Kuter, A. M., & Levine, A. J. (2009). Stem cell biology meets p53. Nat. Biotechnol., 27, 914–915. Quigley, H. A. 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(2004). Differentiation of human and mouse embryonic stem cells along a hepatocyte lineage. Cell Transplant, 13, 197–211. Singhal, S., Lawrence, J. M., Salt, T. E., Khaw, P. T., & Limb, G. A. (2010). Triamcinolone attenuates macrophage/microglia accumulation associated with NMDA-induced RGC death and facilitates survival of Muller stem cell grafts. Exp. Eye Res., 90, 308–315. Spaide, R. F., Laud, K., Fine, H. F., Klancnik, J. M., Jr., Meyerle, C. B., Yannuzzi, L. A., Sorenson, J., Slakter, J., Fisher, Y. L., & Cooney, M. J. (2006). Intravitreal bevacizumab treatment of choroidal neovascularization secondary to age-related macular degeneration. Retina, 26, 383–390. Takahashi, K., & Yamanaka, S. (2006). Induction of pluripotent stem cells from mouse embryonic and adult fibroblast cultures by defined factors. Cell, 126, 663–676. Thompson, M. J., Toomre, J., Anderson, E. R., Antia, H. M., Berthomieu, G., Burtonclay, D., Chitre, S. 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Human Parthenogenetic Pluripotent Stem Cells N Turovets, University of California, Irvine, CA, USA M Csete, Huntington Medical Research Institutes, Pasadena, CA, USA © 2019 Elsevier Inc. All rights reserved.

Parthenogenesis and Parthenogenetic Stem Cells. General Terms Parthenogenesis Names and Acronyms hpSC Line Derivation Human Oocytes Parthenogenetic Activation of Unfertilized Oocytes hpSC Isolation HLA Homozygous and HLA Heterozygous hpSC Lines History of hpSC Line Creation Properties of Undifferentiated hpSC Morphology and Gene Expression Patterns Karyotype Homozygosity of Pericentric Region Pluripotency Differentiation Capacity of hpSC Are hpSC a More Risky Cell Source for Transplantation than Other Pluripotent Cell Types? Acknowledgments References

608 608 609 609 609 609 612 613 615 615 615 615 615 616 616 616 617 617

Glossary Aneuploidy Abnormal number of chromosomes (too many or too few). Aneuploidy can cause tumor development or birth defects. Blastocyst Stage of the early embryo that contains two cell types: (i) inner cell mass (ICM) which (if implanted) forms the embryo and (ii) trophoblast, the cellular source of placenta. The trophoblast surrounds the ICM in a fluid-filled blastocyst cavity. Diploid chromosome set Normal set of 46 chromosomes (human) made of two 23-chromosome sets. Haploid chromosome set Set of 23 chromosomes (human), half the normal chromosome set. HLA Human leukocyte antigens are expressed on almost every cell and immunologically distinguish one person from others. IVF (in vitro fertilization) Medical procedure performed for infertility, in which oocytes are fertilized in a lab (in vitro), grown into early embryos, and then transferred into the uterus. Karyotype The number and arrangement of chromosomes. Normal human karyotype is 46 chromosomes including sex chromosomes XY (male) or XX (female). ‘46,XX’ means ‘cell(s) contain 46 chromosomes including two X chromosomes’. ‘47,XX,þ6’ means ‘cell(s) contain 47 chromosomes including two X chromosomes and one extra (pathologic) number 6 chromosome.’ ‘45,X0’ means ‘cell(s) contains 45 chromosomes where one X chromosome is missing.’ Mosaic karyotype Tissue/organ with cells of different karyotypes. Pericentric chromosome region Chromosome region surrounding the centromere.

Parthenogenesis and Parthenogenetic Stem Cells. General Terms Parthenogenesis Parthenogenesis is a form of asexual reproduction in which an egg (oocyte) develops without fertilization (by spermatozoa) and therefore without male contribution into the embryo genome. Parthenogenetically activated oocytes pass through similar stages of embryonic development as do fertilized eggs: They form one or several pronuclei, undergo cleavage, form morulae, and generate blastocyst-like structures. Human parthenogenetic stem cells (hpSC) are pluripotent stem cells derived from a parthenogenetically activated oocyte.

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Names and Acronyms Because the field of hpSC is relatively new, a consensus name for hpSC has not emerged. In the literature the cells may be called parthenogenetic stem cells, parthenogenetic pluripotent stem cells, parthenogenetic embryonic stem cells, parthenote stem cells, p-cells, or parthenotes, with ‘human parthenogenetic stem cells’ the most common. To emphasize the equivalent pluripotentiality of hpSC with human embryonic stem cell (hESC) and induced pluripotent stem cells (iPSC), hpSC are also commonly called ‘human pluripotent parthenogenetic stem cells.’ Here we use ‘hpSC.’

hpSC Line Derivation Human Oocytes All reported hpSC lines were derived from the cells of p-blastocysts (from parthenogenetically activated oocytes). Unlike hESC, no blastomere- or morula-derived hpSC lines are known. Generation of hpSC lines takes place in three stages: (1) the process of obtaining of human oocytes, (2) parthenogenetic activation of the oocyte and in vitro cultivation to the p-blastocyst stage, and (3) derivation/isolation of hpSC lines from the blastocyst ICM. Obtaining human eggs, similar to donation for research on any human tissue or cells, requires review and approval of legal and ethical protocols and should be done only after approval of a local Institutional Review Board and, in some jurisdictions, an Embryonic Stem Cell Research Oversight (ESCRO) committee, respecting international standards for human subjects protection. The oocyte donor should be healthy and able to tolerate hormone stimulation and anesthesia. Generally, only volunteer oocyte donation without any financial compensation is approved by IRBs in the United States. Human oocytes are collected from preovulatory follicles by ultrasound-guided follicular puncture after controlled ovarian hyperstimulation or a natural menstrual cycle without any or minimal hormone stimulation. The natural cycle approach is not complicated by the massive stimulation of ovaries, but only one oocyte per cycle can be obtained. Ovarian hyperstimulation with hormones is more common for patients undergoing in vitro fertilization (IVF) and yields multiple oocytes per cycle. The daily high-dose hormone injections over long periods of time are often associated with severe side effects and so should only be done in the context of fertility treatments. In spite of a number of attempts to use fertilization-failed oocytes (oocytes that underwent in vitro addition of sperm but did not show signs of sperm penetration), or frozen oocytes to derive hpSC, no hpSC lines have emerged from these alternative egg sources. With the exception of the chHES-32 line (Table 1) from an oocyte that underwent in vitro maturation, all hpSC lines reported originated from fresh (not frozen) metaphase II-arrested (MII) oocytes.

Parthenogenetic Activation of Unfertilized Oocytes Oocyte activation can be induced by a variety of stimuli including electrical, chemical, and mechanical stimuli, and even spontaneous activation can happen. With the exception of two hpSC lines (P-TJ and chHES-32; Table 1) from spontaneously activated oocytes, all other reported hpSC lines originated from oocytes stimulated by chemical agents. (Other lines may have been derived in industry settings or not reported.) The chemicals used are chosen because they initiate transients in calcium concentration in the egg cytoplasm similar to sperm-induced repetitive Ca2þ oscillations (also called calcium waves) that persist for several hours after spermatozoon penetration. These calcium waves are critical for moving the egg out of meiotic arrest and into further development (Figures 1(a) and 2(a)). The most commonly used method of activation is serial treatment of oocytes with ionomycin (a calcium ionophore) and 6dimethylaminopurine (6-DMAP, a protein kinase inhibitor). Short (5 min) calcium ionophore incubation causes the first large calcium wave, and 3–5 h of 6-DMAP treatment stimulates slightly prolonged calcium oscillations (Figure 1(b)). 6-DMAP also blocks extrusion of the second polar body so that the activated oocytes retain all their genetic material, which ultimately contributes to a diploid chromosome set (Figure 2(b)). In 2008 three hpSC lines (hpSC-Hhom-2, hpSC-Hhom-3, and hpSC-Hhom-4; Table 1) were derived using ionomycin and puromycin. Similar to 6-DMAP, puromycin supported oocyte activation, but unlike 6-DMAP, puromycin has no effect on spindle integrity and the second polar body was extruded. The puromycin-activated oocyte therefore contained only half of a set of metaphase II chromosomes but they resulted in diploid stem cells with homozygous genotypes. These cells’ genome contained a duplicated set of half of the genes derived from the oocyte donor, including HLA genes (Figure 2(c)). The exact mechanism and timing of duplication of haploid genetic material following oocyte activation are unclear. More than likely, DNA replication occurs in the absence of cell cleavage or division. Prior studies suggest that 80% of parthenogenetically activated mouse oocytes preserve their haploid state until the morula stage, although stem cell lines derived from these embryos become diploid. Five other HLA homozygous hpSC lines were derived using yet other approaches. The chHES-32 line originates from a spontaneously activated oocyte that extruded its second polar body. hpSC-Hhom-1, pSC, and SCNT-hES-1 lines originate from oocytes activated by ionomycin and 6-DMAP, and FY-phES-018 is from an oocyte activated by ionomycin, 6-DMAP, and trichostatin A (TSA) (a histone deacetylase inhibitor) (Table 1).

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Human pluripotent stem cell lines published by July 2013

Oocyte hpSC line name stage

Origin

Method of activation

Genome status

HLA status

Karyotype

In vivo tumorigenicity

In vitro pluripotency

hpSC derivation method

References

SCNT-hES-1

MII

p-blastocyst

Ionomycin6-DMAP

Hetero

Homo

45,X0a

Teratoma

Yes

Immuno surgery

phESC-1 phESC-3

MII MII

p-blastocyst p-blastocyst

Ionomycin6-DMAP Ionomycin6-DMAP

Hetero Hetero

Hetero Hetero

46,XXc 46,XXd

Teratoma Teratoma

Yes Yes

phESC-4

MII

p-blastocyst

Ionomycin6-DMAP

Hetero

Hetero

46,XXe

Teratoma

Yes

phESC-5

MII

p-blastocyst

Ionomycin6-DMAP

Hetero

Hetero

46,XXf

Teratoma

Yes

phESC-6

MII

p-blastocyst

Ionomycin6-DMAP

Hetero

Hetero

46,XXg

Teratoma

Yes

phESC-7

MII

p-blastocyst

Ionomycin6-DMAP

Hetero

Hetero

Teratoma

Yes

hPES-1

MII

p-blastocyst

Hetero

ND

Teratoma

Yes

Immune surgery, MEF feeder

Mai et al. (2007)

hPES-2

MII

p-blastocyst

Hetero

ND

Mai et al. (2007)

P-blastocyst

Homok

Homol

Failed to form any teratoma Teratoma

Immune surgery, MEF feeder

ND

46,XX, translocationsj 46,XXm

Yes

chHES-32

ElectricalIonomycin6DMAP ElectricalIonomycin6DMAP Spontaneous

Mosaich: 47,XXX; 48,XXX,þ6 46,XXi

Immuno surgery, NSF feeder Whole p-blastocyst plating on NSF Whole p-blastocyst plating on NSF Whole p-blastocyst plating on NSF Whole p-blastocyst plating on NSF Whole p-blastocyst plating on NSF

Hwang et al. (2004) Kim et al. (2007),b Revazova et al. (2007) Revazova et al. (2007)

Yes

Lin et al. (2007)

hpSC-Hhom-1

MII

p-blastocyst

Ionomycin6-DMAP

Hetero

Homo

46,XX

Teratoma

Yes

hpSC-Hhom-2

MII

p-blastocyst

IonomycinPuromycin

Homo

Homo

Teratoma

Yes

hpSC-Hhom-3

MII

p-blastocyst

IonomycinPuromycin

Homo

Homo

Teratoma

Yes

Whole p-blastocyst plating on NSF

Revazova et al. (2008)

hpSC-Hhom-4

MII

p-blastocyst

IonomycinPuromycin

Homo

Homo

Mosaicn: 46,XX 47,XX,þ8 Mosaico: 46,XX 47,XX,þ1 46,XX

Whole p-blastocyst plating on HEF Whole p-blastocyst plating on NSF Whole p-blastocyst plating on NSF

Teratoma

Yes

Revazova et al.

HP1 HP3 P-TJ

MII MII MIIr

p-blastocyst p-blastocyst p-blastocyst

Ionomycin6-DMAP Ionomycin6-DMAP Spontaneous

ND ND ND

ND ND ND

NDp NDq 46,XX

Myofibrosarcoma Myofibrosarcoma ND

Yes Yes Yes

Whole p-blastocyst plating on NSF Microsurgical ICM removal Microsurgical ICM removal Immunesurgery, hFF feeder

Revazova et al. (2007) Revazova et al. (2007) Revazova et al. (2007) Revazova et al. (2007)

Revazova et al. (2008) Revazova et al. (2008)

Brevini et al. (2009) Brevini et al. (2009) Lu et al., 2010

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Table 1

pSC

MII

p-blastocyst

Ionomycin6-DMAP

ND

Homo

46,XX

Teratoma

Yes

FY-phES-018

MII

p-blastocyst

Ionomycin6DMAPTrichostatin A

Homos

Homo

46,XXt

ND

Yes

Whole p-blastocyst plating on HFF Immunesurgery, MEF feeder

Vassena et al., 2012 Liu et al., 2011

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a The original (when line was reported as somatic nuclear transfer originated instead of parthenogenetically originated) karyotype was reported as 46,XX (Hwang, W.S., et al., 2004. Evidence of a pluripotent human embryonic stem cell line derived from a cloned blastocyst. Science 303, 1669–1674.), but independent investigation (Kim, K., Lerou, P., Yabuuchi, A., Lengerke, C., Ng, K., West, J., Kirby, A., Daly, M.J., Daley, G.Q., 26 January 2007. Histocompatible embryonic stem cells by parthenogenesis. Science 315 (5811), 482–486.) demonstrated 45,X0 karyotype for this line. b Line was reported to be of somatic cell nuclear transfer origin (Hwang, W.S., et al., 2004. Evidence of a pluripotent human embryonic stem cell line derived from a cloned blastocyst. Science 303, 1669–1674.), but later after detection homozygosity in pericentric region by single nucleotide polymorphism (SNP) analysis was assumed as parthenogenetically originated (Kim, K., Lerou, P., Yabuuchi, A., Lengerke, C., Ng, K., West, J., Kirby, A., Daly, M.J., Daley, G.Q., 26 January 2007. Histocompatible embryonic stem cells by parthenogenesis. Science 315 (5811), 482–486.). c 12% of cells show X chromosome heteromorphism. d 80% of cells show X chromosome heteromorphism. e 86% of cells show X chromosome heteromorphism. f 42% of cells show X chromosome heteromorphism. g 12% of cells show X chromosome heteromorphism. h Aneuploid mosaic karyotype: 91% of cells have a 47,XXX karyotype and 9% of the cells have a 48,XXX,þ6 karyotype. 70% of cells show X chromosome heteromorphism. i Over 100 passages; karyotyping was performed every 20 passages from passage 20 to 120. j Chromosome translocations beyond passage 50. k The homozygous sites comprised >99% out of SNP sites tested (500 447). l The status of HLA-A, -B, and -DRB loci was determined. m As determined at passage 6 and 49. n Aneuploid mosaic karyotype: 85% of cells have a 46,XX karyotype and 15% of the cells have a 47,XX,þ8 karyotype. o Aneuploid mosaic karyotype: 95.8% of cells have a 46,XX karyotype and 4.2% of the cells have a 47,XX,þ1 karyotype. p At the later work (Brevini et al., 2012) line was described as aneuploid contained (from hypo-haploid to hypo-diploid configurations). q At the later work (Brevini et al., 2012) line was described as aneuploid (from hypo-haploid to hypo-diploid configurations). r In vitro maturation from MI oocyte. s Determined by short tandem repeat analysis. t FY-phES-018 line demonstrates unstable karyotype. Before passage 20: 46, XX. At passage 35, almost all the cells displayed a 45,XO karyotype. At passage 45: the mosaic ratio of 46,XX to 45,XO was 67:33. After passage 60, most cells displayed the 46,XX karyotype with a mosaic ratio of 97:3.

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Figure 1 Parthenogenetic activation simulates spermatozoon-caused calcium waves in oocyte. (a) The penetration of human spermatozoon causes calcium oscillations that are part of signal releasing oocyte from meiotic arrest and permitting embryonic development (oocyte activation). The graph represents time-dependent fluctuations of intracellular free calcium. Spermatozoon-activated human oocyte follows development program through release of second polar body and formation of zygote with male and female pronuclei, subsequent passing the cleavage from two to eight blastomeres, formation of morula and blastocyst. (b) Similar to fertilization, parthenogenetic activation causes calcium oscillations in oocyte (graph represents tendency and not actual oscillations): short-time ionomycin treatment simulates first large calcium oscillations and long-term 6-DMAP treatment supports prolonged calcium waves. Parthenogenetically activated human oocyte follows development program similar to fertilized oocyte. Human parthenogenetic stem cells can be isolated from p-blastocyst. P-zygote, p-morula, and p-blastocyst are the structures that are similar to human zygote, morula, and blastocyst (developed from fertilized oocyte) but originated from parthenogenetically activated human oocyte.

hpSC Isolation Under proper in vitro culture conditions (usually commercially available culture systems developed and used for IVF) parthenogenetically activated oocytes are able to follow the same developmental program as fertilized oocytes, through cleavage and morula stages, followed by development into p-blastocysts (Figures 3(a)–3(c)). The derivation of the stem cell line from the p-blastocyst is similar to derivation of hESC lines from the ICM. Two approaches to isolating pluripotent stem cells from blastocysts have been used. In the first, ICM is isolated by removing the outer layer of the blastocyst or through immuno- or micro- surgery to open the blastocyst for plating ICM cells on a feeder cell layer. Because the quality of p-blastocyts is often very poor and the ICM is not always easy to identify microscopically, most parthenogenetic lines are derived through a second approach in which the whole p-blastocyst is plated on feeder cells. The ICM cells grow out from the plated blastocyst and can be further mechanically isolated and transferred to fresh feeder cells.

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Figure 2 Formation of diploid karyotype in human parthenogenetic stem cells. (a) Normal fertilization leads to the diploid nuclear genome. Human spermatozoon brings male genes (23 chromosomes) to metaphase II-arrested oocyte (MII) that preserves a diploid chromosome set via the spindle. After spermatozoon penetration, the oocyte releases 23 chromosomes with the second polar body. The remaining 23 chromosomes (oocyte) form the female pronucleus, while 23 chromosomes contributed by the spermatozoon form the male pronucleus. The resultant oocyte diploid genome is retained throughout development (and preserved in derived pluripotent cell lines). (b) Formation of diploid heterozygous hpSC. In addition to supporting prolonged calcium oscillations in the parthenogenetically activated oocyte, 6-DMAP also furthers spindle degradation, preventing extrusion of the second polar body. The entire MII oocyte genome (46 chromosomes) resulting from parthenogenetic activation remains in the activated oocyte and through development of the p-blastocyst (as well as in derived hpSC). (c) Formation of diploid homozygous hpSC. The example here results from ionomycin and puromycin activation but presumably the mechanism leading to diploid stem cells from spontaneously activated oocytes is similar. Puromycin treatment does not affect spindle integrity; therefore parthenogenetically activated oocytes release 23 chromosomes with the second polar body and retain a haploid (23) chromosome set. The timing of duplication of these 23 chromosomes is not known. In theory, duplication can occur at any stage (activated oocyte, zygote, during cleavage or during morula or blastocyst formation, or even during stem cell line isolation) after spontaneous activation of oocytes. (Cleavage and morula stages of parthenote development are not shown.)

HLA Homozygous and HLA Heterozygous hpSC Lines Transplantation of allogeneic (from another human) organs or cells initiates an immune response in the host, in which the host immune system recognizes HLA antigens of the donor. The risk of transplant rejection is generally proportional to the degree of cell surface antigen disparity between the donor and recipient cells. Matching donor and recipient tissue for HLA antigens reduces the chance of a cytotoxic T-cell response in the recipient and thus increases the likelihood of transplant survival. Since a normal HLA repertoire has some maternal and paternal antigens, successful immune matching of donor to host requires matching both maternal and paternal antigens, it is easy to see that the complexity of immune matching is greatly reduced if the donor cells are HLA homozygous. Using various activation approaches, it is possible to create hpSC that are HLA heterozygous or HLA homozygous. With HLA homozygous hpSC the immune matching to a potential recipient is easier because there is simply less diversity in the HLA

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Figure 3 Morphology of an MII oocyte, embryos derived from parthenogenetically activated oocytes, and hpSC. (a) Metaphase II-arrested (MII) human oocyte with clearly visible first polar body. To date all hpSC lines have been derived from MII oocytes. (Figure source: Thomas Elliott, http:// www.ivf.net/ivf/stripped-oocyte-o682.html, reproduced with permission.) (b) Cleavage embryo (eight blastomeres) from a parthenogenetically activated human oocyte (source of P-TJ hpSC). (Figure source: Lu, Z., Zhu, W., Yu, Y., Jin, D., Guan, Y., Yao, R., Zhang, Y.A., Zhang, Y., Zhou, Q., June 2010. Derivation and long-term culture of human parthenogenetic embryonic stem cells using human foreskin feeders. J. Assist. Reprod. Genet. 27 (6), 285–291, reproduced with permission.) (c) Blastocyst from a parthenogenetically activated human oocyte. The expanded blastocyst (left) and hatching blastocyst with a clearly visible ICM extruding out of the zona pellucida (right). (Left figure source: Mai, Q., Yu, Y., Li, T., et al., 2007. Derivation of human embryonic stem cell lines from parthenogenetic blastocysts. Cell Res. 17, 1008–1019, right figure source: Lu, Z., Zhu, W., Yu, Y., Jin, D., Guan, Y., Yao, R., Zhang, Y.A., Zhang, Y., Zhou, Q., June 2010. Derivation and long-term culture of human parthenogenetic embryonic stem cells using human foreskin feeders. J. Assist. Reprod. Genet. 27 (6), 285–291, reproduced with permission.) (d) hpSC morphology: Undifferentiated chHES-32 colonies growing on an inactivated human feeder layer. (Inset is higher magnification of one colony). (Figure source: Lin, G., OuYang, Q., Zhou, X., et al., 2007. A highly homozygous and parthenogenetic human embryonic stem cell line derived from a one-pronuclear oocyte following in vitro fertilization procedure. Cell Res. 17, 999–1007, reproduced with permission.) (e) hpSC express normal human pluripotent stem cell markers. OCT4 expression in the hpSC-Hhom-1 line. (Figure source: Revazova, E.S., Turovets, N.A., Kochetkova, O.D., Agapova, L.S., Sebastian, J.L., Pryzhkova, M.V., Smolnikova, V.I., Kuzmichev, L.N., Janus, J.D., March 2008. HLA homozygous stem cell lines derived from human parthenogenetic blastocysts. Cloning Stem Cells 10 (1), 11–24, reproduced with permission. SSEA-3, SSEA-4, TRA-1-60, and TRA-1-81 expression in the P-TJ line. Figures source: Lu, Z., Zhu, W., Yu, Y., Jin, D., Guan, Y., Yao, R., Zhang, Y.A., Zhang, Y., Zhou, Q., June 2010. Derivation and long-term culture of human parthenogenetic embryonic stem cells using human foreskin feeders. J. Assist. Reprod. Genet. 27 (6), 285–291, reproduced with permission.) (f) Normal 46,XX karyotype of hpSC-Hhom-4. (Figure source: Revazova, E.S., Turovets, N.A., Kochetkova, O.D., Agapova, L.S., Sebastian, J.L., Pryzhkova, M.V., Smolnikova, V.I., Kuzmichev, L.N., Janus, J.D., March 2008. HLA homozygous stem cell lines derived from human parthenogenetic blastocysts. Cloning Stem Cells 10 (1), 11–24, reproduced with permission.) (g) Abnormal aneuploid karyotype of hpSC-Hhom-3. This hpSC line has a mosaic karyotype: 95.8% of cells are 46,XX (left) and 4.2% are 47,XX, þ1 (right) with an extra chromosome 1. Figure source: Revazova, E.S., Turovets, N.A., Kochetkova, O.D., Agapova, L.S., Sebastian, J.L., Pryzhkova, M.V., Smolnikova, V.I., Kuzmichev, L.N., Janus, J.D., March 2008. HLA homozygous stem cell lines derived from human parthenogenetic blastocysts. Cloning Stem Cells 10 (1), 11–24, reproduced with permission.

genotype of HLA homozygous hpSC. In theory, this means that HLA homozygous hpSC could be generated to match large segments of the population with significantly fewer lines required for allogeneic matching than would be possible using HLAcharacterized hES cells or iPS cells. (Of course, iPS cells have the potential advantage of autologous use without inducing an immune response.)

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History of hpSC Line Creation The first intentional creation of hpSC lines was described by Revazova et al. in 2007: Six pluripotent hpSC lines were derived from chemically activated human oocytes. Additional lines were derived by this group and by others with 19 lines published to date.

Properties of Undifferentiated hpSC Morphology and Gene Expression Patterns Undifferentiated hpSC are morphologically similar to other types of human pluripotent stem cells such as hESC and hiPSC. Specifically in culture, undifferentiated hpSC form colonies of tightly packed cells. Individual cells of hpSC and other undifferentiated pluripotent cells have characteristic features including prominent nucleoli and a small cytoplasm to nucleus ratio (Figure 3(d)). hpSC demonstrate high levels of alkaline phosphatase activity, a marker of pluripotent stem cells, and express other ‘traditional’ pluripotent stem cell markers including OCT4 (also known as POU5F1), SSEA-3, SSEA-4, TRA-1-60, and TRA-1-81 (Figure 3(e)). In addition, hpSC demonstrate high levels of telomerase activity associated with indefinite proliferation potential, unlike adult stem cells that invariably undergo proliferative senescence in culture. Similar to hESC and hiPSC, hpSC can be maintained in undifferentiated state for a long period of time using growth factors and/or feeder cell layers. hpSC demonstrate similar global gene expression patterns to hESC and hiPSC and DNA methylation patterns, with exception of imprinted genes. Genomic imprinting is the mechanism by which monoallelic gene expression of certain genes (imprinted genes) is achieved in a parent-of-origin-specific manner based on DNA methylation (an epigenetic change, not a change in sequence). Because hpSC originate from oocytes without contribution of male genetic material, hpSC express only maternal imprinted genes. It is important to note that methylation patterns are dynamic and can be altered by in vitro culture conditions.

Karyotype All hpSC lines are diploid and XX (Figure 3(f); Table 1) except one line that had loss of one X chromosome. Some hpSC lines demonstrate aneuploidy including aneuploid mosaic variants (Figure 3(g)). Different degrees of X chromosome heteromorphism are reported for some hpSC lines. Chromosome instability during in vitro cultivation has also been reported. The absence of paternal centriole in the parthenogenetic zygote (normally inherited through the sperm) leads to a centrosome amplification process and may contribute to chromosome instability and aneuploidy.

Homozygosity of Pericentric Region The distinguishing property of hpSC is homozygosity of chromosome pericentric regions adjacent to centromeres. For hpSC lines derived from oocytes that did not release the second polar body after activation, homozygosity of chromosome pericentric regions reflects the failure of independent segregation of the sister chromatids during meiosis and is consistent with meiotic recombination events (Figure 4). Pericentric homozygosity is uniquely found in hpSC lines and is not observed in either hESC or iPSC. The analysis of single nucleotide polymorphisms in genome regions adjacent to the centromere can reliably distinguish hpSC from other cell types.

Figure 4 Formation of the homozygous pericentric region in parthenogenetically activated oocytes. Germinal vesicle (GV) oocyte contains duplicated diploid chromosome set (4n ¼ 2  46); each homologous chromosome contains two sister chromatids joined at the centromere (homozygous sister chromatid). During MI, recombination occurs between both sister chromatid pairs of homologous chromosomes, yielding heterozygosity of the distal ends of sister chromatids, but recombination does not occur at the centromere region which remains homozygous. During formation of MII oocytes (first meiotic division) one of the recombined chromosomes (from a pair of homologous chromosomes) is segregated into the first polar body, resulting in a diploid oocyte (second recombined chromosome from the other homologous chromosome pair) represented as 23 pairs of chromatids still joined at the centromere. During a normal MII in a fertilized egg, the second polar body would be extruded. However, after parthenogenetic activation with ionomycin þ 6-DMAP, the second polar body is not extruded, leaving an oocyte diploid chromosome. After chromatid disjunction each becomes a separate chromosome.

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Pluripotency hpSC lines demonstrate ability to give rise to derivatives of three germ layers (endoderm, ectoderm, and mesoderm) in embryoid body formation assays. Most hpSC are capable of forming teratomas approximately 2–3 months after injection into immunocompromised rodents. Teratomas are benign tumors that contain cell types from all three germ layers. Teratoma formation is the current gold standard confirmatory test of pluripotency. Some hpSC lines, however, fail to form teratomas and other aberrant differentiation patterns have also been observed.

Differentiation Capacity of hpSC Differentiation of hpSC into definitive endoderm (DE) and early hepatocytes has been accomplished in four hpSC lines (Table 2). hpSC are able to respond to well-characterized developmental signals which direct differentiation of pluripotent cells into DE, the precursor for number of cell types including hepatocytes. hpSC (like hESC) exposed to the proper developmental signals demonstrate a similar temporal sequence of gene expression to that which occurs in the course of in vivo DE differentiation during vertebrate gastrulation. However, in the few lines studied, the yield of DE and hepatocytes from hpSC lines is lower than from hESC. Treatment of undifferentiated hpSC by TSA, a potent histone deacetylase inhibitor, before this directed differentiation significantly increases the efficiency of DE differentiation from hpSC, but significant numbers of hpSC do not respond to the differentiation cues and remain undifferentiated. In anticipating clinical manufacture of cells for therapeutic applications, it is undesirable to have undifferentiated cells in the cell graft because of their potential to form teratomas. Differentiation into neural progenitor cells was reported for two hpSC lines: phESC-3 and phESC-5 (Table 2). Again yield was lower in comparison with hESC lines, and differentiation was less efficient. hpSC differentiation into retinal pigment epithelium (RPE) is quite similar to hESC differentiation of RPE (Table 2). Differentiation toward hematopoietic fate was reported by just one group for two hpSC lines (Table 2) with generation of CD34/CD45positive cells that were able to form colonies in methylcellulose after 3 weeks. Lymphoid, erythroid, and myeloid subpopulations were generated. Under appropriate conditions, hpSC lines are also able to generate mesenchymal stem cells (MSC) (Table 2). In addition, MSC derived from hpSC show the usual differentiation potential of MSC, that is, in vitro differentiation of chondrocyte (cartilage), osteocyte (bone), and adipocyte (fat) lineages.

Are hpSC a More Risky Cell Source for Transplantation than Other Pluripotent Cell Types? The answer to this question is not known because the definitive experiments have not yet been performed. Nonetheless, the absence of sperm centriole in parthenotes is a potential risk factor for safe clinical application of hpSC-derived cell therapies. Abnormal number of centrioles as well as aberrant levels of molecules related to mitotic spindle formation and spindle assembly checkpoint have been reported in some hpSC lines and may also contribute to the high incidence of aneuploidy observed in hpSC lines.

Table 2

Differentiation capacity of hpSC

hpSC line name

Specific cell type/direction of differentiation

References

phESC-1

Definitive endoderm Fetal-like hepatocytes Retinal pigment epithelium Definitive endoderm Fetal-like hepatocytes Neural progenitor cells Retinal pigment epithelium Definitive endoderm Fetal-like hepatocytes Neural progenitor cells Retinal pigment epithelium Mesenchymal stem cells Definitive endoderm Fetal-like hepatocytes Hematopoietic fate Hematopoietic fate Retinal pigment epithelium Mesenchymal stem cells

Turovets et al. (2011a) Turovets et al. (2012) Harness et al. (2011) Turovets et al. (2011a) Turovets et al. (2012) Harness et al. (2011) Harness et al. (2011) Turovets et al. (2011a) Turovets et al. (2012) Harness et al. (2011) Harness et al. (2011) Chen et al. (2012) Turovets et al. (2011a) Turovets et al. (2012) Brevini et al. (2009) Brevini et al. (2009) Li et al. (2012) Vassena et al. (2012)

phESC-3

phESC-5

chHES-32 hpSC-Hhom-4 HP1 HP3 P-TJ pSC

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Concerns that these abnormalities will lead to malignant transformation are a major hurdle for hpSC-based therapies. For some applications, though, hpSC may find a place in regenerative medicine if these concerns can be addressed in rigorous preclinical studies.

Acknowledgments We thank Irina Turovets for the translation of our ideas and thoughts into graphics and for the creation illustrations and schemes for this article.

References Brevini, T. A., Pennarossa, G., Maffei, S., Tettamanti, G., Vanelli, A., Isaac, S., Eden, A., Ledda, S., de Eguileor, M., & Gandolfi, F. (2012). Centrosome amplification and chromosomal instability in human and animal parthenogenetic cell lines. Stem Cell Rev., 8(4), 1076–1087. Brevini, T. A., Pennarossa, G., Antonini, S., Paffoni, A., Tettamanti, G., Montemurro, T., Radaelli, E., Lazzari, L., Rebulla, P., Scanziani, E., de Eguileor, M., Benvenisty, N., Ragni, G., & Gandolfi, F. (December 2009). Cell lines derived from human parthenogenetic embryos can display aberrant centriole distribution and altered expression levels of mitotic spindle check-point transcripts. Stem Cell Rev., 5(4), 340–352. Chen, Y., Ai, A., Tang, Z. Y., Zhou, G. D., Liu, W., Cao, Y., & Zhang, W. J. (January 2012). Mesenchymal-like stem cells derived from human parthenogenetic embryonic stem cells. Stem Cells Dev., 21(1), 143–151. Fried, E. P., Ross, P., Zang, G., Divita, A., Cunniff, K., Denaday, F., Salamone, D., Kiessling, A., & Cibelli, J. (April 2008). Human parthenogenetic blastocysts derived from noninseminated cryopreserved human oocytes. Fertil. Steril., 89(4), 943–947. Harness, J. V., Turovets, N. A., Seiler, M. J., Nistor, G., Altun, G., Agapova, L. S., Ferguson, D., Laurent, L. C., Loring, J. F., & Keirstead, H. S. (7 January 2011). Equivalence of conventionally-derived and parthenote-derived human embryonic stem cells. PLoS One, 6(1), e14499. Hwang, W. S., et al. (2004). Evidence of a pluripotent human embryonic stem cell line derived from a cloned blastocyst. Science, 303, 1669–1674. Kim, K., Lerou, P., Yabuuchi, A., Lengerke, C., Ng, K., West, J., Kirby, A., Daly, M. J., & Daley, G. Q. (26 January 2007). Histocompatible embryonic stem cells by parthenogenesis. Science, 315(5811), 482–486. Li, W. B., Zhang, Y. S., Lu, Z. Y., Dong, L. J., Wang, F. E., Dong, R., & Li, X. R. (August 2012). Development of retinal pigment epithelium from human parthenogenetic embryonic stem cells and microRNA signature. Invest Ophthalmol Vis. Sci., 53(9), 5334–5343. Lin, G., OuYang, Q., Zhou, X., et al. (2007). A highly homozygous and parthenogenetic human embryonic stem cell line derived from a one-pronuclear oocyte following in vitro fertilization procedure. Cell Res., 17, 999–1007. Lu, Z., Zhu, W., Yu, Y., Jin, D., Guan, Y., Yao, R., Zhang, Y. A., Zhang, Y., & Zhou, Q. (June 2010). Derivation and long-term culture of human parthenogenetic embryonic stem cells using human foreskin feeders. J. Assist. Reprod. Genet., 27(6), 285–291. Mai, Q., Yu, Y., Li, T., et al. (2007). Derivation of human embryonic stem cell lines from parthenogenetic blastocysts. Cell Res., 17, 1008–1019. Nazor, K. L., Altun, G., Lynch, C., Tran, H., Harness, J. V., Slavin, I., Garitaonandia, I., Müller, F. J., Wang, Y. C., Boscolo, F. S., Fakunle, E., Dumevska, B., Lee, S., Park, H. S., Olee, T., D’Lima, D. D., Semechkin, R., Parast, M. M., Galat, V., Laslett, A. L., Schmidt, U., Keirstead, H. S., Loring, J. F., & Laurent, L. C. (May 4, 2012). Recurrent variations in DNA methylation in human pluripotent stem cells and their differentiated derivatives. Cell Stem Cell, 10(5), 620–634. Revazova, E. S., Turovets, N. A., Kochetkova, O. D., Agapova, L. S., Sebastian, J. L., Pryzhkova, M. V., Smolnikova, V. I., Kuzmichev, L. N., & Janus, J. D. (March 2008). HLA homozygous stem cell lines derived from human parthenogenetic blastocysts. Cloning Stem Cells, 10(1), 11–24. Revazova, E. S., Turovets, N. A., Kochetkova, O. D., Kindarova, L. B., Kuzmichev, L. N., Janus, J. D., & Pryzhkova, M. V. (2007). Patient-specific stem cell lines derived from human parthenogenetic blastocysts. Cloning Stem Cells, 9(3), 432–449. Fall. Turovets, N., D’Amour, K. A., Agapov, V., Turovets, I., Kochetkova, O., Janus, J., Semechkin, A., Moorman, M. A., & Agapova, L. (June 2011a). Human parthenogenetic stem cells produce enriched populations of definitive endoderm cells after trichostatin A pretreatment. Differentiation, 81(5), 292–298. Turovets, N., Semechkin, A., Kuzmichev, L., Janus, J., Agapova, L., & Revazova, E. (2011b). Derivation of human parthenogenetic stem cell lines. Methods Mol. Biol., 767, 37–54. Turovets, N., Fair, J., West, R., Ostrowska, A., Semechkin, R., Janus, J., Cui, L., Agapov, V., Turovets, I., Semechkin, A., Csete, M., & Agapova, L. (2012). Derivation of high-purity definitive endoderm from human parthenogenetic stem cells using an in vitro analog of the primitive streak. Cell Transplant., 21(1), 217–234. Vassena, R., Montserrat, N., Carrasco Canal, B., Aran, B., de Oñate, L., Veiga, A., & Izpisua Belmonte, J. C. (1 August 2012August). Accumulation of instability in serial differentiation and reprogramming of parthenogenetic human cells. Hum. Mol. Genet., 21(15), 3366–3373.

Human Pluripotent Stem Cells P Rajan, The Scripps Research Institute, La Jolla, CA, USA © 2019 Elsevier Inc. All rights reserved.

Introduction Definition of hESCs Unique Properties of hESCs Methods of Generation of hESCs ‘Traditional’ Derivations Somatic Cell Nuclear Transfer Cell Fusion Induced Pluripotent Stem Cells Current Applications of hESCs Legal and Social Aspects of hESC Research Future Perspectives References Relevant Websites

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Glossary Adherent culture Cell cultures which are adherent to the surface of the culture dish. ADME Absorption, distribution, metabolism, excretion. These are characteristics of therapeutic compounds, which need to be defined before clinical application. Allele One of a pair of a specific gene located on sister chromatids. Companion diagnostics Clinical tests which are performed in conjunction with the administration of a therapeutic. Diploid A cell which has the normal cohort of two sister chromatids per chromosome. Germline transmission Transmission of genes to the next generation by presence in sperm and egg. In vitro Experiment performed in a representative experimental platform, which is outside the body of the animal model under study. In vivo Experiment performed within the body of the experimental animal model. Passaging Technique of expanding a culture by removal from original culture dish and replating in new culture dishes. Pharmacogenomics Science that combines information about characteristics of genome of a particular individual with the effect of a pharmaceutical drug. Suspension culture Cell cultures that proliferate while floating in the culture medium. Telomere Ends of chromosomes, which have specific sequence repeats. Teratology The study of mutations and defects in developing embryos. Tetraploid A cell having four sister chromatids per chromosome – twice the normal number. Untransformed Cells that have not undergone a transformation event, which causes uncontrolled proliferation and abnormal maintenance of differentiation phenotype.

Introduction Stem cells are ‘normal,’ untransformed cells which have dual intrinsic properties of proliferation and differentiation. While empirical evidence for the existence of these cells has existed for several decades especially with relation to the study of embryology of invertebrates and lower vertebrates, they have recaptured the imagination of biologists recently due to their enormous potential in human disease therapies, and the establishment of sophisticated techniques by which they may be studied. Broadly, stem cells may be classified as somatic stem cells (SSCs) and embryonic stem cells (ESCs). SSCs are derived from developing and mature tissues in which they are resident and multipotent as they are restricted in their differentiation potential, and can only give rise to some or all of the cells present in their tissue of origin. ESCs are pluripotent as they are capable of differentiation into any cell type of the body, and may be derived from the inner cell mass (ICM) of the blastocyst of a fertilized embryo. The potential of stem cells may be realized for human therapeutics in two ways: firstly, stem cells may be used as a therapeutic themselves and used for regenerative medicine endpoints where derivatives of stem cells functionally substitute for deficits caused by the disease. Secondly, stem cells may be useful in promoting endogenous regenerative processes. Thirdly, stem cells may be used

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to create laboratory platforms, which are used to model diseases and functional tissues. These models serve as experimental platforms to study disease mechanisms, and identify molecules, which reverse disease phenotypes. This article will describe the properties of human embryonic stem cells (hESCs), methods for their derivation and their properties, and the manner in which their enormous potential may be realized for human benefit.

Definition of hESCs Human embryonic stem cells are identified and characterized based on combinations of cellular markers, and the functional criteria of self-renewal and pluripotency. They are created in vitro by manipulations in culture, and most closely resemble epiblast cells in early mouse development which give rise to the three germ layers of ectoderm, mesoderm, and definitive endoderm, rather than the more primitive mouse ESCs which differentiate into the extraembryonic endoderm in addition to the three germ layers mentioned above (Tesar et al., 2007). The cellular markers which are used to define hESCs are Oct4, Nanog, Sox2, SSEA4, Tra1-60, Tra1-81, alkaline phosphatase, and high telomerase activity (Figure 1). Functionally, hESCs are defined by their property of pluripotency, which may be demonstrated in vitro by directed differentiation of the cultures to defined fates representative of the three germ layer lineages, and in vivo by the formation of teratomas. For instance, in vitro differentiation of the culture into neuronal (ectodermal), blood and muscle (mesodermal), and liver and pancreas (endodermal) would have to be demonstrated by the same hESC line. Teratomas are tumors which comprise cells of ectodermal, mesodermal, and endodermal lineages, and serve as a surrogate assay of in vivo differentiation potential, as germline transmission and formation of entire embryos cannot be demonstrated in humans as it has been shown in mice and other animal models due to ethical reasons.

Unique Properties of hESCs There is an enormous effort being made to study the mechanisms by which stem hESCs maintain their dual properties of selfrenewal (proliferation) and pluripotency (differentiation). One hypothesis considers that the pluripotency of a cell arises due to the fact that it is capable of responding to diverse stimuli hence allowing it to differentiate along several lineages. The enhanced responsiveness of an hESC could be due to the expression of a greater cohort of proteins, which gets culled down to a defined set of proteins as commitment to a particular lineage occurs, thus restricting responsiveness. In fact, ESCs have the potential to activate the majority of the gene expression programs present in embryonic and adult cell lineages. The regulation of the cohort of genes expressed within the hESC is brought about by a combination of genetic and epigenetic characteristics, and is realized by inter- and

Figure 1 Colony of hESC (a). Phase contrast micrograph of an hESC colony. Undifferentiated colony stained with SSEA4 (b). Nuclei are labeled in blue (DAPI), and the cell surface SSEA antigen in red. Undifferentiated colony stained with Nanog (c). Nuclei are labeled in blue (DAPI), and the cytoplasmic antigen Nanog in red. MAP2 staining of neuronal rosettes in hESC colony differentiated into neuronal precursors (d). SSEA, stage-specific embryonic antigen; DAPI, diamidino-2-phenylindole, MAP2, microtubule-Associated Protein 2.

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intracellular signaling mechanisms. Interestingly, with the advent of the technology of generating induced pluripotent stem cells (iPSCs), it appears that the limited cohort of proteins expressed in mature cells may be expanded to give rise to cells, which have properties of pluripotency similar to hESCs. Pluripotency in hESCs is maintained by cellular programs activated by the growth factors such as fibroblast growth factor 2 (FGF2) and Activin/Nodal, along with some involvement of the Wnt (wingless-related integration) proteins. The creation of methods to regulate and control pluripotency in hESCs would be advantageous for the maintenance of these lines in culture in an undifferentiated pluripotent state until they are required to differentiate upon stimulation. A pharmacological inhibitor of GSK3b (glycogen synthase kinase), a kinase central to the Wnt signaling pathway, 6-bromoindirubin-30 -oxime is an effector of human ESC self-renewal, and induces the expression of pluripotency markers Oct3/4, Rex1, and Nanog. Epigenetic modifications, defined as heritable changes of the genome which are not coded in the sequence of the DNA, also contribute to the maintenance of pluripotency. At a gross level, differences have been found in the regulation of acetylation and methylation of DNA, histones and transcription factors in ESCs when compared to differentiated cells. DNA methylation of promoter sequences was detected in hESCs, which may account for regulating about 30% of genes in these cells. A related area of study is the mechanisms, and the creation of improved methods, by which stem cells differentiate into cells of interest. This is particularly relevant to regenerative medicine and cell-replacement therapies where hESC derivatives may functionally replace lost cells or tissues. For instance, dopaminergic neurons are lost in the basal ganglia structures of the brains of Parkinson’s disease patients, and insulin-secreting pancreatic b-cells are lost in diabetes type 1 patients. The replacement of these particular cell types derived from hESCs in these patients may serve as an effective therapy. Similar differentiation techniques are also useful for modeling of several diseases in the laboratory, identification and validation of novel targets of disease, and screening for new entities, which ameliorate disease phenotypes.

Methods of Generation of hESCs ESCs of various species including frogs, rodents, livestock, and monkeys have been created and are in use for over 40 years, for cloning and research purposes. However the first human ESC lines were created by culturing human blastocysts in 1998, almost simultaneously with the report of the first cultures of embryonic germ cells (hEGC; Thomson et al., 1998; Shamblott et al., 1998). hEGCs are derived from the gonadal ridge and mesenchyma of 5- to 9-week fetal tissue, and differentiate into the three germ layers in vitro, but unlike hESCs do not form teratomas in vivo. hESCs or hESC-like cells may be developed using four strategies. While the first method described below involves selective culture of hESC cells from a cultured blastocyst, the latter three methods involve the process of nuclear reprogramming. During the process of nuclear reprogramming, a nucleus from a mature cell is reprogrammed to resemble a more stemlike state, presumably by factors in the cytoplasm of the cell/egg into which the nucleus is transplanted, or by the transient overexpression of reprogramming factors using specialized gene expression methods and vectors (Figure 2).

‘Traditional’ Derivations Human blastocysts are obtained from discarded fertilized eggs from in vitro fertilization clinics. Fertilized eggs are cultured to a blastocyst over 5 days. The ICM is released from the blastocyst either by mechanical dissection with a microscalpel or by hatching the blastocyst using acid Tyrode solution (pH 2.2) and immunodissection using alkaline phosphatase and complement-mediated interactions. The ICM is cultured with or without feeder cultures, and colonies of hESCs form at low frequency, which develop into

Figure 2 Schematic of hESC derivation methods. Traditional derivations involve developing a blastocyst from fertilized eggs, and culturing an hESC line from the inner cell mass (ICM). SCNT also involves the establishment of a blastocyst, but it occurs from introducing a nucleus from a donor cell into the egg, and then stimulating it to develop a blastocyst. iPSC allows for development of a cell line without involvement of a blastocyst, but directly from reprogrammed cells (fibroblasts in this case).

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a stable line over several weeks. The hESC lines have been traditionally derived on mouse embryonic fibroblasts (MEFs), which secrete requisite supporting factors and function as a feeder layer. However, several lines have now been derived on human fibroblast feeder layers, or in feeder-free conditions. The latter is preferred for cell lines that have been created with the intention of use in clinical applications, as mouse glycoproteins have been detected on hESCs grown on MEFs. hESC proliferate in colonies, and are usually passaged mechanically or using mild enzymatic treatment with Accutase. hESC lines that have been passaged using harsher enzymes like trypsin appear to progressively acquire mutations. hESC colonies typically are comprised of small cells with high nuclear: cytoplasm ratios, and are positive for the markers mentioned above. They also need to have a stable karyotype in order to give rise to a successful and usable line (Rajan et al., 2007). The hESC lines have also been obtained from human eggs, which have been activated parthenogenetically to give rise to blastocysts. Although these lines are diploid (2n), since they have not arisen from the fusion of a sperm with an egg, they have some homozygosity at HLA loci, which may be advantageous for immunologic matching in the context of regenerative therapies.

Somatic Cell Nuclear Transfer This procedure was used successfully to clone the first sheep Dolly. In this procedure a diploid nucleus from a parent cell is reprogrammed to a stemlike state by introduction into an enucleated egg. After appropriate stimulation, which causes the egg to divide, formation of a blastocyst (with cells containing the donor nucleus) can be cultured as described for traditional hESC derivations. Somatic cell nuclear transfer (SCNT) requires the donation of eggs from healthy human volunteers, and is the platform necessary for “therapeutic cloning”. Therapeutic cloning involves the creation of cells or tissues (but not entire organisms) with known nuclear chromosome input, for research purposes.

Cell Fusion This procedure results in cells, which are tetraploid (4n), and hence derivatives may be used only as research tools. In this procedure a mature cell such as a fibroblast of the desired genotype is chemically fused with an established hESC cell. Reprogramming factors that are presumably in the cytoplasm of the hESC cause the progeny to have hESC-like characteristics. Although 4n, these cells exhibit the pluripotent properties of hESCs and resemble hESCs in gene expression and methylation profiles.

Induced Pluripotent Stem Cells This remarkable procedure which garnered the Nobel Prize for Physiology and Medicine in 2012 (awarded to Dr Shinya Yamanaka) involves the overexpression of four to six proteins which include at least four transcription factors that are usually expressed in hESCs, and leads to reprogramming of cells in cultures derived from mature differentiated tissues to culture with characteristics of hESCs. Oct4, Klf4, Sox2, and Myc are the genes used for induction of pluripotency, in the absence or presence of other genes including SV40 Large-T and Nanog. The exact nature of the ‘mature’ cell is not clear, and iPSC reprogramming appears to be more efficient with stemlike cells are used as the source of iPSC. While the original iPSC experiments were performed with the use of retroviral and lentiviral transduction to overexpress proteins, several refinements have since been reported involving the use of small molecules, microRNAs, protein transductions, plasmid transfections, transposons, adenoviral vectors, and more. These refinements make the procedure more accessible to laboratories around the world. The genetic/epigenetic regulation of reprogramming events is complex and takes place in a series of steps that are still being elucidated. Thus, although the use of these cells in regenerative medicine therapies requires more research, the power of this technology is already being realized for in vitro disease modeling and drug discovery applications. iPSC technology has already been used successfully to model muscular dystrophies, neurological diseases including amyotrophic lateral sclerosis, Huntington’s disease, and Alzheimer’s disease, psychiatric diseases including autism, and metabolic diseases including juvenile diabetes mellitus, Lesch–Nyhan syndrome, and Gaucher’s disease.

Current Applications of hESCs Several hESC lines exist around the globe, and while the majority of them have been created using good laboratory practices procedures, a few have been created under current good manufacturing practices (cGMP). While the former lines are useful in research and development projects, it is only the latter that may be developed for use in regenerative applications. The list of approved cell lines which may be used with US Federal funding is present on the NIH website (see website list), and other similar lists of lines approved by other countries including the UK is also included in the website list. Cell lines specific to other countries are also in the process of being created, notably in India, Brazil, China, and Japan. It is essential that genotypes of most countries are eventually represented in the global library of hESC and iPSC lines, as this will be extremely relevant to present drug discovery efforts, and future regenerative medicine technologies. Currently hESCs (including those derived using iPSC technology) have found excellent use in drug discovery applications, where their in vitro differentiation potential can be appropriately exploited. hESCs are developed into relevant in vitro platforms for validating targets, creating screening platforms, and developing toxicology platforms where they may be differentiated into liver and cardiac tissues. These toxicology platforms are invaluable in creating physiologically relevant laboratory paradigms of these tissues,

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and may eventually be developed into companion diagnostics using pharmacogenomic technologies. This would enable the screening of compounds, which may be particularly toxic or effective in selected populations or groups of patients, based on the genotypes of the hESC platforms created. While the inherent property of pluripotency of hESCs makes them an ideal candidate for regenerative medicine and cellreplacement therapies, the promise of this potential remains to be realized. Due to the possibility of cancer arising from the transplanted hESCs, only partially differentiated hESCs are used as a therapeutic. Geron, USA, spearheaded the effort of using hESC therapy for spinal cord injury with their FDA approved Phase I clinical trial. This program has since been discontinued by the company. Currently, Advanced Cell Technology, Inc. is determining the safety and efficacy of hESCs in macular degeneration in an FDA approved trial, while Viacyte is at the stage of preclinical testing in the development of an hESC-derived therapeutic for diabetes.

Legal and Social Aspects of hESC Research Since the creation of hESC lines involves the use of a fertilized human egg, and hESCs possess the enormous potential to create tissues, organs, or even clone a human, there are considerable social and religious overtones to this field of research. Legislation in the USA had restricted the use of federal funds for the creation of new hESC lines, although this has recently been relaxed. Other countries are more permissive to this research, but regulate the field tightly. The patent landscape was also tightly controlled by the published inventors of the technology, Wisconsin Alumni Research Foundation, and control of access of cell lines by creators of other cell lines have had a somewhat restrictive influence on progress in the field. These situations are also partially due to the fact that appreciable expense in time and effort is expended to create these lines.

Future Perspectives In order for the enormous potential of hESCs to be realized in both regenerative medicine and drug discovery, several aspects of hESC need to be clarified. These include determination of the factors controlling stability of genotype and phenotype, predictability of differentiation phenotype and efficiency of differentiation upon providing defined stimuli, control of length of telomeres and ‘aging’ of the cell line, control of mutation (deletions and duplications) of hot spots in the genome with continuous passaging, control of tumor formation after administration, and need for better assays of pluripotency and efficacy. The use of patient-specific cells for regenerative medicine provides a therapeutic which will not induce classic cytotoxic rejection. However, this is a possibility only when the hESCs do not carry a genetic mutation, which is instrumental in manifestation of the disease. The creation of patient-specific hESCs would likely use iPSC technology, and newer methods need to be developed where no retroviral transduction or oncogene expression is involved. Also, additional methods need to be developed for largescale culture of cGMP grade cells so that ample numbers are generated for differentiation and transplantation into patients. This could involve suspension or adherent culture formats depending on the projected use. Finally, as instances of trials and treatments evolve and increase, a streamlined regulatory process needs to be developed for cellular replacement and regenerative medicine therapies. The power of hESC technology has already been proven for in vitro applications. It has been successfully used for disease modeling as mentioned above, and in selected instances proved to be an effective platform for reversal of disease phenotype (Raya et al., 2009). iPSC technology can be used very effectively for creating models of disease using tissue from patient samples with relevant disease backgrounds. Deficits due to these particular mutations can be studied in dynamic and physiologically relevant laboratory platforms where processes such as differentiation, maturation, survival/apoptosis, proliferation, etc. may be dissected. hESC and iPSC technology will prove extremely useful for identification of novel therapeutic targets of disease, the validation of known and novel targets, and the prosecution of validated targets. Sophisticated disease-specific screening platforms and assays may be developed using high content screening and similar technologies. Targeted transgenic and knockout lines may also be created as required and this will be valuable in target validation. hESC lines will also be used for toxicology and selected in vitro ADME assays. Toxicology may be studied at several levels such as differentiation of functional liver, cardiac and neural platforms for studying mechanisms of toxicity of therapeutic molecules, and teratological screening of molecules on undifferentiated hESC lines. The power of hESC/iPSC technology will be realized when these physiological studies are performed on panels of cultures of known genotypes, so that companion diagnostics may be developed using pharmacogenomic methods. This will allow predictive toxicology of medications on individuals of known genotypes. The question of “are all hESC cultures equal” remains. As mentioned above, hESCs may be created by culture of ICM, SCNT, and iPSC techniques, among others. The similarities, differences, and limitations of each of the stem cell types needs to be further defined. Are the final hESC lines which are formed distinct based on the method of derivation, the inducing factor/s, and the characteristics of the cells, which have undergone selection to create the hESC line? As well, the reasons for different differentiation thresholds and potentials for each individual heSC line requires further definition.

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References Rajan, P., Ross, R., Mackler, A., Smotrich, D., & Loring, J. (2007). Derivation of human embryonic stem cells. In J. Loring (Ed.), Human Stem Cells: A Laboratory Guide (pp. 291– 308). Elsevier Publications. Raya, A., Rodríguez-Pizà, I., Guenechea, G., et al. (2009). Disease-corrected haematopoietic progenitors from Fanconi anaemia induced pluripotent stem cells. Nature, 460, 53–59. Shamblott, M. J., Axelman, J., Wang, S., et al. (1998). Derivation of pluripotent stem cells from cultured human primordial germ cells. Proc. Natl. Acad. Sci. U.S.A., 95, 13726– 13731. Tesar, P. J., Chenoweth, J., Brook, F. A., et al. (2007). New cell lines from mouse epiblast share defining features with human embryonic stem cells. Nature, 448, 196–199. Thomson, J. A., Itskovitz-Eldor, J., Shapiro, S. S., et al. (1998). Embryonic stem cell lines derived from human blastocysts. Science, 282, 1145–1147.

Relevant Websites http://clinicaltrials.gov/ct2/show/NCT01469832?term¼advancedþCellþTherapy&cond¼%22MacularþDegeneration%22&rank¼3 – ACT hESC clinical trial. http://www.roslin.ed.ac.uk/public-interest/dolly-the-sheep/ – Dolly the cloned sheep. http://grants.nih.gov/stem_cells/registry/current.htm – List of NIH approved hESC lines. http://stemcells.nih.gov/Pages/Default.aspx – NIH General Stem Cell information. http://stemcells.nih.gov/info/scireport/pages/chapter3.aspx – NIH hESC information site. http://www.nobelprize.org/nobel_prizes/medicine/laureates/2012/yamanaka-facts.html – Shinya Yamanaka. http://www.ukstemcellbank.org.uk/ – UK Stem Cell Bank. http://viacyte.com/products/vc-01-diabetes-therapy/ – Viacyte diabetes treatment.

Introduction to Regenerative Engineering Manisha Jassal and Radoslaw Junka, Stevens Institute of Technology, Hoboken, NJ, United States Cato T Laurencin, University of Connecticut Health Center, Farmington, CT, United States Xiaojun Yu, Stevens Institute of Technology, Hoboken, NJ, United States © 2019 Elsevier Inc. All rights reserved.

Introduction Regenerative Engineering Clinical Need for Regenerative Engineering Different Strategies for Regenerative Engineering Top-down engineering approach Bottom-up engineering approach Key Elements of Regenerative Engineering Stem Cells: The Fundamental Building Block of New Tissues Morphogenetic Signals: Importance of Transition From Individual Cells to Structured Tissues Biomaterials Regenerative Engineering Application Areas Current Challenges and Future Directions Acknowledgment References Further Reading

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Glossary Blastema Mass of undifferentiated cells that has the capability to develop into an organ or appendage. Inducerons Small-molecule inducers of cell differentiation.

Introduction Regenerative Engineering One of the most fascinating clinical issues being faced by scientists of the 20th century has been the regeneration of complex tissues and organ systems such as a knee or whole limb. Compared to urodele amphibians that exhibit a remarkable ability to completely regrow severed limbs during any time point in their life, the humans do not possess such regenerative ability for amputated limbs. Rehabilitation of with patients amputated limbs has been achieved using prosthetic devices. However, these devices still cannot perform complex functions such as normal gait and movement feedback. Limitations of current biological and engineering approaches toward limb regeneration have led to emergence of a new field called “Regenerative Engineering.” Regenerative Engineering, an interdisciplinary field, is defined as the convergence of advanced Materials Science, Stem Cell Sciences, Physics, Developmental Biology, and Clinical Translation for the regeneration of complex tissues and organ systems. Built constructs are produced using a combination of cells, growth factors, and synthetic scaffolds with an aim to achieve functionality either equivalent to or greater than that of damaged or lost native tissue (El-Amin et al., 2013). The focus of regenerative engineering has advanced from replace/restore function to regenerate the living tissue, making it the future of tissue engineering. To achieve this goal, advances made in gaining knowledge of phenomenon taking place in nanoregime combined with advanced materials science, a mature stem cell science field that gives the opportunity of using stem cells as an everyday tool, and deeper understanding of developmental biology mechanisms has given scientists the tools to work toward regeneration of whole tissue (Laurencin and Khan, 2012). Success would be achieved for regenerative engineering field when elements of science, engineering, and medicine are combined together to create constructs that would ultimately improve body’s ability to regenerate its diseased/lost organs.

Clinical Need for Regenerative Engineering Blood and bone represent two systems inside of a human body that can regenerate. Bone regeneration is limited, because it cannot regenerate beyond a critical sized defect (Nair and Laurencin, 2015). Similarly, liver has a regenerative capacity that is seen in all vertebrate organisms (Michalopoulos, 2007). Most of the other tissues present in human body undergo a repair mechanism instead of regeneration that leads to formation of a scar tissue. This scar tissue lacks necessary biological and mechanical properties that are present in the native tissue, hence leading to functional impairment (Nair and Laurencin, 2015). In order to address the issue of

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functional impairment, efforts were focused on different types of biological grafts, namely autograft, allograft, and xenografts. Autografts are considered the gold standard, because they do not present complications associated with other grafts such as adverse immune response. The limited availability of autografts and complications associated with other grafts led to engineering tissues as substitutes for clinical use to replace diseased organs or heal damaged tissues (Nair and Laurencin, 2015). However, the clinical success of tissue-engineered products is still limited due to various limitations such as insufficient vascularization that leads to lack of nutrients and oxygen and insufficient removal of waste products including scaffold degradation products (Liu et al., 2013; Seliktar et al., 2014). Tissue engineering has given successful outcomes in terms of regenerating small segments of bone and hallow organs, but regeneration of complex structures has still not been realized clinically. This has led the scientific community to focus on developing novel strategies to regenerate complex tissue and organ systems such as a functional knee or whole limb, hence giving rise to a much advanced research field of regenerative engineering (Reichert et al., 2011).

Different Strategies for Regenerative Engineering The success of regenerative engineering will depend on recapitulating the complex architecture of organs and the native arrangement of cells and extracellular matrix (ECM) at the functional level. A comprehensive approach is required for the design of regenerative constructs based on understanding of cellular and morphogenic processes, biological factors, scaffolds, physical forces, and the interrelation between these essential components (Nair and Laurencin, 2015). Regenerative engineering could make complex tissue regeneration possible using the top-down and bottom-up approaches that are described below.

Top-down engineering approach The top-down approach for limb regeneration considers the limb structure as an aggregate of individual specific tissues integrated one by one into a unified structure. The top-down approaches are built on the integration of advanced material science and engineering, cells with high regenerating ability, physical forces, and immune system modulation (Nair and Laurencin, 2015). Cells are seeded on prefabricated porous polymeric scaffolds that serve as temporary template for new tissue growth. Bioreactors can be utilized to simulate physiological conditions so that cells can differentiate and secrete their own ECM (Urciuolo et al., 2013). The main advantage associated with top-down approaches is that these processes are cost-effective, scalable and able to generate clinically relevant nanostructured matrices. Limitations of this approach includes difficulty in recreating the complex microarchitecture of tissues, slow vascularization, and diffusion barriers which restrict the potential of this strategy to be used effectively for regeneration of complex tissues (Urciuolo et al., 2013; Lu et al., 2013).

Bottom-up engineering approach The bottom-up approach is based on the principle of molecular self-assembly. This phenomenon is defined as a spontaneous organization of molecules under near thermodynamic equilibrium conditions to form stable structures. A larger tissue construct can be generated by assembling smaller building blocks via multiple assembling methods, such as surface tension assembly, acoustic assembly or magnetic assembly. The advantage of the bottom-up approach is spacial control over features and composition of individual blocks (Lu et al., 2013). The successful regeneration of complex tissues and organ systems will depend on the ability to integrate both approaches, while optimizing advantages offered by each. Many amphibians have a remarkable ability to regenerate complex tissues and organ systems by the self-assembly of proliferating and differentiating cells in blastema. Humans lack such capability, except repair of a small injury in tissues such as bone or skin (Nair and Laurencin, 2015). The bottom-up approach will help to understand the highly orchestrated process of tissue development as it is unfolding through a tightly controlled spatial and temporal expression of various morphogenetic cues (Nair and Laurencin, 2015). The top-down approach will aids in designing strategies to deliver biologically active effector molecules and inducerons in a tissue-engineered construct that would replicate the complex architecture of the organ system (Nair and Laurencin, 2015).

Key Elements of Regenerative Engineering The epimorphic regeneration exhibited by the urodele amphibians occurs through the formation of a blastema that consists of progenitor cell population with intrinsic morphogenetic cues. The blastema presents a highly complex gene profile, and the unique ECM of the blastema provides directional cues to modulate cellular functions. Some of the pathways that control embryonic limb formation such as fibroblast growth factor (FGF) and Wnt-ß catenin signaling, along with retinoic acid and Shh signaling, are essential in patterning and morphogenesis during limb regeneration. A combination of these signaling pathways combined with the role played by the immune system and innervation gives rise to a favorable regenerative environment in the blastema. The understanding achieved from the regenerative process in urodele amphibians, combined with interdisciplinary scientific approaches, can lead to exciting new solutions that address current challenges to human organ regeneration. A deeper understanding of adult and embryonic stem cells over the past few decades led researchers to appreciate similarities these cells have with blastema cells. Similarly, induced pluripotent stem cells (iPSCs) derived from adult differentiated cells undergo a process of cellular dedifferentiation also present in blastema. Considering blastema cells molecularly similar to a cell in a more undifferentiated state, the possibility of controlling the extent of cellular dedifferentiation may have significant impact on

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developing a translational protocol for limb regeneration in humans. Another key component of regenerative engineering that has advanced significantly in the past two decades is the biomaterials science. The current research focuses on developing advanced biomaterials, wherein the physical, mechanical, and biological properties of the scaffold can be fine-tuned to enhance the natural regenerative process in the body. In the following sections, essential components for regenerative engineering are described in detail including stem cells, morphogenetic signals, and biomaterials. In addition to these, physical forces such as electrical forces influence morphogenesis and patterning, and modulate cellular functions leading to a more permissive microenvironment for tissue regeneration.

Stem Cells: The Fundamental Building Block of New Tissues Regeneration of a damaged tissue largely depends on cells’ ability to repopulate the defect or the substituting scaffolding material and reestablish the native architecture of the tissue. To this end, of a particular interest to engineers are populations of regenerationcompetent cells, called stem cells. Stem cells are phenotypically characterized by their ability to divide continuously through asymmetric mitosis to yield two different populations of cells, daughter stem cells of the same phenotype and progenitor cells that are more differentiated than the original cell (Zhong, 2008). This self-renewal capability allows for stem cells to undergo almost infinite number of replications, generating in the process any mass of tissue needed. Therefore, they can produce populations of cells capable of maintaining this special phenotypic state and simultaneously differentiate into other cell types. These progenitor cells are plastic; meaning they have the ability to alter their genetic profile expression and adopt functional phenotype of cells present in developed tissues through a process of transdifferentiation. Transdifferentiation can result directly through interaction of stem cells with another tissue of a different phenotype, or it can be achieved indirectly via artificially manipulation of the cell environment. The indirect transdifferentiation transforms an adult differentiated cell into a more primitive state, which is then differentiated into another cell type (Lodi et al., 2011). However, stem cells’ plasticity depends on the origin of a tissue type, embryo or the adult tissue. Stem cells are categorized into three populations depending on the tissue of origin, such as embryonic stem cells (ESCs), induced pluripotent stem cells (iPSCs), and postnatal or adult stem cells. ESCs have the greatest potential for self-renewal and expansion and are able to differentiate into any cell type (Lenoir, 2000). While these characteristics make ESCs a promising candidate for the use in tissue regeneration, their implementation is often entangled with serious ethical considerations. iPSCs are derived from adult differentiated cells, for example, skin fibroblasts, by induction of a more primitive state or dedifferentiation (Takahashi and Yamanaka, 2006). They do not have the same self-renewal and differentiation potential; however this is an area of intense ongoing investigation. Although both of these stem cell populations present a tremendous therapeutic potential, their application encompasses serious risks and limitations such as formation of cancerous tissues, immune response to ESCs, and difficulty in eliciting specific genetic profiles in iPSCs (Ben-David and Benvenisty, 2011). Postnatal adult stem cells, on the other hand, do not pose these difficulties in culture. Two of the most widely studied adult stem cell populations are mesenchymal stem cells (MSCs) and adipose-derived stem cells (ASCs). Likewise, these cell populations are limited in their ability to undergo self-renewal and transdifferentiation process, hence they are termed as multipotent and not pluripotent as ESCs (Mizuno et al., 2012). MSCs can be isolated from postnatal organs and connective tissues, such as bone (Song et al., 2005; Choi et al., 2008), synovial membrane (De Bari et al., 2001), skeletal muscle (Dodson et al., 2010), peripheral blood (Shi et al., 2009), periodontal ligament (Seo et al., 2004), and umbilical cord (Baksh et al., 2007; Musina et al., 2005) among other tissues. However, these sources usually yield low number of cells that can be harvested. Thus, to become a viable option for the clinical use, strategies implementing the use of ESCs, iPSCs, or MSCs require labor intensive and expensive ex vivo culture methods. In contrast, ASCs overcome this limitation and hold greater promise for future therapeutic use as adipose tissue can be easily harvested in large quantities with little donor site morbidity (Zuk et al., 2001). This advantage is manifested in a large number of preclinical and clinical studies of injury and disease already underway (Gimble et al., 2007; Tobita et al., 2011; Rosado-de-Castro et al., 2013). Nevertheless, there remain several challenges to translation of stem cell therapies to a wide clinical use. The limited understanding of motility, survival, proliferation, and differentiation in in vivo setting is a major obstacle. Human and animal studies show poor engraftment of stem cells as reflected in only a few percent of cells remaining several weeks after injection (Terrovitis et al., 2010). Likewise, long-term safety and efficacy is a concern as previous studies demonstrate these cells’ ability to form tumors (Hentze et al., 2009). Once implanted, stem cells, progenitor cells derived from them, and the surrounding endogenous cells encounter and exchange a multitude of signals in a highly dynamic niche that affects their phenotype. Thus, greater efforts are necessary to elucidate and control molecular processes that guide expansion and transdifferentiation of these cells in vivo. Recent developments in the science of stem cells and their associated morphogenic cues offer exciting possibilities for the future use of stem cells to regenerate tissues and, eventually, to replace entire organs (Platt, 2004).

Morphogenetic Signals: Importance of Transition From Individual Cells to Structured Tissues Complex interrelationship between mechanical forces and a number of protein families orchestrate the process of tissue morphogenesis during development. The mechanical stimuli are known to affect cells’ size, shape, and phenotype. These changes in gene expression in turn induce cells to secrete a host of protein products that further influence tissue formation (Heisenberg and Bellaïche, 2013). These molecules regulate self-renewal, migration, and differentiation uncommitted stem or progenitor cells. Their secretion is tightly controlled in space and time (Murry and Keller, 2008). Interestingly, many of the same morphogenetic processes

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observed during development also take place in some adult human tissues as they regenerate (Stocum, 1998; Lutolf and Hubbell, 2005). Thus, scientists and engineers aim to augment the self-healing capacity in certain tissues by artificially recapitulating this intricate balance between mechanical forces and cell signaling molecules. Application of one or more of these morphogenic cues guides cells toward a desired tissue formation by enhancing several of the processes important for healing. A category of these signaling molecules, called growth factors, are often added to cultures, delivered via polymeric vehicles, or bound to scaffolds. During development, these factors are actively secreted by stem cells and nearby niche cells. Tissue growth and differentiation occurs as a result of the interpretation of specific concentrations or a gradient of these morphogenetic signals (Discher et al., 2009). Concurrently, local cell–cell interactions between stimulated and unstimulated cells establish boundaries of different cell phenotypes. Functional tissue recovery is thought to be possible only when undifferentiated or genetically modified cells are recruited into a regeneration or controlled growth pathway by the presence of these morphogenetic signals. Taking into consideration each tissue’s own ability to heal, functional recovery might also depend on suppression of inhibitory signals originating from rapid scar formation that curtail the regenerative process. Therefore, to achieve successful tissue regeneration, it is imperative to create and control over time the artificial environment that enables cells to proliferate and differentiate by carefully manipulating timing and concentrations of morphogenic cues. The dependence of successful tissue regeneration on growth factors is perhaps best exemplified in their direct induction of angiogenesis, sprouting of immature capillaries, which supply oxygen and nutrients to cells. This process is necessary to maintain biological functions of cells in immature regenerating tissue of a clinically relevant size (Tabata, 2003). Although biological effects of growth factors are greatly enumerated, successful long-term delivery of these morphogenetic factors evades scientists due to their poor in vivo stability. Direct injections are deemed impractical, because their systemic effects are likely to be short-lived or cause undesired cancerous growth. Therefore, scientists turned toward various drug-delivery systems to achieve sustained long-term targeted delivery of these factors. Despite fervent efforts, application of growth factors in clinical treatment remains limited due to inefficient loading of these large molecules into their carriers and subsequent burst release upon administration. In addition, future studies should bring vast improvement in production of recombinant growth factors to drive down the cost of production. Emphasis on robust preclinical models and detailed pharmacokinetic studies should ensure safety and efficacy of these recombinant factors. Remaining important caveats to be accounted for are the likely required use of a combination of these factors, their corresponding therapeutic concentrations, and the native time sequence at which they are secreted during regeneration (Koria, 2012). To achieve more efficient delivery systems and realize the therapeutic potential of these morphogenic cues requires novel biomaterials with tunable physical and chemical properties.

Biomaterials Biomaterials act as extracellular matrix scaffolding as they provide physical structure to retain cell population after implantation, guidance, and template for cells to lay new extracellular matrix. However, over time, biomaterials have evolved from just functioning as a template where cells can attach and lay their own extracellular matrix to multifunctional systems that actively regulate various aspects of tissue regeneration. Apart from being biocompatible and biodegradable, they can now incorporate biological and structural cues to induce favorable cell response such as enhanced cell attachment and directed cell differentiation to achieve a particular outcome. One of the examples includes biomaterials utilized for bone regeneration that could be made osteoconductive (e.g., by selecting appropriate chemical composition) and osteoinducive (e.g., by varying the microstructure of the material) (Yu et al., 2015). Biomaterials can be classified as natural and synthetic, based on their origin with both offering certain advantages and disadvantages. Natural biomaterials can be subclassified based on their chemical structure into proteins, polysaccharides, and polyesters (Mano et al., 2007). These offer advantages such as being similar to biological macromolecules that might aid in reducing the inflammation and toxicity when implanted in body. Some examples of natural protein biomaterials include collagen, fibrin, fibronectin, and silk. Chitin, chitosan, alginate, and hyaluronic acid are some examples of polysaccharide-based natural biomaterials. Polyhydroxyalkanoates, derived from microorganisms, are examples of natural thermoplastic polyesters that are biocompatible and biodegradable. Decellularized tissue-derived biomaterials can also be considered another category of natural biomaterials that are obtained by the elimination of cellular/nuclear components and retaining just the extracellular matrix components. Examples include decellularized dermis, heart valves, blood vessels, small-intestinal submucosa, and liver (Chen and Liu, 2016). The main limitations of natural biomaterials include insufficient mechanical strength, batch-to batch variation due to complex purification process, possibility of adverse immunogenic reaction, and inability to control in-vivo degradation rate. Synthetic biomaterials, on the other hand, offer advantages such as tunable mechanical and degradation properties by varying the chemical composition of the polymers, and the ability to be fabricated into various shapes with controllable micro and macrostructure. The most extensively studied synthetic biomaterials include polycaprolactone, polyglycolide, polylactide, polyhydroxybutyrate, and their copolymers. Their disadvantages include reduced bioactivity and harmful degradation by-products produced in-vivo. Studies have shown that the bioactivity of the advanced biomaterials can be significantly enhanced using biological proteins/ peptides as well as biologically active effector molecules and inducerons. The three-dimensional structure of the biomaterial scaffold also plays an important role in modulating cellular behavior that could positively impact the regenerative capability of tissues and organs. The developments in the micro- and nanotechnologies have led to creation of novel 3D biomimetic scaffolds and the last decade exhibited significant growth in the fabrication and characterization of nanofibrous 3D structures as biomimetic scaffolds. The latest innovation in this direction is focused on additive manufacturing or 3D printing that can create patient-specific complex

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3D structures. The success of regenerative engineering would depend on our ability to use these biomaterial scaffolds, natural, synthetic, or a blend of the two, as a delivery system for growth factors, adhesion peptides, and cytokines as well as a mechanical support structure with favorable micro- and macrostructure to induce regeneration of specific tissue/organ (O’brien, 2011).

Regenerative Engineering Application Areas Recent progress made in understanding stem cells science and developmental biology, combined with advances made in material science and nanotechnology have brought the scientific community closer to realizing their goal of regenerating complex tissues, organs, or organ systems (Nair and Laurencin, 2015). Regenerative engineering principles have been applied for musculoskeletal tissue, cardiovascular tissue, and neural tissue regeneration. Apart from these, other organs of interest include the liver, lung, kidneys, and pancreas. Regenerative engineering clinical applications based on biodegradable scaffolds include the artificial skin, nerve conduits, bone, articular cartilage, blood vessels, and bladder. Another technique utilized for cell-dense-tissues is cell sheet technology where autologous cells are grown and collected as a contiguous sheet of cells while preserving their natural extracellular matrix (ElloumiHannachi et al., 2010). This technology has been utilized for corneal surface reconstruction, periodontal tissue, and myocardial tissue reconstruction. Various clinical trials have been undertaken to establish the efficacy of different scaffolds for their regeneration potential in-vivo, but only a few have been approved till date. For bone regeneration, PCL scaffolds in various forms and shapes such as Osteoplug, Osteoplug-C, Osteomesh, and Osteomesh-Osteostrip have been approved for craniofacial applications (Liu et al., 2013). Many clinical trials for bone regeneration are currently in progress that involves a combination of stem cells, growth factors, and/or scaffoldbased technologies. For nerve regeneration applications, many natural resorbable and synthetic resorbable devices have been approved by FDA that include Type I collagen based devices (NeuraGenÒ, NeuroflexÔ, NeuromatrixÔ, NeuraWrapÔ, and NeuroMendÔ), polyglycolic acid-based device (NeurotubeÒ), and poly D,L lactide-co-3-carprolactone-based device (NeurolacÒ) (Kehoe et al., 2012). Type I collagen-based devices fall under the category of natural resorbable materials, whereas polyglycolic acid and poly D,L lactide-co-3-carprolactone-based devices fall under synthetic resorbable materials utilized for nerve regeneration applications. Similarly, various trials are underway in the lung, liver, and pancreas regeneration field that utilize stem-cell-based therapies. Examples include infusion of mesenchymal stem cells (MSCs) for lung repair and regeneration (Kotton, 2012) and transplantation of bone-marrowderived stem cells that initiated endogenous pancreatic regeneration (Hess et al., 2003).

Current Challenges and Future Directions Although enormous advances have been made in the application of regenerative engineering principles to regenerate various human tissues, the regenerative field still far from achieving its ultimate goal of regeneration of a whole organ system. One of the major challenges is to recreate the complex architecture of tissues and organs as it directly influences the function of that particular organ/tissue. Various strategies to capture native organ structure and material composition include decellularization of organs followed by recellularization before transplantation or use of biomaterials seeded with cells. The limitation associated with these techniques is inadequate control of cell placement within the scaffolds. 3D bioprinting can address some of these limitations, but they suffer from trade-offs such as feature resolution, cell viability, and print resolution (Mao and Mooney, 2015). Another challenge lies in properly integrating the implanted grafts with the body in terms of vascularization and innervation. Vascularization can either be achieved by exploiting body’s own angiogenic response via presentation of angiogenic growth factors through controlled delivery or prevascularization of the graft. Similarly, innervation of engineered tissues/biomaterials can also be induced by use of growth factors (Mao and Mooney, 2015). Modulating the immune system in order to avoid rejection of the implanted biomaterial/decellularized tissue presents complication that can be addressed by either engineering the responses of immune cells or changing the properties of the implanted material such as increasing hydrophilicity or modifying the surface with adhesion ligands that promote cell attachment and proliferation (Mao and Mooney, 2015). The future success of regenerative engineering holds promise for patients whose tissue/organ systems cannot recover from disease or trauma by currently available drugs or therapeutic methods. The advances made in design of biomaterials that mimic the natural architecture of tissues along with the ability to deliver biomolecules in a controlled manner, better understanding of developmental biology along with advances made in stem cell research would pave the way to regenerate complex tissues and organs to heal patients.

Acknowledgment This work was partly supported by the Assistant Secretary of Defense for Health Affairs, through the Peer Reviewed Medical Research Program under Award No. W81XWH-16-1-0132.

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Further Reading Laurencin, C. T., & Khan, Y. S. (Eds.). (2013). Regenerative engineering. Florida: CRC Press. Laurencin, C. T., & Nair, L. S. (Eds.). (2015). Nanotechnology and regenerative engineering: The Scaffold. Florida, USA: CRC Press. Liu, S. Q. (2007). Bioregenerative engineering: Principles and applications. New Jersey: Wiley. Berthiaume, F., Maguire, T. J., & Yarmush, M. L. (2011). Tissue engineering and regenerative medicine: History, progress, and challenges. Annual Review of Chemical and Biomolecular Engineering, 2, 403–430. Orlando, G., Wood, K. J., Stratta, R. J., Yoo, J. J., Atala, A., & Soker, S. (2011). Regenerative medicine and organ transplantation: Past, present, and future. Transplantation, 91(12), 1310–1317. Dawson, J. I., & Oreffo, R. O. (2008). Bridging the regeneration gap: Stem cells, biomaterials and clinical translation in bone tissue engineering. Archives of Biochemistry and Biophysics, 473(2), 124–131. Huey, D. J., Hu, J. C., & Athanasiou, K. A. (2012). Unlike bone, cartilage regeneration remains elusive. Science, 338(6109), 917–921.

Nanoelectronics for Neuroscience Sahil Kumar Rastogi and Tzahi Cohen-Karni, Carnegie Mellon University, Pittsburgh, PA, United States © 2019 Elsevier Inc. All rights reserved.

Introduction Bioelectrical Signals Recording Extracellular Recordings MEAs Planar passive electrodes Flexible passive electrodes FETs 2D FETs 3D FETs Injectable Electronics Intracellular Recordings MEA Functionalized gold-spine electrode (FGSE) Vertical Si nanowire (SiNW) electrode array (VNEA) Pt nanopillars FETs 3D-kinked NW probes Branched intracellular nanotube-FET (BIT-FET) Active Si nanotube transistor (ANTT) Conclusion References Further Reading

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Glossary Action potential Localized change in the electric potential across the cell membrane of an electrically active cell, from about
70 mV to þ 30 mV. Electrocorticography (ECoG) Technique of recording brain electrical field potentials with metal electrodes placed directly on the cortical surface. Electroencephalogram (EEG) Technique of recording brain electrical field potentials with metal electrodes placed on the scalp. Electrophysiology Study of the electrical properties of electrically active cells. Electroporation A technique in which an electrical field is applied to cells to increase the permeability of the cell membrane. Excitatory postsynaptic potentials (EPSP) Depolarization of the membrane potential of a postsynaptic neuron caused by the binding of an excitatory neurotransmitter from a presynaptic cell to a postsynaptic receptor. Extracellular recording Monitoring the change in the potential across cell membrane of cells from outside the cell. Field-effect transistor (FET) A type of electrical device where the current across source and drain terminals is controlled by the potential applied to a third terminal named gate. fMRI Functional neuroimaging procedure using magnetic resonance imaging (MRI) technology that measures brain activity by detecting changes associated with blood flow. Glia Nonneuronal cells that maintain homeostasis, form myelin, and provide support and protection for neurons in the central and peripheral nervous systems. Graphene One atom thick layer of sp2-hybridized carbon atoms arranged in hexagonal lattice. Hyperpolarization Change in the membrane potential of the cell making it more negative. Inhibitory postsynaptic potentials (IPSP) Hyperpolarization of the membrane potential of a postsynaptic neuron caused by the binding of an inhibitory neurotransmitter from a presynaptic cell to a postsynaptic receptor. Intracellular recordings Monitoring electrical activity of cells from within the cell cytosol. Ion channels Proteins that act as pores in a cell membrane and permit the selective passage of ions such as potassium, sodium, and calcium ions. Microelectrode array (MEA) An array of passive electrodes, ranging from 10 mm to 500 mm in dimensions, that enable simultaneous multisite recording of electrical signals from the electrically active cells.

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Neuron An electrically excitable cell in the nervous system that processes and transmits information through electrical and chemical signals. Neurotransmitter A chemical released from an electrically excited presynaptic neuron that leads to either excitation or inhibition of the postsynaptic neuron. Optogenetics A technique in which genes for light-sensitive proteins are introduced into neurons to monitor and control their activity precisely using light signals. Patch clamp Type of intracellular recording that involves attaching a glass pipette to the outer membrane of a single cell and then recording the activity of ion channels in the cell membrane. Phospholipid Amphipathic compounds with hydrophilic head and hydrophobic lipophilic tail that serve as a major structural component of cell membranes. Positron emission tomography Imaging technique that measures emissions from radioactively labeled metabolically active chemicals that have been injected into the bloodstream. Resting membrane potential (RM) Voltage difference across the cell membrane when the cell is at a resting or quiescent state. Seal resistance Resistance generated in the cleft between the cell and the sensing element. Subthreshold potential A stimulus that leads to change in membrane potential of a neuron with a magnitude smaller than the threshold that is required to initiate an action potential. Synapse Junction that permits a neuron to pass an electrical or chemical signal to another neuron.

Introduction Nervous system, the most complex network in the body, consists of two classes of specialized cells called glia and neurons. Glia are cells that mechanically support the neural growth, form the immune system, supply the neurons with nutrients, and electrically isolate the neurons from each other (Kandel et al., 2000). Neurons, on the other hand, are electrically excitable cells that receive signals, process them, and transmit information throughout the body (Kandel et al., 2000). Typical morphology of a neuron consists of a cell body called soma, thin branched outgrowths called dendrites, and a filament that extends away from the cell body called axon, as represented in Fig. 1A. Dendrites are the main apparatus for receiving incoming signals from other nerve cells, whereas axon is the main conducting unit for carrying signals to other neurons. An axon can convey electrical signals, called action potentials (APs), along the distances ranging from 0.1 mm to 3 m (Kandel et al., 2000). Neurons communicate with one another at junctions called synapses. The communication between neurons, also called neurotransmission, is accomplished by the movement of chemicals across the chemical synapse or electrical signals across the electrical synapse. At electrical synapses, two neurons are physically connected to one another through gap junctions (Nicholls et al., 2001). Gap junctions permit changes in the electrical properties of one neuron to effect the other, and vice versa. In chemical neurotransmission, the presynaptic neuron and the postsynaptic neuron are separated by a small gap called the synaptic cleft. The AP in the transmitting neuron releases a chemical messenger (a neurotransmitter) across the synaptic cleft. The neurotransmitter binds to receptor proteins on the receiving neuron, and the resulting reaction transduces the potential chemical energy of the transmitter into electrical energy (Nicholls et al., 2001). The cellular membrane is composed of a phospholipid bilayer that separates cytoplasm of the cell from the extracellular environment. In general, the major ionic constituents of the external tissue fluid are Naþ and Cl
ions, whereas there is high Kþ ion concentration inside the cell. Since the ions are charged, they cannot pass directly through the hydrophobic lipid regions of the membrane. Therefore, in order to cross the membrane, they have to use specialized ion-specific channel proteins such as (i) voltage-gated ion channels that are activated by changes in the membrane electrical potential, (ii) leak ion channels that are always permeable to the ions, and (iii) ATP-dependant Naþ–Kþ pumps that pump Naþ and Kþ ions against their concentration gradients (Fig. 1B). The movement of the ions across the cell membrane is influenced by two energetic factors: the concentration gradient and the electrical potential difference (caused due to the charge on ions). When the electrical potential difference across the cell membrane exactly balances the concentration gradient for an ion, it is known as the equilibrium potential of that ion. When the cell is at rest, the potential across the neuronal membrane is referred to as resting membrane potential (VM). Resting VM is determined by the uneven distribution of ions between the inside and the outside of the cell, and by the different permeability of the membrane to different types of ions. At the resting state, the voltage-gated ion channels are closed, and the mode the ions can cross the membrane is through leak ion channels. Since the resting membrane comprises of high number of leak channels specific to Kþ ions, therefore the resting membrane potential is closer to the Kþ equilibrium potential than it is to the Naþ equilibrium potential, and lies around
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55 mV, it initiates an AP which gets conducted down the axon without failure or distortion at rates of 1–100 ms
1. Initiation of an AP involves opening of voltage-gated Naþ channels (Fig. 1C). When the channels open, Naþ ions flow inside the cell and depolarize the cell, thus generating a positive intracellular and a negative extracellular potential change (Fig. 1D). The reversed VM results in the closing of Naþ channels and activation of the voltage-gated Kþ channels. The outflow of Kþ ions repolarizes the cell which eventually crosses the resting VM leading to hyperpolarization. With the help of

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Fig. 1 Morphology and electrical activity of neurons. (A) Schematic illustration of a mammalian neuron (Reprinted with permission from Koeppen, B. M. and Stanton, B. A. (2009). Berne & Levy physiology. Elsevier Health Sciences. Copyright 2009 Mosby, Elsevier). (B) Neuronal membrane containing voltage-gated Kþ ion (blue) and Naþ ion (orange) channels, as well as Kþ–Naþ pumps (green). (C) Coarse of the membrane potential and the channel conductance of Naþ and Kþ ion channels during an action potential. Red arrow indicates the threshold value and blue arrow the after hyperpolarization (Reprinted with permission from Bezanilla, F. (2006). The action potential: From voltage-gated conductances to molecular structures. Biological Research 39 (3), 425–435. Copyright 2006 BioMed Central). (D) Intracellular (top) and extracellular (bottom) action potentials measured simultaneously using a patch clamp (intra), and a Si-FET (extra) interfaced to the axon of a leech neuron (Reprinted with permission from Fromherz, P. (2002). Electrical interfacing of nerve cells and semiconductor chips. ChemPhysChem 3 (3), 276–284 Copyright 2002 John Wiley & Sons).

the Naþ–Kþ pumps that exchange three intracellular Naþ for two extracellular Kþ, the cell gets back to its resting VM. Once a single AP spike is initiated on a small section of the membrane, it triggers an AP in the subsequent section, ensuring the signal propagation along the membrane. At the same time a transient negative shift due to the hyperpolarization prevents the back-propagation of the signal. The amplitude of an AP remains constant while traveling down the axon because the AP is an all-or-none impulse that is regenerated at regular intervals along the axon (Nicholls et al., 2001). APs lead to multitude of functions from signal propagation to information encoded in their firing frequency and pattern (Nicholls et al., 2001). A stimulus that is insufficient to initiate an AP is known as a subthreshold stimulus. It can either lead to depolarization of the cells called excitatory postsynaptic potentials (EPSPs) or hyperpolarization of the cells called inhibitory postsynaptic potentials (IPSPs).

Bioelectrical Signals Recording In the 18th century, Luigi Galvani discovered that the function of the nervous system is intrinsically linked to its electrical activity (Galvani, 1791). By the late 1930s, researchers started to understand the conduction of electrical signals within neurons. Kenneth Cole and Howard Curtis demonstrated that the AP is associated with a large increase in membrane conductance (Cole and Curtis, 1939). During the same time, Alan Hodgkin and Andrew Huxley demonstrated the first intracellular recording of an AP in a giant squid axon model (Hodgkin and Huxley, 1939; Hodgkin and Katz, 1949; Hodgkin and Huxley, 1952a; Hodgkin and Huxley, 1945; Hodgkin and Huxley, 1952a; Hodgkin and Huxley, 1952b; Hodgkin and Huxley, 1952c; Hodgkin et al., 1952). Their mathematical model explained the initiation and propagation of AP and the flow of currents across the cell membrane (Hodgkin and Huxley, 1952d). Hodgkin and Huxley’s work allowed researchers a step-by-step view of the processes involved in an AP. The effect of their work was tremendous, leading to an increase in the interest in electrophysiology. Since then, generations of investigators have invested immense effort into building and developing tools capable of recording and controlling the electrical activity of neurons to enable better understanding of the brain in health and disease, enable early diagnosis and treatment of neurodegenerative disorders, and help restore lost neuronal function associated with neurological diseases and injuries (Marconi et al., 2012; Kanagasabapathi et al., 2012; Ruaro et al., 2005; Liu et al., 2007; Sun, 2002; DeLorenzo et al., 2000; Yuan et al., 2016; Marg and Adams, 1967; Evarts, 1966; Wolpaw et al., 2002; Wolpaw and McFarland, 2004; Lebedev and Nicolelis, 2006). The recording platforms that have been developed until now can be categorized into three groups. First, neuroimaging approaches, such as functional magnetic resonance imaging (fMRI) and positron emission tomography (PET) (Kornblum et al., 2000; Poldrack and Farah, 2015). Their advantages include noninvasive nature and successful long-term monitoring of brain

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activity. However, the detected signal represents a low bandwidth superimposition of the activity of large neuron populations leading to low spatiotemporal resolutions which limit the investigations and study of circuit dynamics/neural circuitry at the cellular level (Poldrack and Farah, 2015). Second, optical imaging techniques, such as voltage and Ca2 þ sensitive dyes (Hamel et al., 2015) and optogenetics (Hochbaum et al., 2014). These techniques use reporter chemicals or proteins that convert changes in VM to an optical signal. Their advantages include outstanding spatial resolution, allowing signals in even the smallest neuronal structures to be resolved, and the possibility for simultaneous measurement from a wide range of spatial locations that have enabled mapping at a tissue as well as cellular level. However, these optical imaging techniques are limited by the penetration depth of light in the tissue, the toxicity of the fluorescent dyes, genetic incorporation of the proteins, and slow time-resolution in macroscopic three-dimensional (3D) imaging (Hamel et al., 2015). Third category involves devices that can interface directly with the cells or tissues and enable electrical recording and stimulation. The primary strength of this technique is its combination of high temporal resolution and sensitivity with higher signal-to-noise ratio (SNR) (Scanziani and Häusser, 2009). It has been the preferred means of analyzing brain activity due to the ability to capture a wide range of neural phenomena, from the spiking activity of individual neurons to the slower network oscillations of small populations (Llinás, 1988; Contreras, 2004; Assad et al., 2014). In this article, we will discuss the advancements made in the development of the nanomaterials-based platforms that lie in the third category. Over the years, several groups came up with different kinds of materials, platforms, and designs to improve the capabilities of these recording platforms. In principle, a device could be considered as an ideal recording system if it has the following capabilities (Scanziani and Häusser, 2009; Spira and Hai, 2013; Fattahi et al., 2014): (i) it is biocompatible and mimics properties of biological tissue to minimize any mechanical mismatch, immune response, scar tissue formation, and tissue disintegration, (ii) it supports integration with neurons to avoid shear motion at the cell-electrode interface, and remain functional for long period of time, (iii) it has low impedance to enable high SNR, (iv) it enables readout that covers the entire spectrum of membrane potential events, that is, APs, subthreshold potentials (EPSPs and IPSPs) and subthreshold membrane oscillations, (v) it provides high spatial and temporal resolution to allow recordings at single cell level with submillisecond resolution, and (vi) it simultaneously records/ stimulates multiple neurons to enable the study of network dynamics and signal propagation in a neuronal network. In addition to recording signals, seamless integration of these biomaterials with engineered or native 3D tissues is essential. This would not only enable monitoring but also simultaneously controlling the functional or electrical activity of the cells. This two-way interaction would provide a platform for both understanding and promoting treatment of disorders associated with neural circuitry and injuries in the cells. The development of such a platform would be an important advancement for the fusion of nanoelectronics with the field of regenerative engineering. Currently two common recording paradigms exist: passive devices, that is, microelectrode arrays (MEAs) and active devices, that is, field-effect transistors (FETs). Depending on the design of the device, the electrical activity can either be recorded extracellularly or intracellularly.

Extracellular Recordings The change in potential across the neuronal cell membrane gives rise to transmembrane current in the extracellular medium. The current contributions of neurons superimpose in the extracellular medium and generate a potential with respect to a reference potential. This potential gradient across the brain tissue results in an electrical field, which can be monitored by extracellularly placed electrodes. The characteristics of the detected electrical signal, such as amplitude and frequency, depend on the proportional contribution of the multiple sources, the properties of the brain area, and the distance of the recording electrode from the source. Depending on the electrode position, the recorded potentials can be categorized into four groups: First, electroencephalogram (EEG) signals, which are recorded from the scalp, are slow rhythms (5–300 mV, < 100 Hz); second, electrocorticogram (ECoG) signals, which are recorded from the cortical surface, are medium rhythms (0.01–5 mV, < 200 Hz); third, local field potentials (LFP), which are generated through the superimposition of all ionic processes, including action potentials, synaptic activity, and Ca2 þ spikes, are fast rhythms (few hundred mV–1 mV, < 200 Hz); and fourth, AP, which is measured at a single cell level, can be up to ca. 500 mV at 0.1–7 kHz when measured extracellularly (Fig. 2A,B) (Fattahi et al., 2014; Buzsáki et al., 2012). Extracellular recording techniques are less invasive and therefore, are compatible for long-term measurements, enabling simultaneous recording and stimulation of large populations of excitable cells for days and months without inflicting mechanical damage to the neuronal cell membrane. Reducing the dimension of electrical transducers to the micron and nanometer scale significantly improves the spatial resolution and enables high scalability and multifunctional platforms.

MEAs The flow of ions across the neuronal cell membrane gives rise to an extracellular current that spreads across the gap between the cell and the surface of the electrode called cleft with resistance Rcleft. This results in an extracellular potential difference, Vextra with respect to the bath solution. Vextra can be detected by MEAs since the extracellular potential change leads to a charge transfer at the electrode surface, resulting in a small capacitive/faradaic current through the microelectrode owing to the electrochemical impedance (Spira and Hai, 2013).

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Fig. 2 Detection of neuronal field potentials and single spike events. (A) Schematic of sensor position with corresponding neural signals. (B) Different categories of neural signals and their properties (Reprinted with permission from Fattahi, P., et al. (2014). A review of organic and inorganic biomaterials for neural interfaces. Advanced Materials 26 (12), 1846–1885. Copyright 2014 John Wiley & Sons).

MEA was first introduced in 1972 as a new platform for studying cultured cardiac myocytes (Thomas et al., 1972). Within few years, more platforms were developed to record from cultured neurons (Pine, 1980; Novak and Wheeler, 1986). Since then it has become a popular experimental platform for electrophysiological studies of neural networks. MEAs allow for the simultaneous recording of signals from multiple neurons at multiple locations, thereby improving the understanding of signal propagation in a neural network. Each electrode in an MEA can be individually addressed for signal amplification and processing. Also, the electrodes can be used to apply a voltage transient to depolarize the cell membrane and thus, stimulate electrical activity of neurons. These capabilities make MEAs a powerful tool to investigate the spatiotemporal neuronal network dynamics as well as allow active influence and control of neuronal activity both in vitro and in vivo (Wise et al., 2008).

Planar passive electrodes Typically, planar MEAs are based on recording sites made of metallic conductors, such as Au (Fig. 3A) (Oka et al., 1999) and Pt (Thiebaud et al., 1997). Sputtering of materials on metal electrodes, such as IrOx (Fig. 3B) (Gawad et al., 2009) and TiN (Janders et al., 1996), has also been demonstrated to enhance the capacitive properties of the electrodes. These platforms enable study of neurons at a network scale by enabling multisite recording and stimulation of dissociated cultured neurons or brain slices. However, being on a flat, stiff two-dimensional (2D) surface these are restricted to in vitro experiments involving cultured cells. To enable multisite recording and stimulation of neurons in in vivo, numerous attempts have been made to obtain recordings from the brain tissue using arrays of microwires (Schwartz, 2004; Nicolelis et al., 2003; Yuen and Agnew, 1995; Kralik et al., 2001). These electrodes consist of fine wires 20–50 mm in diameter that are mainly composed of metallic conductors, such as Pt, Au, Ir, stainless steel, and W. Spacing between wires is usually around 100–300 mm and is maintained either with polyethylene glycol or methyl methacrylate. These probes can enable recording from up to 2 mm below the brain tissue surface. However, postinsertion, a key issue associated with these probes in the change in the relative position of the electrode tip from the recording site, thus moving it away from the required site in the brain (Schwartz, 2004). To overcome the technical challenges associated with microwire technology, penetrating electrodes, such as Utah MEAs and Michigan arrays, have been developed (Kipke et al., 2008; Wise and Najafi, 1991; Rousche and Normann, 1998; Maynard et al., 1999; Hatsopoulos et al., 1998; Hochberg et al., 2006). Utah electrode arrays (Fig. 3C) (Hochberg et al., 2006) are composed of 100 Si needle electrodes with exposed tips of diameter 10–30 mm, and Michigan electrodes (Fig. 3D) (Kipke et al., 2008) comprise of multiple recording sites with the surface area of 100–400 mm2, arranged along a flattened shank of dimensions 15 mm thick, 3 mm long, and 90 mm wide at the base. While these electrodes allow localized recordings from specific brain areas, such as motor or visual cortex at high spatial resolution, their application in neuroscience is limited due to the highly invasive nature of the recordings. These probes are fabricated from rigid materials such as Si with different mechanical properties from the brain tissue, leading to mechanical mismatch (Schwartz, 2004; Polikov et al., 2005; Lagoa et al., 2006; Leach et al., 2010; Biran et al., 2005). Furthermore, the poor biocompatibility of these electrodes’ materials leads to chronic immune responses (Maynard et al., 1999; Biran et al., 2005; Szarowski et al., 2003; Turner et al., 1999; Edell et al., 1992; Biran et al., 2007). The mechanical mismatch and poor biocompatibility result in shear motion, glial scar formation, and neuron depletion at electrode-tissue interfaces leading to degradation of recording and stimulation capabilities over time (Kralik et al., 2001; Rousche and Normann, 1998; Maynard et al., 1999; Hatsopoulos et al., 1998; Polikov et al., 2005; Leach et al., 2010; Edell et al., 1992; Ludwig et al., 2006; Kim et al., 2009).

Flexible passive electrodes In implantable MEA technology, one of the main focuses lies in the development of stable electrode/tissue interfaces with reduced chronic inflammatory response. Some strategies for bio-integrated electronics have been incorporated to overcome challenges

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Fig. 3 Extracellular recordings using passive planar electrodes. (A) Au planar microelectrode array (MEA) with 64 microelectrodes of 50 mm  50 mm area each arranged in an 8  8 array with an interpolar distance of 150 mm. Scale bar: 150 mm (Reprinted with permission from Oka, H., et al. (1999). A new planar multielectrode array for extracellular recording: application to hippocampal acute slice. Journal of Neuroscience Methods 93 (1), 61–67. Copyright 1999 Elsevier). (B) IrOx planar MEA. (I) Top view of the full MEA assembled on a PCB. (II) Close-up of the substrate electrode array showing 16 electrode sites and the insulated leads. Each microelectrode is 30 mm in diameter, with a 100 mm spacing. Scale bars: (I) 6 mm, (II) 100 mm (Reprinted with permission from Gawad, S., et al. (2009). Substrate arrays of iridium oxide microelectrodes for in vitro neuronal interfacing. Frontiers in Neuroengineering 2, 1. Copyright 2009 Frontiers Media). (C) Utah array with 100 Si-based electrode sensors for neural recording. Individual electrodes are 1 mm long and spaced 400 mm apart, in a 10  10 grid. Scale bar: 1 mm (Reprinted with permission from Hochberg, L. R. et al. (2006). Neuronal ensemble control of prosthetic devices by a human with tetraplegia. Nature 442 (7099), 164–171. Copyright 2006 Macmillan Publishers). (D) Michigan probes. Photograph of (I) a four-shank Si neural probe having four electrode sites arranged near the tip, each terminated in a bond pad at the tab, (II) four different types of sites layouts for specialized interfaces, and (III) modular 128-site, three-dimensional array made from several multishank planar arrays (Reprinted with permission from Kipke, D. R. et al. (2008). Advanced neurotechnologies for chronic neural interfaces: New horizons and clinical opportunities. The Journal of Neuroscience 28 (46), 11830–11838. Copyright 2008 Society for Neuroscience).

associated with the mechanical mismatch between the hard, planar surfaces of the recording platform and the soft, curvilinear nature of biological systems. Silk-based MEAs One such strategy to overcome the challenges associated with rigid substrates relies on the fabrication of ultrathin Au electrode array supported by silk fibroin substrate which is highly flexible, biocompatible, biodegradable, transparent, and mechanically robust (Fig. 4A-I) (Kim et al., 2010a). Mounting such devices on tissue and then allowing the silk to dissolve and resorb initiates a spontaneous, conformal wrapping process driven by capillary forces at the biotic/abiotic interface (Fig. 4A-II). No evidence of immune response or inflammation was observed even after 4 weeks enabling long-term stable integration with the brain tissue. The high degree of conformal contact with the neural tissue enables recording of field potentials with good signal amplitude and high SNR (Fig. 4A-III). However, the spatial resolution is compromised due to the large electrode size of 500 mm  500 mm. And, scaling down the dimensions of the microelectrode while maintaining a high enough SNR has been a major challenge. Graphene-based MEAs To better decode the functions of individual circuit elements, it is important to achieve recording from a network level to a single cell level at high temporal resolution. Performing electrophysiology and optical imaging simultaneously could leverage the temporal and spatial resolution advantages of both techniques. However, metal-based MEAs that are commonly used for recording neural activity are opaque in nature because of which they generate optical shadows and are prone to producing light-induced artifacts in the recordings (Kuzum et al., 2014). Using a transparent material as electrodes can enable simultaneous optical and electrophysiology studies. Indium tin oxide (ITO), with transmittance of  80%, can be used for transparent neural electrode array. However, ITO does not have a flat transmittance across the visible spectrum, and the brittleness of ITO may limit the conformability of the device to the brain surface (Kuzum et al., 2014). Graphene (Novoselov et al., 2004; Geim, 2009), a one-atom-thick 2D honeycombed arrangement of sp2 hybridized carbon lattice, has been explored in as a building block in bioelectronics owing to its transparency, electrical and thermal conductivity, flexibility, and biocompatibility (Kuzum et al., 2014; Geim, 2009; Park et al., 2016; Park et al., 2014; Rastogi et al., 2017).

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Fig. 4 Extracellular recordings using passive flexible electrodes. (A) MEA on biodegradable silk. Image of electrode arrays (I) wrapped onto a glass hemisphere and (II) on a feline brain. Scale bars: 2 mm (Reprinted with permission from Kim, D.-H., et al. (2010). Dissolvable films of silk fibroin for ultrathin conformal bio-integrated electronics. Nature Materials 9 (6), 511–517. Copyright 2010 Macmillan Publishers). (B) Graphene-based MEA. (I) Schematic illustration of a flexible graphene neural electrode array. (II) Photograph of a 50  50 mm2 graphene electrode and 500  500 mm2 Au electrode placed on the cortical surface of the left and right hemispheres, respectively. The inset shows the flexibility of the electrodes. (III) Interictallike spiking activity recorded by doped-graphene (black) and Au electrodes (red) (Reprinted with permission from Kuzum, D., et al. (2014). Transparent and flexible low noise graphene electrodes for simultaneous electrophysiology and neuroimaging. Nature Communications, 5. Copyright 2014 Macmillan Publishers). (C) Stretchable MEA (SMEA). (I) Top view of the prototype SMEA array with 11 recording electrodes and 1 reference electrode. (II) Hippocampal slice placed on the SMEA; the parallel horizontal lines are part of a mesh that holds down the tissue. (III) Spontaneous and (IV) evoked field potentials recorded from the hippocampal slice. Scale bars: (I) 8 mm, (II) 0.5 mm (Reprinted with permission from Graudejus, O., et al. (2009). Characterization of an elastically stretchable microelectrode array and its application to neural field potential recordings. Journal of the Electrochemical Society 156 (6), P85–P94. Copyright 2009 The Electrochemical Society).

Graphene-based flexible and transparent electrodes enable simultaneous optical and electrophysiology studies as demonstrated in the reference (Kuzum et al., 2014). To fabricate such electrodes, the chemical vapor deposition-synthesized graphene was transferred on to flexible polyimide (PI) films and patterned to obtain 50 mm  50 mm active electrode area (Fig. 4B-I). The graphene electrode was then compared with Au electrode by recording from the cortical surface of the left and right hemispheres simultaneously (Fig. 4B-II). When compared to Au electrodes, graphene electrodes showed five- to sixfold improvement in SNR and 100-fold reduction in electrical interference noise (Fig. 4B-III). The in vivo neural recording experiments performed using graphene electrodes in conjunction with optical recordings demonstrated that graphene electrodes can detect interictal and ictal activity and fast population spikes with durations < 5 ms, and the Ca2 þ transients within the electrode area captured by the confocal microscope showed an increase in Ca2 þ signal coinciding with the interictal-like event recorded by the graphene electrode. In summary, the temporal resolution of the graphene electrode recordings enabled detection of high-frequency population discharges, which could not be resolved by the Ca2 þ fluorescence responses. In contrast, Ca2 þ imaging responses were able to capture complex network contributions of individual neurons, which were not evident in the electrical recordings using graphene electrodes. Therefore, combination of both techniques revealed temporal and spatial characteristics of high-frequency bursting activity and synaptic

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potentials in hippocampal slices with high precision. In addition to low noise and transparency, graphene-based electrodes were shown to be highly robust, stable for long term, and resistive to corrosion, thus enabling long-term stable recordings from brain tissues at high temporal resolution. However, the major drawback that still remains is the inability to perform long-term recordings at high spatial resolution from the deep regions in the brain due to the limitations of optical dyes. Stretchable MEAs Fabricating MEAs on PI-based substrate (Kuzum et al., 2014) reduce the problems associated with mechanical mismatch of the MEA materials with the underlying soft neural tissue as the elastic modulus of PI is ca. 2.8 GPa which is much lower than the Si-based rigid electrodes with elastic modulus of > 100 GPa. However, it is still several orders of magnitude higher than the neural tissue which has an elastic modulus of < 1 kPa. To further improve on the mechanical mismatch between the electrodes and the underlying tissue, stretchable MEAs (SMEAs) have been fabricated on soft, elastically stretchable, silicone membranes with an elastic modulus of  1 MPa (Fig. 4C-I) (Graudejus et al., 2009). The purpose behind fabrication of stretchable array was to enable the study of traumatic brain injury (TBI) which requires stable integration of MEAs with the tissue and allow direct comparison of electrical activity from the same location on the tissue before and after traumatic injury (or a similar mechanical stimulation). To develop SMEA, < 100 mm wide Au electrodes were fabricated on PDMS that could be stretched repeatedly and reproducibly by > 30% while remaining electrically conducting. The recording sites of SMEA were electroplated with Pt to reduce the impedance of the electrode/electrolyte interface. This leads to reduction in the noise during recording and the stimulus artifacts during stimulation. The SMEAs were then used to record extracellular spontaneous and evoked electrophysiological activity of neurons in an organotypic hippocampal slice culture while the electrodes were under biaxial strain (Fig. 4C-II). The SMEAs enable recordings of both spontaneous and evoked responses (Fig. 4C-III, IV). Also, the electrical attributes of the electrodes did not change even by severe stretching (up to 30% strain) of the substrate, thus enabling stable and functional integration without any motion at the electrode-tissue interface. Even though MEAs have been modified to achieve stable integration with neural tissues at high temporal resolutions, nonetheless, achieving high spatial resolution still remains a challenge. Ideally, for high spatial resolution, the site of a microelectrode should have a small geometric area to enable communication with individual neurons. And for high sensitivity, the device should have a low impedance and high injection charge density during recording and stimulation (Kovacs et al., 1994; Paik and Park, 2003; Cogan, 2008). However, in the case of planar MEAs, there is a trade-off between spatial resolution and sensitivity since decreasing the geometric area of a recording site causes an increase in the impedance and a decrease in the capacity of the injection charge density of neural microelectrodes (Scanziani and Häusser, 2009; Kovacs et al., 1994; Drake et al., 1988; Robinson, 1968; Frey et al., 2009). Reducing the impedance by surface modification of the electrodes has led to an improvement in the performance of the MEAs. Pt black has been used to coat the metal electrodes since its porous structure is effective for impedance reduction (Novak and Wheeler, 1986; Chang et al., 2000; Mathieson et al., 2004; Maher et al., 1999). Various Au nanostructures have also been reported that involve deposition of nanoflake (Kim et al., 2010b), nanograin (Kim et al., 2013), and fuzzy Au (Cui and Martin, 2003) structures. The rough surface formed due to the nanostructures lead to increase in the net surface area, thus reducing the impedance of the microelectrodes. To increase the effective surface area, the microelectrodes have also been modified with carbon nanotubes (CNTs) by direct synthesis on the electrode surface (Gabay et al., 2007), electroplating (Suzuki et al., 2013; Keefer et al., 2008), and microcontact printing (Fuchsberger et al., 2011) methods. Application of conductive polymers to MEA such as poly(3,4ethylenedioxythiophene) (PEDOT) and polypyrrole has also been demonstrated to reduce the impedance of the electrodes (Cui et al., 2001; Ludwig et al., 2011; Venkatraman et al., 2011). In addition, the use of PEDOT/CNT composites (Zhou et al., 2013) and polypyrrole/graphene oxide composites (Deng et al., 2011) has shown to increase stability and charge injection capacity of the microelectrodes (> 200 mCcm
2) by two orders of magnitude when compared with planar Pt electrodes (1.4 mCcm
2). Even though the surface modifications improve the sensitivity of the MEAs by reducing the surface impedance and increasing the SNR, the electrodes geometry of tens of microns limits the capability for recording signals from subcellular structures, such as single axon and dendrites (Scanziani and Häusser, 2009). In addition, due to the large-size geometry of MEAs, the recorded signals are often comprised of average signals from multiple neurons and thus, complicating the interpretation of recorded signals (Spira and Hai, 2013).

FETs FET is a semiconductor device that exhibits a conductance change in response to variations in the charge or potential at the surface of the channel region. A standard planar FET consists of three terminals: (i) source, through which the carriers enter the channel; (ii) drain, through which the carriers leave the channel; and (iii) gate, the terminal that modulates the channel conductivity. The gate electrode is capacitively coupled to the semiconductor channel by an insulating oxide layer. When there is no gate voltage, the FET is in its off state as there is no flow of current. When gate voltage exceeds a threshold voltage, charge carriers (e.g., holes for p-Si and electrons for n-Si) are induced at the semiconductor–oxide interface, and the potential barrier of the channel drops, resulting in the current flow. Therefore, the conductance of the semiconductor channel between the source and drain regions can be switched from off to on and modulated in the on-state by the potential at the gate electrode. When a FET is integrated with a neuron, the potential change across the neuronal membrane acts as a gate and modulates the conductivity of the transistor channel by changing the charge carrier concentration. This conductance change can be easily detected by recording the net current through the FET channel at a constant bias voltage. The detected current change depends on the

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transconductance of the device, and thus, is independent of gate impedance unlike MEAs. Development of a FET device for neurophysiological measurements was first demonstrated in 1970s (Bergveld, 1970; Bergveld et al., 1976). Since then, many studies have demonstrated the use of planar FETs to enable recordings from electrogenic cells (Weis et al., 1996; Patolsky et al., 2006; Lambacher et al., 2004; Fromherz, 2002; Fromherz and Offenhausser, 1991; Voelker and Fromherz, 2005; Merz and Fromherz, 2005; Besl and Fromherz, 2002; Vassanelli and Fromherz, 1999; Vassanelli and Fromherz, 1998; Offenhäusser and Knoll, 2001; Offenhäusser et al., 1997; Ingebrandt et al., 2005). Some of the recent advances in the development of FETs for cellular recordings are discussed later.

2D FETs Since FETs are independent of gate impedance, it makes it possible to go down to nanoscale, thus making them one of the most promising nanoscale bioelectronics devices for recording with subcellular resolution (Patolsky et al., 2006; Duan et al., 2012; Cohen-Karni et al., 2010; Tian et al., 2012; Zhang and Lieber, 2016; Duan et al., 2013; Gao et al., 2012; Lieber, 2011; Tian and Lieber, 2013; Dai et al., 2016; Tian et al., 2010). Graphene and nanowire (NW) FET To achieve single cell resolution with high SNR, graphene and SiNW FETs have been interfaced with cultured cardiomyocytes (Fig. 5A-I,II) (Cohen-Karni et al., 2010). The advantage of using NWs in FETs lies in the 1D nanoscale morphology of NWs that enhances the net surface area of the device active region. High surface area leads to enhanced interactions of NW FETs with the cell membrane and thus, high sensitivity when compared with their planar counterparts. On the other hand, graphene FETs have an advantage due to their ambipolar behavior that enables both n- and p-type recording with the same device. This characteristic was demonstrated by signal shape flip of recorded extracellular potentials across the Dirac point. In addition, when compared with other planar structures, one atom thick graphene FETs demonstrate better performance by yielding well-defined extracellular signals with SNR > 4, exceeding typical values for other planar devices (Cohen-Karni et al., 2010). However, the measured signal is dependent on the size of the graphene flake. A large graphene FET with active channel of 20.8 mm  9.8 mm recorded signals with peak-to-peak width of 1.31  0.04 ms (Fig. 5A-III). Whereas signals recorded from a much smaller graphene FET with active channel dimensions of 2.4 mm  3.4 mm yielded peak-to-peak width of 0.73  0.04 ms, which is almost a factor of 2 smaller than that obtained from the larger device (Fig. 5A-IV). Signals recorded from a 0.07 mm2 active area SiNW-FET had a peakto-peak width of 0.76  0.04 ms, which is similar to the value for the smaller graphene device (Fig. 5A-IV). These results indicate that the signals recorded with the larger graphene device represent an average of the extracellular potential from sufficiently distinct sources of the beating cell. Although NW FET device yielded similar peak-to-peak widths as the  100 bigger area graphene FET, NW devices still have an advantage for spatially resolved multiplexed measurements in high-density device arrays.

3D FETs When compared with cells cultured in 2D, 3D cellular networks mimic more to an in vivo like environment: spherical cell morphology, high cell-to-cell interaction, and neurite outgrowth in all directions (Irons et al., 2008). Coupling recording platforms with 3D networks would represent a powerful in vitro model capable of better emulating in vivo physiology. Coupling of electronics with the 3D tissues using flexible (Kim et al., 2010a; Kuzum et al., 2014) and stretchable planar devices (Graudejus et al., 2009) that conform to natural tissue surfaces was recently reported. However, these planar devices have been used to probe electrical activities near surfaces of the tissue and lack seamless 3D integration of the electronics with the tissue. NW nanoelectronic scaffolds (NanoES) To enable recordings from cells in 3D, nanoelectronics and synthetic tissues have been integrated with macroporous nanoelectronic scaffolds (nanoES) (Tian et al., 2012). The aims behind building such recording platform involved: developing macroporous electronic structures to enable 3D interpenetration with biomaterials; nanometer to micrometer scale features of electronic network comparable to biomaterial scaffolds; and 3D interconnected electronic network with mechanical properties similar to biomaterials. The electronic scaffold was based on SiNW FETs to achieve subcellular resolutions. The fabrication of the nanoES involved the following steps: first, chemically synthesized SiNWs were deposited, and then NW FET devices were lithographically patterned to create two types of nanoES designsdReticular and Mesh. Reticular nanoES were made to mimic the size scale and morphology of submicron ECM features, such as the fibrous meshwork of brain ECM. Open mesh nanoES were made with a regular structure, similar to the ECM of the ventricular myocardium. These scaffolds were fabricated on sacrificial layers, which were subsequently removed, yielding freestanding macroporous scaffolds (Fig. 5B-I). The nanoES were designed to be 3D and mimic ECM structures, having nanometer to micrometer features with high (> 99%) porosity, and to be highly flexible and biocompatible. NanoES were then combined with synthetic or natural macroporous ECMs providing ECMs with electrical sensory function and nanoES with biochemical environments suitable for tissue culture. Finally, cells were cultured within the nanoES to yield 3D hybrid nanoelectronics tissue constructs (Fig. 5B-II). The emphasis on the nanoscale and biomimetic bottom-up pathway allows minimally invasive integration of electronic devices with cells and ECM components at the subcellular level in 3D. Simultaneous recordings from four NW FETs with separations up to 6.8 mm in a nanoES-cardiac construct demonstrated multiplexed sensing of a coherently beating cardiac patch (Fig. 5B-III). The data revealed submillisecond time resolution with regularly spaced spikes with a frequency of  1 Hz, calibrated potential change of  2–3 mV and  2 ms width, and SNR of  3. This platform enabled good integration with engineered tissue allowing access to the cellular response with single-shot submillisecond time resolution. However, the size and the design of the device precluded 3D tissue mapping.

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Fig. 5 Extracellular recordings using field-effect transistors (FETs). (A) Si nanowire (NW) and graphene (Gra)-FETs. (I) Schematic illustrating the cardiomyocyte cell interfaced with Gra-FET and SiNW-FET devices. (II) Optical microscope image of PDMS/cells interfaced with devices. Whitedashed line, red and blue arrows represent graphene flake, Gra-FET and SiNW-FET devices, respectively. (III) Electrical signal recorded from the GraFET (active channel of 20.8 mm  9.8 mm). (IV) Electrical signals recorded from the Gra-FET (red, active channel of 2.4 mm  3.4 mm) and SiNW-FET (blue, active channel area of 0.07 mm2). Scale bar: (II) 13.6 mm (Reprinted with permission from Cohen-Karni, T., et al. (2010). Graphene and nanowire transistors for cellular interfaces and electrical recording. Nano letters 10 (3), 1098–1102. Copyright 2010 American Chemical Society). (B) NW nanoelectronic scaffolds (NanoES). (I) Device fabrication schematics of (i) reticular NW FET devices and (ii) mesh NW FET devices. Light blue, dark blue, green, and yellow represent SiO2 substrates, nickel sacrificial layers, nanoES, and individual NW FETs, respectively. (II) 3D reconstructed confocal images (x: 127 mm; y: 127 mm; z: 68 mm) of rat hippocampal neurons after a 2-week culture in Matrigel on reticular nanoES. Red (Alexa Fluor 546) and yellow (rhodamine 6G) represent neuronal b-tubulin and epoxy ribbons, respectively. The white arrow highlights a neurite passing through a ring-like structure supporting a NW FET. (III) Multiplex electrical recording from four NW FETs in a mesh nanoES (Reprinted with permission from Tian, B., et al. (2012). Macroporous nanowire nanoelectronic scaffolds for synthetic tissues. Nature Materials 11 (11), 986–994. Copyright 2012 Macmillan Publishers). (C) 3D-folded scaffold integrated NW FET. Schematics of (I) freestanding macroporous nanoelectronic scaffold with NW FET arrays (red dots); Inset: one NW FET, (II) folded 3D freestanding scaffolds with four layers of individually addressable FET sensors, and (III) nanoelectronic scaffold/cardiac tissue resulting from the culturing of cardiac cells within the 3D folded scaffold; Inset: the nanoelectronic sensors (blue circles) innervate the 3D cell network. (IV) Simultaneous traces recorded from 16 sensors in the top layer (L1) of the nanoelectronics–cardiac tissue. The (x,y) coordinates correspond to each FET sensor in the 4  4 array (Reprinted with permission from Dai, X., et al. (2016). Threedimensional mapping and regulation of action potential propagation in nanoelectronics-innervated tissues. Nature Nanotechnology. Copyright 2016 Macmillan Publishers).

3D folded scaffold integrated NW FET Real-time 3D spatiotemporal mapping of APs in engineered cardiac tissues with submillisecond temporal resolution was achieved using nanoelectronics that mimic tissue scaffolds (Dai et al., 2016). In this work, 3D macroporous nanoelectronic networks were fabricated using high-density SiNW FETs, with miniaturized features designed to achieve ca. 2 mm dimensions and ca. 0.29– 2.8  10
16 N m
2 bending stiffness values that are comparable to conventional electro spun fiber tissue scaffolds (Fig. 5C-I). The freestanding macroporous nanoelectronic networks were then folded into 3D scaffolds (Fig. 5C-II) and neonatal rat cardiac cells were cultured within the scaffolds to yield nanoelectronics cardiac tissues (Fig. 5C-III). Extracellular cardiac AP signals recorded from 4  4 FET sensors in a single layer across a 5  5 mm2 domain show a synchronized beating rate of 1.8 Hz, an amplitude of 1–2 mV, and a peak width of  1 ms from all 16 channels. Higher-resolution examination of these peaks reveals a submillisecond time latency between any given set of AP peaks recorded by the 16 FET sensors, where the intrinsic temporal resolution of the device is 0.01–0.05 ms (Fig. 5C-IV). This platform enables 3D electrophysiology mapping at high temporal resolution.

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Integrating the FET sensors in 3D scaffolds did allow recording from cells that mimic morphology as in vivo models, however these platforms are incapable of achieving measurements from intact brain tissue which involves much more complicated neuronal circuitry when compared with the 3D tissue engineered for the earlier-mentioned devices.

Injectable Electronics To address the challenge of integrating electronics with the 3D tissue in vivo, the syringe injection of macroporous mesh electronics in the brain tissue was recently demonstrated (Liu et al., 2015). The seamless and minimally invasive delivery of macroporous NW nanoelectronic probes into tissue enables long-term and intimate electronics-tissue interface due to their flexibility, nonplanar structure, and tissue-scaffold mimicking properties. The syringe injection concept involves (i) loading the macroporous mesh electronics into a syringe needle, (ii) inserting the needle into the biological tissue such as the brain, and (iii) injecting the mesh nanoelectronics and delivering the input/output region of the mesh outside the material for subsequent bonding and measurements (Fig. 6A). To enable injection of the electronic mesh into the tissue, the structure of mesh network was designed with transverse bending stiffness values sufficiently small and with flexibility closer to that of tissue, to allow the mesh electronics to roll up and smoothly go through the needle during injection. To demonstrate the capabilities of the mesh electronics, it was injected stereotaxically into the lateral ventricle and hippocampus of live rodents for in vivo brain activity recordings (Fig. 6B). The mesh was able to record signals with modulation amplitude of 200–400 mV and dominant modulation frequency of 1–4 Hz which are characteristic of d- wave LFPs (Fig. 6C). The recorded spike had an average duration of  2 ms and peak-to-peak amplitude of  70 mV which is a characteristic of single-unit APs (Fig. 6D). These results suggested substantial promise of using injectable electronics to mobilize and monitor neural networks in in vivo models. Those examples illustrate how extracellular recording platforms have been modified and improved over the years in order to enable long-term measurements as well as simultaneous recording and stimulating large populations of excitable cells for days and months at high spatiotemporal resolution. However, the major limitation of these technologies is the extraction of less detailed information from the excitable cells and their network since these electrodes record spikes with 2–3 orders of magnitude lower than the true AP across the cell membrane. In addition, the extracellular recordings are blind to the subthreshold potentials such as EPSPs, IPSPs, and membrane oscillations. These potentials play a significant role in the functioning of neuronal circuit, and a great deal of neuroplasticity is associated with changes in the amplitude of synaptic potentials. Hence, it is crucial to enable recordings of the entire dynamic range of transmembrane voltage changes to better understand the functioning and connectivity of neuronal circuit (Spira and Hai, 2013).

Fig. 6 Injectable electronics. (A) Schematics showing the (I) insertion and (II) retraction of the needle and (III) placement of the mesh electronics in the cavity. (B) Optical image of the stereotaxic injection of mesh electronics into the brain of an anesthetized mouse. Scale bar: 5 mm. (C) (I) A 16channel recording with the mesh electronics following injection into the brain of a live mouse. (II) Superimposed single-unit neural recordings from one channel (Reprinted with permission from Liu, J., et al. (2015). Syringe-injectable electronics. Nature Nanotechnology 10 (7), 629–636. Copyright 2015 Macmillan Publishers).

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Intracellular Recordings Unlike extracellular recording, intracellular recording can enable accurate readout of the entire dynamic range of transmembrane voltage changes with little signal distortion and therefore, can monitor subthreshold events and very slow changes of transmembrane potentials due to the synaptic interactions (Spira and Hai, 2013; Purves, 1981). As a result, intracellular recording can provide faithful amplitude and shape information of APs in individual cells, as well as help in the investigations of neuronal circuits. The glass micropipette patch-clamp technique developed in the late 1970s has been the most widely used approach to record intracellularly (Molleman, 2003). This technique monitors the voltage across the cell membrane by either creating a tight seal between the cell membrane and micropipette, referred to as cell-attached patch clamp, or accessing the cell interior by puncturing the cell membrane using a glass micropipette, referred to as whole-cell patch clamp (Molleman, 2003). However, there are some limitations associated with these patch clamp techniques (Spira and Hai, 2013; Edwards et al., 1989). The patch-clamp measurement involves a bulky setup, including 3D manipulators for precise and delicate interfacing between a micropipette and individual cells. This impedes the development of multiplexed intracellular recording (Spira and Hai, 2013). In addition, the whole-cell patch technique leads to mixing of cell cytosol and the exogenous solution filling the pipette, which results in an irreversible change to the cells, making long-term measurements difficult (Edwards et al., 1989). To address these limitations, substantial effort has been placed on the development of micro- and nanoscale 3D electrode arrays to enable long-term, multiplexed intracellular recording and stimulation from cells in cellular networks.

MEA When neurons are cultured on a recording device, a gap or cleft exists between the electrode and cell membrane where this cleft determines the seal resistance. Enhancing the adhesion between electrode and cell surface leads to an increase in electrode-cell seal resistance. A number of approaches such as surface modification, substrate modulation, and electrode shape control have been focused on to increase the neuron-microelectrode electrical coupling coefficient and decrease the cleft size, thus increasing the seal resistance as a means to enhance signals and achieve intracellular-like recordings (Spira and Hai, 2013; Hai et al., 2010; Hai and Spira, 2012; Hai et al., 2009; Xie et al., 2012; Robinson et al., 2012).

Functionalized gold-spine electrode (FGSE) An enhancement in tight coupling of the electrodes with the cell membrane was achieved by fabrication of functionalized microscale 3D mushroom-shaped Au protrusions, also referred to as functionalized gold-spine electrode (FGSE) (Fig. 7A-I) (Spira and Hai, 2013; Hai et al., 2010; Hai et al., 2009). These Au electrodes were fabricated with a stem height of approximately 1 mm, a stem diameter of approximately 800 nm, and a mushroom-shaped cap of 1.8–3 mm in diameter (Fig. 7A-II). The shape and the dimension of the electrodes were selected to mimic the geometry and dimensions of dendritic spines. The electrodes were chemically functionalized by multiple Arg-Gly-Asp (RGD) peptides and engulfment-promoting peptide (EPP) to facilitate the engulfment by the cultured neurons and other types of cells, as observed from the transmission electron microscopic analysis of the interfaces formed between cells and mushroom-like microelectrodes (Fig. 7A-III). The use of an FGSE as the sensing electrode led to the activation of endocytotic-like mechanisms in the neurons that increased the neuron-microelectrode electrical coupling coefficient to 50% as compared to approximately 0.1% as recorded by a planar extracellular MEA. The generation of high Rseal ( 100 MU) between the neuronal membrane and the engulfed FGSE, and the increased junctional membrane conductance led to intracellular-like recording enabling successful monitoring of APs in the range of tens of mV (as compared to 100 mV recorded by planar Au electrodes) as well as subthreshold synaptic potentials (Fig. 7A-IV).

Vertical Si nanowire (SiNW) electrode array (VNEA) A scalable multiplexed intracellular nanoelectrode platform based on arrays of vertical nanopillars (VNEA) was recently demonstrated (Robinson et al., 2012). In this work, 3  3 arrays of 9 nanopillars, 150 nm in diameter, 3 mm in height at 2 mm pitch were grown on 16 sensing pads (Fig. 7B-I). The geometry of each NW array (a 4 mm square) was chosen to be smaller than the size of a typical neuronal cell body so as to increase the probability of single-neuron coupling. The entire device is fabricated from a Si-on-insulator (SOI) substrate so that each pad could be independently addressed electrically. Each NW in the array is composed of a doped Si core, a Ti/Au tip, and an insulating SiO2 shell. The Si core and metal tip provide electrical access to the interior of the cell, and the glass shell plays the dual role of preventing current leakage through the NW sidewalls and serving as a material with which to make tight seals to the cell membrane. Because of its planar integrated geometry, the VNEA is well suited to studying in vitro dissociated neuronal circuits and ex vivo preparations, such as brain slices. When embryonic rat cortical neurons were cultured on the VNEAs, 50% of the VNEAs spontaneously penetrated through the plasma membrane of the cells and created a high seal resistance of 100–500 MU (Fig. 7B-II). In cases where spontaneous penetration of the membrane were not evident, an electroporating pulse (approximately  3 V, 100 ms) was applied to permeabilize the cell membrane and promote NW penetration. Consistent with the intracellular positioning of the VNEA, all recorded APs were positive monophasic and intracellular-like with amplitude of ca. 5 mV (Fig. 7B-III). The nanoscale structure of the electrodes enables recording and stimulation of individual neurons. Even though a single sensing pad carries multiple nanopillars,

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Fig. 7 Intracellular-like recordings using passive electrodes. (A) Functionalized gold-spine electrode (FGSE). (I) Confocal microscopy image of three Aplysia neurons cultured on an FGSE array. Green lines represent the conducting lines. (II) Scanning electron microscopic image of FGSE fabricated on a glass surface. (III) Transmission electron microscopic image of electrode/cell interface. (IV) Recording of Aplysia LUQ neuron action potentials using FGSE (blue) and patch pipette (red). Scale bars: (I) 50 mm, (II) 500 nm, (III) 2 mm (Reprinted with permission from Hai, A., J. Shappir, and M.E. Spira (2010). In-cell recordings by extracellular microelectrodes. Nature Methods 7 (3), 200–202 and Hai, A., et al. (2009). Spine-shaped gold protrusions improve the adherence and electrical coupling of neurons with the surface of microelectronic devices. Journal of the Royal Society Interface rsif20090087. Copyright 2009, 2010 Macmillan Publishers). (B) Vertical Si nanowire (SiNW) electrode array (VNEA). SEM image of (I) nine SiNW array with metal-coated tips (gray) and insulating silicon oxide (blue), and (II) rat cortical cell (3 DIV) on top of a VNEA pad. Inset represents the cell/electrode interface. (III) Action potentials stimulated and recorded using a patch pipette (blue) and recorded by the NW array (magenta). Scale bars: (I) 10 mm, (II) 2.5 mm (Reprinted with permission from Robinson, J. T., et al. (2012). Vertical nanowire electrode arrays as a scalable platform for intracellular interfacing to neuronal circuits. Nature Nanotechnology 7 (3), 180–184. Copyright 2012 Macmillan Publishers). (C) Pt nanopillars. (I) SEM image of an array of five vertical nanopillar electrodes with dimensions of 1.5 mm in height and 150 nm in width; Inset: schematic of a nanopillar electrode. (II) SEM image of HL-1 cell/nanopillar electrode interface. Electrical recording by Pt nanopillars (III) before and (IV) after electroporation. Scale bars: (I) 2 mm, (II) 200 nm (Reprinted with permission from Xie, C., et al. (2012). Intracellular recording of action potentials by nanopillar electroporation. Nature Nanotechnology 7 (3), 185–190. Copyright 2012 Macmillan Publishers).

and a number of them penetrate the plasma membrane, the electrodes impedance is too high to enable recording of subthreshold potentials.

Pt nanopillars In parallel, the use of vertical Pt nanopillar electrodes (150 nm in diameter and 1–2 mm in height) (Fig. 7C-I) for intracellular-like recording from cultured HL-1 cardiomyocytes (Fig. 7C-II) was also demonstrated (Xie et al., 2012). The aim was to increase the Rseal by forming tight junctions between cell membranes and vertically aligned nanopillar electrodes. The study showed that transient electroporation using the nanopillars leads to a change from biphasic extracellular signature (Fig. 7C-III) to two orders of magnitude higher monophasic intracellular-like signal signature (Fig. 7C-IV). The drastically improved signal is attributed to the reduction in impedance between the electrode and the cell interior due to the formation of nanopores in the cell membrane by electroporation. However, after electroporation, the amplitude of the recorded AP signals continuously decays and reduces to 30% of its original amplitude in 120 s due to the subsequent self-sealing of electroporation-generated pores in the cell membrane. This transient nature of the electrical coupling and the attenuation of the APs indicate that electroporation could not be used to obtain long-term stable recordings from the cells. Modifying the shape of the passive electrodes, as discussed earlier, does enable recording of spikes with 1–2 orders of magnitude higher than their planar counterparts, thus providing more information about the electrical activity. However, the high impedance

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associated with the small size of these electrodes prevents the recording of the full-amplitude AP which is an order of magnitude higher than what these electrodes can record.

FETs NW FETs have a potential to be ideal for intracellular probes because (i) their nanoscale dimensions allow for minimally invasive insertion and (ii) the FET device can enable true intracellular recording with high SNR since the size of the device is not dependent on interfacial impedance in contrast to passive electrode techniques.

3D-kinked NW probes The overall size of nanoFETs is larger than the active FET component due to the source and drain electrical contacts. The necessity of having two contacts makes minimally invasive insertion of a nanoFET into cells challenging and less feasible. One of the methods to overcome the geometry-size constraint was achieved by developing nanoFETs based on kinked NWs (Tian et al., 2010). The method involved synthesis of SiNWs with two 120o kinks. Leveraging on the bottom-up synthesis approach, contacts (source/drain) to the nanoprobe were defined by in situ degenerately n-doping the SiNW arms. In addition, a nanoprobe was synthetically encoded close to the kinks site by changing the dopant to lightly n-doped. The SiNWs were transferred on SU8 supported on a sacrificial layer and mechanical supports (SU8) and metallization lines (Cr/Pd) were defined using electron beam lithography. The interfacial stress between materials was used to bend the probe upward off the surface leading to a freestanding nanoFET (Fig. 8A-I). The acuteangle kinked NW geometry and extended source/drain arms spatially separate the functional nanoscale FET from the bulky metallic interconnects, thus allowing better interface of the active region of the device with the cell with minimal interference from the source/drain contacts. An interesting feature of these nanoFET probes was modification with phospholipid bilayers to promote spontaneous cellular internalization without external forces. Contact of cultured cardiomyocyte cells to a 3D-kinked SiNW bioprobe showed three distinguishable recording stages during internalization: extracellular (Fig. 8A-II), extracellular to intracellular transition (Fig. 8A-III), and stead-state intracellular recordings (Fig. 8A-IV). The initial extracellular signals had an amplitude of 3–5 mV and a submillisecond width. Without applying external force, the initial extracellular signal gradually disappeared with a concomitant emergence of intracellular peaks featured with an opposite polarization, a much larger amplitude of  80 mV, and a longer duration of  200 ms. This study demonstrated for the first time intracellular electrical recording platform based on FET nanoprobes that are capable of recording true AP with high spatial and temporal resolution and high SNR. However, the kink configuration and device design places limits on the probe size and the potential for multiplexing.

Branched intracellular nanotube-FET (BIT-FET) Overcoming the challenge of multiplexing of the FET probes to achieve intracellular recordings was demonstrated by using branched intracellular nanotube FET (BIT-FET) (Duan et al., 2012). BIT-FET is a combination of a SiNW FET and an electrically insulating SiO2 nanotube. An NW FET is fabricated on the chip as the recording transducer and a vertical nonconductive nanotube is created directly on active channel region of the nanowire FET to interface the FET to the cell interior (cytosol) for intracellular recording (Fig. 8B-I). The fabrication of the BIT-FET probe involved combination of bottom-up synthesis/assembly and topdown lithographic patterning. The steps involved growth of Germanium NW (GeNW) branches on top of SiNWs using the Aunanocluster-catalyzed vapor–liquid–solid mechanism. The GeNWs were then coated with conformal thin SiO2 layer. Following selective removal of the topmost part of the SiO2 shell and etching off the Ge nanowire core, a vertical hollow SiO2 nanotube was created on the SiNW FET. To facilitate internalization of BIT-FET probes in cells for intracellular recording, the probes were modified with phospholipids. The capability of BIT-FET probes was demonstrated by recording intracellular signals from spontaneously beating cardiomyocyte cells cultured on the PDMS sheets. As shown in Fig. 8B-II, two BIT-FET devices can be interfaced with a single cardiomyocyte cell to achieve simultaneous two-channel recording from the same cell (Fig. 8B-III). The nanoscale dimensions of the probe along with the modification with the phospholipid led to easy access to the cell interior, leading to intracellular measurements with full amplitude of 75–100 mV. This suggests the potential to implement large arrays of BIT-FET probes for multiplexed intracellular recording in cellular networks.

Active Si nanotube transistor (ANTT) Complementing the BIT-FET probe, needle-shaped intracellular nanoprobes based on an active Si nanotube transistor (ANTT) were demonstrated (Gao et al., 2012). The ANTT probe consists of a single semiconductor nanotube. The source/drain contacts of the nanotube transistor are defined on one end of the Si nanotube and insulated from external solution so that the active channel regions of nanotube FET can sense the transmembrane potential changes through the solution filling the nanotube. Therefore, if the free end of an ANTT probe is inserted into the interior of an electrogenic cell, the time-dependent changes associated with an AP spike will give rise to time-varying conductance signal that maps the intracellular AP (Fig. 8C-I). The fabrication of ANTT probes involved synthesis of GeNWs followed by p-type Si shell deposition, and contact printing of these Ge/Si NWs onto a prebaked SU-8 layer, which was initially deposited on a sacrificial Ni sacrificial layer. The next steps involved registration of positions of Ge/Si NWs, definition of the bottom SU-8 layer, and definition of source/drain metal contacts followed by the top SU-8 passivation layer (Fig. 8C-II). Finally, etching the Ni sacrificial layer and Ge core of the Ge/Si NW yielded the p-type Si ANTT probe (Fig. 8CIII,IV). The capability of ANTT probes was demonstrated by recording intracellular signals from spontaneously beating cardiomyocytes cultured on PDMS sheets. To enable better seal between the ANTT probes and the cells for intracellular recording, the probes

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Fig. 8 Intracellular recordings using FETs. (A) 3D-kinked NW probes. (I) Schematics of cellular recording from the cardiomyocyte monolayer on PDMS (left) and highlight of extracellular (middle) and intracellular (right) NW/cell interfaces. Inset: SEM image of the kinked NW device. Purple lines represent the cell membrane and NW lipid coatings. Plots corresponding to (II) extracellular, (III) extracellular to intracellular transition, and (IV) steady-state intracellular recording. Green and pink stars denote the peak positions of intracellular and extracellular signal components, respectively. Scale bar: (I) 5 mm (Reprinted with permission from Tian, B., et al. (2010). Three-dimensional, flexible nanoscale field-effect transistors as localized bioprobes. Science 329 (5993), 830–834. Copyright 2010 American Association for the Advancement of Science). (B) Branched intracellular nanotube-FET (BIT-FET). (I) Schematic of a cell coupled to a BIT-FET. S and D indicate source and drain electrodes, respectively. (II) Differential interference contrast (DIC) image of two BIT-FET devices (positions marked with dots) coupled to a single cardiomyocyte cell. Yellow-dashed line represents the cell boundary. (III) Representative trace of stable intracellular action potentials recorded 120 s after internalization of the device. Scale bar: (II) 10 mm (Reprinted with permission from Duan, X., et al., Intracellular recordings of action potentials by an extracellular nanoscale fieldeffect transistor. Nature Nanotechnology 7 (3), 174–179. Copyright 2012 Macmillan Publishers). (C) Active Si nanotube transistor (ANTT). (I) Schematic of an ANTT probe inserted into a cell. (II) Overview of the steps used for ANTT probe fabrication: (1) Transfer of NWs to a SU-8 layer on a sacrificial layer (silver), (2) registration of positions of NWs and definition of the bottom SU-8 layer, and (3) definition of S/D metal contacts followed by the top SU-8 passivation layer. (III) Schematic of the ANTT probe released from the substrate. (IV) SEM image of an ANTT probe; Inset: zoom in of the probe tip from the dashed red box. Representative potential versus time data recorded (V) immediately following contact between the ANTT probe and a single cardiomyocyte, (VI) ca. 100 s following contact, and (VII) ca. 5 min following trace VI. The tick marks in V-VII correspond to 1 s. Scale bars: (IV) 10 mm, (inset) 100 nm (Reprinted with permission from Gao, R., et al. (2012). Outside looking in: Nanotube transistor intracellular sensors. Nano Letters 12 (6), 3329–3333. Copyright 2012 American Chemical Society).

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were modified with phospholipid. Bringing the cell in contact with the probe led to regularly spaced spikes with a frequency of 1.8 Hz that is correlated with cell beating (Fig. 8C-V). These peaks had widths of  0.7 ms and amplitudes up to 10 mV that are consistent with extracellular cardiomyocyte APs. After ca. 100 s following contact between the ANTT probe and PDMSsupported cell, the recorded periodic signals changed substantially with an increase in amplitude and duration to 40–50 mV and ca. 200 ms, respectively (Fig. 8C-VI). Over a period of several minutes, the probe was able to achieve stable intracellular recordings with amplitude and duration of ca. 80 mV and 200 ms, respectively (Fig. 8C-VII). The nanoscale dimensions of the probe allow stable interface with the cell. The straightforward fabrication of the devices enables multiple ANTTs at the end of single probes, demonstrating the potential to achieve multiplexed recording of full-amplitude intracellular APs from single cells as well as cellular networks.

Conclusion In this article, we have discussed the current state of the recording platforms and the significant advances in approaches that took place over the last several years. Various groups around the globe have been working on (i) modifying the design and optimizing the shape, size, number, tip geometry, and flexibility of the electrodes to gain access to the cell surface or interior and maximize interactions with large cell population; (ii) using materials such as metals, semiconductors, composites, and polymers to minimize the impedance and making the devices more sensitive and achieve high signal-to-noise ratio; and (iii) modifying the surface of the devices with bioactive coatings to mimic the biological environment and thus, enhancing electrode-tissue interface and minimize immune response. As stated previously in the introduction, a device could be considered as an ideal recording system if it: (i) is biocompatible and mimics properties of biological tissue to minimize any mechanical mismatch, immune response, scar tissue formation, and tissue disintegration; (ii) supports seamless integration with neurons to avoid shear motion at the cell-electrode interface, and remain functional for long period of time; (iii) enables readout that covers the entire spectrum of membrane potential events, that is, APs, subthreshold potentials (EPSPs and IPSPs), and subthreshold membrane oscillations; (iv) provides high spatial and temporal resolution to allow recordings at single cell level with submillisecond resolution; (v) simultaneously records/stimulates multiple neurons to enable the study of network dynamics and signal propagation in a neuronal network (Scanziani and Häusser, 2009; Spira and Hai, 2013; Fattahi et al., 2014). Using flexible substrates, such as polyimide films, hydrogels and silk, have minimized mechanical mismatch of MEAs with the brain tissue, thus improving the integration of devices with the cortical surface. However, the limitations that remain with these systems include limited access to deeper brain regions, inability to record the entire spectrum of membrane potential events, recording attenuated AP signals, and inability to achieve subcellular spatial resolution owing to the size of the electrodes. Designing passive electrodes in 3D such as Au mushroom electrodes, vertical NWs and nanopillars lead to enhanced surface area and high seal resistance between the probes and cell membrane, thus enabling multiplex stimulation and intracellulartype recording with amplitude of one to two orders of magnitude higher compared with planar MEAs. However, the high impedance associated with the small size of these electrodes prevents the recording of the full-amplitude AP which is an order of magnitude higher than what these electrodes can record. Also, even though these platforms show promise for multisite recordings and stimulations of multiple cells, they have not been explored to achieve multiplex recordings from neuronal circuits in 3D or in vivo model. Using extracellular nanoprobes, such as Si FETs, has enabled recording with submillisecond temporal resolution and subcellular spatial resolution, and integrating the FET sensors in 3D scaffolds allowed recording from cells that mimic morphology as in vivo models. However, these platforms are incapable of achieving measurements from intact brain tissue which involves much more complex neuronal circuitry. Using injectable electronics have enabled access to the neuronal network deep in the brain tissue. However, not being able to access the interior of the cells prevents these devices to record potentials of entire spectrum. Modifying the nanoprobe design to access interior of the cells has enabled recording of full-amplitude intracellular AP. However, there are some limitations associated with these platforms: there is a need to mechanically manipulate the cultured cells and the substrate on which they grow into physical contact with the electrodes; these probes show promise of multiplex recordings but have not been demonstrated to record intracellular signals from a large neuronal network; the design of these probes might not enable recording from in vivo models, especially from deep regions of the brain; and these devices cannot be used for stimulation purposes. We believe that combining the best features of recently developed techniques will lead to new hybrid devices with multifunctional capabilities that could enable long-term stable two-way interaction, that is, stimulation and mapping of in vivo brain activity at high spatial and temporal resolution. This two-way interaction would not just allow understanding of neuronal circuitry but also enable regenerative engineering where these platforms will aid in treatment of disorders and injuries in the electrically active tissues. Another crucial aspect is to create smart nanomaterials by utilizing bio-inspired designs and biologically derived materials, such as receptors and proteins. This would enable seamless integration of the devices into the complex cellular circuitry causing minimal interference to the functioning of the brain. A promising route in this research field is the combination of the nanoelectronic devices with other modalities, such as optogenetics to study the functional roles of different classes of neurons by manipulating the activity of specific cellular subpopulations, and chemical sensing to study the biochemistry involved in the neuronal function. The advancements in the development of smart hybrid electronic devices in combination with other modalities will enable greater insight into the multitude of neural interactions, understanding complex biological systems and disease progression, and potential new therapeutic directions.

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Further Reading Buzsáki, G., Anastassiou, C. A., & Koch, C. (2012). The origin of extracellular fields and currentsdEEG, ECoG, LFP and spikes. Nature Reviews Neuroscience, 13(6), 407–420. Hodgkin, A. L., & Huxley, A. F. (1952). A quantitative description of membrane current and its application to conduction and excitation in nerve. The Journal of Physiology, 117(4), 500. Kandel, E. R., Schwartz, J. H., Jessell, T. M., Siegelbaum, S. A., & Hudspeth, A. J. (2000). Principles of neural science (Vol. 4). New York: McGraw-Hill. Nicholls, J. G., Martin, A. R., Wallace, B. G., & Fuchs, P. A. (2001). From neuron to brain (vol. 271). Sunderland, MA: Sinauer Associates. Offenhäusser, A., & Knoll, W. (2001). Cell-transistor hybrid systems and their potential applications. Trends in Biotechnology, 19(2), 62–66. Spira, M. E., & Hai, A. (2013). Multi-electrode array technologies for neuroscience and cardiology. Nature Nanotechnology, 8(2), 83–94. Zhang, A., & Lieber, C. M. (2016). Nano-bioelectronics. Chemical Reviews, 116(1), 215. Bezanilla, F. (2006). The action potential: from voltage-gated conductances to molecular structures. Biological Research, 39(3), 425–435. Koeppen, B. M., & Stanton, B. A. (2009). Berne and Levy physiology. St. Louis: Elsevier Health Sciences.

Neural Crest Stem Cells T Hochgreb-Ha¨gele and ME Bronner, California Institute of Technology, Pasadena, CA, USA © 2019 Elsevier Inc. All rights reserved.

Origins of NC Cells Epithelial to Mesenchymal Transition Developmental Potential of NC Cells Molecular Mechanisms of NC Formation NC-Related Birth Defects and Cancers NC Stem Cells and the Potential to Treat Disease Acknowledgments References Further Reading

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Glossary CNS Central nervous system. EMT Epithelial to mesenchymal transition. GRN Gene regulatory network. hESC Human embryonic stem cells. NC Neural crest.

Stem cells are functionally defined as cells that are able to self-renew as well as capable of generating several distinct differentiated cell types. Given their important therapeutic potential, there has been growing interest and steady progress in understanding the biology of stem cells. Embryonic stem cells are totipotent cells, i.e., able to give rise to all cell types, and are generally derived from the inner cell mass of the blastocyst. More restricted stem cell populations have been identified in association with differentiated tissues or organs, such as skeletal muscle (Collins and Partridge, 2005; Zammit et al., 2006), heart (Laugwitz et al., 2005), skin (Wong et al., 2006), and hematopoietic tissue (Bryder et al., 2006). These are more limited in their differentiation potential than totipotent stem cells, but have greater significance for maintenance of tissue homeostasis. These tissue- or organ-specific populations constitute important local sources of multipotent, proliferating cells, which can be activated and recruited in response to tissue damage. In addition, these precursors play a constitutive role in replenishing differentiated tissue with newly generated cells throughout life. The embryonic neural crest (NC) is a unique population of highly migratory and multipotent cells that generate a variety of derivatives, including peripheral neurons and associated glia, Schwann cells lining the peripheral nerves, and melanocytes responsible for the pigmentation of our bodies. They also form endocrine cells of the adrenal medulla and thyroid calcitonin-producing cells, as well as fibroblasts, myoblasts, adipocytes, and angioblasts (Le Douarin, 1982; Le Douarin and Kalcheim, 1999). In the head region, cranial NC cells give rise to mesenchymal cells that differentiate into connective tissue and skeletal cells, such as chondrocytes and osteocytes, which contribute to cranial cartilage and bone. The latter comprises most of the skull and much of the facial skeleton, most notably the jaws, as well as the odontoblasts of the tooth primordia. From an evolutionary perspective, NCs are unique to vertebrates and one of their defining features. Evolutionarily NC is thought to have facilitated predatory ability by forming a very efficient and powerful jaw. This in turn allowed vertebrate brains to get ever larger, making them the most efficient predators on the planet. Cell lineage analysis of NC cell fate in vivo as well as clonal analysis in vitro have revealed that some individual NC cells indeed are multipotent and able to contribute to several derivatives. Clonal studies in tissue culture show that some multipotent NC progenitors can differentiate into glia, neurons, melanocytes, and cartilage, or combinations of these cell types. In addition, NC cells have a limited capacity for self-renewal, meaning that individual cells can divide and form progeny comprised of similar self-renewing NC cells, plus sister cells that can differentiate into several types of derivatives (Trentin et al., 2004). In vivo, they retain the ability to self-renew at least for a limited time. For example, crest-derived sympathoadrenal cells give rise to sympathetic neurons or adrenomedullary cells, but remain mitotic even after expressing adrenergic markers (Morrison et al., 1999). Unlike true stem cells, however, the self-renewal ability of NC cells is transient. Thus, it is more appropriate to call them ‘stemlike’ cells rather than true stem cells.

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Origins of NC Cells The early embryo contains three layers: the ectoderm and the endoderm lining the outer and inner surfaces of the embryo, respectively, and the mesoderm localized in between. The ectoderm can be further subdivided in two regions: a middle portion which becomes a thickened neural plate that will give rise to the central nervous system (CNS), while the adjacent tissue will form the epidermis of the skin (Figure 1(a) and 1(a’)). Between the presumptive epidermis and neural ectoderm, ‘inductive’ signals define the neural plate border, where induction of NC takes place (LaBonne and Bronner-Fraser, 1999). By definition, induction is an event whereby two tissues are brought together and in response to an interaction between them, they form a new tissue. This is generally mediated by the cross-talk of signaling molecules that travel between the two tissues. The neural plate is initially a flat swathe of ectoderm that subsequently folds in upon itself during ‘neurulation.’ During this event, the sheet of cells starts to fold in upon itself in a process called invagination. Concurrently, the neural folds containing the NC precursors begin to elevate (Figure 1(b) and 1(b’)). These morphogenetic movements bring the neural folds together (Figure 1(c) and 1(c’)) and their fusion in the dorsal aspect of the embryo results in closure of the neural tube, as a cylindrical epithelium that becomes internalized and covered by the overlying epidermis (Figure 1(d)). The neural tube gives rise to the entire CNS, forming the brain in the head region and spinal cord more posteriorly. Closure of the neural tube initiates in the head and then proceeds progressively tailward in the embryo. Migration of NC cells is closely associated with neurulation, and in most vertebrates, cells start migrating from the neuroepithelium shortly after neural tube closure. NC cells first emigrate in the cranial region, then at vagal levels, and subsequently in the trunk, similar to the rostral to caudal direction of neural tube closure. NC cells exhibit different patterns of migration and form distinct derivatives depending upon their site of origin in the neural tube. At caudal forebrain and midbrain levels, cranial NC cells migrate as a broad and uniform sheet of cells that expands away from the neural tube (Figure 1(d’)) and invades the cranial mesenchyme, where they populate the cranial ganglia and frontonasal process. These cells contribute to neurons, glia, and cartilage of the face. At the hindbrain level, NC cell migration is segmental, resulting in several distinct streams of migrating cells that invade the branchial arches. These cells then differentiate into portions of the jaw and bones of the neck. The vagal NC undergoes the most extensive migrations of any embryonic cell type. These cells migrate along the entire length of the gut, where they form the enteric ganglia that control gut motility. Trunk NC cells undergo segmental migration through the somites. Some condense adjacent to the neural tube to form dorsal root ganglia, whereas others migrate further ventrally to form sympathetic ganglia and the adrenal medulla. Melanocytes arise from NC cells at all axial levels.

Epithelial to Mesenchymal Transition The neural tube is comprised of polarized epithelial cells with adherens and tight junctions establishing their intercellular connections. At the end of neurulation, the premigratory NC cells reside within the dorsal portion of the neuroepithelium, and thus initially are part of the CNS. Subsequently, NC cells lose their polarity and undergo changes of adhesion. Their tight junctions are disrupted, the cytoskeleton is reorganized, and they transition from a columnar epithelial arrangement to a migratory mesenchymal cell type. This process is known as epithelial to mesenchymal transition (EMT), and it involves great changes in cell morphology, as well as in the repertoire of adhesion and recognition molecules that are expressed on the surface of the cells (Figure 2). The cellular changes of EMT imbue NC cells with the ability to delaminate from neural tube and migrate extensively through the embryo along well-established pathways. They then differentiate into a variety of cell types in the head and trunk. Although it is an essential step in the development of NC cells, EMT is not unique to this population but also occurs in several types of embryonic cells. Importantly, it is characteristic of metastatic transformation of cancer cells. Interestingly, NC and other stem cells share many biological properties with tumor cells: the ability to undergo EMT; their capacity to migrate extensively; and their ability to differentiate into numerous cells.

Developmental Potential of NC Cells NC cells have the ability to form derivatives as diverse and distinct as neurons, pigment cells, and cartilage. Initially, premigratory NC cells appear to be multipotent. However, as these cells migrate along specific pathways, they encounter diverse environments and are exposed to inductive signals that are differentially distributed within the embryo. These signals direct NC cells to assume diverse identities and gradually restrict their developmental potential, ultimately resulting in acquisition of their differentiated state. As a result, migrating NC cells are a heterogeneous population, including various types of intermediate precursors and highly multipotent cells, some of which retain the capacity for self-renewal. The latter can become both neuronal and nonneuronal derivatives, including melanocytes, neurons and support cells of the peripheral nervous system, smooth muscle cells, and facial cartilage and bones (Figure 3). The kinds of derivatives that NC cells form depend upon their axial level from which they originate. For example, in vivo lineage analysis has shown that the site of origin along the neural axis leads to generation of some distinct derivatives. Whereas melanocytes are derived from NC cells from all axial levels, cranial NC cells have the unique potential to contribute to the bones and cartilage of the face, in addition to neurons and glia of the cranial sensory ganglia. Even within the head region, NC cells contribute to different derivatives. Those arising from the midbrain, rhombomeres (r) 1 and 2 contribute to the upper and lower jaws, and trigeminal

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Figure 1 The process of neural tube and neural crest (NC) formation. (Left) Schematic diagram illustrating the process of neural tube closure and NC formation from the open neural plate stage to the closed neural tube. (Right) Whole mount in situ hybridization of bird embryos with gene markers involved in NC development. (a) At the open neural plate stage, the neural plate border (green) is induced at the edges of the neural plate (gray) and epidermal ectoderm (blue), as well as influenced by the underlying mesoderm (yellow) in some vertebrates. (a0 ) Expression of the neural plate border gene, msx1 outlines the neural plate border, as seen from a dorsal view of an embryo at the open neural plate stage. (b) With time, the neural folds (green) begin to elevate as the first step in invagination of the presumptive neural tube. (b0 ) Msx1 expression is retained on the elevating neural folds as seen by whole mount in situ hybridization. (c) As the neural folds appose, bona fide NC markers like foxd3 (c0 ) initiate expression in the newly closed neural tube. (d) The overlying epidermis (blue) closes over the neural tube. Premigratory NC cells are contained within the central nervous system but some have already undergone an epithelial to mesenchymal transition to become migrating NC cells that migrate around and through the mesodermal somites (yellow). (d0 ) Dorsal view of avian embryo expressing sox10 in migrating neural crest.

ganglion; those from r4 contribute to the proximal facial ganglion and the hyoid bone. The vagal NC forms the enteric nervous system, the cardiac septum, and components of the aortic arch. In the trunk, NC cells form all of the peripheral ganglia as well as the chromaffin cells in the adrenal medulla. However, unlike cranial crest, trunk NC never contribute to cartilage and bone, even if transplanted to the head. These observations suggest that, although NC cells from all axial levels appear to have multiple developmental potentials, the types of derivatives formed vary somewhat accordingly to axial level of origin.

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Figure 2 The epithelial to mesenchymal transition (EMT) during neural crest (NC) emigration from the neural tube. After neural tube closure, premigratory NC cells (green) reside within the dorsal or back portion of the neural tube. In this location, individual premigratory NC cells are polarized epithelial cells with close connections mediated by tight junctions and adhesion molecules. As these cells undergo EMT, they lose their junctional connections, depart from the neural tube, delaminate from the neural tube, and become migratory mesenchymal cells. This process of NC EMT is similar to events occurring during cancer metastasis.

Figure 3 Neural crest (NC) cells are multipotent and form many distinct derivatives. The NC stem cell has the ability to give rise to progeny cells that contribute to multiple and diverse lineages. For example, cranial NC cells can form bone and cartilage, some smooth muscle, as well as neurons and glia of cranial ganglia. Vagal NC cells give rise to mesenchymal cells of the cardiac septum and to enteric ganglia that innervate the gut. Trunk NC cells also give rise to neurons and glia of the dorsal root ganglia and sympathetic ganglia, and to chromaffin cells of the adrenal medulla. At all axial levels, NC cells contribute to melanocytes.

The timing of migration from the neural tube also influences the type of derivative formed by NC cells at a particular axial level. In the head, for example, the early migrating NC cells populate the branchial arches, where they contribute to bone, cartilage, and connective tissue of the craniofacial skeleton. In contrast, the later wave of migrating NC stays close to the CNS and forms the glia and some neurons of the cranial ganglia. In the trunk, earliest migrating NC cells contribute to ventrally located sympathetic ganglia, while later migrating NC cells form more dorsal derivatives, such as dorsal root ganglia and melanocytes. However, if one ‘switches’

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the positions of these early and late populations by performing transplants experiments in the embryos, they can behave appropriately for their new location. Thus, it may be the environmental cues from the spots that are open to them rather than the inherent knowledge that leads some cells to their final resting sites.

Molecular Mechanisms of NC Formation Embryonic NC cells have a fascinating ability to maintain a stem-cell-like state prior to differentiating into very diverse derivatives. This ability is retained as tissue-specific stem cells from the NC lineage are identified in niches of stem cells in adult tissues such as the cornea and the dental pulp. Therefore, there has been great interest in uncovering the molecular mechanisms underlying NC induction, how they maintain multipotency and acquire migratory ability, and finally what signals instruct them to choose their final fates. Answers to these interesting topics are potentially useful in stem cell therapy and regenerative medicine. The complex sequence of events, cooperating to transform naïve ectodermal cells into NC cells with broad differentiation potential, has been assembled into a putative NC gene regulatory network (GRN). This is essentially a circuit diagram describing the molecular pathways that guide development of NC (Meulemans and Bronner-Fraser, 2004; Sauka-Spengler and Bronner-Fraser, 2008; Betancur et al., 2010). Each step is represented as distinct module that plays a role in (1) induction of the neural plate border, (2) specification of neural plate border, (3) specification of NC progenitors, and (4) differentiation of NC derivatives. During this complex set of interactions, factors operating at different hierarchical levels of the NC GRN work in concert to establish the NC transcriptional state. Formation of NC cells initiates during gastrulation, when signals mediated by diffusible growth factors emanating from the adjacent epidermal ectoderm and mesoderm are integrated to induce a presumptive NC territory at the neural plate border (Figure 4(a) and 4(a’)). NC induction requires intermediate levels of BMPs (Bone Morphogenetic Proteins), secreted proteins that are members of the TGF-b superfamily involved in many important developmental events such as dorsoventral patterning during early embryonic development (Knecht et al., 1995; Piccolo et al., 1996; Dale et al., 1992). In the ectoderm, BMPs play an early role in the induction of the neural plate and NC (LaBonne and Bronner-Fraser, 1998; Marchant et al., 1998; Wilson and Hemmati-Brivanlou, 1995; Liem et al., 1995). Subsequently, expression of BMP inhibitors such as chordin, noggin, and follistatin establishes the intermediate levels of BMP necessary for induction of NC cells (Lamb et al., 1993; Sasai et al., 1994; Piccolo et al., 1996; Hemmati-Brivanlou et al., 1994). However, BMPs alone are not sufficient to induce NC progenitors, and other signaling systems also contribute to NC formation (Streit et al., 1998; Marchant et al., 1998; LaBonne and Bronner-Fraser, 1998). For example, it is known that a combination of factors like FGF (Fibroblast Growth Factors), members of the Wnt signaling pathway and Notch-Delta is also necessary to induce an NC-forming territory (Garcia-Castro et al., 2002; Mayor et al., 1995; Mayor et al., 1997; Monsoro-Burq et al., 2003; Endo et al., 2002; Glavic et al., 2004; Lewis et al., 2004; LaBonne and Bronner-Fraser, 1998). FGFs are secreted from the paraxial mesoderm, and Wnt signals emanate from the nonneural ectoderm and/or the paraxial mesoderm. Notch is expressed in the neural plate, with higher levels in NC cells, while its ligand Delta is expressed in the epidermal ectoderm. Because of these signaling events, the close interaction between the neuroepithelium, nonneural ectoderm, and paraxial mesoderm is critical for the establishment of a presumptive NC territory. Integration of these various environmental signals is processed by the cells of the neural plate border and manifested by expression of transcription factors (neural plate border specifier genes), whose overlapping expression at the junction between neural and nonneural ectoderm specifies this region as the neural plate border. The neural plate border specifiers include transcription factors such as Msx1/2, Dlx3/5, Pax3/7, Gbx2, as well as Zic (Figure 4(b)) (Monsoro-Burq et al., 2005; Tribulo et al., 2003; Sato et al., 2005; Monsoro-Burq et al., 2003; Khudyakov and Bronner-Fraser, 2009). Their regions of overlapping expression constitute a molecular signature of the neural plate border (Figure 4(b’)). Cells within this territory are imbued with multipotency and, later, with the ability to respond to NC specifying signals. With time, positional information supplied by gradients of signaling molecules dictates the transcriptional state of NC precursors within the neural plate border. Prospective NC cells integrate these signaling inputs to become premigratory NC cells. As the neural folds elevate and position these cells within the dorsal portion of the developing neural tube, induction of bona fide NC cells is characterized by expression of another group of transcription factors termed NC specifier genes (Figure 4(c) and 4(c’)). These include transcription factors such as Snail2, Sox10, and FoxD3, as well as AP-2, Sox9, and c-Myc (Nikitina et al., 2008; Sauka-Spengler et al., 2007). Expression of the NC specifier genes in premigratory and delaminating cell reflects the fact that these cells have been specified to an NC cell fate. Functionally, these transcription factors control the expression of effector genes that confer unique migratory and multipotent characteristics via changes in adhesion, shape, motility, and signaling repertoire of the NC precursors. NC specifier genes are turned on at different phases during NC specification and directly activate gene batteries controlling cellular processes like EMT, thus enabling cells to delaminate from the neural tube and migrate extensively throughout the embryo to then differentiate into diverse derivatives (Figure 4(d) and 4(d’)) (Sauka-Spengler and Bronner-Fraser, 2008; Meulemans and Bronner-Fraser, 2004). For example, expression of the transcriptional repressor Snail2 initiates during NC specification and plays an important role in promoting EMT via repression of cell adhesion molecules called cadherins (Taneyhill et al., 2007; Batlle et al., 2000; Cano et al., 2000). In addition, Snail influences the expression of proteins involved in cell motility and assembly of junctions (Nieto, 2002; Ikenouchi et al., 2003). Whereas Snail2 is expressed transiently, Sox10 expression persists in migrating crest, as well as in subsets of differentiating NC cells. Other early NC specifier genes like Id and c-Myc have been implicated in the maintenance of multipotency of NC cells by controlling expression of genes involved in cell cycle and cell fate decision (Bellmeyer et al.,

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Figure 4 Gene regulatory network underlying progressive development of the neural crest (NC). A combination of signaling and transcriptional events operates at progressive stages of NC development. For hypothetical purposes, these can be viewed as organized into distinct modules that operate at defined stages of NC development. Collectively, this gene regulatory network describes the molecular pathways that guide development of the NC from the open neural plate stage to the formation of differentiated derivatives. (a, a0 ) Signaling molecules like Wnts, BMPs, FGFs, and their inhibitors transit inductive cues between the neural plate, nonneural ectoderm, and underlying mesoderm, resulting in formation of the neural plate border at the junction between neural and nonneural ectoderm. (b, b0 ) These signals ultimately result in establishment of the neural plate border territory via upregulation of transcription factors, referred to as neural plate border specifier genes (e.g., Msx1/2, Pax3/7, Zic1, etc.), whose overlapping expression defines the border region. (c, c0 ) The neural plate border specifiers cooperate with inductive signals to activate NC specifier genes (e.g., Snail, Ets1, Sox10, FoxD3, etc.) in the elevating neural folds and dorsal neural tube. (d, d0 ) After neural tube closure, the NC specifier genes, in turn, influence expression of various effector genes that are involved in the process of epithelial to mesenchymal transition that allows emigration from the neural tube and creates a population of migratory NC cells. (e, e0 ) NC specifier genes in combination with environmental factors also lead to activation of lineage-specific differentiation programs, facilitating the formation of various NC derivatives like neurons, glia, melanocytes, and craniofacial cartilage.

2003; Kee and Bronner-Fraser, 2005; Light et al., 2005). FoxD3 also appears to maintain multipotency, by preventing early differentiation of crest cells (Mundell and Labosky, 2011; Teng et al., 2008). Interestingly, FoxD3 is not only important in NC development but also plays a role in maintaining pluripotency during early embryonic development and in stem cells (Hanna et al., 2002; Tompers et al., 2005). Finally, NC cells transition from their migratory and multipotent state and begin to differentiate into defined and diverse derivatives like neurons and glia of peripheral ganglia, cartilage and bones of the face, or pigment cells (Figure 4(e)). After NC cells reach their final destinations, expression of most early NC specifiers is downregulated. However, expression of certain genes persists in subsets of derivatives, as in the case of FoxD3 for the neural and glial precursors in the dorsal root ganglia, Sox9 in the NCderived chondrocytes, as well as Sox10 in melanoblasts and elements of the peripheral nervous system. Effector genes involved in differentiation of NC derivatives sometimes function to specify cell fate. For example, Mitf expression in melanoblasts, together

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with Sox10, directly regulates expression of dopachrome tautomerase, an enzyme necessary for melanin synthesis (Ludwig et al., 2004); in chondrocytes, Sox9 directly regulates expression of the chondrocyte matrix protein as Collagen type II (Lefebvre et al., 1997).

NC-Related Birth Defects and Cancers As a population, the NC and its derivatives are highly susceptible to errors at multiple steps, from specification to migration, proliferation, and differentiation into distinct derivatives. In fact, pathologies that affect NC-derived tissues are among the most common causes of birth defects in humans. For example, craniofacial defects represent 10% of human birth defects in term births. These include cleft lip and palate, defects in dentition, and serious skull malformations such as missing or drastically reduced bones. Such defects are critically important, as the frontal bones are necessary to protect the brain. TGF-b signaling has been shown to be important during formation of the palate in both mouse and humans (Proetzel et al., 1995; Sanford et al., 1997; Kaartinen et al., 1995). Some of these anomalies, including craniofacial and cardiac septation defects, are traditionally corrected surgically, it is clear that the causes are complex and involve both gene–gene and/or gene–environment interactions that may alter NC specification, migration, or differentiation. Only some of the anomalies can be addressed surgically. In addition, the NC is known to be important for coordinating head and brain development, since removal of cranial crest results in severe brain abnormalities (Le Douarin et al., 2007). Some NC-related birth defects have their roots in mutations in transcription factors that play an important role in the NC GRN. For example, human mutations in the transcription factor AP2, a NC specifier gene, cause severe defects in facial development (Satoda et al., 2000). Children with this disorder, called Char syndrome, typically have characteristic facial features with a ‘duckbill’ appearance resulting from a flattened midface, wide-set eyes, and flat nasal bridge and tip of the nose. Defects in another NC specifier gene, Ets1, cause a mutation in cardiac NC development that causes a cardiovascular septation defect (Gao et al., 2010; Ye et al., 2010). And several other mutations in genes affecting the NC cause colonic agangliogenesis or Hirschsprung’s disease (Obermayr et al., 2013; Iwashita et al., 2003). A less well-characterized NC disorder is familial dysautonomia, which affects the development and survival of sensory, sympathetic, and some parasympathetic neurons. This causes many debilitating symptoms, including insensitivity to pain, poor growth and fluctuating blood pressure. The disorder appears to be caused by incomplete development of NC-derived sensory and autonomic neurons (Nordborg et al., 1981; Lee et al., 2009). These are only a few examples of the many birth defects that have their origin in defects in NC development. Collectively, these are referred to as neurocristopathies. Perhaps because NC cells are highly migratory and invasive by nature, many of their derivatives are prone to metastasis, giving rise to several common types of cancer. These include melanoma, neuroblastoma, and neurofibromatosis (Elephant man’s disease). Mutations in genes associated with EMT within the developing or adult organism often result in tumor development and metastasis. These same genes are often critical for normal NC development and their successful EMT. Accordingly, genes like Snail2 and Sox10, which function as important NC specifier genes also, are highly elevated in many types of adult cancer cells (Shakhova et al., 2012; Chakrabarti et al., 2012). This raises the hopeful prospect that understanding normal NC development may lead to targets of therapeutic intervention to help prevent metastasis of several types of cancers.

NC Stem Cells and the Potential to Treat Disease Work by several investigators has led to the identification and purification of NC stem cells that have some ability to self-renew and can also give rise to diverse derivatives. Clonal analysis of migrating cranial NC cells has demonstrated that many of these progenitors are multipotent and can give rise to bone, neural, and pigment cell types. On the other hand, NC stem cells can be driven by environmental factors to adopt specific fates under proper conditions. Neuregulin, for example, mediates development of NC stem cells into Schwann cells and glia, whereas BMP-2 promotes neuronal differentiation, and TGF-b1 favors development of smooth muscle cells (Shah et al., 1994; Shah et al., 1996; Shah and Anderson, 1997). Transient activation of Notch also promotes glial production by NC at the expense of neurogenesis (Morrison et al., 2000). In addition to the embryo, NC stem cells can be isolated from several NC derivatives including the gut, peripheral nerve, skin, and ganglia, as well as several craniofacial tissues such as the cornea and dental pulp, from both the fetus and adult (Li et al., 2007; Nagoshi et al., 2008; Yoshida et al., 2006; Wong et al., 2006; Gronthos et al., 2000; Morrison et al., 1999; Kruger et al., 2002). Under proper culture conditions, these cells can self-renew and differentiate into neurons, glia, and smooth muscle cells within single colonies, demonstrating that they retain multipotency. Because many birth defects are caused by abnormal NC development, it is hoped that understanding development of NC stem cells may allow investigators to ‘replace’ abnormal NC derivatives for potential use in regenerative medicine. A critical first step is to obtain sufficient numbers of human NC cells that can be differentiated into the appropriate cell type. Embryonic NC cells have the ability to generate a vast array of distinct cell types, such as bones and cartilage, neurons and glia, melanocytes, as well as endocrine, connective and adipose tissues. Thus, there has been great interest in characterizing the prospective stem cell properties of human NC populations and differentiating them into particular derivatives. NC cells can be derived in tissue culture from human embryonic stem cells (hESCs). The basic process has been to induce hESCs into neural stem cells in either adherent or suspension culture and further induce them into NC cells using either coculture with

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other cell types or defined medium with added factors (Lee et al., 2007, 2010). The cells located at the periphery of these cultures often express NC specifier genes like Snail2 and Sox10. Because the efficiencies tend to be low, it is often necessary to purify NC cells by fluorescence-activated cell sorting and/or perform clonal analysis to examine differentiation in detail. Ongoing efforts are focused on differentiating these into particular types of derivatives like sensory or autonomic neurons, for treatment of diseases like familial dysautonomia, in which peripheral neurons are defective. NC cells derived from hESC have been shown to differentiate into a wide range of NC derivatives, including sensory and autonomic neurons, Schwann cells, myofibroblasts, adipocytes, cartilage, and bone cells (Lee et al., 2007). Potential problems with hESC are their general lack of availability and the fact that they may be rejected by the patients’ immune system. With recent advances in stem cell technology, it is now possible to reprogram somatic cells from adult tissue into induced pluripotent stem (iPS) cells, thus expanding the possibilities of cell therapy. The knowledge obtained from basic research in various organisms to build the NC GRN can be applied to direct and monitor differentiation of NC cells and their derivatives from both stem and iPS cells. This opens many interesting new opportunities to establish cell lineages from tissue obtained from patients with genetic syndromes and it may help elucidate the molecular mechanisms underlying the genetic syndromes affecting NC cells, as well as helping devise efficient therapies.

Acknowledgments We thank Dr. Marcos Simões Costa for helpful discussion and for providing the images of avian embryos and help with the diagram of NC derivative (Figure 3).

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Requirement for Foxd3 in the maintenance of neural crest progenitors. Development, 135, 1615–1624. Tompers, D. M., Foreman, R. K., Wang, Q., Kumanova, M., & Labosky, P. A. (2005). Foxd3 is required in the trophoblast progenitor cell lineage of the mouse embryo. Dev. Biol., 285, 126–137. Trentin, A., Glavieux-Pardanaud, C., Le Douarin, N. M., & Dupin, E. (2004). Self-renewal capacity is a widespread property of various types of neural crest precursor cells. Proc. Natl. Acad. Sci. U.S.A., 101, 4495–4500. Tribulo, C., Aybar, M. J., Nguyen, V. H., Mullins, M. C., & Mayor, R. (2003). Regulation of Msx genes by a Bmp gradient is essential for neural crest specification. Development, 130, 6441–6452. Wilson, P. A., & Hemmati-Brivanlou, A. (1995). Induction of epidermis and inhibition of neural fate by Bmp-4. Nature, 376, 331–333. Wong, C. E., Paratore, C., Dours-Zimmermann, M. T., Rochat, A., Pietri, T., Suter, U., Zimmermann, D. R., Dufour, S., Thiery, J. 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Ye, M., Coldren, C., Liang, X., Mattina, T., Goldmuntz, E., Benson, D. W., Ivy, D., Perryman, M. B., Garrett-Sinha, L. A., & Grossfeld, P. (2010). Deletion of ETS-1, a gene in the Jacobsen syndrome critical region, causes ventricular septal defects and abnormal ventricular morphology in mice. Hum. Mol. Genet., 19, 648–656. Yoshida, S., Shimmura, S., Nagoshi, N., Fukuda, K., Matsuzaki, Y., Okano, H., & Tsubota, K. (2006). Isolation of multipotent neural crest-derived stem cells from the adult mouse cornea. Stem Cells, 24, 2714–2722. Zammit, P. S., Partridge, T. A., & Yablonka-Reuveni, Z. (2006). The skeletal muscle satellite cell: the stem cell that came in from the cold. J. Histochem. Cytochem., 54, 1177–1191.

Further Reading Achilleos, A., & Trainor, P. A. (2012). Neural crest stem cells: discovery, properties and potential for therapy. Cell Res., 22, 288–304. Bailey, C. M., Morrison, J. A., & Kulesa, P. M. (2012). Melanoma revives an embryonic migration program to promote plasticity and invasion. Pigment Cell Melanoma Res., 25, 573–583. Bellin, M., Marchetto, M. C., Gage, F. H., & Mummery, C. L. (2012). Induced pluripotent stem cells: the new patient? Nat. Rev. Mol. Cell Biol., 13, 713–726. Bronner, M. E., & Le Douarin, N. M. (2012). Development and evolution of the neural crest: an overview. Dev. Biol., 366, 2–9. Dupin, E., & Sommer, L. (2012). Neural crest progenitors and stem cells: from early development to adulthood. Dev. Biol., 366, 83–95. Hall, B. K., & Gillis, J. A. (2013). Incremental evolution of the neural crest, neural crest cells and neural crest-derived skeletal tissues. J. Anat., 222(1), 19–31. Le Douarin, N. M., & Dupin, E. (2012). The neural crest in vertebrate evolution. Curr. Opin. Genet. Dev., 22, 381–389. Milet, C., & Monsoro-Burq, A. H. (2012). Embryonic stem cell strategies to explore neural crest development in human embryos. Dev. Biol., 366, 96–99.

Osteoarthritis at the Cellular Level: Mechanisms, Clinical Perspectives, and Insights From Development Melanie Fisher, Tyler Ackley, Kelsey Richard, Bridget Oei, and Caroline N Dealy, UConn Health, Farmington, CT, United States © 2019 Elsevier Inc. All rights reserved.

Introduction The Osteoarthritis Epidemic Articular Cartilage Structure Articular Cartilage is Built to Last Mechanisms of Osteoarthritis Multiple OA Phenotypes OA Is a “Whole-Joint” Disease Anabolic Versus Catabolic Phases of OA Chondrocyte Proliferation and Matrix Turnover: Catabolic or Anabolic Signs? Chondrocyte Hypertrophy in OA Articular Cartilage Development Adult Articular Cartilage Progenitor Cells Cell-Based Osteoarthritis Treatments Cartilage Grafts Autologous Chondrocyte Implantation Microfracture MSC Implants MSC Injection Donor Versus Host? Challenges in MSC-Based OA Interventions Alternative Cell-Based Strategies Future Directions References Further Reading

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Introduction Osteoarthritis is a painful and debilitating degenerative joint disease present in epidemic proportions worldwide. Osteoarthritis occurs when the articular cartilage of the joint surfaces degrades and is lost. Although historically, articular cartilage loss has been attributed to passive wear-and-tear, research informs us that osteoarthritis is an active disease, triggered by injury, inflammation, metabolic disorder, or cellular senescence; and leading to a complex and dynamic series of events that disrupts homeostasis of the entire joint, and ultimately results in net destructive loss of the articular cartilage. Because widespread articular cartilage loss is likely irreversible, the current clinical focus is to repair focal loss before it progresses to overt osteoarthritis. Based on the dogma that articular cartilage cannot repair by itself, clinical approaches have attempted articular cartilage repair using various kinds of exogenous cells, especially progenitor cells, either implanted directly into focal articular cartilage defects, or even just injected into the damaged joint. Although clinical efficacy of these approaches is not yet unequivocally demonstrated, trends towards positive outcomes have prompted mechanistic investigation to understand and ultimately optimize progenitor cell-based interventions as a way to slow or halt osteoarthritic disease. Surprisingly, these studies are revealing unsuspected potential for endogenous repair capacity by adult articular cartilage in response to as-yet-undefined signals provided by exogenous progenitor cells. This review discusses recent insights on cellular mechanisms of osteoarthritis; cellbased interventions in the clinic; and our current understanding of articular cartilage development which is informing growing appreciation of its natural potential for progenitor-mediated self-repair. The findings that are discussed emphasize the continued need for basic research to understand how cartilage and joint tissue responds to the triggers that cause osteoarthritis, and how these responses might be reversed and re-directed to support cell-mediated cartilage restoration instead of cartilage destruction.

The Osteoarthritis Epidemic Osteoarthritis is a painful and disabling degenerative joint disease characterized by progressive loss of the articular cartilage, which is the layer of connective tissue that covers the ends of the long bones, providing a smooth gliding surface and functional weight

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bearing (Martel-Pelletier et al., 2016). Osteoarthritis has been a problem for humankind for thousands of years, and examination of the pre-historic record links OA occurrence to the same risk factors experienced today; namely intense physical activity, age, diet, and/or genetics (Dequeker and Luyten, 2008; Cheverko and Bartelink, 2017; Zhang et al., 2017). Today, OA is present in epidemic proportions around the world, with an estimated 27–30 million adults affected with OA in the US alone (Lawrence et al., 2008; Johnson and Hunter, 2014; Cisternas et al., 2016). The incidence of OA is expected to rise to 67 million by 2030 (Hootman and Helmick, 2006) due to aging of the population (Holt et al., 2011; Anderson and Loeser, 2010) and increasing prevalence of obesity (Sridhar et al., 2012). Specific groups at particular high risk for osteoarthritis include women, who have a three times higher osteoarthritis prevalence than men (Johnson and Hunter 2014), athletes (Gouttebarge et al., 2015) and military service members (Cameron et al., 2016; Showery et al., 2016), and populations with bias for the disease due to genetic ethnicity (Reynard and Loughlin, 2013). OA is a major cause of disability and decreased quality of life in adults (Ma et al., 2014). The societal loss of productivity due to pain and disability, along with the health costs of caring for those affected with OA, is estimated at $67– 185 billion annually in the United States (Kotlarz et al., 2009; Losina et al., 2015). Despite intensive basic research and clinical investigation, the mechanisms that cause OA are still poorly understood, and an effective treatment has yet to be found, making osteoarthritis an “intractable disease.”

Articular Cartilage Structure Articular cartilage has a unique structure and physiology that reflects its physical demands (Fox et al., 2009). Normal human articular cartilage is 2–4 mm thick (Shepherd and Seedhom, 1999); 10% of the tissue is cells (chondrocytes), 10%–25% is the extracellular matrix the chondrocytes synthesize (Eyre, 2002; Archer and Francis-West, 2003; Quinn et al., 2013; MartelPelletier et al., 2016), and 65%–80% of the tissue is water (Fox et al., 2009). Articular cartilage matrix is rich in collagen (especially collagen type II), proteoglycans (especially aggrecan), and glycosaminoglycans, especially hyaluronan, a water-retentive molecule which is responsible for the high water content of articular cartilage (Leyett et al., 2014). The layered structure of articular cartilage is diagrammed in Fig. 1. The uppermost layer of the articular cartilage is known as the superficial zone, which comprises about 10%–20% of the total articular cartilage thickness, and contains linearly-arranged collagen fibers interspersed with occasional flattened chondrocytes. This layer provides a smooth, durable surface for articulation, and also secretes abundant hyaluronan and lubricin, which are essential for joint lubrication (Seror et al., 2015). The middle-zone comprises 40%–60% of the articular cartilage. The upper portion of the middle zone contains randomly (isometrically) arranged collagen fibrils, and rounded chondrocytes that are sparsely but fairly evenly distributed. The lower portion of the middle zone contains linear collagen fibrils arranged perpendicular to the surface, with rounded chondrocytes arranged in short stacks in alignment with the fibers (Fox et al., 2009). The extracellular matrix of the middle zone is rich in proteoglycans, especially aggrecan. The main function of the middle zone is shock absorption. The remaining 30% of the articular cartilage is the deep zone. The deep zone is divided by the tidemark into an upper hyaline cartilage portion that is continuous with the lower middle zone, and extends its linear arrangement of collagen fibrils and chondrocyte stacks; and a lower region which is calcified (mineralized). The mineralized matrix of the deep zone provides a gradual transition in mechanical stiffness of the cartilage tissue as it merges with the supporting underlying subchondral bone. Hypertrophic chondrocytes, which secrete mineralized matrix, are present in the deepest region of the deep zone, where it meets the subchondral bone. The junction where cartilage and bone meet is called the chondro-osseous junction or osteochondral interface. The subchondral bone contains blood vessels and nerves; while the articular cartilage itself is avascular and aneural (Fox et al., 2009).

Superficial Zone

Middle Zone

Deep Zone Tidemark Chondro-osseous Junction

Subchondral Bone Fig. 1 The structure and cellular features of adult articular cartilage are uniquely designed to create and maintain a tissue with maximal durability and stability.

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Articular Cartilage is Built to Last The natural design of hyaline articular cartilage tissue is optimized for strength, durability, and stability. Some individual structural components of articular cartilage are so stable that they are essentially permanent. For example, estimates of the half-life of collagen, the primary structural component of articular cartilage, range from 117 to 400 years (Verzijl et al., 2000; Heinemeier et al., 2016), and aggrecan, the most abundant proteoglycan in articular cartilage, has a half-life of at least 25 years (Maroudas et al., 1998). The cross-layered fiber arrangement of articular cartilage gives it great resistance to compressive and torsional forces, while its avascular nature contributes its very low metabolic state and high overall stability. Clearly, the structure of articular cartilage is meant to enable it to last the length of a human lifetime, or more. However, it is also apparent that the stability of articular cartilage tissue comes at the expense of a relative lack of response to insult or injury, since once damaged, articular cartilage has little native ability to heal itself. The lack of repair response of articular cartilage is what makes osteoarthritis an enormous clinical challenge.

Mechanisms of Osteoarthritis Multiple OA Phenotypes It is becoming apparent that osteoarthritis is not a single disease. Multiple clinical OA phenotypes exist, which vary by age at onset, severity of signs, and rate of degeneration. Idiopathic/primary OA is the classic progressive, chronic disease phenotype associated with the elderly, and is the most common form of OA, with some estimates placing 70% of the population over the age of 60 as being affected (Lawrence et al., 2008). Idiopathic/primary OA is the most common form of OA, and has no defined root cause (hence idiopathic, or “unknown”). Since disease signs appear late in life, and slowly but progressively worsen with time, idiopathic/primary OA has been considered an unavoidable consequence of a lifetime of mobility, causing cartilage wear over time. More likely, idiopathic/primary OA is related to the effects of earlier, un-diagnosed cartilage and joint injury. For instance, in one study, undiagnosed articular cartilage defects were detected in the majority (60%–67%) of 25,124 patients receiving arthroscopies, 90% of whom were over age 50 (Widuchowski et al., 2007). The presence of un-diagnosed joint damage in idiopathic/ primary OA is a probable contributing factor to the low-grade, chronic joint inflammation observed in these patients, which is now recognized as a key factor in OA pathology (Robinson et al., 2016). A less common but better-understood form of OA is known as post-traumatic osteoarthritis (PTOA) (Anderson et al., 2011). PTOA comprises about 12% of all OA cases (Brown et al., 2006). PTOA results from a known acute traumatic injury to the joint, typically sustained in sports, combat or accidents (von Porat et al., 2004). Compared to idiopathic/primary OA, a distinguishing feature of PTOA is its rapid onset, which is in the range of 10–15 years from injury to severe disease. Another feature of PTOA that distinguishes it from idiopathic/primary OA is that it occurs in young and otherwise healthy individuals. These features decouple the factor of chronological age as a requisite for joint degeneration. Why then, does PTOA have such a rapid progression? One possibility is that the rapid onset of PTOA is related to sudden load imbalance caused by the injury, which accelerates wear and leads to faster cartilage loss (Du et al., 2016; Hsia et al., 2017). Indeed, some cases of PTOA involve discrete injuries to the articular cartilage itself, which serve as focal sites of further cartilage loss (Martin et al., 2017). However, many PTOA cases can be traced back to ligament and/or meniscal injury, especially ACL tears. In one study, 80% of patients with traumatic ACL tear progressed to radiographic PTOA within 12– 14 years (Svoboda, 2014). We now know that a major cause of PTOA is the acute inflammatory response that occurs with injury to the joint (Lieberthal et al., 2015), and that it is likely that the rapidity of disease onset and severity of signs is directly related to the magnitude of this response. Compelling evidence for inflammation as a key mediator of PTOA is intriguing lack of evidence that ACL reconstruction after injury improves joint structural outcomes or delays OA onset (Heard et al., 2013; Svoboda, 2014). What we have learned by comparing the etiology and disease course of idiopathic/primary OA and PTOA is that factors like age and load-induced wear, which appear to be obvious disease triggers, may not by themselves be directly sufficient to induce OA disease, nor are they necessarily even required for OA disease to occur. Rather, these factors lead to a change in joint homeostasis that shifts the balance in the joint towards unstable pathologic phenotypes. This is illustrated when considering the role of obesity in OA development. At first glance increased load-bearing and accompanying damage to the joints due to weight gain in obese individuals would logically seem to be responsible for the correlation between rising rates of obesity and rising rates of OA (Sridhar et al., 2012). However, recent studies suggest that the direct agents of change in obesity-related OA are more likely to be metabolic disturbances associated with obesity, including metabolic syndrome, diabetes, or even changes in the joint microbiome (Li et al., 2016a; Mobasheri et al., 2017). In fact, obesity-related OA has been suggested to comprise a third OA phenotype, known as metabolic OA (Kluze et al., 2015). What is common to each of these phenotypes is profound loss of joint homeostasis, causing the cells of the various tissues of the joint to acquire inappropriate and ultimately harmful behaviors, which eventually converge on irreversible cartilage loss. It is clear that the modern view of osteoarthritis etiology is much more complex than previously thought, and it will be necessary to understand how the unique mechanisms underlying each phenotype cause OA, in order to develop successful approaches to prevent or treat it.

OA Is a “Whole-Joint” Disease Although osteoarthritis is characterized by degeneration and loss of the articular cartilage (Martel-Pelletier et al., 2016), osteoarthritis is a “whole joint” disease, with articular cartilage loss being accompanied by morphological changes to other joint tissues.

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For example, after joint injury, the subchondral bone beneath the articular cartilage undergoes rapid changes including increased resorption leading to loss of trabecular bone, increased vascularization, and transient thinning of the supportive subchondral bone plate (Botter et al., 2011). The loss of underlying bony support may predispose the articular cartilage to further damage due to imbalanced load support, and the possible presence of trauma-induced bone microcracks and increased subchondral vascularization may facilitate vascular breach of the osteochondral junction, allowing transfer of cartilage-degrading signals from the bone to the cartilage layer (Goldring, 2012a; Zhen and Cao, 2014; Findlay and Kuliwaba, 2016). In an over-compensatory reaction to the initial bone loss triggered by joint injury, subsequent changes in bone lead to overall subchondral bone sclerosis (thickening), presence of subchondral bone cysts (Neogi, 2012) and formation of osteophytes within the joint space (van der Kraan and van den Berg, 2007). Additionally, other joint tissues are also affected in osteoarthritis including mineralization/hardening of ligaments and menisci (Tsujii et al., 2017); and thickening/hyperplasia of the synovial joint lining (Scanzello and Goldring, 2012). The morphological changes that characterize OA are mediated by active behaviors on the part of the cells including increased or decreased proliferation or survival; loss of appropriate stable phenotypes, changes in metabolic status, and changes in synthesis of anabolic (tissue building) or catabolic (tissue destructive) genes and factors (Ramos et al., 2014; Remst et al., 2014; Mueller et al., 2017; Steinberg et al., 2017). In order to develop rational OA interventions, it will be necessary to understand the nature of these cellular behaviors during normal joint homeostasis, as well as in response to OA-inducing stimuli, so we can 1 day harness or re-direct these behaviors towards prevention and/or treatment of OA.

Anabolic Versus Catabolic Phases of OA Much of our knowledge of OA mechanisms has been gained from studies in animal models of post-traumatic osteoarthritis, in which disease can be conveniently triggered by a known injury event and followed over a fairly rapid (weeks/months) and predictable time course (Little and Hunter, 2013). Many of these models involve surgical transection of the anterior cruciate ligament and/ or other ligaments of the knee, which can be imposed on small animals like mice or rats, or large animals including goats, sheep, dogs, minipigs or horses (Moran et al., 2016). In rodents, non-invasive, non-surgical, mechanical load-induced ACL rupture or articular fracture (Christiansen et al., 2015) models are also available. Studies in these PTOA models have revealed a phasic progression of cell-mediated catabolic (tissue-destructive) and anabolic (tissue-building) events that occur after joint trauma and during disease progression (Anderson et al., 2011). The first phase of PT-OA, is a transient early and acute catabolic phase, occurring in the days following joint injury, and characterized by pro-inflammatory macrophage infiltration of the synovium and joint, causing synovial hyperplasia and release of inflammatory cytokines, and increased expression of degradative enzymes (MMPs) by joint tissues that mediate cartilage cell death and matrix breakdown. The tissues in the joint that secrete the highest levels of destructive enzymes are not the articular cartilage itself, but rather the ligaments, tendons, menisci, and synovium (Hausler et al., 2013), emphasizing the need to consider the whole joint when trying to understand and therapeutically-manipulate OA phenotypes/outcomes. The second phase in PTOA is a transient anabolic phase in which articular chondrocytes and/or chondroprogenitors proliferate and increase synthesis of cartilage matrix proteins. The thickness of articular cartilage as a whole even after ACL or ligament disruption increases slightly and transiently during this phase (Anderson et al., 2011). Unfortunately, however, the anabolic response and increased articular cartilage thickness it causes is not sufficient to reverse OA progression, and after a few weeks is overcome by a prolonged and catabolic third phase in which continued joint inflammation and cartilage degradation leads to late-stage progressive articular cartilage loss accompanied by development of other associated abnormalities such as osteophytes and subchondral bone changes (Anderson et al., 2011). It is important to note that this phasic disease progression was validated in PTOA models, thus it is not known to what extent phasic progression of cellular behavior changes occurs in other OA phenotypes. Nonetheless, interventions that may be common to all phenotypes may include ways to inhibit catabolic responses in the joint (such as blocking acute or chronic inflammation), or ways to promote anabolic responses (such as stimulating cartilage growth and matrix synthesis).

Chondrocyte Proliferation and Matrix Turnover: Catabolic or Anabolic Signs? While increased articular cartilage thickness is an obvious beneficial outcome of the transient anabolic phase of PTOA, the cellular behaviors that precede cartilage thickeningdwhich include cell proliferation and matrix turnoverdhave not always been viewed as beneficial, or even anabolic. Chondrocyte proliferation is a necessary part of cartilage growth, but in OA, cell proliferation is known as “chondrocyte cloning” and is considered a morphological hallmark of disease. Chondrocyte cloning is the presence of numerous clusters of daughter cells, typically located in the middle zone of the degenerating articular cartilage, which are each surrounded by an immediately-localized region of territorial cartilage matrix synthesized by the dividing cells. Although chondrocyte proliferation and new matrix synthesis are anabolic cartilage behaviors, aspects of these responses are not exclusively beneficial in articular cartilage. For instance, the territorial matrix synthesized in OA chondrocyte clones is different in composition than newly-synthesized territorial matrix in young, healthy, growing cartilage, in that it is rich in collagen types III and VI, which are otherwise not abundant articular cartilage collagens (Pullig et al., 1999; Hosseininia et al., 2016). Paradoxically, chondrocyte clones also produce copious amounts of matrix degradative enzymes such as matrix metalloproteinases (MMPs) and A Disintegrin and Metalloproteinase with Thrombospondin Motifs (ADAMTS), which are known to be major effectors of cartilage matrix degradation (Li et al., 2017a; Yang et al., 2017). Together, these observations indicate that chondrocyte clones are sites of active, localized, matrix remodeling in OA. Indeed, the progressive destruction of articular cartilage in OA has been suggested to involve continual catabolic matrix degradation, coupled with simultaneous but ultimately insufficient anabolic matrix synthesis (Goldring, 2012b).

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Chondrocyte Hypertrophy in OA Another cellular characteristic of OA cartilage is an increased number of hypertrophic chondrocytes in the deepest cartilage layer. Chondrocyte hypertrophy is a normal chondrocyte fate in the growth plate, where hypertrophic cells carry out the essential function of synthesizing the mineralized cartilage that will become the substrate for bone deposition by osteoblasts (Sun and Beier, 2014). Growth plate hypertrophic chondrocytes also secrete factors that stimulate vascular entry, which brings clasts/osteoblasts into the region to remodel the mineralized matrix and replace it with bone. Hypertrophic chondrocytes may even directly differentiate into osteoblasts (Zhou et al., 2014). The role of hypertrophic chondrocytes in articular cartilage, and the role of increased hypertrophy in OA, is not yet clear (van der Kraan and van den Berg, 2012; Pesesse et al., 2013). In healthy as well as OA articular cartilage, hypertrophic chondrocytes are present beneath the tidemark, the distinct line that demarcates the upper nonmineralized cartilage layers from the deepest, mineralized cartilage layer (Fox et al., 2009). One of the characteristics of idiopathic/primary OA is that the tidemark line is closer to the surface, which has been attributed to increased thickening of the underlying mineralized layer and/or increased thickening of the subchondral bone (Deng et al., 2016), both of which are consistent with the presence of greater numbers of hypertrophic chondrocytes below the tidemark. As the function of the mineralized cartilage layer is to transition mechanical strength in a gradual fashion from relatively flexible cartilage to the stiff underlying bone, increased tissue mineralization could exacerbate cartilage damage caused by loading (Schultz et al., 2015). Excessive chondrocyte hypertrophy in OA might also facilitate the abnormal neovascularization that occurs across the chondro-osseous junction (Pesesse et al., 2013). While it is not yet clear if it is a consequence, or a cause, the tendency of chondrocytes within the injured or osteoarthritic joint to assume unstable hypertrophic phenotypes is a key pathologic feature of articular cartilage disease (Caron et al., 2015; Guidotti et al., 2015; Yahara et al., 2016). Indeed, a classic morphological sign of OA is formation of osteophytes (bone spurs), which are preceded by localized regions of inappropriate chondrocyte hypertrophy and subsequent endochondral ossification (Gelse et al., 2003; van der Kraan and van den Berg, 2007). Inappropriate chondrocyte hypertrophy in OA is not limited to the articular cartilage. The fibrocartilage of the menisci also commonly hypertrophies and mineralizes in end-stage idiopathic/primary OA (Abraham et al., 2014) and in PTOA models (Sun and Mauerhan, 2012). The fibrocartilage of the joint ligament entheses also undergoes hypertrophy and mineralizes in response to joint destabilizing injury, and the former fibrocartilage cells eventually make major contributions to formation of osteophytes as OA progresses (Dyment et al., 2015). The presence of excessive hypertrophic chondrocytes in OA has been suggested to reflect an overall shift from a stable, permanent articular cartilage phenotype, to an unstable phenotype which favors inappropriate formation of mineralized tissues and bone within the joint via endochondral ossification.

Articular Cartilage Development Since chondrocyte hypertrophy, a developmental phenotype, is a key phenotypic switch in OA disease, better understanding of how joint tissues, including the articular cartilage, form during development may help us decipher the cellular mechanisms that go awry in osteoarthritis, and may even inform future cellular approaches to repair or regenerate damaged articular cartilage tissue (Iwamoto et al., 2013). Studies using cell lineage tracing in transgenic mice have revealed that the tissues of the adult joint, including articular cartilage, tendon, ligament, synovium and menisci, all originate from one embryonic mesenchyme cell population that is collectively known as the “joint interzone” (Archer et al., 2003). The joint interzone comprises the localized region of mesenchyme in the developing limb that is found in the space between the ends of the early skeletal elements (Archer et al., 2003). Until recently it was thought that the joint interzone formed as a result of localized cartilage de-differentiation occurring at sites of presumptive joints along the cartilage models (Hyde et al., 2007). The de-differentiated interzone cells re-acquired distinct joint tissue fates based on their location within the interzone region, with cells close to the cartilage elements contributing to articular cartilage, and cells in the middle of the interzone contributing to the rest of the joint tissues (Koyama et al., 2008). More recent studies have re-explored the process of joint development using sophisticated lineage-tracking approaches, which have refined our understanding of this complex process and shed new insight on the origin and fate of the cells that comprise the different joint tissues (Li et al., 2017a,b; Shwartz et al., 2017). While there remains agreement that there is a common lineage to all of the structures of the joint that is represented by their embryonic shared expression of the marker GDF5, we now know that the cells that contribute to joint do not predominantly arise from the cartilage model. Rather, the interzone region appears to be continuously populated by cells that migrate into the area from an adjacent, as yet undetermined source, and which then transit out of the interzone to form the different joint tissues (Shwartz et al., 2017). By pulsing the interzone cells with genetic lineage label at different times, it was shown that early transiting interzone cells contributed mainly to menisci, ligaments and the epiphyseal cartilage at the ends of the cartilage models (destined to become growth plate, not articular cartilage). Cells that transited through the interzone later on contributed to articular cartilage, as well as menisci and ligaments, but not to the epiphysis (Shwartz et al., 2017). Thus, the articular cartilage in the embryo appears to form from a mixture of growth plate and joint progenitor cells (Shwartz et al., 2017). Surprisingly however, only descendants of the joint progenitor cells are retained in adult articular cartilage. This was shown by tracing the fate of individual cells in the superficial zone of the developing articular cartilage (Li et al., 2017a,b). These cells express lubricin, the product of the Prg4 gene, and an essential protein for lubrication of the joint surfaces (Rhee et al., 2005). The progeny of the Prg4-expressing cells in Prg4-GFP mice all ended up either

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in the superficial zone or in the deeper layers of the adult articular cartilage, but none of the cells were found in the adult growth plate (Li et al., 2017a,b). Conversely, Col2-lineage tagged cells, which were initially co-mingled with Prg-4 expressing cells in the superficial zone of the articular cartilage, were not found in any region of the adult articular cartilage (Li et al., 2017a,b). The restriction of growth plate versus articular cartilage lineages may involve constraints provided by local sources of growth factors, such as BMPs or Wnt (Ray et al., 2015). Intriguingly, close examination of the pattern of cell division by the Prg4-expressing superficial zone cells indicated that most daughter cells from each division populated the deeper cartilage layers, but some daughter cells were retained within the superficial zone itself (Li et al., 2017a,b). The behavior of articular cartilage superficial cells is an example of population asymmetry, where some stem cells can renew themselves symmetrically (one daughter cell becomes a replacement stem cell, and the other becomes a differentiated cell) and others are lost through asymmetric differentiation (in which both cells become differentiated cells). This behavior is typical for stem cell niches in adult organisms, and suggests the superficial zone could be a self-renewing articular cartilage stem cell niche (Li et al., 2017a,b). Consistent with this possibility, the surface zone cells are a slow-cycling population, which is another characteristic of stem cell niches (Hsu and Fuchs, 2012).

Adult Articular Cartilage Progenitor Cells The observation that the superficial zone of adult articular cartilage may be a stem cell niche is important, as it suggests that during adulthood, the superficial zone might be able to supply new chondrocytes to participate in and/or augment cartilage repair. Whether or not superficial zone cells participate in articular cartilage repair in vivo is not yet clear (Chagin and Medvedeva, 2017). Evidence that superficial zone cells might have capacity to repair articular cartilage is suggested by their ability to readily undergo chondrogenesis in vitro (Dowthwaite et al., 2004; Hattori et al., 2007; Yu et al., 2014) and their ability to migrate to sites of cartilage injury (Soel et al., 2012). This was shown in an in vitro study, in which the surface zone progenitors of bovine articular cartilage explants were fluorescently-labeled prior to subjecting the explant to a cutting injury (Soel et al., 2012). Remarkably, the labeled superficial zone cells were found to be surprisingly mobile, responding to the injury by undergoing directed chemotaxis towards the site of damage (Soel et al., 2012). However, in a transgenic Prg4-reporter mouse model, articular cartilage defects made in the trochlear groove were observed to be filled 7 days after surgery with cells that originated from the synovium, rather than from the adjacent superficial zone articular cartilage (Decker et al., 2017). Furthermore, surprisingly, ablation of the superficial zone cells by Prg4-Cre-directed expression of diphtheria toxin in a mouse model was found to improve, rather than exacerbate, subsequent surgically-induced PTOA, suggesting that in response to injury, superficial zone cells acquire harmful catabolic activities, rather than helpful anabolic ones (Zhang et al., 2016). Further studies are necessary to define the role, participation, and behavior of superficial zone cells in articular cartilage repair. Progenitor-like cells have also been found in other regions of the articular cartilage besides the superficial zone (Alsalameh et al., 2004; Grogan et al., 2009; Pretzel et al., 2011). These cells typically display mesenchymal stem cell (MSC)dlike features like multipotency (the ability to differentiate into cartilage, bone or fat lineages); clonicity (an indicator of proliferative ability) and migratory capacity. For example, the middle (and superficial) zone contains chondrocytes that express the mesenchymal stem cell markers CD105 and CD166 (Pretzel et al., 2011). FACS-isolated CD105/CD166 þ cells have been isolated from both healthy and OA human articular cartilage (Alsalameh et al., 2004; Pretzel et al., 2011), and were found to be surprisingly abundant, comprising from 5% to 15% of all chondrocytes (Alsalameh et al., 2004; Pretzel et al., 2011). The CD105/CD166 þ cells migrated on matrices in vitro and displayed enhanced chondrogenic potential in in vitro assays (Alsalameh et al., 2004; Pretzel et al., 2011). Indeed, articular chondrogenic progenitors isolated from equine cartilage not only had enhanced in vitro chondrogenic capability compared to bone marrow-derived MSC, but also underwent chondrogenesis without progression to hypertrophy (McCarthy et al., 2012). A population of stem-cell like progenitors has also been isolated from the deep zone of human osteoarthritic articular cartilage (Koelling et al., 2009). This multipotent population was isolated by its ability to migrate out of OA explants in vitro, and in intact tissue, the cells were observed to enter the articular cartilage deep zone via vascular breach from the subchondral bone (Koelling et al., 2009). The location of these progenitors in the deep zone, and the association of their presence with vascularization of this layer during disease, suggests that these progenitors were likely bone marrow stem cells originating from the subchondral bone, rather than chondroprogenitors originating from the superficial or middle zones. Although one explanation for the relative inability of articular cartilage to repair itself could be a lack of a sufficient population of chondrogenic progenitor cells, several studies have reported that the number of progenitor cells in osteoarthritic cartilage is actually increased (Alsalameh et al., 2004; Pretzel et al., 2011). However, a subset of the progenitors in osteoarthritic cartilage has been found to display characteristics of enhanced senescence (aging) including telomere shortening (Fellows et al., 2017), which could suggest that their regenerative ability is impaired (Fellows et al., 2017). Chondrocyte senescence may result from cumulative oxidative stress causing DNA damage or from disruption of cellular processes involving mitochondrial function which can lead to apoptosis (Carames et al., 2010). Consistent with this possibility, chondrocytes with senescent characteristics and disrupted mitochondrial function accumulate in the articular cartilage of aged mice, consistent with generally decreased regenerative potential of tissues as aging occurs (Jeon et al., 2017). Adding to the complexity of the role of cellular senescence in articular cartilage degeneration is the observation that senescent cells also accumulate in the articular cartilage of the joints of mice after ACL transection-induced post-traumatic OA, demonstrating that chondrocyte senescence is not just an aging response, but also is an injury response (Jeon et al., 2017). The complication of senescence among chondroprogenitor cells in osteoarthritis may mean

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that selective removal of these cells, and/or suppression of their activity, may be necessary to facilitate repair by endogenous progenitors that still retain chondrogenic potential. There is excitement in the possibility that endogenous progenitor cells naturally present in the joint and/or articular cartilage will 1 day prove to be suitable sources for repair of articular cartilage damage and disease. More research is needed to understand the activity and nature of these progenitor cells, and the signals that influence them, in order to realize this potential. Meanwhile, since it is clear that endogenous progenitor cells are not by themselves sufficient to repair articular cartilage damage, current clinical approaches are focusing on strategies to replace damaged or lost articular cartilage with exogenously-supplied cartilage cells or progenitors.

Cell-Based Osteoarthritis Treatments There are several approaches in use and/or under investigation, in which replacement of lost cartilage is attempted with cartilage grafts or cells. One of the challenges of these approaches is that often, it is only at the end-stage of OA when articular cartilage loss encompasses large areas of the joint surfaces that patients report to their doctors, who have little to offer at that point other than total joint replacement with a prosthetic joint. Because the prosthetics have a limited lifespan and revision surgeries tend to have poor outcomes, patients are encouraged to wait as long as possible before having joint replacement surgery, meanwhile enduring pain and disability that compromises their quality of life. Accordingly, a way to prevent widespread articular cartilage loss by correcting it before it spreads would be an important clinical advance. Unfortunately, there are no available biomarker assays that are validated to detect early cartilage damage. Cell-based repair of focal articular cartilage defects as an approach to prevent or slow OA progression is therefore only feasible at the present time for patients who have experienced a known traumatic joint injury and are at risk for PTOA.

Cartilage Grafts Various kinds of approaches have been developed for articular cartilage repair that use cartilage tissue grafts. In these approaches, to ensure that the inserted graft does not fall out, the defect in the patient’s articular cartilage is extended by the surgeon down into the subchondral bone, which provides a better anchoring site than the articular cartilage. The grafted plug consists of both articular cartilage and underlying bone (known as an “osteochondral graft”), and is inserted and press-fit into the defect. Osteochondral Allograft Transfer System (OATS) is the only procedure available for attempted repair of both large and small articular cartilage defects. In this procedure, the donor osteochondral plugs are obtained from a cadaver and inserted into the milled defect in the patient’s articular cartilage. Reports of success of OATS are variable, with a 2015 systematic review of 11 clinical trials of allografts reporting graft survival of nearly 90% at 5 years, with return to activity and overall good functional outcomes (De Caro et al., 2015). Problems with the approach include risk of disease transmission from the donor, difficulty in finding matching histocompatible grafts to minimize tissue rejection, and lack of integration of the cartilage portion of the graft with the adjacent articular cartilage (De Caro et al., 2015). In the best of cases in which patient follow-up histology was performed, integration with the articular cartilage was only via a layer of fibrocartilage “grouting”; in the worst cases, there was no cartilage integration at all (De Caro et al., 2015). In contrast, integration of the bone was typically robust (De Caro et al., 2015). The failure of articular cartilage integration in these clinical reports is consistent with animal studies showing that osteochondral autografts in sheep and goats remain as inert structures within the cartilage defects and fail to induce integrative cartilage repair (Lane et al., 2004; Gelse et al., 2014).

Autologous Chondrocyte Implantation Because of the problems with graft tissue source and lack of integration at the graft/articular cartilage interface, alternatives to osteochondral grafts have been explored including the use of dissociated articular chondrocytes obtained from the patient’s own cartilage, which are implanted into the defect region. This procedure is known as autologous chondrocyte implantation (ACI) and was approved in 1998 under the name Carticel for the repair of relatively small focal defects in young adult patients. In this approach, a small biopsy of articular cartilage is harvested arthroscopically from a non-load bearing region of the patient’s own joint, and digested to release the chondrocytes within, which are then expanded in vitro for 6 weeks, implanted in an open surgery into the prepared defect, and covered with a periosteal membrane (Madeira et al., 2015). While initial reports following the Carticel procedure were favorable and histological biopsy suggested formation of hyaline-like cartilage in some cases (McCarthy and Roberts, 2013), issues with durability and persistence of repair have been frequently reported and large, randomized, double bind trials assessing outcome 14–15 years after ACI have been systematically reviewed with the conclusion that Carticel was not effective in half of the patients with OA, and in fact more patients in the ACI group ended up needing joint replacement than controls (Knutsen et al., 2016). Carticel was replaced in 2017 by MACI (Matrix-assisted ACI), a modified procedure in which the chondrocytes are cultured in a hydrogel sponge, which is then implanted into the defect site using a membrane to secure the cells in place. While functional improvement after MACI was noted in small trials at 5 years (Eber et al., 2017) and at 15 years (Gille et al., 2016), it is concerning that in a study examining 150 biopsies taken within 1½ years of the procedure, nearly 28% contained hypertrophic cells, revealing formation of inappropriate growth plate like cartilage instead of articular cartilage (Eber et al., 2015). A systematic

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review of 15 MACI trials noted variations in surgical procedures and post-surgical therapy, but concluded overall that any improvement in function tended to stop improving and decline at 24 months (Shaikh et al., 2017). The clinical disappointment of Carticel/ MACI has been attributed to the propensity of the harvested articular chondrocytes to de-differentiate during in vitro expansion, even when maintained in the 3D sponge, losing their articular cartilage chondrogenic characteristics, so that they subsequently form fibrocartilage in vivo, which is structurally weaker than hyaline cartilage, or growth plate cartilage, which tends to form bone.

Microfracture The ineffectiveness of adult articular chondrocytes as a cell source for cartilage defect repair has prompted a search for alternate cell sources, with particular interest in progenitor cells with chondrogenic potential, especially mesenchymal stem cells (MSCs). The most commonly-used source of MSCs is adult bone marrow, and the most direct way in which MSCs have been used for articular cartilage defect repair is through a procedure known as microfracture (Mithoefer et al., 2009). In this procedure, which can be done arthroscopically, small holes are punched through the patient’s damaged articular cartilage down into the subchondral bone, allowing blood and bone marrow (containing the MSCs) to flow into the joint, pooling in the holes and eventually filling them with tissue. Systematic review of 28 clinical studies suggested microfracture offers short term improvement in function for patients (Mithoefer et al., 2009), however, as is the case with Carticel/MACI, the tissue formed in microfracture is not hyaline cartilage, but is instead the less-durable fibrocartilage (Mithoefer et al., 2009). Fibrocartilage lacks the proteoglycan-rich matrix characteristic of hyaline/articular cartilage, and contains high amounts of collagen type I, rather than collagen type II. In the human knee joint, fibrocartilage is found in the tendons and menisci, but is not normally found in the articular cartilage (Fox et al., 2009). The fibrocartilage nature of microfracture repair has been conclusively shown using T2-MRI mapping, which detects the zonal architecture of collagen fibers (Welsch et al., 2008). Microfracture is still commonly performed in the clinic, and there are reports that encourage its use (Davatchi et al., 2016; Soler et al., 2016) particularly on a cost: benefit basis (Schrock et al., 2017); however, randomized trials do not support long-term benefit of the procedure (Knutsen et al., 2016). A recent systematic meta-analysis comparing 3–6 clinical trials each of OATS, ACI, and microfracture, with a total of 765 patients, at 2-year follow-up, concluded there is no significant difference in functional outcomes among any of the treatment or control groups (Mundi et al., 2016). A 2017 study compared clinical effectiveness of ACI compared to microfracture (Mistry et al., 2017) in four randomized, controlled trials. The trials compared included large studies such as ACTIVE (Autologous Chondrocyte Implantation/Transplantation Versus Existing Treatment); and SUMMIT (Superiority of Matrix Induced ACI versus MF for treatment of symptomatic articular cartilage defects); with 5-year follow ups. In one study, MACI gave better outcomes than MF, but in another, there was no difference (Mistry et al., 2017). These examples illustrate how rigorous statistical comparisons through systematic meta-analysis of large, randomized, double blind, multi-site, trials make it clear that none of the current clinical approaches (microfracture, OATS and ACI/MACI) are as effective as once believed (Mistry et al., 2017).

MSC Implants Although bone marrow is the most commonly-used source of MSCs for clinical use, it actually contains very few MSCsda 1995 study calculated that only 0.01% of cells in a bone marrow aspirate are MSCs (Jones et al., 2002). The relative paucity of MSCs in bone marrow may be one reason that microfracture forms fibrocartilage instead of hyaline cartilage. To address this, protocols have been developed to enrich for MSCs using cell surface markers and FACS, or through their predilection to adhere to tissue culture plastic. Adipose tissue, which is readily obtained and contains more MSCs than bone marrow (Ruetze and Richter, 2014) has been considered as an alternate MSC source to bone marrow. In a recent study, patient outcomes were compared after microfracture with and without addition of adipose-derived MSCs (Koh et al., 2016). At 2-year follow-up, there were improved clinical scores (patient-reported pain, mobility) with the adipose-MSC-supplemented group, and a slight improvement in repair by second-look arthroscopy, and in morphology by tissue biopsy (Koh et al., 2016). Infrapatellar fat pad and cord-blood are also being investigated as MSC sources (do Amaral et al., 2017); cord blood appear particularly promising but is not that easy to obtain (Zhang et al., 2011). Another source of MSCs that is gaining attention is the joint synovium (De Bari et al., 2001; Sakaguchi et al., 2005). Synovial cells are highly chondrogenic (Huang et al., 2017), and since they can be harvested fairly easily through arthroscopy from the patient’s joint without donor site morbidity, they may offer a clinically-feasible autologous MSC source for articular cartilage repair (De Sousa et al., 2014) Hundreds of studies have been carried out testing the repair ability of MSCs obtained from various sources, and implanted into articular cartilage defects in small animals like rats, mice, and rodents; and in large animals such as the horse, dog, minipig, and sheep (e.g., Wilke et al., 2007; Ha et al., 2015; Zorzi et al., 2015; Kazemi et al., 2017). Clinical efficacy of MSC-mediated articular cartilage defect repair has also been investigated in patients, mostly using implants of bone-marrow or adipose-derived MSCs, but more recently also using synovial-derived MSCs (Bornes et al., 2014; Sekiya et al., 2015). Systematic reviews have attempted to collate the multiplicity of these studies. In Pastides et al. (2013), 36 preclinical studies including 21 small animal and 15 large animal, and 15 human clinical trials, were compared, which overall displayed relatively positive outcomes, including positive functional outcomes reported at 12–48 months after MSC implantation in articular cartilage defects in large animals and humans (Pastides et al., 2013). However, in Bornes et al. (2014), systematic review of 11 human clinical trials produced only limited evidence showing benefit of MSC implantation in articular cartilage defects in humans (Bornes et al., 2014). Both studies noted wide variation in cell preparation, surgical implant methodology, presence or absence of scaffolds or growth factors, study design,

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follow-ups and criteria for functional outcomes reporting; and called for standardization of approaches so that efficacy of MSC implants for articular cartilage repair can be clearly examined.

MSC Injection While implantation of cells directly into articular cartilage defects may offer a therapeutic approach for repair of localized damage, this approach is not feasible in cases of late stage OA, where cartilage loss is widespread throughout the joint. Interest in developing a non-invasive (non-surgical) means for articular cartilage restoration in overt OA has prompted consideration of MSCs as an injectable therapy delivered directly into the joint space. Some of the first use of intra-articular MSC injections were carried out in veterinary practice, where the approach is regularly used on horses and dogs (Cryanoski, 2013). Mostly anecdotal outcomes suggested benefit of the approach in terms of reducing pain and increasing mobility, encouraging consideration of intra-articular autologous MSC injections in human OA patients. To date, several clinical trials have been carried out to examine the safety and efficacy of knee MSC intra-articular injections (Koh et al., 2013; Jo et al., 2014; Afiza and Hu, 2016; Davatchi et al., 2016; Soler et al., 2016; Jo et al., 2017; Cui et al., 2016), finding them generally safe and leading to short-term improved patient scores on subjective criteria and on cartilage morphology as assessed by MRI or biopsy. In Koh et al. (2013), 18 OA patients received intra-articular injections of autologous MSC derived from infrapatellar fat pad, and were followed up at an average of 24 months later, reporting improved pain and function scores (Koh et al., 2013). In 2016 a study was reported with 15 patients who each received a single injection of bone marrow derived MSC (Soler et al., 2016). At 6 months there was significant patient reported reductions in pain, and increased functional scores, which persisted for up to 12 months, at which time T2 MRI mapping also revealed signs of cartilage thickening (Soler et al., 2016). In another 2016 study of four patients with moderate to severe OA, improvement in mobility was noted 6 months after autologous MSC intra-articular injection, but this improvement declined thereafter, and by 5 years, there was no improvement in range of motion between the injected or non-injected knee (Davatchi et al., 2016). The only study that employed biopsy as an outcome was Jo et al. (2014, 2017), in which 18 OA patients with a mean age of 61 years received a single injection of autologous MSC derived from buttock adipose tissue, at either a high, medium or low dose, and were followed for 2 years. There was significant improvement in functional scores at 6 months for the high dose, which was paralleled by MRI structural outcome, and biopsies taken from the region of damage before and after MSC injection showed thickening of the articular cartilage and synthesis of new matrix following the MSC treatment (Jo et al., 2014). While promising, the biopsies also revealed increased collagen type I production consistent with potential fibrocartilage formation (Jo et al., 2014). At 12 month follow-up, functional outcomes declined in the lower dose groups, and improved pain and function scores in the high dose group plateaued until final followup at 24 months. Disturbingly, the size of the defects increased regardless of the treatment dose, some by as much as 78% by 2 years (Jo et al., 2017), raising concerns over durability of the repair tissue. Since this study had no control group, it is impossible to determine if the defect size also increased in the absence of MSC treatment (Jo et al., 2017), but this study emphasizes the need for longterm follow-up and tissue biopsy as monitoring assays Systematic mega-analyses of studies reporting effects of MSC intra-articular injections for OA treatment have been carried out to attempt to obtain meaningful consensus on efficacy outcomes (Afiza and Hu, 2016; Cui et al., 2016; Goldberg et al., 2017). A metaanalysis of 18 clinical studies concluded that MSC treatment generally ameliorated overall outcomes of patients with knee OA, including pain and functional evaluations, particularly at 12 and 24 months follow-up (Cui et al., 2016). Another systematic review of clinical studies concluded that safety of MSC intra-articular injections is clearly established (Afiza and Hu, 2016), suggesting that the focus should now be on optimizing MSC efficacy for OA treatment. However, the largest systematic review to date, which analyzed 252 studies (100 in vitro studies; 111 animal studies, and 31 clinical studies) on the use of MSCs for cartilage repair and regeneration as either implant or intra-articular injection, found tremendous variability in the source of MSCs used, their preparation and dose, and choice of allogenic or autologous cells, which combined to make drawing definitive conclusions about the efficacy of MSC treatment for cartilage restoration difficult, calling for a return to basic science and better communication among pre-clinical and clinical stakeholders in order to move the field forward in a meaningful way (Goldberg et al., 2017).

Donor Versus Host? An example of how basic science can (and should) inform the clinic is highlighted by studies in which MSCs that were injected into the osteoarthritic joints of animals were tracked to determine where they end up and how long they persist. Although this is a logical question to ask before injecting MSCs into the joints of patients, in actuality it was mainly after results started to surface in the clinic that analysis of fluorescently or otherwise labeled MSC, which can be tracked in the joints of animals, was rigorously pursued. Some studies reported robust colonization of the joint following MSC intra-articular injection. For example, in the Harley spontaneous OA guinea pig model, labeled MSCs in a hyaluronan gel were injected into the diseased joints, and donor MSCs were detected histologically 3 weeks later in the upper and middle zones of the articular cartilage, where they appeared to integrate and make new cartilage matrix (Sato et al., 2012). Another study examined distribution of fluorescently-labeled human adipose-derived MSCs injected into the joints of rats following OA-inducing surgical disruption of the knee ligaments, using IVIS, and found strong signal present in the joints at 35 days, which diminished markedly thereafter, although histologically, some labeled cells persisted in the menisci and cartilage for up to 10 weeks (Li et al., 2016a,b). These studies suggested that injected MSCs do have the potential to engraft after inter-articular injection, however, most studies revealed surprisingly weak engraftment of intra-articularly-injected MSCs, especially considering that the studies uniformly reported improved articular cartilage healing. For example, in a rat model in which OA was

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induced by injection of mono-iodoacetate, signal from radiolabeled MSCs decreased just 3 days after injection, and was virtually absent 3 weeks later (van Buul et al., 2014). Moreover, in another study, synovial-derived GFP-labeled MSCs were injected into the joints of a mouse surgical OA model, and were detected in only isolated locations at the injury site at 2 weeks, and not at all at 4 weeks (Mak et al., 2016). Similarly, dye-labeled MSCs injected into the joints of a mouse surgical OA model were virtually absent when examined histologically at day 3 and day 7, despite impressive cartilage restoration when examined 21 days later (Diekman et al., 2013). It is particularly interesting that in most cases, if engraftment of intra-articularly-injected MSCs occurred, it was most likely to be in non-cartilage joint tissue such as synovium, menisci or ligament. For example in the study by Diekman et al. (2013), injected MSC appeared only in the synovium, ligaments and menisci, not cartilage; and similarly, in another study, labeled MSCs injected in the joints of a sheep OA model were found in the synovial lining, dorsal fat pad, and meniscus, but not in cartilage or bone (Delling et al., 2015). How then, do we reconcile the inconsistent engraftment of intra-articularly-injected MSCs, with consistent clinical reports of improved joint function, and in animal studies, even improved cartilage morphology? One possible explanation is that the injected MSCs are acting as a general suppressor of inflammation. MSCs have long been known as immunomodulatory (Contreras et al., 2016), and release of anti-inflammatory compounds like IL-10 may serve to suppress the inflammation that occurs coincident with OA signs, restoring homeostasis within the joint, and establishing an environment which is conducive to anabolic articular cartilage responses. This activity is consistent with the “whole joint” view of OA disease, and the idea that tissues other than the articular cartilage are responsible for secreting the harmful signals and factors that create the toxic environment within the diseased joint. Indeed, Hausler et al. (2013) showed that ligaments, menisci and especially synovium were the source of 80% of the degradative enzymes produced in the joint in response to trauma. It is probably not a coincidence that in the cases where MSC integration is observed in the OA joint after injection, it is within these tissues, raising the possibility that the MSCs actively “home” to these regions in response to the pro-inflammatory chemokines the tissue express. Consistent with this possibility, in the study of Li et al. (2016a,b), MSC engraftment after intra-articular injection was weaker in joints that did not experience surgery, suggesting injury stimulated the engraftment response (Li et al., 2016a,b). Remarkably, a recent study using a non-human primate (monkey) model suggests MSCs may even be able to home to the surgically-manipulated joint from the bloodstream (Fernandez-Pernas et al., 2017). Perhaps the most surprising studies correlating MSC engraftment with cartilage outcome are those that have examined the persistence of donor MSC’s after direct implantation into articular cartilage defects, where homing or retention in the joint are not at issue. In Ostrander et al. (2001), using presence of the Y-chromosome in male cells as a marker, female rabbit perichondrial-derived MSCs were implanted into an articular cartilage defect in a male animal, and followed over time. After 28 days, the defect region showed some healing, but less than 15% of the original population of implanted male donor cells remained (Ostrander et al., 2001). In another study, rabbit MSCs were labeled with a fluorescent dye and implanted into a full thickness defect within a polymer scaffold. Formation of new collagen type II-containing tissue was initiated, which contained labeled donor cells, but this tissue was slowly replaced over a 2-month period by tissue that did not contain labeled donor cells (Tatebe et al., 2005). A similar result was reported by Niemietz et al. (2014) using human articular chondrocytes implanted into minipig articular cartilage defects, which were 95% gone after 6 weeks despite repair of the defect region by host cells (Niemietz et al., 2014). Using a sophisticated double transgenic labeling system, Zwolanek et al. (2017) demonstrated that healing of articular cartilage defects in the joints of rats 6 months following intra-articular injection of immune-compatible MSCs was exclusively the result of host cells, not donor cells. Thus, MSC implantation into articular cartilage defects suggest a conundrum similar to that presented by MSC intra-articular injection into the OA joint: some healing occurs, but few repair cells are present in the healed cartilage tissue. The interpretation is that the introduced cells are providing a source of factor(s) that somehow facilitates the ability of the host to repair damaged tissue. In the case of intra-articularly injected MSCs, it is possible that this factor could be related to the immune-suppressive activity of MSCs (Contreras et al., 2016) which is likely responsible for their chondroprotective activity when injected into the chronicallyosteoarthritic joint. In the case of implanted MSCs, the factor(s) in question could be some type of growth factor that acts in a paracrine fashion (Xu et al., 2016) to locally recruit and/or simulate cells in the injured region, which subsequently carry out repair.

Challenges in MSC-Based OA Interventions While MSC-based approaches for future articular cartilage repair or treatment of advanced OA may offer potential, evidence for their efficacy in the clinic or the lab remains inconsistent, and inconclusive. This is due in part to the inability to systematically compare published studies in a meaningful way because of tremendous variability in study design, rigor and/or reporting. Some of these variabilities include the species being studied (human, or large or small animal, and if animal, what model? -spontaneous OA, ACL rupture or transection, or osteochondral defect); the OA condition (idiopathic/primary OA or injury-induced PTOA); the delivery approach (intra-articular injection, grafts, scaffolds, gels); and the outcomes (qualitative patient scores, MRI, histology). Studies examining clinical outcomes tend to use small cohorts which may not have controls, and typically are only a few years in follow-up. Greater rigor in the form of large, controlled, randomized, multi-center, double-blind clinical trials are needed to make definitive conclusions. Long-term follow-up over 10–15 years is required to satisfactorily assess durability. A particular challenge is the need for second-look arthroscopy and tissue biopsy, invasive procedures which are difficult to justify (particularly biopsy), but which are essential to track and monitor early stages in potential repair that might dictate future directions. Only biopsy can provide information about the nature of the tissue structure and its molecular profile, in order to conclusively show that the repair tissue made is comparable to native articular cartilage.

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Additional confounders in experimental and clinical studies using MSCs as a stem-cell based OA intervention relate to the nature and biology of MSCs themselves. Consensus has not emerged on which is the optimal MSC tissue sourcedbone-marrow, adipose tissue, cord-blood, or synoviumdand each has advantages and disadvantages in terms of chondrogenic potential and availability. MSC are relatively rare cells, and enrichment is likely necessary to obtain sufficient MSCs from these sources, but selection protocols use different cell surface markers (if MSC are isolated using FACS) or different passage number (if MSC are isolated via adherence to tissue culture plastic). The age and health status of the donor is also a factor. MSCs tend to proliferate poorly in vitro, especially when obtained from older patients (Payne et al., 2010), making it difficult to expand autologous cells to a number sufficient for clinical use, and the culture conditions used for expansion vary greatly (e.g., different growth factors, cell density, etc.). These issues result in highly variable MSC preparations from patient to patient, or even from the same patient. A major disadvantage of MSCs is their inherently unstable nature, and their tendency to undergo fibrochondrogenesis and/or terminal hypertrophic differentiation, which leads to formation of repair tissue that lacks the necessary durability of true articular cartilage. Ideally, in the future, a source of MSCs for cartilage repair or OA treatment will be developed that enables generation of large numbers of homogeneous progenitors with enhanced chondrogenic potential, uniformly characterized for dose and activity, that could be used by physicians as an “off-the-shelf” product, perhaps supplied from an allogenic cell bank.

Alternative Cell-Based Strategies Because of their unlimited capacity for self-renewal, human pluripotent stem cells (embryonic stem cells, ESC, or induced pluripotent stem cells, iPSC) could provide a readily obtainable allogenic cell source for articular cartilage repair or OA treatment (Mobasheri et al., 2014; Lietman, 2016; Murphy et al., 2017). Estimates are surprisingly low of the number of cell lines ( 200) needed to provide allogenic cells to suit most of world’s MHC haploptypes (Taylor et al., 2005; Nakajima et al., 2007; Lin et al., 2009). Moreover, protocols have been developed for efficient generation of chondroprogenitor cells from pluripotent stem cells, including some that create cells with enhanced potential for forming stable cartilage tissue (Gong et al., 2010; Oldershaw et al., 2010; Sternberg et al., 2012; Koyama et al., 2013; Toh et al., 2010; Yamashita et al., 2015; Guzzo et al., 2014; Craft et al., 2015; Nam et al., 2017). Preclinical studies in small animals have begun to assess the safety and efficacy of implanting pluripotentderived MSCs or chondroprogenitors into articular cartilage defects, and/or injecting them into the joints of OA models, finding the cells do not form tumors and generally there are reported improvements in cartilage tissue histology (Toh et al., 2010; Uto et al., 2013; Cheng et al., 2014; Yamashita et al., 2015; Gibson et al., 2016). Several studies have examined persistence of the human-derived cells, with variable but often little long-term persistence within the joint (Toh et al., 2010; Cheng et al., 2014; Yamashita et al., 2015; Gibson et al., 2016). This suggests ESC-derived cells may act via a similar paracrine mechanism of action as suspected for native allogenic or autologous MSCs. Pluripotent-derived chondroprogenitor cells may offer a readilyobtainable, infinitely-expandable exogenous cell source for OA intervention, with unique advantages including the potential to differentiate the cells directly into articular-chondrocyte progenitors, and/or to screen and select cell lines with exceptional chondrogenic ability. These advantages emphasize that pluripotent-based cell sources, as compared to native MSCs, should continue to be considered as an alternativedor even potentially preferabledcell source for OA intervention.

Future Directions As no therapy yet exists to treat widespread, overt OA, the key to OA management will be prevention, by restoring local articular cartilage defects before they progress to irreversible disease. Current clinical methods for articular cartilage defect repair (OATS, ACI/ MACI, MF) have not met this need. Approaches using progenitor cells such as MSCs, or MSC-derived chondroprogenitors have some promise, but there is a lack of consistency in reported efficacy studies which makes it difficult to determine which of the many different approaches is most likely to succeed. The concept of using the body’s own articular chondrocytes to regenerate lost articular cartilage, rather than just repair it, is exciting; however, it is clear that endogenous cells cannot effectively heal articular cartilage on their own, and that some stimulating signal(s) is required. The complexity of cellular behaviors suggests that multiple signals will likely be needed to stimulate native chondroprogenitors to a sufficient extent that a successful regenerative response is obtained. Until the identity of each of the necessary signals is known, the best way to provide these signals to the articular chondroprogenitors in situ may well be through paracrine activity from locally-delivered MSCs, implanted or injected into the joint. In this regard, it is also essential to consider the environment of the joint and the participation of other joint tissues including synovium, menisci and ligaments as either hostile or nurturing to articular neochondrogenesis. Modifying agents will likely be needed to restore joint homeostasis before any approach for articular cartilage defect repair or regeneration will be able to proceed successfully. Delivery of such agents will need to be precisely timed to achieve desired effect, with consideration of the natural anabolic and catabolic phases of OA progression after injury. Until this complex “interactome” of factors, stimuli and modifiers is understood and defined, there will be a role for exogenous cells to supply the requisite pro-reparative cues. However, it will be necessary to find an exogenous cell source that is reproducible and predictable in its behavior, so that it can be safely dosed and controlled; as well as readily obtained, so that costs can be minimized; and a uniform protocol for handling and delivering the cells will need to be established and validated, so that the therapy can be “off-the-shelf” and available to all patients. Serious gaps in our knowledge of articular cartilage biology and development, and the mechanisms by which cartilage responds to injury, have contributed to establishment of dogmas that have misdirected progress in cartilage repair and

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regeneration strategies. Historically, the anecdotal success of MSC injections into the joints of large animals, especially in veterinary practice, led to trials of autologous MSC-based injections or implantations in human patients even prior to validation of the approach at the histological level in small animals. The observation that articular cartilage morphology appeared to be improved after MSC injection or implantation in human or animal studies led to the assumption that the exogenous MSCs formed the new cartilage tissue, reinforcing a dogma that articular cartilage cannot heal itself, and that accordingly, any viable cartilage repair treatment would require exogenous cells. It was years before this dogma was questioned as the result of cell-tracking studies in small animal models that definitively showed that exogenously-introduced cells, even when implanted directly into the articular cartilage, eventually disappear and are replaced by host cells. This observation prompted a re-examination of articular cartilage development which in turn revealed unexpected potential lability in the superficial zone of adult articular cartilagedthereby launching an entire new research direction examining native chondroprogenitor self-healing. This history illustrates the hazards of pre-conceived bias in mis-guiding and delaying research progress, and emphasizes the importance of mechanistic understanding and validation gained by basic science, prior to human clinical testing. Going forward, and as diagrammed in Fig. 2, continued studies in genomics, development, immunology and cartilage biology will be needed to fully understand the cellular mechanisms that underlie OA disease. Then, to translate this basic information into feasible therapeutic interventions that can be tested in preclinical studies, input from the field of bioengineering will be needed to develop scaffolds and materials for support or delivery of cells and/or co-factors, and to better monitor and evaluate cartilage restoration and tissue integrity using biomechanical testing. Careful and rigorous pre-clinical studies will be essential to test the most promising interventions, with standardization of methods and outcomes so that studies can be meaningfully compared. In particular, there is a need for discussion of what doesn’t work, which is just as important as what doesdbut often is not reported. Most importantly, in order to efficiently translate research advances to the clinic, better communication among all stakeholders is needed to integrate knowledge in basic biology and applied engineering towards the common goal of a preventative therapy, or even a lasting cure, for the many individuals at risk for, or already afflicted by, osteoarthritis.

RESEARCH

CLINIC

Fig. 2 Input from fields of molecular biology, cartilage biology, and developmental biology are needed to inform research in normal joint homeostasis, osteoarthritis disease mechanisms, and cartilage injury responses, in order to successfully and rationally translate discovery to clinical cures.

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Further Reading Goldberg, A., Mitchell, K., Soans, J., Kim, L., & Zaidi, R. (2017). The use of mesenchymal stem cells for cartilage repair and regeneration: A systematic review. Journal of Orthopedic Surgery and Research, 12, 39. Li, L., Newton, P., Bouderlique, T., Sejnohova, M., Zikmund, T., Kozhemyakina, E., Xie, M., Krivanek, J., Kaiser, J., Qiang, H., Dyachuj, V., Lassar, A. B., Warman, M. L., Barenius, B., Adameyko, I., & Chagrin, A. S. (2017). Superficial cells are self-renewing chondrocyte progenitors, which form the articular cartilage in juvenile mice. FASEB Journal, 31, 1067–1084. Jo, C. H., Lee, Y. G., Shin, W. H., Kim, H., Chai, J. W., Jeong, E. C., Kim, J. E., Shim, H., Shin, J. S., Shin, I. S., Ra, J. C., Oh, S., & Yoon, K. S. (2014). Intra-articular injection of mesenchymal stem cells for the treatment of osteoarthritis of the knee: A proof-of-concept clinical trial. Stem Cells, 32, 1254–1266. Jo, C. H., Chai, J. W., Jeong, E. C., Oh, S., Shin, J. S., Hackjoon, S., & Yoon, K. S. (2017). Intra-articular injection of mesenchymal stem cells for the treatment of osteoarthritis of the knee. A two-year follow up story. American Journal of Sports Medicine, 45, 2774–2783. Carbone, A., & Rodeo, S. (2017). Review of current understanding of post-traumatic osteoarthritis resulting from sports injuries. Journal of Orthopaedic Research, 35, 397–405. Courties, A., Sellam, J., & Berenbaum, F. (2017). Metabolic syndrome-associated osteoarthritis. Current Opinion in Rheumatology, 29, 214–222. Feng, C., Luo, X., Xia, H., Lv, X., Zhang, X., Li, D., Wang, F., He, J., Zhang, L., Lin, X., Lin, L., Yin, H., He, J., Wang, J., Cao, W., Wang, R., Zhou, G., & Wang, W. (2017). Efficacy and persistence of allogeneic adipose-derived mesenchymal stem cells combined with hyaluronic acid in osteoarthritis after intra-articular injection in a sheep model. Tissue Engineering Part A, 24, 219–233. https://doi.org/10.1089/ten.tea. Jiang, Y., Cai, Y., Zhang, W., Yin, Z., Hu, C., Tong, T., Lu, P., Zhang, S., Neculai, D., Tuan, R. S., & Ouyeang, H. W. (2016). Human cartilage-derived progenitor cells from committed chondrocytes for efficient cartilage repair and regeneration. Stem Cells Translational Medicine, 5, 733–744. Mirando, A. J., Liu, Z., Moore, T., Lang, A., Kohn, A., Osinski, A. M., OKeefe, R. J., Mooney, R. A., Zuscik, M. J., & Hilton, M. J. (2013). RBPjk-dependent notch signaling is required for articular cartilage and joint maintenance. Arthritis and Rheumatism, 65, 2623–2633. Ozturk, A., Ozdemir, M. R., & Ozkan, Y. (2006). Osteochondral autografting (mosaicplasty) in grade IV cartilage defects in the knee joint: 2- to 7-year results. International Orthopaedics, 30, 200–204. Toh, W. S., & Cao, T. (2014). Derivation of chondrogenic cells from human embryonic stem cells for cartilage tissue engineering. In K. Turksen (Ed.), Methods in molecular biology: vol. 1307. Human embryonic stem cell protocols. New York, NY: Humana Press.

Reproductive Technologies, Assisted D Pergament, Case Western Reserve University School of Law, Cleveland, OH, USA; and Children’s Law Group, LLC, Chicago, IL, USA © 2019 Elsevier Inc. All rights reserved.

Religious Bioethics and ARTs Roman Catholicism Protestantism Judaism Islam Greek Orthodoxy Access to ARTs Access to ARTs for Gays, Lesbians, and Unmarried Persons Access to IVF for Postmenopausal Women Gamete Donation Donor Anonymity Embryo Donation Three-Person IVF Surrogacy Intrafamilial Collaborative Reproduction Fertility Preservation Posthumous Collection and Use of Reproductive Tissues Risks of ARTs Multifetal Pregnancies Birth Defects Preimplantation Genetic Diagnosis References Relevant Websites

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Glossary Artificial insemination (AI) The deliberate introduction of semen into a female vagina or oviduct for the purpose of achieving a pregnancy through fertilization by means other than natural insemination. Assisted reproductive technologies (ARTs) Methods to achieve pregnancy by artificial or partially artificial means. Autonomy, reproductive Autonomy is a moral, political, and bioethical philosophical construct. Within the context of human reproduction, it is the capacity of a rational individual to make an informed, uncoerced decision. Beneficence Beneficence is action that is done for the benefit of others. Collaborative reproduction Achieving pregnancy with the help of a third party to provide gametes or the uterus; reproduction involving more than two biogenetic parents. Cross-border reproductive care The movement of people across national borders in search of fertility and related treatments in other jurisdictions to obtain services often at lower costs or to circumvent legal restrictions on the treatment being sought. The terms transnational reproduction, reproductive travel, and procreative tourism are also used. Cryopreservation Freezing at very low temperatures, such as in liquid nitrogen (
196  C), to keep embryos, oocytes, or sperm viable. Fertility preservation Fertility preservation is the effort to help patients undergoing treatment with gonadotoxic agents retain their fertility or ability to procreate. Fertility preservation is also used by women choosing to delay childbearing for social reasons to extend their fertility. Gestational carrier or gestational surrogate A woman who carries a pregnancy for another couple or person. The pregnancy is derived from the egg and sperm of the couple or gamete donors. Although she carries the pregnancy to term, she does not have a genetic relationship with the resulting child. Heterologous insemination Refers to artificial insemination in which semen is used from a donor. Homologous insemination Refers to artificial insemination in which semen is used from the woman’s husband or partner, also called therapeutic donor insemination.

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Human leukocyte antigen (HLA) The major histocompatibility complex in humans. In some diseases treated by hematopoietic stem cell transplantation, preimplantation genetic diagnosis may be used to assess embryos in an attempt to conceive a child with matching HLA to serve as donor for a sibling. Intended parents People who use assisted reproduction technologies to create a child whom they intend to parent, whether or not they have a genetic or biological relationship to the child; also called contracting or commissioning parents. Intrafamilial collaborative reproduction Achieving pregnancy with the help of a third party to provide gametes or the uterus provided by a family member; also called intrafamilial medically assisted reproduction. Intracytoplasmic sperm injection (ICSI) A micromanipulation procedure in which a single sperm is injected directly into an oocyte to attempt fertilization, used with male infertility or couples with prior IVF fertilization failure. In vitro fertilization A process in which an egg and sperm are combined in a laboratory dish to facilitate fertilization. If fertilized, the resulting embryo is transferred to the uterus. Microepididymal sperm aspiration (MESA) A microsurgical procedure used to collect sperm in men with blockage of the male reproductive ducts such as prior vasectomy or absence of the vas deferens. Used in IVF/ICSI procedures. Multifetal pregnancy reduction A procedure to reduce the number of fetuses in the uterus. This procedure is sometimes performed on women who are pregnant with multiple fetuses who are at increased risk of late miscarriage or premature labor. These risks increase with the number of fetuses; also known as selective reduction. Next-generation DNA sequencing A colloquial term that is used to describe techniques to rapidly compare genetic sequences among multiple genomes and identify germline and somatic variants of interest, such as single nucleotide polymorphisms (SNPs), insertions and deletions (indels), copy number variants (CNVs), and other structural variations. Nonmaleficence Nonmaleficence means to do no harm by refraining from providing ineffective or harmful treatments as determined by whether the benefits of treatment outweigh the burdens. Ovarian stimulation The administration of hormone medications that stimulate the ovaries to produce multiple oocytes. This is sometimes called controlled ovarian hyperstimulation, ovarian induction, or follicular recruitment. Preeclampsia Is a condition characterized by high blood pressure and significant amounts of protein in the urine of a pregnant woman. If left untreated, it may develop into eclampsia, a life-threatening occurrence of seizures during pregnancy. Preimplantation genetic diagnosis (PGD) Refers to the genetic analysis of embryos prior to implantation. It is performed by an embryologist using a variety of methods for removal of one or two cells from an embryo, which are then screened for genetic abnormalities or HLA type. PGD is performed in conjunction with IVF. Surrogacy Is an arrangement in which a woman carries and delivers a child for another couple or person. The surrogate may be the child’s genetic mother (traditional surrogacy) or she may be genetically unrelated to the child (gestational surrogacy). Testicular sperm extraction (TESE) Operative removal of testicular tissue in an attempt to collect living sperm for use in an IVF/ICSI procedure.

Assisted reproductive technologies (ARTs) are methods to achieve pregnancy by artificial or partially artificial means. Although artificial insemination (AI) had been used in humans since the eighteenth century, infertility was historically an area of medicine with few treatment options. This dramatically changed on 25 July 1978 with the announcement of the birth in England of a baby girl conceived through in vitro fertilization (IVF); she is popularly referred to as the world’s first test tube baby. In the decades since, ARTs have become commonplace technologies. It is estimated that around 5 million babies have been born since the first IVF baby was born (Sandin et al., 2013). The use of ARTs also enables individuals to extend or to preserve their fertility or single individuals and same-sex couples to conceive their own biologically related offspring. Other forms of ARTs are used to facilitate preimplantation genetic diagnosis (PGD) of embryos. The development of ARTs raises questions about whether it is proper for science to interfere, and to what degree, with reproduction. The use of ARTs generates myriad ethical, religious, legal and regulatory, and social issues concerning the nature of human life and reproduction, justice and equality, as well as genetic and social familial relationships. The use of ARTs separates human conception from the sexual act, thereby challenging traditional paradigms regarding gender roles and family relationships. The use of donor gametes and embryos raises questions about family structures and how lineage, inheritance, and citizenship are determined. In many modern societies, reproductive choices are most often personal decisions that individuals make in accordance with their own ethics and beliefs. Whether and how ARTs are used in an effort to have a child is affected by a confluence of laws, medical guidelines, economic factors, and cultural and religious norms. There is wide variability in access to ARTs and the laws regulating the use of ARTs among different countries and even states/provinces within a single country. In countries with publicly funded health care systems, reproductive choices, including the use of ARTs, may not be considered purely personal decisions, but may be viewed as public health issues. The number and variety of ARTs is remarkable, therefore a detailed discussion of all technologies and the ethical issues that arise from them is precluded in this article. Accordingly, the focus is on AI, IVF, intracytoplasmic sperm injection (ICSI), cryopreservation, and PGD.

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Religious Bioethics and ARTs The use of ARTs challenges many religious doctrines governing martial relationships, family formation, and parenthood. The fact that IVF and other forms of ARTs involve the creation and potential destruction of embryos has also generated discourse on the moral status of embryos. Religious doctrines play an essential role in considering many bioethical challenges raised by ARTs as these views play a role in the decisions individuals may make with regard to using ARTs and have influenced legislative and regulatory responses.

Roman Catholicism One of the major forces of opposition to ARTs is the Roman Catholic Church. The Catholic Church views embryos as individual human beings with the right to life and that they should be defended against anything that may threaten their human dignity. The Catholic Church opposes IVF based on the argument that it separates procreation from the unitive and procreative elements of reproduction within a marriage. The Catholic Church rejects surrogacy, gamete donation, and posthumous reproduction as contrary to the values of marriage, the unity of spouses, and the dignity proper to parents and child. The Catholic Church considers homologous insemination less reprehensible than other ARTs that require the use of donor sperm or a surrogate mother. Yet, even homologous insemination remains morally unacceptable because of the dissociation of the sexual act from procreation except for situations in which technical means are not a substitute for the conjugal act but help facilitate conception (Vatican, 1987).

Protestantism Mainstream Protestant denominations view embryos as having the potential for personhood and therefore deserving of respect. The Anglican Church supports IVF; however, it takes the position that no embryo should be used, selected against, or destroyed for trivial reasons. Moreover, the welfare of the conceived child is paramount above the procreative interests of the adults involved. The major synods of the Lutheran Church generally accept IVF and agree that marriage partners may use IVF to conceive. The Methodist and Presbyterian Churches accept IVF but advocate that embryos should not be created for research or with the intention of destroying them, but that for those embryos that would have been destroyed, research is permissible. Evangelical and fundamentalist Protestants typically believe that embryos are human beings, equate the potential destruction of embryos with abortion, and advocate for public policies mandating the transfer of excess embryos resulting from IVF to adoptive families.

Judaism Many rabbinical authorities agree that ARTs are permitted under Jewish law. IVF is widely accepted among Reform, Reconstructionist, Conservative, and Orthodox branches of Judaism. There are debates among Orthodox rabbinic authorities concerning whether multifetal pregnancy reduction (MPR) is permitted and when it may be performed (Schenker, 2008). There are also disputes about gamete donation and surrogacy. Traditional interpretations of Jewish law hold that Jewishness is conferred through matrilineal descent particularly from the acts of gestating and birthing the baby. Most Orthodox rabbis prefer that non-Jewish donor sperm be used, to prevent adultery between a Jewish man and a Jewish woman. In Israel, because it is a country with a small population, another concern is preventing genetic incest among the offspring of anonymous donors. This has resulted in widespread use of sperm imported from other countries. Orthodox authorities who permit surrogacy prefer single Jewish women as surrogates to avoid the implications of adultery for married surrogate women and to ensure that the baby is Jewish.

Islam Both Sunni and Shia authorities agree that the use of ARTs should be limited to those with a valid marriage contract. IVF is permitted using the gametes of a married couple, if the procedure is indicated for a medical reason and performed by an expert physician. Embryos may be cryopreserved. They are considered the property of the married couple but they can only be transferred to the woman during the duration of the marital contract. MPR is allowed only if the prospect of carrying a high-order pregnancy is very small or to protect the mother. The use of homologous AI is accepted as it involves a married couple. There is a significant difference between Sunni and Shia Islam regarding gamete donation. Sunni authorities prohibit heterologous insemination because it is outside of the marriage contract. In the late 1990s, some Shia authorities began to permit egg donation under the parameters of a form of temporary marriage between a man and an unmarried woman that is permissible under some Shia interpretations of Islamic law (Inhorn, 2006).

Greek Orthodoxy The Greek Orthodox Church posits that the exact moment human life begins is unknown but views embryos as having a human identity as a person under development, therefore rejecting embryo research and cyropreservation. The Orthodox Church has reservations about homologous and heterologous AI and IVF, but does not prohibit married infertile couples from using these techniques. The Orthodox Church rejects surrogacy, gamete donation, selective fetal reduction, posthumous reproduction, the

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use of IVF by postmenopausal women, as well as the use of IVF by same-sex couples and single women, and expresses reservations toward ICSI and PGD (Nikolaos, 2008).

Access to ARTs Some feminist bioethicists argue that because women carry greater burdens with regard to biological and social reproduction, bioethical issues surrounding ARTs are essentially debates over women’s rights to bodily integrity and reproductive autonomy. Related concerns also include how the use of ARTs may affect women’s overall health and what impact ARTs have on women’s efforts to achieve economic and political equality. Several feminists criticize ARTs as another means of subjugating women by emphasizing their social roles as mothers. Yet, others view ARTs as useful tools in achieving women’s liberation and autonomy by allowing them to control their reproduction. Other bioethicists argue that issues surrounding procreation involve human rights, regardless of the gender, because the desire to have children and to participate in family life is a fundamental biological and social aspect of being human. Reproductive freedom or procreative liberty is essentially the individual right to have or avoid having children and to have the information and means to do so. The right to procreate can be construed as a negative or positive right. As a negative right, it is a right against coercive interference with decisions regarding procreation and reproduction. As a positive right, it provides an entitlement to assistance in reproduction. One of the most obvious and long-standing ethical challenges surrounding ARTs is the inequitable access to treatments and whether high costs interfere with procreative rights that may include access to ARTs. Some argue that the financial and psychological costs of infertility enable the development of a fertility industry that profits by exploiting desperate patients, but actually limits their reproductive autonomy by providing what may be ineffectual and sometimes even dangerous treatments at exorbitant costs. If there are rights to ARTs, how should potentially scarce resources be allocated becomes an essential question; specifically, does the state have an obligation to mandate insurance coverage for treatment or fund the use of ARTs through the public health system. Some argue that access to publicly funded ARTs is justified by the social advantages resulting from facilitating reproductive choice. A related argument is that state funding of ARTs fulfills the state’s obligation to promote good health, including reproductive health. Another, more practical, argument is that public funding of ARTs will help halt declining fertility rates and promote the addition of productive members of society. This may be an increasingly acceptable justification for public funding of ARTs in countries experiencing a negative or flat population growth rate and seeking to develop public policies to alleviate the financial dilemmas of an aging population (Hoorens et al., 2007). The funding structure for ARTs is highly variable among different countries. Many European countries with generous public financing of health care tend to restrict access to ARTs to a greater degree by imposing specific requirements governing age limits and relationship status, and imposing other restrictions such as refusing treatment to smokers or obese patients. In these countries, patients denied publicly funded treatment may use private facilities or travel abroad to engage in cross-border reproductive care to find more affordable options. In the United States, 15 states mandate insurance coverage for some forms of infertility diagnosis and ARTs and federal tax laws allow deductions for some expenses associated with ARTs. The recently enacted Affordable Care Act does not clearly mandate infertility treatment as an essential health benefit and individual state health exchanges can decide whether or not to pay for ARTs (Omurtag and Adamson, 2013).

Access to ARTs for Gays, Lesbians, and Unmarried Persons Whether it is acceptable for individuals or couples to reproduce regardless of their sexual orientation or marital status raises questions about the reproductive rights and interests of unmarried individuals and same-sex couples and the value of practicing nondiscrimination. It also involves questions regarding the welfare of any potential offspring. The answers vary widely throughout the world and often depend on culturally specific norms regarding single motherhood, homosexuality, and nontraditional families. Access to ARTs may depend on what types of family units are eligible for publicly funded treatments, whether gamete and embryo donation and gestational surrogacy is permitted, and whether same-sex civil unions or marriages are recognized. The decision to use ARTs may also be influenced by the legal status of resulting offspring; for example, some countries reject the argument that the citizenship of the intended parents should determine citizenship and deny citizenship to children born to surrogate mothers in other countries. Other countries or states/regions only recognize one legal parent in same-sex families. In comparison, other countries or states/regions have created various legal procedures such as second parent adoptions that result in the legal recognition of both parents. In the United States, private providers may claim that the exercise of professional autonomy justifies limiting treatment to married heterosexual couples. However, professional practice guidelines advise that refusing treatment to patients based solely on sexual orientation and marital status may constitute illegal discrimination and does not have a sound ethical basis. A longstanding argument against allowing single individuals and same-sex couples to use ARTs was the potential harm to children if they are raised in single-parent or same-sex households. Professional guidelines advise that there is no pervasive evidence that children raised by single parents or gays or lesbians are harmed or disadvantaged by that fact alone. The same factors that would prevent anyone from receiving services, such as serious doubt about the ability to parent for reasons unrelated to sexual

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orientation or the fact that the clinic does not offer anyone a particular service, would be the only justification for denying treatment to gays, lesbians, or single people (ASRM, 2009a).

Access to IVF for Postmenopausal Women IVF has allowed women in their 50s and 60s to become pregnant using donor oocytes or embryos. In the near future, there may be greater numbers of women who use their own cryopreserved eggs to become pregnant after they enter menopause. Although lifestyle choices and medical care enable many older women to remain fit and active, statistically speaking, an older mother is more likely to die sooner than a younger mother. The use of ARTs by postmenopausal women involves the dichotomy between the right of society to exert control over private behaviors, including actions undertaken by both the mothers and treatment providers. An analysis considering the ethical principles of beneficence and nonmalfeasance raises additional questions about the rights of potential offspring. The essential question is whether the woman’s right to reproductive autonomy is outweighed by any potential suffering that her children may experience from potentially being left motherless at a young age. This also generates questions about the potential cost to society resulting from the burden on public funds if the children become dependent on government resources after the death of their primary caretaker. However, a consideration of the balance of personal rights and costs to society must also address issues related to gender equality and justice. Some argue that it is not equal or just to deny postmenopausal women access to ARTs if men of the same age are able to father biological offspring, given the sole difference in their access to parenting results from biological and not social factors. Some countries limit public funding of IVF based on age. However, such limits may not resolve any ethical challenges or avoid the potential burden on public resources created by any resulting and later dependent offspring. These patients may simply turn to private providers or cross-border reproductive care to obtain IVF and then return to their home countries. In the United States, physicians are permitted to decide whether to offer treatment to postmenopausal patients based on personal judgment. Some may decide that a healthy and legally competent patient, regardless of age, has an absolute right to reproductive autonomy and that she should determine the best interest of any of her children. Other treatment providers may use the ethical principles of beneficence and nonmalfeasance to evaluate the risks, burdens, and benefits of all parties including potential offspring.

Gamete Donation Both men and women choose to donate gametes for many reasons including compensation, altruism, or wanting biological offspring without the responsibility of raising them. The commodification of human gametes is considered by some as an inherently immoral result of ARTs. This is a position that has been advocated by influential political, religious, and social institutions that view embryos not as property but human life. While some legislative enactments reflect these arguments, Western jurisprudence has largely embraced the view that gametes are property. The use of donor sperm predates the advent of the modern ART revolution by almost a century, as the use of human donor sperm was reported in the nineteenth century. In the 1980s and the 1990s, innovations in ovulation stimulation and oocyte retrieval made egg donation possible. There is little controversy over whether men should be compensated as sperm donors; paying small amounts of compensation to these donors is a long-standing practice. In sharp contrast, the morality of egg donation especially when donors are paid remains controversial. Both feminists and social conservatives, who usually have diametrically opposite opinions about reproductive autonomy, have argued that allowing women to profit from their bodies through egg donation is exploitative and tantamount to prostitution. A less politically charged explanation for the controversy is that biological differences make oocyte donation riskier. Women who donate oocytes must undergo IVF and are exposed to the risks of ovarian stimulation and oocyte retrieval. It is a fundamental ethical prerequisite that oocyte donors participate voluntarily and without coercion or undue influence. Donors may proceed against their own best interests, given the risks involved, because of the potential for lucrative payments. Although the immediate risks are well established, information about the long-term physical risks and the emotional benefits and risks of donation is still being developed. In the absence of robust data regarding the long-term health consequences of IVF, the debate will continue over whether it is ethical, just, and fair to place young fertile donors at risk for the benefit of older, infertile women.

Donor Anonymity The ethical consideration of gamete donor anonymity requires an examination of the distinct and sometimes competing interests and rights of donors, recipients, and offspring. Recipients are interested in having healthy offspring and control over the resulting childrearing. This includes choosing the gametes that will be used and receiving health information about the donor. They may also want protection from later involvement with the donor or the ability to control the parameters of that contact. Offspring may be interested in basic health issues or have more complex concerns about personal identity as information about genetic history is universally important in the formation of self-identity. Gamete donors are interested in their ability to donate, being protected

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during that process, being treated fairly if interests occur, and not having obligations imposed upon them without their consent. Donors also have an interest in having or not having contact with offspring (Ethics Committee ASRM, 2009b). Arguments for allowing either donors or their offspring to break anonymity include the potential advantages of sharing medical information with their genetic offspring, in the case of the donor, or learning about their genetic history directly, in the case of the offspring. Other arguments include the inherent right or even emotional need to meet and to develop a relationship with a biological relative. The jurisprudence and legislation addressing these issues vary from country to country and parameters governing donor anonymity are continuing to evolve. Some states and countries have laws that mandate donor registries and create formal systems of third-party intermediaries. Moreover, as the use of ARTs becomes more prevalent and more children are born from donor gametes, social conventions regarding the disclosure of donor information and the formation of relationships among recipients, donors, and offspring will also evolve and may resolve many of these debates.

Embryo Donation There are four possible approaches to the disposition of surplus embryos that often result from IVF cycles: thawing and discarding, donating to research, indefinite storage, and donating the embryos for transfer to another woman. All these approaches have attracted staunch supporters and detractors resulting from long-standing debates over the moral and legal status of embryos. The use of embryos for research purposes, particularly as it relates to human stem cells and possible cloning, has been the source of intense debate and has resulted in substantial regulation that varies from nation to nation. Embryo donation is recognized and regulated in some countries as part of the comprehensive system that regulates all aspects of ARTs. In the United States, the Food and Drug Administration oversees the process through regulations that apply to all types of donated human tissue.

Three-Person IVF Researchers are now perfecting mitochondrial transfer (popularly called ‘three-person IVF’) to prevent mitochondrial disease. These techniques inject DNA from the intended parents into a donor egg with normal mitochondria (mitochondrial DNA (mtDNA)) that has had its nucleus removed. The resulting embryo carries the DNA from the intended parents and healthy mtDNA from the donor. Ethical and legal guidelines are currently being developed in the United Kingdom and debated in other countries. Ethical concerns include the unknown effects of germline DNA alteration on subsequent generations and whether the technique will lead to other forms of genetic modification (Parliamentary Office of Science and Technology, 2013).

Surrogacy Although sometimes used as interchangeable terms, there is an important distinction between traditional surrogacy and gestational surrogacy. Surrogacy is a broad term used to describe situations when a woman agrees to carry a pregnancy for someone else with the intent of giving custody over to the intended parents. In a traditional surrogacy arrangement, the surrogate is the child’s genetic mother as she agreed to conceive through AI and deliver a child for the intended parents. In gestational surrogacy, an embryo is transferred to a woman who has agreed to carry the pregnancy and deliver the child. The embryo may have been created from sperm and/or oocytes from the intended parents or donor gametes. Surrogacy arrangements are characterized as altruistic surrogacy or commercial surrogacy. Altruistic surrogacy arrangements often occur between relatives or friends such as when a mother or a sister carries a pregnancy. Commercial surrogacy is often arranged by a fertility clinic or another private resource and typically involves the formation of a legal contract between the women and intended parents. Both surrogates and gestational carriers are subject to medical and emotional risks from carrying a pregnancy and undergoing a delivery. Commercial and traditional surrogacies have generated the most heated controversy. Some opponents focus on psychosocial concerns and take issue with the entire concept of surrogacy as contrary to the definition of motherhood, which they believe cannot be altered by a legal contract. They argue that a legal contract cannot undermine the biological and psychological attachments formed during pregnancy. The woman carrying the fetus is the mother regardless of whether donor gametes or embryos were used. There are some feminist scholars who are proponents of surrogacy believing that it is the ultimate exercise of reproductive autonomy, but only if the woman retains the right to terminate the pregnancy and revoke any agreement regarding the pregnancy. These arguments reject the idea that women should be compelled to carry out specific performance elements of any surrogacy contract such as allowing the intended parents to decide whether a pregnancy should continue or not or whether the woman should undergo multifetal reduction. This view also rejects prohibitions on surrogacy as sexist and paternalistic. Others argue that commercial surrogacy results in the commodification of the body and the commercialization of parenthood. Surrogacy is compared to prostitution without the stigma of prostitution because there is no sexual act. Under this rubric, it is argued that men will inevitably control the surrogacy process and use it solely for their own economic interests. A corollary argument focuses on the potential harms to children viewing commercial surrogacy as a form of human trafficking, disrespectful of their

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inherent moral value as individual human beings. Another less extreme argument addresses issues of children’s best interests. These include the harm that may come to a child subject to a custody dispute or continued uncertainty about identity and origin because of the potential dividing of the genetic, biological, and social roles of mother among three different women. Regulation of surrogacy varies widely among different countries and even between states or within regions of individual countries. Some laws facilitate these practices. Other jurisdictions refuse to enforce them, and if a contractual or custody dispute arises, they will recognize the surrogate as a legal parent. The compensation paid to surrogates also varies throughout the world; some surrogates are only allowed to receive payments for their medical care and basic living expenses, while other surrogates are also compensated for carrying the pregnancy. The differences in regulations and costs have resulted in the development of international surrogacy. Such arrangements typically involve intended parents from the developed countries and gestational carriers in developing countries. This form of cross-border reproductive care has intensified already difficult questions about exploitation, commodification, stigmatization, and coercion and informed consent. These questions are particularly challenging because of significant differences in wealth and bargaining power between surrogates and intended parents and the ostracism that surrogates may experience in socially conservative cultures that consider reproduction as only acceptable within marriage.

Intrafamilial Collaborative Reproduction Intrafamilial collaborative reproduction is when a donor or surrogate is related. Some countries categorically prohibit collaborative reproduction or impose the ‘anonymity rule’ in the context of collaborative reproduction, meaning that any collaborator should be unknown to the prospective parents. Therefore, these arrangements are illegal (ASRM, 2012). In countries that permit such arrangements, the ethical questions that must be addressed include whether there is undue pressure on family members to participate as a result of emotional or financial dependency. Another concern is the potential that the child may have identity problems because of growing up in an unconventional familial environment or when there is role confusion or conflict between genetic and social parents. Other issues include questions about apparent incest or consanguinity. In the United States, professional guidelines advise that gamete arrangements in which the child would have the same genetic relationship to the participants as would children of incestuous or consanguineous unions between first-degree relatives (including adopted and stepchildren) are considered unethical (ASRM, 2012).

Fertility Preservation Exposure to gonadotoxic agents during chemotherapy or other medical treatment can jeopardize the prospects of genetic parenthood. Patients undergoing these treatments may be offered the chance to preserve their gametes with cryopreservation. This is called fertility preservation (FP). FP for patients with life-threatening illnesses raises the question whether it is ethical to delay treatment to obtain reproductive tissues safely. Patients must also confront questions about the disposition of stored gametes, embryos, or gonadal tissue in the event of death. Ethical questions also surround the use of FP for pediatric patients because they cannot legally consent to treatment and would normally not face reproductive decisions until they reach adulthood. Current ethical guidelines recommend that treatment providers discuss these issues and involve children in deciding on FP. Because only a parent or guardian can consent, they must determine what is in the child’s best interests related to current treatment and their future as adults. Guidelines also suggest that parents should consider their child’s opinion, the details of the procedures involved, and whether such procedures are proven or experimental (ASRM, 2005). With increasing numbers of women delaying parenthood until their late 30s and 40s, healthy women wishing to guard against age-related infertility and having the option of delaying childbearing to pursue education and career goals or until they find the right partner with whom they want to share their life are undergoing FP. This has been termed social egg freezing. The contraceptive technology revolution gave women unprecedented ability to prevent pregnancies and resulted in many changes in women’s social roles. The use of FP may similarly challenge existing gender conventions, social structures, and even economic systems. However, FP is currently only available to women able to pay for IVF and long-term cryopreservation. This will likely result in future ethical and social issues related to the inequalities in access to this ART.

Posthumous Collection and Use of Reproductive Tissues Posthumous reproduction typically involves the use of sperm after a man’s death to initiate a pregnancy in his surviving female partner. Currently, techniques such as stimulated ejaculation, microepididymal sperm aspiration, or testicular sperm extraction can be used to procure sperm from a dead or brain-dead individual. Physicians have also been asked to harvest oocytes and ovarian tissue from comatose and brain-dead women for later transfer to a gestational surrogate. These practices raise questions about whether it is ethical to initiate a pregnancy using reproductive materials from a person who may not have consented to its harvesting or use.

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Posthumous reproductive treatment generates a conflict between the rights of reproductive autonomy of the deceased person and surviving partner. In the United States, professional guidelines advise that posthumous gamete (sperm or oocyte) procurement and reproduction are ethically justifiable if written documentation from the deceased authorizing the procedure is available or when such requests are initiated by the surviving spouse or life partner (ASRM, 2009b). Others reject the right of the spouse to initiate posthumous reproduction without written documentation stating the deceased person’s wishes even if the posthumous reproduction will use gametes that the deceased had preserved before death. While the existence of cryopreserved gametes or embryos demonstrates intent to reproduce, it does clearly demonstrate that the deceased may have wished posthumous reproduction to occur. Posthumous reproduction also raises ethical concerns about potential offspring. These questions include whether the circumstances surrounding their birth will have any long-term negative consequences for the resulting children’s psychosocial development. The number of children who have been conceived in this manner is few in number and many are still in early childhood, limiting the ability to research this issue; it may be some time before their experiences are fully appreciated. More practical concerns surround questions about the resulting children’s legal status with regard to lineage, citizenship, rights of inheritance, and entitlements to government benefit programs. In 2012, the United States Supreme Court addressed some of these issues in Astrue v. Capato. The Court determined that twins conceived and born after their genetic father’s death were not entitled to Social Security Survivor benefits based upon an interpretation of a Florida law that did not recognize the children as legal heirs as they were conceived after his death. This decision is applicable to children conceived from oocytes used from deceased women. Some states have made changes to their inheritance laws, but there is no uniformity with regard to the rights of posthumously conceived children.

Risks of ARTs Multifetal Pregnancies A major risk of IVF is multifetal pregnancies. They often lead to premature delivery and its accompanying problems, cerebral palsy and other disabilities, lung and gastrointestinal problems, and even neonatal death. The mother is at risk for hypertension, preeclampsia, gestational diabetes, and death. To reduce the incidence of multiple gestations, some countries have passed laws about the number of embryos that can be transferred during IVF cycles. This protocol is called single embryo transfer (SET). In the United States, there is no legal requirement to use SET and professional guidelines emphasize the right to reproductive autonomy. Therefore, it is for patients to decide whether to use SET based on consultations with their doctors. Some women choose to transfer multiple embryos believing that becoming pregnant outweighs the risks or they choose to undergo MPR. Others reject this procedure because of opposition to abortion. Many argue that either one of these decisions is a reasonable exercise of reproductive choice. However, even among physicians who perform MPR, controversy surrounds the reduction of twins. Some believe that twin reduction is unethical because it involves selecting one fetus over another when either one is equally wanted. In addition, twin reductions are viewed as crossing the ethical line between doing the procedure for medical indications versus social indications such as to avoid the financial and social stresses of caring for two or more same-aged children.

Birth Defects There is a general consensus that IVF creates a small but statistically significant increased risk for congenital abnormalities including rare epigenetic and anatomical abnormalities (Brezina and Zhao, 2012; Hansen et al., 2013). There is also evidence that IVF is associated with a small but statistically significant risk of mental retardation and IVF with ICSI performed because of paternal infertility is associated with a small increase in the relative risk for autism (Sandin et al., 2013). Questions remain as to whether any possible increased risk is related to the parents’ underlying infertility or is caused by the actual process of IVF (Davies et al., 2012). The examination of the lasting effects that ARTs may have on early development will continue as the population of people born through IVF increases and ages. The uncertainties that continue to surround the use of ARTs will likely not dissuade patients from utilizing them in their efforts to become parents. Therefore, physicians and patients will continue to struggle with balancing patient autonomy and other ethical principles such as beneficence and nonmaleficence when deciding to use ARTs. Each will have to consider whether the chance to parent outweighs the risk of having a child who may have disability, at least in part, because of the means used to achieve the pregnancy. The risks associated with ARTs may also generate debate about the public health implications including whether the benefits of widespread access to ARTs outweigh the potential costs associated with children who may be born with disabilities as a result of the use of ARTs.

Preimplantation Genetic Diagnosis PGD is a fusion of genomics and ARTs. It allows would-be parents to screen and select the genetic characteristics of their potential offspring, to a limited but growing degree. Objections to PGD echo debates over the moral status of embryos. People who believe that the embryo or fetus is a person also object to creating and destroying embryos including those involved in PGD. Others believe

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embryos are too rudimentary in development to have interests or rights but they are entitled to special respect as the first stage toward human development. Under this view, PGD is ethically acceptable when done for such reasons as preventing offspring with serious genetic disease or for obviating the need for selective abortion to prevent those diseases. Other ethical concerns include the use of PGD to screen for adult-onset conditions such as breast cancer that may be life limiting but might not develop until middle age. Another concern is the use of PGD for sex selection of embryos unrelated to diagnosis of Xlinked genetic disorders but for gender preference or family balancing. Several countries prohibit gender selection for reasons other than the detection of disease and professional guidelines in the United States discourage the use of PGD solely for nondiagnostic sex selection. While some ethicists argue that such bans are unwarranted intrusions on reproductive autonomy, others argue that the principles of equal justice and the need to promote gender equality and ensure a gender balanced population ratio mandate such restrictions. Another ethical concern is the use of PGD for the deliberate selection of embryos with disabling conditions (e.g., genetic forms of deafness or dwarfism) that the parents would share with their offspring and they believe forms a part of their family or cultural identity. Some ethicists argue that despite the principle of reproductive autonomy, children possess a unique class of rights called rights in trust – rights that they cannot yet exercise, but which they will be able to exercise when they reach maturity. Based on this principle, it is unethical for parents to take deliberate steps to create children with disabilities because doing so denies the child’s right to an open future, which is the right to have available as an adult the widest variety of life choices possible. Another controversial use of PGD is for the creation of the so-called ‘savior siblings.’ In these cases, the aim is to create embryos without the same disorder that are human leukocyte antigen matches for an existing child that requires stem cell or bone marrow transplant. Ethical issues include whether it is ethical to create a child not just for himself or herself but with the responsibility to save a sibling. There is the related concern that the parents may prioritize the interests of the first child over the second and allow that child to suffer physically and/or emotionally. Some argue that PGD for ‘savior siblings’ creates a slide toward the acceptance of screening embryos for nonmedical traits with the aim of selecting better children. Some believe that emerging technologies such as the recent introduction of next-generation DNA sequencing for PGD could result in screening using whole genome sequencing of all IVF embryos prior to transfer. This raises issues about what are the consequences of ARTs that permit parents not only to choose to have children but also to choose the kind of children they will have. These developments will further challenge the principles of reproductive autonomy and the right of the child to an open future, and may increase the responsibility of clinicians toward the welfare of the future child.

References American Society for Reproductive Medicine. (2005). Fertility preservation and reproduction in cancer patients. Fertil. Steril., 83, 1622–1628. American Society for Reproductive Medicine. (2009a). Access to fertility treatment by gays, lesbians, and unmarried persons. Fertil. Steril., 92, 1190–1193. American Society for Reproductive Medicine. (2009b). Interests, obligations, and rights of the donor in gamete donation. Fertil. Steril., 91, 22–27. American Society for Reproductive Medicine. (2012). Using family members as gamete donors or surrogates. Fertil. Steril., 98, 797–803. Astrue v. Capato, 566 U.S. _______ 2012. Brezina, P. R., & Zhao, Y. (2012). The ethical, legal, and social issues impacted by modern assisted reproductive technologies. Obstet. Gynecol. Int., 2012, 1–7. Davies, M. J., Moore, V. M., Wilson, K. J., et al. (2012). Reproductive technologies and the risk of birth defects. N. Engl. J. Med., 366, 1803–1813. Hansen, M., Kurinczuk, J. J., Milne, E., de Klerk, N., & Bower, C. (2013). Assisted reproductive technology and birth defects: a systematic review and meta-analysis. Hum. Reprod. Update, 19, 330–353. Hoorens, S., Gallo, F., Cave, J. A. K., & Grant, J. C. (2007). Can assisted reproductive technologies help to offset population ageing? an assessment of the demographic and economic impact of ART in Denmark and UK. Hum. Reprod., 22, 2471–2475. Inhorn, M. C. (2006). Making Muslim babies: IVF and gamete donation in Sunni versus Shi’a Islam. Cult. Med. Psychiatry, 30, 427–450. Nikolaos, M. (2008). The Greek orthodox position on the ethics of assisted reproduction. Reprod. Biomed. Online, 17, 25–33. Omurtag, K., & Adamson, G. D. (2013). The Affordable Care Act’s impact on fertility care. Fertil. Steril., 99, 652–655. Parliamentary Office of Science and Technology. (2013). Preventing mitochondrial disease. PostNote, 431, 1–4. Sandin, S., Nygren, K. G., Iliadou, A., Hultman, C. M., & Reichenberg, A. (2013). Autism and mental retardation among offspring born after in vitro fertilization. JAMA, 310, 75–84. Schenker, J. G. (2008). The beginning of human life: status of embryo. perspectives in Halakha (Jewish religious law). J. Assist. Reprod. Genet., 25, 271–276. Vatican. (1987). Sacred congregation for the doctrine of the faith. In Lawful and Illicit Uses of New Techniques in Human Embryology.

Relevant Websites www.asrm.org – American Society for Reproductive Medicine. www.cdc.gov/art – Centers for Disease Control and Prevention. www.eshre.eu – European Society of Human Reproduction and Embryology. www.resolve.org – Resolve: The National Infertility Association. www.sart.org – Society of Assisted Reproductive Technology.

Tooth Regenerative Therapy: Tooth Tissue Repair and Whole Tooth Replacement M Oshima, K Ishida, R Morita, and M Saito, Tokyo University of Science, Noda, Chiba, Japan T Tsuji, Tokyo University of Science, Noda, Chiba, Japan; and Organ Technologies Inc., Tokyo, Japan © 2019 Elsevier Inc. All rights reserved.

Introduction to and Concepts for Tooth Regenerative Therapy Tooth Development Tissue Repair Using Tooth Tissue-Derived Stem Cells and Cytokines The Use of Apical Papilla Stem Cells for Dental Pulp Regeneration Periodontal Tissue-Derived and Dental Follicle Stem Cells and Their Application to Periodontal Tissue Regeneration Bioengineered Root Regeneration Using Dental Stem Cell Populations and Tissue Engineering Technology Whole Tooth Regeneration as a Future Organ Replacement Regenerative Therapy Generation of a Bioengineered Tooth Germ Using a Biodegradable Scaffold Bioengineering of Tooth Germ Using Cell Aggregation Methods Novel Three-Dimensional Cell Manipulation Methods for Bioengineered Tooth Germ: The ‘Organ Germ Method’ Functional Tooth Replacement Therapy Transplantation of Bioengineered Tooth Germ or Bioengineered Mature Tooth Unit as a Tooth Replacement Strategy Bioengineered Tooth Response to Mechanical Stress Perceptive Neuronal Potential of Bioengineered Teeth Future Perspectives for the Tooth Regeneration Concluding Remarks References

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Glossary Bioengineered tooth germ Regenerated tooth germ reconstituted by the organ germ method between tooth-germ-derived epithelial and mesenchymal cells can develop correctly, both orthotopically and ectopically. Bioengineered tooth unit They are composed of bioengineered mature tooth, periodontal ligament, and alveolar bone and are generated by ectopic transplantation. This unit can be transplanted into a properly sized hole in the alveolar bone through bone integration. Epithelial–mesenchymal interaction Tooth development is regulated by sequential and reciprocal interactions between epithelial and mesenchymal tissues. These interactions are mediated by multigene signal families, such as bone morphogenetic protein (BMP), fibroblast growth factor (FGF), sonic hedgehog (Shh), and Wnt. Organ germ method The organ germ method was developed as a technology to generate bioengineered organ germ that properly reproduces the three-dimensional cell processes that occur during embryonic development to yield bioengineered organ germ in vitro (Nakao et al., 2007). Tooth germ All organs develop from organ germs that are generated during the embryonic period. The organ germ of tooth is referred to as tooth germ, which is composed of two tissues: the oral epithelium and neural-crest-derived mesenchyme.

Introduction to and Concepts for Tooth Regenerative Therapy Oral functions such as enunciation, mastication, and occlusion are important components of a healthy life. The tooth is an ectodermal organ regulated by reciprocal epithelial–mesenchymal interactions (Thesleff, 2003) and has distinctive hard tissues that include enamel, dentin, cementum, and alveolar bone. Teeth also have soft connective tissue such as pulp and periodontal ligament that contain nerve fibers and blood vessels to maintain tooth homeostasis (Avery, 2002). Damage, loss, and disease in teeth, including dental caries and periodontal disease, cause fundamental problems for oral function and are associated with a number of health issues (Proffit et al., 2004). Conventionally, the restoration of tooth functions in these circumstances involves replacement with dentures or dental implants. Although these artificial therapies are very effective, there have been recent improvements that enhance the biological functions underlying tooth movement through bone remodeling (Huang et al., 2009). The current advances in future regenerative therapies have been influenced by many previous studies of embryonic development, stem cell biology, and tissue engineering technologies (Brockes and Kumar, 2005). To restore the partial loss of organ functions and to repair damaged tissues, attractive regenerative therapy concepts include stem cell transplantation into various tissues and organs

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Figure 1 Concepts for tooth regenerative therapy. Approaches to developing technologies for tooth regenerative therapy have included tissue repair, tissue engineering, and whole tooth organ replacement.

and cytokine therapy to activate tissue stem/progenitor cells. Tooth tissue stem cells and the cytokine network that regulates tooth development have been well characterized at the molecular level (Thesleff, 2003). These advances can be applied to the repair of dental pulp and periodontal tissues, including the alveolar bone (Figure 1). The ultimate goal of regenerative therapy is to develop fully functional bioengineered tissues that can replace lost or damaged organs following disease, injury, or aging (Ikeda and Tsuji, 2008). In the dental field, this therapy would involve replacement of a lost or damaged tooth with a bioengineered tooth built with stem cells that has the capacity to become a functional unit, including the whole tooth and periodontal tissue surrounding alveolar bone (Yen and Sharpe, 2006). It is expected that regenerative tooth replacement therapy will be established in the near future as a novel and successful biological treatment that will provide essential functional recovery of lost teeth to satisfy esthetic and physiological requirements (Sharpe and Young, 2005). Many approaches to functionally replace missing teeth have been evaluated in the past three decades, including three-dimensionally bioengineered teeth and tooth germ generation using biodegradable materials and cell aggregation methods (Duailibi et al., 2006; Ikeda and Tsuji, 2008; Yen and Sharpe, 2006). The first reports of fully functioning bioengineered tooth replacement with correct tissue orientation, masticatory potential, responsiveness to mechanical stress, and perceptive potential following transplantation into a lost tooth region were recently published (Nakao et al., 2007; Ikeda et al., 2009). In this article, the most recent findings and technologies for tooth tissue repair and whole tooth replacement, also known as tooth regenerative therapy, which have the potential to provide essential functional recovery and ultimately replace currently used artificial materials, are discussed.

Tooth Development Organs such as hair, glands, kidneys, and teeth arise from their respective germs, which are induced by reciprocal interactions between epithelial and mesenchymal tissues in the developing embryo (Thesleff, 2003). The principle mechanisms of tooth organogenesis are also regulated by reciprocal epithelial and mesenchymal interactions, particularly those involved in stem cell, signaling molecule, and transcription factor pathways (Figure 2). In tooth germ development, the dental lamina first thickens (lamina stage), followed by epithelial thickening (placode stage) at the sites of future teeth and subsequent epithelial budding to the underlying neural crest-derived ectomesenchyme. Tooth germ formation is initiated on embryonic days (ED) 10– 11 in mice by epithelial signals such as fibroblast growth factor 8 (FGF8), bone morphogenetic protein 4 (BMP4), sonic hedgehog (Shh), tumor necrosis factor, and Wnt10b. These signals induce the expression of transcription factors such as Barx1, Dlx1/2, Lhx6, Lhx7 Msx1, Pax9, and Gli in the dental mesenchyme, which then condense around the developing epithelial bud (bud stage) (Thesleff, 2003). At ED14 (cap stage), a transient epithelial signaling center known as an enamel knot, which expresses several signaling molecules, including Shh, BMP family molecules, FGFs, and Wnts, is thought to regulate individual cell fates and epithelial–mesenchymal interactions. The terminal differentiation of dental epithelial and dental mesenchyme cells into ameloblasts and odontoblasts, respectively, occurs during ED16–18 (bell stage) to perinatal stage. These differentiated cells secrete enamel proteins or a collagenous extracellular matrix that mineralizes into the enamel and dentin matrix at the epithelium–mesenchyme interface. A portion of the dental mesenchymal cells forms the apical papilla, which is capable of differentiating odontoblasts and pulp cells following their migration to the root apex side (Sonoyama et al., 2006). The outer mesenchymal cells outside of the tooth germ, which comprises an epithelium and condensed mesenchymal cells, forms the dental follicle that can generate periodontal tissue, including cementum, periodontal ligament, and alveolar bone. After tooth development, immature cells seem to be maintained as adult tissue stem cells that are thought to act as a self-repair system for dental tissues and supply a wide variety of dental cell types, both under steady-state conditions and after dental tissue injury. Tooth morphogenesis begins with cusp formation in the early bell stage to form the tooth crown. Tooth size and shape are thought to be regulated by signaling molecules such as the BMPs and FGF4 emanating from the secondary enamel knots, which regulate the cusp pattern of the mature natural tooth (Thesleff, 2003). Tooth root formation is initiated after tooth crown formation and is followed by tooth eruption in the oral cavity.

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Figure 2 Schematic of tooth germ development. Tooth germ development begins at the lamina stage and proceeds to the bud stage, during which the dental epithelial bud and neural-crest-derived ectomesenchyme are formed. Subsequent morphogenesis occurs at the cap stage during the development of dental epithelium and dental mesenchyme, which can later diverge into the papilla and the follicle. Tooth crown is formed from the early bell stage to the late bell stage. During tooth eruption, the root is developed and dental follicle cells differentiate into periodontal tissue to attach the tooth root and jawbone (adult tooth).

Tissue Repair Using Tooth Tissue-Derived Stem Cells and Cytokines Recent studies of stem/progenitor cells and organogenesis have provided considerable new insights that have furthered our understanding of tooth tissue-derived stem cells, which can differentiate into various dental cell lineages such as odontoblasts, pulp cells, periodontal ligament cells, cementoblasts, and osteoblasts (Huang et al., 2009). Dental stem cells will be useful for developing stem cell transplantation therapy, one of the more promising concepts in regenerative therapy, to restore the partial loss of organ function by replacing enriched and purified stem cells and thereby achieving dental tissue repair (Figure 3(a)).

The Use of Apical Papilla Stem Cells for Dental Pulp Regeneration Dental pulp is composed of connective tissue, blood vessels, nerves, fibroblasts, and odontoblasts and develops from the dental papillae after being encased by dentin tissue (Huang et al., 2009). Dental pulp stem cells (DPSCs) and stem cells from human exfoliated deciduous teeth (SHED) have been isolated from the dental pulp of human permanent third molars and exfoliated deciduous teeth, respectively. DPSCs and SHED express CD146/STRO-1 and are thought to be tooth tissue-derived stem cells with high proliferative capacities and sufficient potency to develop into odontoblasts, adipocytes, and neurallike cells. Therefore, these stem cells may be a good resource for stem cell-mediated tissue repair, including dentin or pulp regeneration. As the origin of root and pulp development, the dental papilla is located apically to the developing pulp and is thus known as the apical papilla, which is less vascular and contains cellular, gelatinous soft tissue. Apical papilla contain unique stem cells, known as stem cells from apical papilla (SCAP), which have a high proliferative potential that is reflected by high levels of telomerase activity and multipotent differentiation capability that produces odontoblasts and adipocytes (Sonoyama et al., 2006). SCAP cells can also generate typical dentin structures after transplantation in vivo and may offer a promising avenue for cell-based therapies for tissue repair and tissue engineering. Dental decay is the most commonly observed tooth pathology, and the current standard treatment involves the substitution of the natural/physiological dental tissue with artificial material. A molecular medical approach using growth factors involved in tooth development and based on epithelial–mesenchymal interactions is anticipated to enable tooth tissue repair, such as dentin regeneration via the addition of BMPs and transforming growth factor (TGF)-b1 (Huang et al., 2009). Tooth tissue-derived stem cells, such as DPSCs and pulp stem cell subfractions, CD31/CD146 side population (SP) cells and CD105þ cells, which can generate pulp tissue, may also be useful for tooth tissue repair and dental pulp regeneration. It is feasible that growth factors and tooth tissuederived stem cells will one day be used to repair damaged dental pulp tissue.

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Figure 3 Dental tissue repair and engineering. (a) Dental tissue treatment against caries, pulp injury, and periodontal disease. Tooth regenerative therapy, stem cell transplantation, and cytokine therapies are regarded as attractive approaches for repairing tissue damaged by dental caries and periodontal disease. For dental caries and pulp injury, the transplantation of dental stem cells including DPSCs, SHED, and SCAP, which can differentiate into odontogenic progenitors and pulp cells, has been examined (right panel). Cytokines that have the potential to activate and differentiate dental stem cells such as BMPs, TGF-b, NGF, VEGF, GDNF, and BDNF are also expected to mediate tissue repair. For periodontal tissue repair, the application of PDLSCs and DFSCs for stem cell transplantation and cytokines including PDGF, IGF, BDNF, bFGF, EMP, and ADAMTSL6b have the potential to regenerate periodontal tissue (left panel). (b) Bioengineered tooth root regeneration by tissue engineering. The bioengineered root structure is generated using a root-shaped HA/TCP carrier loaded with SCAP that is covered with gelfoam containing PDLSCs and is inserted onto a porcelain crown (left panel). The HA/SCAP-gelfoam/PDLSC implant successfully regenerates dentin and periodontal ligament tissue on the outside of the implant after transplantation into the alveolar bone, resulting in normal tooth function (right panel). PDLSCs, periodontal ligament stem cells; DFSCs, dental follicle stem cells; PDGF, platelet-derived growth factor; IGFs, insulinlike growth factors; BDNF, brain-derived neurotrophic factor; bFGF, basic fibroblast growth factor; EMP, enamel matrix protein; ADAMTSL6b, a disintegrin-like metalloprotease domain with thrombospondin type I motifs like 6b; DPSCs, dental pulp stem cells; SHED, stem cells from human exfoliated deciduous teeth; SCAP, stem cells from apical papilla; BMPs, bone morphogenetic proteins; TGF-b, transforming growth factor-b; NGF, nerve growth factor; VEGF, vascular endothelial growth factor; GDNF, glial cell line-derived neurotrophic factor.

Periodontal Tissue-Derived and Dental Follicle Stem Cells and Their Application to Periodontal Tissue Regeneration Periodontal tissue is a tooth-supporting connective tissue that acts as a shock absorber for occlusal force. Periodontal tissue structure, including the periodontal ligaments, cementum, and alveolar bone, can be irreversibly damaged by periodontitis, a chronic inflammatory disease. No reliable treatment for regenerating periodontium has been established. Periodontal tissues, including cementoblasts, periodontal ligaments, and osteoblasts, are derived from dental follicle stem cells (DFSCs) (Yen and Sharpe, 2006), which migrate onto tooth root surfaces to form periodontal tissue during the tooth root-forming stages. Periodontal ligament-derived mesenchymal stem cells (PDLSCs), which have also been identified in adult human periodontal ligaments and can be cultured as stem cells in vitro, can differentiate into all periodontal cell types after transplantation in vivo. DFSCs have the ability to reproduce periodontal formations and are thought to be stored as PDLSCs in adult periodontal ligaments after tooth development. Using cell sheet engineering in conjunction with the stem cell transplantation therapies described above, stem cell sheets are now being developed for clinical use in periodontal tissue regeneration. Recently developed molecular treatments can be administered via the local application of human recombinant cytokines such as platelet-derived growth factor, insulinlike growth

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factor-I, brain-derived neurotropic factor, and basic FGF to accelerate periodontal regeneration. Furthermore, periodontal tissue is composed of fibrillar extracellular matrices such as collagen fibrils and microfibrils, which play a critical role in periodontal tissue formation and in the onset of severe periodontitis. Recently, it was reported that the local administration of fibrillin-1-associated protein ADAMTSL6b effectively accelerates wound healing in periodontal tissues by restoring microfibrils (Saito et al., 2011). This molecular medical therapy involving microfibril reinforcement is a novel approach to periodontal tissue regeneration, in addition to stem cell transplantation and cytokine treatment.

Bioengineered Root Regeneration Using Dental Stem Cell Populations and Tissue Engineering Technology Dental implants, which require direct integration into the alveolar bone, have recently gained momentum as a valid therapy for replacing missing teeth and an alternative to fixed or removable dentures. Although dental implants have been the mainstay of dental therapy over the past few decades, they are associated with an absence of periodontal ligaments, and are therefore deficient in essential tooth function and natural structural relationship with the tooth root and alveolar bone. To regenerate the tooth root and its associated periodontal tissues, both of which are necessary for maintaining physiological tooth function, a tissue engineering application has been anticipated for stem cell-based root regeneration (Figure 3(b)). A unique approach for tooth root regeneration employing a root-shaped hydroxyapatite/tricalcium phosphate carrier loaded with gelfoam/PDLSC-covered SCAP cells has been reported to form a rootlike structure to which a porcelain crown can be attached, resulting in normal tooth function (Sonoyama et al., 2006). This tissue engineering technology produced a bioengineered root by using a combination of root and periodontal tissues and may underlie the next generation of regenerative medical technologies that integrate stem cell-mediated tissue regeneration strategies, engineered materials as structural components, and current dental crown technologies. The root regeneration concept also has the potential for earlier clinical application compared to whole tooth regeneration.

Whole Tooth Regeneration as a Future Organ Replacement Regenerative Therapy The current approach to generating ectodermal organs such as teeth, hair follicles, and salivary glands is to recreate organogenesis through epithelial–mesenchymal interactions that occur in the developing embryo and thereby develop fully functioning bioengineered organs from bioengineered organ germ generated from immature stem cells via three-dimensional cell manipulation in vitro (Ikeda and Tsuji, 2008; Sharpe and Young, 2005). For tooth regeneration, a concept has been proposed in which a bioengineered tooth germ will be transplanted into recipient jaw and develop into a functional mature tooth (Figure 4, upper panel). It is also expected that it will be possible to transplant a bioengineered tooth unit that includes mature tooth, periodontal ligament, and alveolar bone, which will achieve engraftment through the physiological bone integration of the recipient’s jaw (Figure 4, lower panel; Oshima et al., 2011). To realize whole tooth replacement, the first major issue is to develop a three-dimensional cell

Figure 4 Strategies for whole tooth replacement via regenerative therapies. Functional teeth can now be regenerated by transplanting bioengineered tooth germ reconstituted from epithelial and mesenchymal cells via the organ germ method or by transplanting bioengineered tooth units with periodontal ligament and alveolar bone developed from bioengineered tooth germ.

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manipulation technology using completely dissociated epithelial and mesenchymal cells in vitro. Two conventional approaches and a novel cell manipulation method are currently being investigated for the purpose of generating bioengineered tooth germ or mature tooth and are described below:

Generation of a Bioengineered Tooth Germ Using a Biodegradable Scaffold The scaffold technology has proven useful in three-dimensional tissue engineering methods to regenerate suitably shaped tissues by seeding single cells onto degradable materials fabricated from either natural ingredients or synthetic polymers. This technique can be used to grow tissues with a uniform cellular distribution and has been applied in clinical bone and cartilage regenerative therapies. Several previous studies that employed polyglycolic acid and poly-L-lactate-co-glycolide copolymer or a collagen sponge as scaffolding materials reported that a tooth-shaped scaffold onto which epithelial and mesenchymal cells isolated from porcine unerupted third molars or a rat tooth bud were seeded could generate small tooth structures containing all of the correct tooth tissue components, including pulp, dentin, and enamel, but not periodontal tissue (Sharpe and Young, 2005; Yen and Sharpe, 2006). Although the scaffold method could be useful for producing the desired tooth shape and size, it has critical limitations, including tooth formation frequency, appropriate replication of tooth tissue structures such as enamel–dentin complex formation, and proper arrangements of ameloblast and odontoblast cell lineages, which are achieved through epithelial–mesenchymal interactions that mimic natural tooth development (Avery, 2002; Thesleff, 2003).

Bioengineering of Tooth Germ Using Cell Aggregation Methods Cell aggregation is a typical bioengineering method that aims to reconstitute bioengineered organ germ that will properly reproduce epithelial–mesenchymal interactions and organogenesis (Sharpe and Young, 2005; Yen and Sharpe, 2006). Previously, transplantations of bioengineered cell aggregates using hair follicles and mammary-gland-derived stem cells have shown promise for regenerating correct organ structures with proper arrangements of different cell types in vivo. Artificial tooth germ created from precipitates separated into dental epithelial and mesenchymal cells with cellular centrifugation has also been shown to be capable of generating the appropriate tooth formation. Mixed cell aggregates of dissociated epithelial and mesenchymal cells isolated from molar tooth germ also have a demonstrated capacity to develop into a tooth with the correct structure, following epithelial cell sorting and subsequent epithelial and mesenchymal cell self-reorganization. Although this technique replicated organogenesis, the frequency of tooth development and correct tissue formation was insufficient.

Novel Three-Dimensional Cell Manipulation Methods for Bioengineered Tooth Germ: The ‘Organ Germ Method’ To achieve large-scale replication of organogenesis, an in vitro three-dimensional novel cell manipulation method designated as a bioengineered organ germ method has been developed. This method is accomplished with cell compartmentalization between epithelial and mesenchymal cells at a high cell density in type I collagen gel (Figure 5(a)). Bioengineered tooth germ created by this technique could allow for large-scale organ development and mimics multicellular assembly and epithelial–mesenchymal interactions, as well as natural tooth development. The bioengineered tooth germ generates a structurally correct tooth after transplantation in an organ culture in vitro but also following placement into a subrenal capsule in vivo. The isolated single bioengineered tooth germ also developed in the oral cavity and formed the correct tooth structure (Nakao et al., 2007). Furthermore, there is a unique technology that can successfully generate a size-controlled bioengineered tooth unit comprising mature tooth, periodontal ligament, and alveolar bone that can also be generated by transplantation into the subrenal capsule (Figure 5(b); Oshima et al., 2011). These technologies have the potential to be adapted for successful functional tooth replacement in vivo and are expected to represent an innovative advance in bioengineered organ replacement regenerative therapies.

Functional Tooth Replacement Therapy Organ replacement regenerative therapy holds great promise for the future replacement of dysfunctional organs with a bioengineered reconstruction created using three-dimensional cell manipulation in vitro. Regenerated organs should be able to complete organ-intrinsic functions in cooperation with the surrounding environment in the recipient. In previous reports, however, artificial organs constructed with various cells and artificial materials could not restore functionality and thus were not viable options for long-term organ replacement in vivo. However, bioengineered organs can be grown in vivo in amphibian models; activin-treated cell aggregates could form a secondary heart with pumping function and also regenerate eyes that were light responsive and connected with the host nervous system. Oral functions such as mastication, pronunciation, and facial esthetics have an important influence on the quality of life because they facilitate both oral communication and nutritional intake (Proffit et al., 2004). These functions are achieved with the teeth, masticatory muscles, and the temporomandibular joint, under control of the central nervous system. For the realization of tooth replacement regenerative therapy, a tooth developed from bioengineered germ or a transplanted bioengineered mature tooth unit must be capable of properly engrafting in the lost tooth region in an adult oral environment and acquiring full functionality, including sufficient masticatory performance, biochemical cooperation with periodontal tissues, and afferent responsiveness to noxious stimulations via neurons in the maxillofacial region (Proffit et al., 2004).

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Figure 5 The organ germ method: a three-dimensional cell processing system. (a) A high density of dissociated mesenchymal cells are injected into the center of a collagen drop (left panel). Dissociated tooth germ-derived epithelial cells are subsequently injected into the drop adjacent to the mesenchymal cell aggregate (right panel). (b) Transplanting bioengineered tooth germ into a subrenal capsule for 30 days can yield a bioengineered tooth unit composed of a mature tooth, periodontal ligament, and alveolar bone with the correct structural components, such as enamel (c), dentin (d), periodontal ligament (PDL), and alveolar bone (AB).

Transplantation of Bioengineered Tooth Germ or Bioengineered Mature Tooth Unit as a Tooth Replacement Strategy A successful tooth regenerative therapy via the transplantation of bioengineered tooth germ into the lost tooth region requires that the transplanted germ erupt and properly occlude with the opposing tooth in an adult jawbone (Figure 4, Upper panel). Tooth eruption is a corporative regulatory mechanism that involves the tooth germ cell component and the surrounding alveolar/ jawbone. It has been shown that follicle cells migrate from near the surface of the enamel organs and dental papillae to give rise to the cementum, periodontal ligament, and alveolar bone. Because these cells affect overlying bone resorption and can cause enzymatic degeneration in growth and bodily movement of the teeth, they most likely play a role in tooth eruption. Follicle cells may contribute to the formation of tissues surrounding and underlying the teeth during tooth eruption. So far, it has been demonstrated that transplanted natural tooth germ erupts in a murine toothless diastema region. Furthermore, bioengineered tooth germ can only develop correct tooth structure in an oral cavity (Nakao et al., 2007) and can successfully erupt 37 days after transplantation. The bioengineered tooth subsequently reaches the occlusal plane and achieves and maintains occlusion with the opposing tooth from 49 days onward (Figure 6(a) and 6(b); Ikeda et al., 2009). In the case of a transplanted bioengineered mature tooth unit comprising a mature tooth, periodontal ligament, and alveolar bone, the most critical consideration is whether the unit can be engrafted into the tooth loss region through bone integration, which involves natural bone remodeling in the recipient. A bioengineered tooth unit transplanted at a position reaching the occlusal plane with the opposing upper first molar was successfully engrafted at 40 days and thereafter maintained the periodontal ligament originated from the bioengineered tooth unit through successful bone integration (Figure 6(b); Oshima et al., 2011). The enamel and dentin hardness of bioengineered tooth components were in the normal range when analyzed by the Knoop hardness test (Ikeda

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Figure 6 Bioengineered tooth regeneration by the transplantation of bioengineered tooth germ or a tooth unit. (a) A transplanted bioengineered tooth germ erupted and reached the occlusal plane with the opposing tooth 49 days after transplantation (left panels). The GFP-labeled tooth erupted in an adult oral environment (right panel). (b) The erupted bioengineered tooth occluded with the opposing natural tooth (left panels). The bioengineered tooth unit was engrafted by bone integration at a position reaching the occlusal plane 40 days post transplantation (right panels). :: Bioengineered tooth.

et al., 2009; Oshima et al., 2011). These approaches demonstrate the potential to successfully recover masticatory performance and natural tooth tissue through new technologies.

Bioengineered Tooth Response to Mechanical Stress Loss of teeth and functional disorders in periodontal ligaments cause fundamental problems for oral functions, including pronunciation, mastication, occlusion, and associated health issues (Proffit et al., 2004). The periodontal ligament plays important roles in the pathogenic and physiologic responses of teeth to extreme mechanical forces from bone remodeling accompanied by orthodontic tooth movement. Autologous tooth transplantations following a traumatic dental injury have indicated that natural periodontal tissue remaining on the root could successfully restore physiological tooth function, including bone remodeling, and effectively prevent ankylosis. In contrast, the absence of a periodontal ligament in dental implants is associated with deficiencies in essential tooth functions and in the natural structural relationship with the root and alveolar bone (Huang et al., 2009). The periodontal ligaments of bioengineered teeth that erupted following the transplantation of bioengineered tooth germ and mature tooth units achieved functional tooth movement as well as that of natural tooth. They successfully reproduced bone remodeling via the proper localization of osteoclasts and osteoblasts in response to mechanical stress, indicating that a bioengineered tooth can replicate critical dental functions through the restoration and reestablishment of cooperation with the surrounding jawbone (Ikeda et al., 2009; Oshima et al., 2011).

Perceptive Neuronal Potential of Bioengineered Teeth The peripheral nervous system is established by the growth of axons that navigate and establish connections with developing target organs during embryogenesis. The perceptive potential for noxious stimulation, including mechanical stress and pain, is important for proper organ function. It is also thought that the recovery of the nervous system, which is associated with the reentry of nerve

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fibers following organ transplantation, is critical for organ function reconstitution. Teeth are a peripheral target organ for the sensory trigeminal and sympathetic nerves, both of which play important roles in tooth function and protection. It is anticipated that tooth regenerative therapies will be able to recover the neuronal perceptive ability of mechanical forces that are lacking in implant patients. Sensory and sympathetic nerve fibers innervate both the pulp and periodontal ligament of a bioengineered tooth following its eruption (Ikeda et al., 2009). These bioengineered teeth display appropriate perceptive potentials for nociceptive pain stimulations, such as pulp stimulation and orthodontic treatment, and they can properly transduce these events to the central nervous system through c-Fos immunoreactive neurons (Ikeda et al., 2009; Oshima et al., 2011). In this way, bioengineered teeth can indeed restore the perceptive potential for noxious stimuli in cooperation with the maxillofacial region.

Future Perspectives for the Tooth Regeneration There are several problems that must be solved before bioengineered teeth become feasible. To fully realize the practical clinical application of tooth regenerative, suitable cell sources must be identified. Ideally, tooth regenerative therapy should employ the patient’s own cells to avoid immunological rejection (Ikeda and Tsuji, 2008). Recent studies of stem cells and organogenesis have led to considerable advances in our understanding of the potential of cell sources for tissue repair and organ reconstitution, including tooth regenerative therapy. Tooth tissue-derived stem cells such as DPSCs, SHED, SCAP, PDLSCs, and DFSCs can differentiate into dental cell lineages and contribute to the turnover and supply of various cell populations. These lineages will be useful cell sources for stem cell transplantation therapy for dental tissue repair of dental caries and periodontal disease (Ikeda and Tsuji, 2008; Huang et al., 2009), but human teeth do not regenerate like hair or skin, and there are no stem cell niches in teeth or surrounding tissue that maintain progenitors for whole tooth reproduction. Current whole tooth regenerative therapy research is attempting to induce bioengineered tooth germ to develop a fully functioning tooth using embryonic tooth germ-derived epithelial and mesenchymal cells via the organ germ method (Nakao et al., 2007; Ikeda et al., 2009; Oshima et al., 2011). In the future, it will be important to identify cell sources from somatic dental and nondental tissue-derived stem cell populations isolated from patients that have the potential to reproduce the epithelial and mesenchymal interactions that occur during organogenesis, as well as tooth-forming ability. It will also be critical to recognize inductive genes with the potential to initiate tooth organogenesis, reconstitute bioengineered tooth germ, and ultimately to develop a functional bioengineered tooth. Candidate cell sources for whole tooth regeneration also include embryonic stem cells and induced pluripotent stem (iPS) cells, which are capable of differentiating into endoderm, ectoderm, and mesoderm (Takahashi and Yamanaka, 2006). Recently, iPS cells have been established from various oral tissues. However, iPS cells could be applied with reprogrammed progenitor cell lines; iPS programming procedures are required to establish dental epithelium and mesenchyme fates. Although there are multiple candidates for tooth developmental genes that promote significant expression of dental epithelial and mesenchymal cells, the master genes responsible for tooth development, remain to be discovered. The important task for determining a future tooth regenerative cell source is the identification of specific combinations of factors capable of reprogramming nondental cells to dental epithelium and mesenchyme. Tooth types such as incisors, canines, premolars, and molars have unique morphological features that are programmed at predetermined sites in the oral cavity during tooth development. Several studies have proposed molecular mechanisms for tooth morphology regulation (Ishida et al., 2011). Tooth size as well as crown and root shape are important considerations when generating a bioengineered regenerated tooth with proper functional occlusion and esthetics.

Concluding Remarks The progress of regenerative technology is remarkable, and many patients look forward to the implementation of tooth regenerative therapy. Further studies are required to establish bioengineering technologies that can control tooth morphology, including tissue engineering using scaffolds, the identity of morphogenesis-related genes, and appropriate cytokines to use to guide morphogenesis. Tooth regenerative therapy is now regarded as a crucial study model for future replacement regenerative therapies that can be applied to more complex organs and will contribute substantially to the knowledge and technology required to regenerate organs (Ikeda and Tsuji, 2008; Sharpe and Young, 2005).

References Avery, J. K. (2002). Oral Development and Histology. New York: Thieme Press. Brockes, J. P., & Kumar, A. (2005). Appendage regeneration in adult vertebrates and implications for regenerative medicine. Science, 310, 1919–1923. Duailibi, S. E., Duaibili, M. T., Vacanti, J. P., et al. (2006). Prospects for tooth regeneration. Periodontol., 2000(41), 177–187. Huang, G. T., Gronthos, S., & Shi, S. (2009). Mesenchymal stem cells derived from dental tissues vs. those from other sources: their biology and role in regenerative medicine. J. Dent. Res., 88, 792–806. Ikeda, E., & Tsuji, T. (2008). Growing bioengineered teeth from single cells: potential for dental regenerative medicine. Expert Opin. Biol. Ther., 8, 735–744. Ikeda, E., Morita, R., Nakao, K., et al. (2009). Fully functional bioengineered tooth replacement as an organ replacement therapy. Proc. Natl. Acad. Sci. U.S.A., 106, 13475–13480.

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Ishida, K., Murofushi, M., Nakao, K., et al. (2011). The regulation of tooth morphologenesis is associated with epithelial cell proliferation and the expression of Sonic hedgehog through epithelial-mesenchymal interactions. Biochemi. Biophys. Res. Commun., 405, 455–461. Nakao, K., Morita, R., Saji, Y., et al. (2007). The development of a bioengineered organ germ method. Nat. Methods, 4, 227–230. Oshima, M., Mizuno, M., Imamura, A., et al. (2011). Functional tooth regeneration using a bioengineered tooth unit as a mature organ replacement regenerative therapy. PloS One, 6, e21531. Proffit, W. R., Fields, H. W., & Sarver, D. M., Jr. (2004). Contemporary Orthodontics. St. Louis, MO: Mosby Press. Saito, M., Kurokawa, M., Oda, M., et al. (2011). ADAMTSL6b rescues fibrillin-1 microfibril disorder in Marfan syndrome mouse model through the promotion of fibrillin-1 assembly. J. Biol. Chem., 286(44), 38602–38613. Sharpe, P. T., & Young, C. S. (2005). Test-tube teeth. Sci. Am., 293, 34–41. Sonoyama, W., Liu, Y., Fang, D., et al. (2006). Mesenchymal stem cell-mediated functional tooth regeneration in swine. PloS One, 1, e79. Takahashi, K., & Yamanaka, S. (2006). Induction of pluripotent stem cells from mouse embryonic and adult fibroblast cultures by defined factors. Cell, 126(4), 663–676. Thesleff, I. (2003). Epithelial-mesenchymal signalling regulating tooth morphogenesis. J.Cell Sci., 116, 1647–1648. Yen, A. H., & Sharpe, P. T. (2006). Regeneration of teeth using stem cell-based tissue engineering. Expert Opin. Biol. Ther., 6, 9–16.

Vascularized Tissue Regenerative Engineering Using 3D Bioprinting Technology Sungwoo Kim, Arnaud Bruyas, Chi-Chun Pan, Alexander Martin Stahl, and Yunzhi Yang, Stanford University, Stanford, CA, United States © 2019 Elsevier Inc. All rights reserved.

The Importance of Prevascularization in Tissue Engineering Processes by Which Implants Vascularize In Vivo Strategies to Create an In Vitro Vascularization Model System Biomaterials for Vascularization In Vitro Endothelialization Angiogenic Growth Factors and Biomolecules Mechanobiology Perfusion Systems Vascular Maturation in Coculture Systems 3D Printing Methods for Prevascularized Tissue Constructs General Process of Forming a 3D Printed Scaffold Forming acellular scaffolds for supporting the vascularized scaffold Manufacturing the cell-laden constructs Challenges of 3D Printed Tissue Construct Applications in Vascularization in Bone and Cardiac Tissue Engineering Applications in Bone Tissue Engineering Challenges for engineered bone grafts Strategies for Developing Vascularized Bone Grafts Applications in Cardiac Tissue Engineering Challenges for engineered cardiac grafts Strategies for developing vascularized cardiac tissues Summary Further Reading

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Glossary Additive manufacturing (AM) A processes used to synthesize a three-dimensional object in which successive layers of material are formed under computer control to create an object. Free flap A term used to describe the transplantation of tissue with intact vascular bed from one site of the body to another in order to reconstruct an existing defect with the circulation in the tissue reestablished by anastomosis of arteries and veins. Lumen The inside space of a tubular structure in which blood flows. Matrix metalloproteinases (MMPs) The calcium-dependent zinc-containing endopeptidases. These enzymes are capable of degrading all kinds of extracellular matrix proteins, but also can process a number of bioactive molecules. Mechanobiology An emerging field of science at the interface of biology and engineering that focuses on how physical forces and changes in the mechanical properties of cells and tissues contribute to development, cell differentiation, physiology, and disease. Neovascularization The natural formation of new blood vessels, usually in the form of functional microvascular networks, capable of perfusion by red blood cells. Osteoinductivity A property of graft material in which it induces de novo bone growth with biomimetic substances, such as bone morphogenetic proteins. Photo-crosslinking The photo-induced formation of a covalent bond between two macromolecules or between two different parts of one macromolecule.

The Importance of Prevascularization in Tissue Engineering Nearly all metabolically active tissues in the body contain a hierarchical network of blood vessels that is essential for maintaining tissue viability and function. These vascular networks are responsible for delivering oxygen and nutrients to the cells that make up tissues and for removing the byproducts of cellular metabolism. In damaged tissues, reestablishment of a functional vascular system

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is an essential step for wound healing and in the event a tissue graft is required, achieving integration of the implant with the surrounding host tissues depends on the formation of a functional vasculature across the graft. Insufficient blood flow to a tissue or implant leads to hypoxia, which in turn leads to cell death in the affected region if not resolved quickly enough. Providing blood flow throughout an engineered tissue in a timely manner is thus essential for preventing graft failure and depends on the successful formation of a vascular network within the implant as well as the connection of such a network to the surrounding host circulation. Currently, the challenge of achieving complete and timely vascularization across large-scale tissue constructs limits the maximum volume at which engineered tissues can remain viable and impedes their translation to clinical practice. A functional vascular bed must furnish an entire tissue with sufficient oxygen and nutrients at the cellular level. In order to meet these demands, the native vasculature is structured hierarchically, with large arteries carrying the necessary volume of blood to the target region and then branching into multiple smaller arterioles in order to achieve more even distribution throughout the tissue. Venules collect the deoxygenated blood and converge into large veins that carry the blood away from the tissue. Capillaries, which bridge the space between arterioles and venules, constitute the smallest blood vessels in the vascular system with an average diameter of just 8–10 mm. In most metabolically active tissues, the distance between capillaries and their target cells is restricted to 100– 200 mm in order to meet the high cellular demands for oxygen, which imposes a need for extremely dense capillary networks. Due to the difficulty of reproducing the inherent complexity of native biological tissues and their vasculature, tissue-engineered grafts often exhibit simplified structures, with the expectation that more complete integration and remodeling by the host will follow. As such, the majority of current scaffolds used in tissue engineering lack a functioning vascular network prior to implantation, relying instead on the surrounding host vasculature to invade the graft. Unfortunately, in many cases, the extent of vascular ingrowth by host vessels into implanted tissue constructs is insufficient, leading to failure of the grafts. Achieving complete revascularization is a key step in the implementation of engineered tissue constructs in clinical settings. The following review will introduce the biological processes involved in vascularization in vivo and strategies to reproduce these processes in vitro, followed by a description of current 3D printing technologies and their potential application for engineering vascularized tissue grafts, with particular focus on applications in bone and cardiac tissue engineering.

Processes by Which Implants Vascularize In Vivo The restoration of a functional vascular network within engineered tissue constructs is known as vascularization. There are two processes by which the formation of new blood vessels can occur: vasculogenesis and angiogenesis. Vasculogenesis is the process of forming de novo blood vessels through the differentiation of progenitor cells to endothelial cells (ECs) and their organization into a capillary network within previously avascular tissue (Fig. 1A). Vasculogenesis primarily occurs during prenatal development, where a primitive capillary network initially forms prior to being remodeled into the more complex hierarchical networks of the mature vasculature. In adults, vasculogenic processes have been observed to contribute to the revascularization of damaged tissues: cues such as hypoxia, growth factors, and cytokines recruit endothelial progenitor cells from the bone marrow, augmenting the population of circulating endothelial progenitor cells. These cells home to the site of vascularization by integrin interactions and engage in tissue healing through a combination of self-assembly into new vessels and growth factor production. Angiogenesis refers to the sprouting or splitting of new capillaries from existing blood vessels and represents the chief mechanism by which new blood vessels form in adults (Fig. 1B). Angiogenic sprouting is the primary route for vascular formation in implanted tissue grafts, with new capillaries growing from surrounding blood vessels to invade nearby engineered tissue constructs.

(A)

Fig. 1

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(C)

The different routes to vascularization in vivo (A) vasculogenesis; (B) angiogenesis; (C) vasculogenesis and angiogenesis.

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The process is initiated by the release of vascular signaling factors from parenchymal cells in poorly perfused tissues in response to hypoxic conditions. In response to these signalling molecules, certain endothelial cells in the surrounding tissues transition into tip cells, which put forth long cellular processes called filopodia to detect gradients in the angiogenic factors and guide endothelial migration in the direction of the hypoxic tissues. The tip cell secretes proteolytic enzymes to clear a path through the surrounding extracellular matrix (ECM) and endothelial stalk cells proliferate behind the tip cell, allowing the developing sprout to extend toward the hypoxic region. The stalk cells produce internal vacuoles that fuse to form a central lumen hollowing out the growing capillary. A host vascular network will generally invade an implanted tissue graft by angiogenic sprouting as a natural response to the presence of the foreign construct. However, due to the slow progression of the tip cells and proliferating stalk cells, in larger implants complete vascularization can take weeks, if it occurs at all. Vasculogenesis and angiogenesis are not mutually exclusive modes of vascularization. Endothelial progenitor cells are often implicated in both processes and might adopt one of several complimentary roles in the formation of new vascular networks within regenerating tissues (Fig. 1C). A single developing microvascular network might be simultaneously expanded by the vasculogenic differentiation and assembly of endothelial progenitor cells which interconnect and merge with the angiogenic sprouting of existing vessels. Once an initial capillary network forms, mural cells (MCs) are recruited to support the developing vasculature. For capillaries, these cells are called pericytes. In larger vessels, smooth muscle cells perform this role. The presence of MCs is essential for stabilizing the newly formed vessel beds, preventing vascular regression and allowing for further capillary reorganization and maturation into the native hierarchical vascular framework. The invasion of ECs and their supporting cells into an engineered tissue construct to form a functional vasculature can take days to weeks, as the rate of capillary ingrowth is generally limited to several tenths of micrometers per day. While this may be sufficient to meet the demand for oxygen and nutrients in thin or avascular tissue grafts such as skin and cartilage, large-scale grafts for most engineered tissue constructs are at high risk of nutrient deficiency and hypoxia, with cells at the center of such implants typically experiencing extreme anoxic stress that can lead to graft necrosis. Some strategies to overcome this challenge seek to include preformed channel networks within the scaffolds in an attempt to guide endothelial invasion either with or without direct surgical attachment to the host vasculature. Other strategies seed grafts with ECs or endothelial progenitor cells in order to eliminate the need for slow vascular invasion via angiogenic sprouting. Preseeding constructs with ECs additionally enable primitive endothelial network formation in vitro so that a provisional microvasculature already exists within the scaffold prior to implantation.

Strategies to Create an In Vitro Vascularization Model System Prevascularization aims to form functional vascular networks within an engineered tissue construct in vitro under controlled conditions prior to implantation. This can improve cell survival and functionality in vivo via rapid establishment of perfusion by integration with the host vasculature, which can generate more physiologically relevant tissues. There are numerous approaches to engineer tissue constructs with vascular networks. Each strategy has unique advantages and disadvantages. Generally, the capillary-like networks must be self-assembled and self-organized by inclusion of relevant cell types because they are too small to be fabricated. Thus, the strategy of vascularization in vitro needs to mimic the several stages of the biological mechanisms involved in vascularization in vivo including cell–cell interaction, elongation, network formation, and maturation. Below, we will discuss general strategies for creating in vitro model systems of vascularization.

Biomaterials for Vascularization The selection of appropriate biomaterials is the first important step for the prevascularization of the tissue construct. A variety of biomaterials have been studied to mimic structural and functional properties of the natural ECM. Whether natural or synthetic, the ideal material should contain several key properties such as biocompatibility, biodegradability, tunable mechanical strength and swelling behavior, and binding sites that interact with cells to enable attachment and growth. The biomaterials must not be toxic and should provide a 3D matrix for cells to grow, proliferate, and function normally. Any degraded by-products should not interrupt cell behavior or tissue regeneration. The degradation rate should also be controllable so that cells can move freely and organize themselves within the three-dimensional (3D) matrix. In this regard, hydrogel-based materials have been widely used to create a 3D microenvironment that allows ECs to self-assemble and organize into functional vascular networks. Although ECs can form cord-like structures in 2D culture, they can only form functional tubular structures with lumens when cultured in a 3D microenvironment. However, cells in a 3D environment will not lead to capillary-like tube formation unless the matrix has binding or cleavage sites for cell attachment and degradation, respectively. Many natural materials have RGD (Arg-Gly-Asp) binding and matrix metalloproteinase (MMP)-degradable sites, but synthetic materials need to be incorporated with binding and cleavable sites. Natural materials that have been used to provide a matrix for prevascularization of tissue constructs include collagen, fibrin, matrigel, decellularized extracellular matrix, silk, and so on. Among them, collagen is the most abundant protein in mammalian extracellular matrix, and has been widely used for prevascularization. It degrades faster than synthetic materials, and its physicochemical properties can be readily manipulated to promote angiogenesis by mixing with other biomaterials such as chitosan, fibrin, and hyaluronic acid. Fibrin plays important roles in blood clotting and wound healing and provides binding for cell adhesion and growth factor attachment. Fibrin can also promote angiogenesis by incorporation of growth factors or coculturing ECs with MCs. Matrigel is

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one of the most widely used natural materials extracted from mouse sarcoma cells. It is pro-angiogenic, but is challenging to control batch-to-batch reproducibility with regard to chemical composition that makes it difficult to apply to clinical practice. Synthetic materials, on the other hand, can be made via controllable chemical reactions from different precursors, and are easily reproduced at large scales. They have tunable mechanical properties, degradation rates, and porosity. However, they do not have cell binding sites, and thus require additional chemical modification to promote cell–material interactions. Many synthetic materials that have been used for the prevascularization of tissue constructs include polyethylene glycol (PEG), polylactic acid (PLA), polyglycolic acid (PGA), polylactic-glycolic acid (PLGA), polycaprolactone (PCL), polyurethane, and so on. Among them, PEG-based hydrogels have been widely studied for vascularization applications. The PEG terminal hydroxyl groups can be modified with various biomolecules to improve cell adhesion, growth, and functionality. The degradation of PEG, which naturally occurs very slowly, can be accelerated by incorporating other materials such as collagen, fibrin, and MMP-sensitive molecules. Its mechanical properties can be easily modulated by using different concentrations, molecular weights, and cross-linking densities. Appropriate biomaterial selection plays an important role in the success of prevascularized tissue constructs. Ideal biomaterials need to maintain the mechanical integrity of the tissue construct without interfering with preformed vascular networks during the remodeling process.

In Vitro Endothelialization The next important issue to be discussed is the alignment and organization of ECs within an in vitro system. There are various cell sources for prevascularized tissue constructs, including autologous circulating endothelial progenitor cells (EPCs), postnatal stem cells, and induced pluripotent stem cells (iPSC) that can act as precursors to ECs or other vascular cell phenotypes. In numerous studies, the addition of ECs or EPCs into engineered tissue constructs has yielded increased vascularization and perfusion both in vitro and in vivo. ECs are a key cell type that plays an important role in formation of structural and functional blood vessels. EPCs have also showed angiogenic potential with enhanced proliferation and survival rate. However, randomly organized ECs or EPCs in vitro will induce uncontrolled vascular networks that may not supply larger constructs with sufficient nutrients or oxygen after implantation. This is because the randomly formed in vitro networks may not be primed for spontaneous anastomosis, leading to a delay or failure of blood perfusion in vivo. Generally, the in vitro formation of vascular networks in hydrogels takes time, and regression after a few days of culture is common. Thus, the in vitro organization of ECs needs to be predesigned and controlled by patterning the vasculature within the hydrogels. One strategy is to culture ECs in hollow microchannels to avoid initial capillary formation. The EC lined channels are then used as a template for further vessel formation. In addition, different cell types can be seeded separately in the surrounding biomaterial to encourage vascular maturation. There are various approaches to create the hollow microchannels and patterns using laser drills, silicon molds, sacrificial materials, and so on. However, formation of highly organized networks is usually limited by the resolution of the channel-forming methods, and nondegradable dense hydrogel layers in in vitro model systems restrain further vascular remodeling. Another strategy is to directly pattern the endothelial cells within a controlled 2D construct under in vitro conditions using PDMS molds, photopatterning, and cell sheets to enhance the initial EC alignment and promote organization of vascular networks. More complex 3D vascular structures can be obtained by combining these 2D constructs or using 3D bioprinting technology. 3D bioprinters can deposit bioinks containing cells or cell aggregates at a designated area with a desired local density, high resolution, and anatomical shape. However, the initial organization of the vascular structures within the in vitro system is easily changed during vascular remodeling. Therefore, further strategies are necessary to maintain the long-term functionality of the preformed vascular networks.

Angiogenic Growth Factors and Biomolecules Sustained and localized release of growth factors is a well-defined chemotactic cue to control vascular organization and angiogenesis in engineered tissue constructs. Since vascular organization consists of multiple phases including the initial network formation and maturation, the patterning of multiple growth factors in space and time is important to control both stages. Vascular endothelial growth factor (VEGF) has proangiogenic functions such as inducing EC elongation, network formation, and branching. The combination of VEGF with other growth factors such as platelet-derived growth factor (PDGF) and Angiopoietin 1(Ang1) can improve both vascular structure formation and maturation. PDGF that is released by ECs in the later stage of vessel formation recruits MCs such as pericytes, and plays an important role in their proliferation and migration toward the preformed vessel networks. Ang1 released by MCs and ECs–MCs contact significantly affect EC quiescence and stabilization. As such, patterning and release of multiple growth factors over different time periods is a major challenge to promote both local vascular formation and stabilization. One approach is to use a dual growth factor delivery system consisting of a polymer scaffold and microspheres. VEGF can be incorporated into a polymer scaffold and PDGF can be encapsulated into polymer microspheres that are mixed with the polymer scaffold. Thus, the release of VEGF and PDGF with different kinetics can promote rapid vessel formation via sustained release of VEGF and vessel maturation by recruited pericytes via release of PDGF. Another approach is to coculture ECs with supporting cells such as fibroblasts, pericytes, and mesenchymal stem cells (MSCs) without additional chemical stimulus. The supporting cells can modulate the expression of multiple angiogenic factors including VEGF, Ang1, basic fibroblast growth factor (bFGF), transforming growth factor beta (TGF-b), laminins, and integrins to produce biologically regulated cues based on their current needs. Thus, controlling cross-talk between different cell types is one of important ways to enhance vessel formation and maturation.

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Mechanobiology Cell behaviors correspond to the mechanical properties of local microenvironments. Thus, another important requirement for vascular organization is to control local stiffness of the biomaterials as measured by the compression modulus. The optimal matrix stiffness varies according to the type of biomaterials used for vascular formation, but it is generally agreed that decreasing stiffness of the materials is likely to induce more EC capillary formation. It has been reported that human umbilical vein endothelial cells (HUVECs) expressed higher levels of functional VEGF receptor-2 protein when low stiffness collagen (3kPa) was used for the culture matrix as compared with high stiffness collagen (30kPa). However, optimal stiffness of the materials must often be determined by taking into consideration conflicting needs in a tissue construct. For example, in designing a prevascularized bone tissue, a soft matrix is required for vascularization while a rigid scaffold is preferred for bone formation. To reach a suitable compromise, it is often a promising approach to place different cell types in separate materials with optimal properties for each desired characteristic. But it is not easy to achieve high spatial resolution between regions of significantly different local stiffness during the creation of a prevascularized tissue in vitro. It is challenging to find mechanobiology strategies capable of combining materials of different properties in order to balance the various needs of the model system.

Perfusion Systems Most cells in the body are known to be located within 100–200 mm from the nearest capillaries that transport nutrients, oxygen, and waste products to support cell and tissue viability. Similarly, the engineered tissue constructs need to be cultured with appropriate perfusion systems providing media flow that enables mass transfer of oxygen and soluble factors to promote cell growth and distribution. In addition, the media flow in perfusion systems can be directed through hollow microchannels within the tissue construct to mimic hemodynamic forces and pressures. The media flow creates shear stress on the surface of the cells, which is an important signal to guide the vascular organization and maturation of ECs and EPCs. The mechanical cues provided by perfusion play an important role in stimulating the formation of ECM molecules such as elastin and collagen, regulate EC proliferation, and improve vascular mechanical properties in vitro.

Vascular Maturation in Coculture Systems Successful prevascularization of engineered tissue constructs relies on stabilization of the vasculature in vitro prior to implantation. Previous studies have demonstrated that immature microvessels are not likely to induce anastomosis with the host vessels and are prone to regression. This is because vascular networks that are not stabilized during vascular remodeling become disorganized and leaky. Coculture with supporting cells such as MCs and MSCs can directly promote stabilization of EC capillary formations. Pericytes can help maintain newly formed capillary tubes by releasing tissue inhibitor of metalloproteinase 2 and 3 (TIMP-2/3). Pericytes also induce ECs to deposit ECM components including collagen IV, fibronectin, laminin, perlecan, and so on. In addition, the coculture of mural precursor cells such as embryonic fibroblasts or MSCs with ECs has been shown to increase the luminal content of the endothelial structures, which is an indicator of the functionalization and maturation of vascular networks. Additionally, the coculture enhances expression of smooth muscle actin (a-SMA) of MSCs and promotes their differentiation into MCs. However, coculturing multiple cell types in the same system is challenging because the optimal culture conditions for ECs may not be appropriate for other cell types. There are numerous factors that can affect cell behavior in a coculture system including the cell sources, cell ratios, seeding densities, culture media formulations, cell seeding methods, and the types of biomaterials used. Balancing all these conditions is difficult and may not be sufficient to achieve the desired functionality of the end tissue construct. Thus, an in vitro coculture model system should be designed by taking into consideration the purpose of the study. If the different cell types are mixed simultaneously, then the homogeneous mixture can be distributed throughout the in vitro model system. This method can enhance cell–cell contact and cell functionality if these cell types are closely located with one another in the tissue. However, the timing of maturation is also important during in vitro prevascularization because vessel maturation usually starts by suppressing EC growth. Initiation of the maturation process should occur after ECs have formed vascular networks that cover the entire tissue construct. But if it is left too late, then the preformed networks will regress. To avoid this problem, the different cell types can be seeded sequentially. Culturing ECs first will provide enough time to induce capillary-like network formation, and the subsequent addition of MCs or MSCs will support the preformed networks in terms of stability and functionality. However, the sequential seeding method is technically challenging, as the second type of cells must be incorporated into the in vitro model system after capillary-like network has already been formed. In order for this to be achieved, the model system should include porous structures, hollow microchannels, or provide chemical stimuli to promote ingrowth of the second type of cells. The immense complexity of native tissues is challenging to mimic in tissue engineered constructs, which must attempt to harmonize a myriad of biomolecular conditions and micro to macroscale architectural designs using only nontoxic materials that can remodel and integrate with host tissues. In order to manufacture such advanced model systems and tissue grafts, 3D printing techniques have gained prominence within tissue engineering as a means to achieve high control over construct composition and spatial organization that would not be possible to replicate by traditional fabrication processes.

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3D Printing Methods for Prevascularized Tissue Constructs 3D printing technologies have led to significant advancements in tissue engineering and regenerative medicine. With 3D biofabrication techniques, researchers have generated prevascularzied tissue constructs with complex 3D architectures to mimic the microenvironment and biological components of native tissues. Compared with traditional manufacturing processes, such as particulate leaching and solvent casting, 3D printing provides greater precision and automation with higher processing speeds for engineering the tissue constructs. A variety of 3D printing technologies exist to manufacture different 3D printed tissues with tailored applications and material properties. Thus, different 3D printing techniques must sometimes be combined to most effectively print specific tissue components within a complex construct. Including a preformed vasculature within 3D printed constructs necessitates selecting a printing technique suitable for both ECs and any secondary cells required for the tissue of interest. If such conditions cannot be found, as is often the case, then a combination of 3D printing technologies might be selected to recreate the various properties of the constituent tissue parts. Several 3D printing techniques are introduced in the following sections, categorized based on whether the printing materials can include cells or not (acellular or cell-laden materials).

General Process of Forming a 3D Printed Scaffold The 3D printing process can generally be divided into 4 steps: generating 3D computer-aid-designed (CAD) models, preprocessing preparation, manufacturing, and postprocessing treatment (Fig. 2). The 3D CAD models can be designed by specialized software such as SolidWorks and AutoCAD, or can be acquired from computed tomography (CT) or magnetic resonance imaging (MRI). The next step is to convert the CAD model into commands for the 3D printer. The third step is to run the manufacturing process; based on the preprocessing files, the product is then printed in a layer-by-layer fashion. Sometimes postprocessing modifications to the printed product might be necessary, for instance, after fused deposition modeling (FDM) production, it may be required to remove supporting materials, and constructs manufactured by selective laser sintering (SLS) must be baked in order to stabilize the object through crosslinking. In tissue engineering, 3D printed scaffolds have been used to provide mechanical support, encapsulate drugs, and position cells in a 3D environment. 3D printing technologies can be differentiated into two categories, acellular and cell-laden, depending on whether the incorporation of cells in the scaffold is compatible with the 3D printing process.

Forming acellular scaffolds for supporting the vascularized scaffold Due to limitations such as high temperatures or extruding pressures during the manufacturing process, acellular scaffold technologies cannot be used for encapsulating cells. Instead, with high stiffness, these scaffolds are usually used as supporting structures within engineered tissues. FDM and SLS are the general methods to fabricate acellular scaffolds. 1. Fused deposition modeling (FDM) and syringe-based extrusion FDM is the most common technology for commercialized 3D printers due to its low cost and the availability of printing materials. FDM 3D printers mainly consist of three parts including the printing filament, an extruder, and a printing stage (Fig. 3A). The FDM 3D printer extruder is able to move in the XY plane, and the printing stage moves along the Z-axis. The 3D printing filament is the raw material that forms the final scaffold, and is typically made of polymer-based materials such as polycaprolactone (PCL), polylactic acid (PLA), or composite materials such as polycaprolactone-tricalcium phosphate (PCLTCP). During the extrusion process, a motor in the extruder drives the filament through the heater at the hot end, the filament in the heating zone is then melted, and the hot end is thus filled with molten material. Motors in the extruder rotate as prescribed by the G-codes and drive the cold filament which is stiff to push out the molten material at the hot end. To print complex shapes of a scaffold such as suspended bridges, supporting structure is required. In addition, when nozzle moves in the air, the redundant material from the last move is attached and floated with the nozzle; hence some redundant strands are formed. After the printing process, any redundant strands or supporting materials are removed from the scaffold. The G-code programming language provides the commands to control the FDM 3D printer. The user is not required to generate all the G-code by themselves; instead, open source G-code generators such as Slic3r (http://slic3r.org/) can help the users to

G1 ×1 Y1 E0.01 G1 ×1 Y1.2 E0.01 G1 ×1.2 Y1.2 E0.02 G1 ×1.2 Y1.4 E0.01 G1 ×1 Y1.4 E0.01 .. .

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Fig. 2 The 3D Printing Process. The first step of 3D printing is designing the 3D CAD model using specialized software, and then translating the model into commands to the printer during the preprocessing preparation. During the manufacturing process which runs based on the code generated from the preprocessing preparation, the construct is printed. Some processes might be required for postprocessing modification.

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Fig. 3 The 3D printing process. (A) Fused deposition modeling (FDM). (B) Syringe-based extrusion printing. (C) Selective laser sintering (SLS). (D) Stereolithography or digital light processing-based stereolithography (DLP-SLA). (E) Inkjet-based printing. (F) Laser-based printing.

convert their 3D model into G-code if given the appropriate defined parameters, such as porosity, nozzle diameter, and printing pattern. Syringe-based extrusion is a similar 3D printing technique to FDM. While FDM heats up the filament to extrude the strand, syringe-based extrusion extrudes viscous biomaterials by pneumatic pump (Fig. 3B). Since the material is stored in the syringe under low temperature, the syringe-based extrusion method can be used for printing cell-laden materials. 2. Selective laser sintering (SLS) SLS printing mixes powders made of polymer, ceramics, or composites with crosslinkers and uses this mixture to fill a pool (Fig. 3C). The CAD model is converted to a path based on the desired dimensions and porosity. The manufacturing process starts with moving the printing stage downward until it is below the level of the of the powder mixture surface by a fixed distance, then the roller pushes the powder and crosslinker mixture onto the stage. A laser beam runs through the desired paths to initiate the crosslinking. After a layer is crosslinked, the stage with the powders moves downward again, and the roller fills the printing stage with powders. The manufacturing process repeats until all the layers are printed. The SLS printed construct, however, must be further baked during postprocessing to stabilize the crosslinked structures.

Manufacturing the cell-laden constructs Stereolithography (SLA), Digital-light-processing-based SLA (DLP-SLA), and inkjet-based methods are the most common 3D printing methods to manufacture cell-laden hydrogels. The printing must be performed under sterile conditions when the cellladen hydrogels are printed, and the printing materials should be mostly water-based and biocompatible. 1. Stereolithography (SLA) and digital light-based stereolithography (DLP-SLA) The SLA and DLP-SLA methods are used to form photo-crosslinked polymer networks when placed under a light source. During the preprocessing preparation of SLA, the 3D CAD model is sliced into several layers based on the desired layer thickness. When it starts printing, after the stage moves up a layer in Z-axis, the laser beam runs through the designed paths and forms a photocrosslinked hydrogel layer containing cells. After a layer is formed, the stage moves to the next layer, and these processes repeat until an engineered tissue construct is built. On the other hand, DLP-SLA method projects each cross-section image directly into the polymer solution to fabricate the tissue constructs at once instead of using a laser beam (Fig. 3D). For engineering cell-laden tissue constructs, the printing materials for SLA and DLP-SLA should be biocompatible, and can be made from either synthetic or natural polymers such as polyethylene glycol (PEG), gelatine, and collagen, or the composites such as gelatin methacryloyl (GelMA). The properties of the printing materials such as porosity, mechanical stiffness, and degradation are easily adjustable by controlling the light exposure times, the concentrations of the polymers, and the photoinitiators. The longer the exposure time is, the higher the degree of crosslinking within the engineered tissue construct. Under the same exposure time, with higher concentration of the polymers or the photoinitiators, the tissue construct becomes denser and stiffer. In order to print prevascularized tissue constructs, the EC and GFs can be homogeneously mixed in the prepolymer solution to form the structure by SLA. 2. Inkjet-based deposition Inkjet-based 3D printers are designed based on the traditional 2D inkjet paper printer (Fig. 3E). The printing material is stored inside the chamber at the print head under low temperature to maintain low viscosity of the material, such as thermally sensitive

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gelatin- or collagen-based polymer solution. When depositing the droplet, the heater heats up the vapor bubble inside the print head to increase its volume, and this droplet cuts the stream of the material and pushes out the droplet physically. In the preprocessing preparation, the CAD model is sliced into layers, and each layer is converted into array of droplets. During the manufacturing process, the print head ejects the droplets onto the print stage under the low temperature, layer by layer. The printed tissue construct is then moved to an incubator to cure the droplets. As mentioned, the printing material should be biodegradable and biocompatible so that living cells can be directly involved inside the droplets during the printing process. The printer can also be programmed to place multiple printing materials and cell types in a specific pattern in order to pattern prevascular networks within larger constructs. 3. Laser-induced forward transfer (LIFT) Similar to the inkjet-based 3D printing, LIFT also forms tissue constructs by depositing droplets (Fig. 3F). There are three main components for the LIFT 3D printer including a laser source, a glass coated with a laser absorbing layer, and a printing material. To deposit the printing material onto the printing stage, the laser focuses on the desired point at the laser absorbing layer, then the focused region is heated and increases its volume to push a droplet of material onto the stage. The construct can be printed by repeatedly depositing droplets with desired patterns in a layer-by-layer fashion.

Challenges of 3D Printed Tissue Construct The acellular scaffolds formed by FDM and SLS are mostly used to provide tissue constructs with higher mechanical stiffness. On the other hand, the cell-laden hydrogel methods are used to provide the supporting cellular matrix and print living cells or drugs in a specific pattern. SLA, DLP-SLA, inkjet- and laser-based 3D printing methods are the common 3D printing processes used to make cell-laden hydrogels. To maintain cell viability, the 3D printing materials should form water-swollen hydrogel networks that remain insoluble after the 3D printing process. These hydrogels can be formed via photo-, chemical-, and thermalcrosslinking during 3D printing process. Although 3D printing enables researchers to create tissue constructs more easily, forming vascular networks within the engineered tissues is still challenging due to limited integration between different materials and cell types. A combinatory 3D printing platform for FDM and DLP-SLA can be used as one strategy to continuously print rigid, porous scaffolds supporting soft, biodegradable hydrogels. This hybrid printing platform can print different cell types in separate materials with different local stiffness at high resolution. Another challenge is the achievement of permeability throughout the large-scale tissue constructs and the design of perfusable channel systems that provide sufficient exchange of soluble factors such as oxygen, nutrients, and wastes throughout the entire area. The diffusion distance is generally restricted within 300 mm in the engineered tissue constructs; thus all embedded cells in a tissue construct thicker than 300 mm may not have immediate access to the mass exchanges, affecting cell survival and functional tissue formation. Thus, the engineered tissue should contain porous structures and perfusable channels for sufficient media supply. Approaches to designing perfusable channel systems use water-soluble sacrificial polymers such as carbohydrate glass, Pluronic F127, sugar, gelatin, polyvinyl alcohol (PVA), and so on. Using 3D printing techniques, the sacrificial materials can be printed alongside cell-laden hydrogels and dissolved upon infusion of aqueous media or temperature changes. The formed channels can provide effective perfusion of culture media and soluble growth factors that enhance cell viability and facilitating vessel formation. In addition, ECs that are perfused through the channels can attach on the surface of the hydrogels and migrate into the hydrogel construct. Different types of cells such as pericytes and fibroblasts can be perfused and cocultured with ECs or preformed vascular networks in the hydrogels. Another issue to be considered is 3D printing resolution for manufacturing the small-scale tissue constructs. To form microvascular networks, resolution of 3D printing should be as low as 100 um. With the different 3D printing techniques, there are various resolutions to form hydrogel-based materials depending on the materials, crosslinking methods, and types of tissue constructs. The printing resolution of photo-crosslinkable polymers is highly affected by the polymer and photoinitiator concentrations, and exposure times of light sources such as visible or UV lamps during the photo-crosslinking process. The chemicals used as photoinitiators are generally toxic, yet higher concentrations of the photoinitiator can enable printing at higher resolution. Thus, researchers need to balance in the tradeoff between the cytotoxicity and the resolution based on the purpose of the study and types of tissue constructs. The polymer concentration and exposure time also need to be considered regarding mechanical stiffness and material degradation, as higher polymer concentrations and longer exposure times produce stiffer and more slowly degrading hydrogel matrices, but they may delay formation of vascular networks in the engineered tissues.

Applications in Vascularization in Bone and Cardiac Tissue Engineering Strategies for engineering vascularized tissues may vary according to the application of the engineered implant. Functional vascularization enables the development of thicker tissue constructs and allows for longer testing periods. Applications for engineered vascularized tissues are numerous both in vitro and in vivo. In this section, we use bone and cardiac tissue grafts as two models for the prevascularization of engineered tissue constructs.

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Applications in Bone Tissue Engineering Challenges for engineered bone grafts More than half a million patients receive bone grafts annually in the United States. Autograft is the gold-standard procedure, consisting in harvesting bone tissues from one site in the patient and transplanting it to the site of need. However, one of the major drawbacks is the creation of a second defect at the harvesting site. Allograft, which consists in transplanting bone tissue from a donor, is the second most common technique. But it presents a reduced osteoinductivity and risks of immune reactions. Thus, there is a great need to improve the current clinical treatments for bone repair and regeneration using engineered functional bone grafts. In addition, bone tissues are highly vascularized and present a complex intraosseous vascular network that ensures tissue survival and plays a key role in the bone’s self-healing capability. But mimicking this vascular network is to date one of the major hurdles that delays the development of functional bone tissues. Given the potential impact of engineered bone grafts, numerous studies have been conducted for the development of vascularized bone tissues.

Strategies for Developing Vascularized Bone Grafts Upon implantation, the main functions of engineered bone grafts are to temporarily provide sufficient structural and mechanical support to adjacent host tissue and create microenvironments for bone cell development and tissue regeneration. Early studies have proposed the use of rigid scaffolds with porous structures in order to promote the growth of bone forming cells. The geometry of the scaffolds displays interconnected open pores, with a combination of macro- and micropores (respectively with sizes over 100 mm and under 20 mm). These porous structures enable the development of bone cells after implantation, but the graft porosity must be balanced with its mechanical properties. The porous scaffolds are mostly made from ceramics such as hydroxyapatite (HA) and tricalcium phosphate (TCP), and polymers such as polyglycolic acid (PGA), polylactic-glycolic acid (PLGA), and polycaprolactone (PCL) for biocompatibility and osteoinductivity. Applying these porous scaffolds alone to patients with bone defects may lead to lengthy times required for healing and regeneration after implantation, and is proven to be of limited interest for large-scale defects (length over 5 cm). In order to reduce the healing time, scaffolds need to be optimized for bone and blood vessel development. To date, the development of a fully functional vasculature inside the scaffolds is still an ongoing process. Many different strategies for in vitro or in vivo prevascularization have been developed to overcome the technical limitations. The goal of in vitro prevascularization is to improve porous engineered bone scaffolds in order to promote conditions amenable to rapid vascularization prior to implantation (Fig. 4). The prevascularized bone scaffolds will directly anastomose with host vessels, reducing the time for ECs to form vascular networks after implantation. Additive manufacturing (AM) techniques are great tools to advance prevascularized bone scaffolds. In bone tissue engineering, FDM is used to manufacture porous rigid scaffolds with reproducible microstructure, and syringe-based deposition or SLA is augmentatively used to deposit ECM-like hydrogels containing different growth factors that support EC sprouting and neovessel formation. Patient-specific bone grafts can be designed using imaging technologies in order to facilitate the implantation surgery and enhances the graft compatibility after insertion. Growth factors such as VEGF and bFGF and cells such as ECs and fibroblasts can be directly incorporated into the printing materials and deposited with a predefined pattern to accelerate vascular network formation. Inkjet- and laser-based methods enable researchers to print droplets of suitable microscopic size and control the deposit location with biologically relevant patterns. Alternatively, the growth factors can be encapsulated in particles such as fibers or microspheres that are embedded in the hydrogel constructs. The rigid scaffold can also be directly coated with polymer solutions containing growth factors that are released simultaneously or sequentially. Alternatively, in vivo prevascularization strategies consist in implanting the engineered scaffold into a region of the patient adjacent to blood vessels that are suitable for surgical transfer. The in vivo incubation period results in the formation of microvascular networks inside the scaffold. After several weeks, the construct can be harvested with the accompanying blood vessels as a free flap, (A)

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Engineered extracellular matrix Neovessel developed through in vivo or in vitro prevascularization Extracellular matrix with growth factors Fig. 4 Schematic drawing of engineering prevascularized bone grafts for large-size bone defect. (A) bone defect; (B) engineered bone scaffold; (C) strategies for constructing prevascularized bone graft.

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and implanted in the defect site. Although it requires two surgical procedures, this strategy can induce neovascularization with direct, instantaneous perfusion through the vascular axis after implantation. Different manufacturing methods have been explored to control the porosity of the scaffold for blood vessel growth. 3DP such as FDM or SLS has proven to be very effective because it allows accurate control of the pore size and an excellent reproducibility of microporous geometry. A summary of the different strategies mentioned above is presented in Fig. 4, as applied to large bone defect repair. More than simply providing a rigid porous scaffold, these recent bone grafting strategies aim to develop a vascular network inside the scaffold before implantation, in order to reduce the healing time after implantation and improve the chances of the grafting success.

Applications in Cardiac Tissue Engineering Challenges for engineered cardiac grafts Heart disease and strokes, which are the principal cardiovascular diseases, account for more than 40% of all deaths in the United States. Following myocardial infarction, macrophages remove dead myocytes as part of an inflammatory response, which after several weeks leads to the development of thick and stiff granulation tissue that then turns into scar tissue. Such tissues reduce the contractile ability of the heart and can potentially lead to heart failure. By helping to regenerate cardiac tissues, tissue engineering could hence play a key role in the recovery of patients after myocardial infarction. Heart walls are composed of a high density of cells such as myocytes and fibroblasts which experience a high metabolic demand because of their constant contractions. A dense vascular network is therefore necessary in the heart wall for oxygen supply. To be clinically relevant, an engineered cardiac patch should present a thickness of several millimeters. Such a thickness implies that the integration of a vascular network in the graft is essential for the cells to survive. Because of the beating nature of the heart, engineered cardiac grafts should present mechanical, electrical, and functional integration into the organ architecture, and therefore, the constructs should be contractile, electrophysiologically stable, mechanically robust yet flexible, and vascularized or at least quickly vascularized after implantation.

Strategies for developing vascularized cardiac tissues Hydrogels are typically used as the main material of cardiac patches, because they present adapted mechanical properties and biocompatibility. To manufacture such materials, many different processes are used. 3DP has proven to be very efficient, offering a high resolution and an important freedom of shape. Syringe-based extrusion, inkjet, and laser deposition can be used for cell and bioagent deposition. Similar to engineering bone grafts, three major factors can be integrated to guide the establishment of vascular networks including a preformed vascular structure, ECM-like hydrogels, and ECs. Variability in the integration of these components prior to implantation leads to the development of different approaches. One strategy is to create a scaffold with a preformed vascular structure integrated into the scaffold geometry. This strategy is to decellularize a vascularized biologic tissue in order to obtain a tissue construct with intact 3D geometry and vascular networks. This construct is then used as a scaffold for repopulation by myocytes and ECs, leading to the formation of a fully contractile cardiac tissue. Another approach aims at building the scaffold from synthetic materials. The geometry of the vascular networks can be designed using 3D printing technology to manufacture the cardiac grafts with improved freedom of shape and high resolution. Sacrificial materials such as carbohydrate glass, sugar, and gelatin are used to create a vascular template within the constructs and removed during a postprocessing phase. ECs can be incorporated inside the channels to facilitate vascular network formation. Coculture with fibroblasts can improve EC organization and myocyte proliferation. In vitro incubation can be performed in order to prevascularize the constructs to enable quick anastomosis of the microvessels and the host coronary artery. The main issue in this approach remains the ability to culture the three kinds of cells together in vitro. Thus, cell compatible manufacturing processes have to be employed. Indeed, constraints in terms of temperature, internal stress, and cytotoxicity have to be considered. Syringe-based or laser-based deposition can be adapted processes, as well as SLA using visible and UV light. Angiogenic growth factors can be incorporated into the cardiac grafts to promote vascularization. The concentration and diffusion rates of the growth factors are the key parameters to control, with the goal being to recreate levels as close as possible to biological values. Some techniques that can be employed to control those parameters include selecting different kinds of hydrogels with varying densities or degradation rates. Indeed, a proper use of growth factors can stimulate proliferation of fibroblasts and thereby significantly improve the development of nearby vascular networks. Fig. 5 summarizes an overview of the different approaches. Even though 3D printing presents advantages for manufacturing cardiac grafts, novel printing processes are still under investigation to increase the potential of this technology for engineering vascularized tissues. The development of vascularized cardiac grafts is an ongoing process, and the ideal strategy still remains to be found. It is still not entirely clear which of the components (vascular structure, cells, and growth factors) is needed early on, and to which extent the vascular network should be developed prior to implantation. Achieving reliable vascularization remains a challenge for cardiac grafts and, indeed, for nearly all large-scale tissue grafts to date. Recent manufacturing technologies such as 3D printing can construct previously inaccessible implant designs and architectures, which will both advance our comprehension of the underlying biological mechanisms required for tissue regeneration and permit the fabrication of more appropriate biomimetic tissue implants in the coming years.

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Myocytes and fibroblasts Endothelial cells forming neovessels through a prevascularization step Scaffold with embedded growth factors Fig. 5 Schematic drawing of engineering prevascularized cardiac patch. (A) heart with scar tissues after myocardiac infarction; (B) engineered cardiac patch; (C) strategies for constructing prevascularized cardiac patch.

Summary Numerous prevascularization strategies have been studied to form functional vascular networks within an engineered tissue construct in vitro under controlled conditions prior to implantation. However, it is difficult but necessary to balance all the conditions, including material selection, alignment, and organization of ECs within an in vitro system, localized chemotactic cues, optimal stiffness and degradation of the materials, appropriate perfusion systems, and stabilization of the vasculature with coculture techniques. These conditions can be incorporated into 3D bioprinting technologies for higher precision and speed. Therefore, it is important for researchers to understand the interrelationship between all the conditions to design appropriate experimental in vitro and in vivo approaches for the creation of the desirable vascularized tissue constructs and to further the successful translation of research into clinical settings.

Further Reading Auger, F. A., Gibot, L., & Lacroix, D. (2013). The pivotal role of vascularization in tissue engineering. Annual Review of Biomedical Engineering, 15, 177–200. Bose, S., Roy, M., & Bandyopadhyay, A. (2012). Recent advances in bone tissue engineering scaffolds. Trends in Biotechnology, 30(10), 546–554. Davis, G. E., Bayless, K. J., & Mavila, A. (2002). Molecular basis of endothelial cell morphogenesis in three-dimensional extracellular matrices. The Anatomical Record, 268(3), 252–275. Davis, G. E., Stratman, A. N., Sacharidou, A., & Koh, W. (2011). Molecular basis for endothelial lumen formation and tubulogenesis during vasculogenesis and angiogenic sprouting. International Review of Cell and Molecular Biology, 288, 101–165. Elomaa, L., & Yang, Y. (2017). Additive manufacturing of vascular grafts and vascularized tissue constructs. Tissue Engineering, Part B: Reviews. https://doi.org/10.1089/ ten.teb.2016.0348. Lovett, M., Lee, K., Edwards, A., & Kaplan, D. L. (2009). Vascularization strategies for tissue engineering. Tissue Engineering, Part B: Reviews, 15(3), 353–370. Lutolf, M. P., & Hubbell, J. A. (2005). Synthetic biomaterials as instructive extracellular microenvironments for morphogenesis in tissue engineering. Nature Biotechnology, 23, 47–55. Mistry, A. S., & Mikos, A. G. (2005). Tissue engineering strategies for bone regeneration. Regenerative medicine II. Heidelberg: Springer, 1–22. Murphy, S. V., & Atala, A. (2014). 3D bioprinting of tissues and organs. Nature Biotechnology, 32(8), 773–785. Novosel, E. C., Kleinhans, C., & Kluger, P. J. (2011). Vascularization is the key challenge in tissue engineering. Advanced Drug Delivery Reviews, 63(4), 300–311. Patan, S. (2000). Vasculogenesis and angiogenesis as mechanisms of vascular network formation, growth and remodeling. Journal of Neuro-Oncology, 50(1), 1–15. Ribatti, D., Vacca, A., Nico, B., Roncali, L., & Dammacco, F. (2001). Postnatal vasculogenesis. Mechanisms of Development, 100(2), 157–163. Shanjani, Y., Pan, C. C., Elomaa, L., & Yang, Y. (2015). A novel bioprinting method and system for forming hybrid tissue engineering constructs. Biofabrication, 7(4), 1–16. Vunjak-Novakovic, G., Tandon, N., Godier, A., Maidhof, R., Marsano, A., Martens, T. P., & Radisic, M. (2009). Challenges in cardiac tissue engineering. Tissue Engineering, Part B: Reviews, 16(2), 169–187. Ucuzian, A. A., & Greisler, H. P. (2007). In vitro models of angiogenesis. World Journal of Surgery, 31(4), 654–663.

Wound Healing and the Host Response in Regenerative Engineering Daniel Chester, Ethan A Marrow, Michael A Daniele, and Ashley C Brown, North Carolina State University and University of North Carolina at Chapel Hill, Raleigh, NC, United States; and North Carolina State University, Raleigh, NC, United States © 2019 Elsevier Inc. All rights reserved.

Overview of the Wound Healing Process The Four Phases of Wound Healing Hemostatic Phase Inflammatory Phase Proliferation Phase Remodeling Phase Host Responses in Wound Healing Abnormal Wound Healing Host Responses to Biomaterials Biomaterials Used in Wound Healing Natural Biomaterials Synthetic Biomaterials Engineered Growth Factors/Growth Factor Delivery Systems Emerging Materials Conclusions Further Reading

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Glossary Biomaterials Any material that has been engineered specifically to interact with biological systems in order to treat a disorder, augment a natural process, or repair/ replace a tissue or function of the body. Chronic Nonhealing Wounds A wound that shows no improvement after four weeks or does not heal completely in eight weeks. Decellularization The process by which the extracellular matrix (ECM) of a tissue is isolated from its inhabiting cells resulting in a scaffold of the original tissue. Fibrosis The accumulation of excess fibrous scar tissue at the site of an injury. Fibrous Encapsulation The layer of fibrous connective tissue that forms around an implant sequestering it from the surrounding tissue. Foreign Body Response The host’s response to a foreign material occurring at the end stage of the inflammatory process involving macrophages, foreign body giant cells, and the fibrous encapsulation of the material. Integration The full acceptance of a biomaterial into native tissue involving the migration and proliferation of cells into the biomaterial, with minimal scarring. Natural Biomaterials Biomaterials derived from natural sources and used for their similar structures and composition to native tissue. Synthetic Biomaterials Biomaterials manufactured with specific mechanical and chemical properties to match their intended use in the body. Wound Healing The four-stage process responsible for maintaining the homeostatic structure and function of tissues following injury.

Abbreviations ECM Extracellular matrix EGF Epithelial growth factor FACIT Fibril-associated collagens with interrupted triple helices FBGC Foreign body giant cells FGF Fibroblast growth factor MMP Matrix metalloproteinase NO Nitric oxide PDGF Platelet-derived growth factor

Encyclopedia of Biomedical Engineering, Volume 1

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PlGF Placenta growth factor TGF-a Transforming growth factor alpha TGF-b Transforming growth factor beta TNF-a Tumor necrosis factor alpha VEGF Vascular endothelial growth factor

Overview of the Wound Healing Process The wound healing process is responsible for maintaining and reestablishing the homeostatic structure, function, and properties of tissues. It is a highly orchestrated, complex process modulated by biochemical signals, such as cytokines, and biophysical stimuli generated from cell–cell or cell–extra cellular matrix (ECM) interactions. Important cell types involved in the wound healing process include platelets, leukocytes, fibroblasts, keratinocytes, and endothelial cells. Cellular responses are modulated by key cytokines including transforming growth factor beta (TGF-b), platelet-derived growth factor (PDGF), tumor necrosis factor alpha (TNF-a), vascular endothelial growth factor (VEGF), and fibroblast growth factor (FGF). Cellular responses can also be modulated by the ECM components fibrin/fibrinogen, fibronectin, and collagen. Together, cytokines and ECM components help to regulate the inflammatory response and control cell proliferation, migration, differentiation, and ECM production. The ideal result of the wound healing process is the reestablishment of homeostasis through the replacement of the damaged tissue with new tissue that is structurally and mechanically similar to the native tissue. This makes the wound healing process a natural regenerative process. As such, understanding how tissue is naturally regenerated will allow for the development of strategies or materials to mimic or greatly augment the natural regeneration process. This junction of understanding the wound healing process and developing new techniques or materials to control tissue repair embodies the concept of regenerative engineering. The end result of the wound healing process can be influenced by the host’s response to the wound environment; therefore, many regenerative engineering strategies investigate ways to precisely control the host’s response in order to prevent abnormal or deficient tissue repair.

The Four Phases of Wound Healing The wound healing process can be separated into four temporarily overlapping phases, the hemostatic, inflammatory, proliferation, and remodeling phases, which vary in time-scale from several hours to a few weeks. Each phase involves the coordination of specific cytokines and ECM interactions in order to direct cellular responses. The end result of the wound healing process, when it proceeds normally, is new tissue with structural and mechanical properties similar to the native tissue. A schematic of the wound healing process can be seen in Fig. 1 and each phase is also detailed in subsequent sections.

Hemostatic Phase The hemostatic phase begins immediately following tissue damage with the initiation of the coagulation cascade. The coagulation cascade results in the accumulation of platelets at the wound site and the formation of a fibrin clot which stems blood flow in order to limit the amount of blood lost. Platelets bind to the exposed collagen found at the wound site, thereby activating the platelets and amplifying the recruitment of additional platelets. Activated platelets have a higher affinity for fibrinogen; therefore platelet activation leads to fibrinogen accumulation at the wound site catalyzing the formation of the fibrin clot. Fibrinogen is then cleaved by a-thrombin and spontaneously forms insoluble fibrin fibers increasing the stability of the provisional fibrin clot. Thrombin also activates the transglutaminase enzyme factor XIIIa, which crosslinks the fibrin fibers overtime to further increase clot stability. Factor XIIIa also allows for the covalent incorporation of soluble fibronectin into the provisional fibrin matrix. The resulting matrix of the fibrin clot then becomes the scaffold on which the remaining phases of the wound repair process proceed. Along with acting as the scaffold of the wound healing process, fibrin, fibrinogen, and fibronectin also play an important role in modulating the wound healing process. Fibrin, fibrinogen, and fibronectin are capable of interacting with platelets, leukocytes, fibroblasts, epithelial cells, and keratinocytes as well as binding with FGF, PDGF, TGF-b, VEGF, and TNF-a. This makes the provisional matrix proteins an important site for cellular interactions as well as a reservoir for growth factors that can provide additional cues to direct cell fate throughout the wound healing process. As the formation of the fibrin clot progresses, the bound, activated platelets will begin to release PDGF and TGF-b. PDGF acts as a chemotaxis agent and will recruit additional platelets to the site of injury along with neutrophils, macrophages, smooth muscle cells, and fibroblasts. The accumulation and incorporation of platelets into the fibrin clot is a key component in determining the extent of fibrin clot collapse. Clot collapse occurs due to the contraction of activated platelets incorporated throughout the fibrin clot and is an important factor in determining the stability of the resulting clot. TGF-b also aides in the chemotaxis of macrophages,

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Fig. 1 The four stages of wound healing: The wound healing process is separated into four temporarily overlapping phases, the hemostatic (A), inflammatory (B), proliferation (C), and remodeling phase (D). (A) The hemostatic phase begins immediately following injury and is responsible for the creation of a fibrin clot in order to stem blood flow and limit the amount of blood lost. Platelets bind to exposed collagen and begin secreting PDGF and TGF-b in order to recruit additional platelets, macrophages, and cells to the injury site. (B) Approximately 24 h following injury the inflammatory phase begins with the accumulation of neutrophils and macrophages at the wound site. The primary purpose of neutrophils and macrophages is the removal of any damaged tissue or foreign materials in the wound site. (C) Two to 3 days after injury, the proliferation phase begins. The proliferation phase is characterized by the accumulation and proliferation of fibroblasts at the site of injury and the production of fibrous extracellular matrix. (D) The final stage of the wound healing process is the remodeling phase where fibroblasts will begin to remodel the newly deposited collagen matrix into newly healed tissue. This phase can last for months or years, but when concluded results in the formation of new tissue with similar structure and mechanics to the native tissue.

fibroblasts, and smooth muscle cells to the wound site along with controlling several other important processes that occur in the subsequent stages of the wound repair process.

Inflammatory Phase Approximately 24 h after the initiation of the hemostatic phase, neutrophils begin to accumulate at the wound site due to PDGF secreted by activated platelets within the fibrin clot. The accumulation of neutrophils into the wound site marks the beginning of the next phase of the wound healing processdthe inflammatory phase. The primary role of the neutrophil is the removal of foreign materials or cell debris that resulted from the injury through phagocytosis. Phagocytosis is the process by which neutrophils will engulf foreign material until they no longer have room for other debris. Neutrophils also secrete VEGF which is an important growth factor in angiogenesis. Approximately 48 h after injury, macrophages accumulate at the wound site due to chemotaxis from the PDGF and TGF-b secreted by the activated platelets in the fibrin clot. Upon entering the wound site, macrophages begin to phagocytose any remaining debris that was too large for the neutrophils and any completely filled neutrophils. Macrophages are also responsible for secreting additional PDGF, FGF, TNF-a, and TGF-b which leads to an increase in the number of macrophages at the wound site and attracts fibroblasts, endothelial cells, and smooth muscle cells to the wound site. Macrophages additionally produce several reactive radicals including nitric oxide (NO), oxygen, and peroxide which are important antimicrobial agents. NO can also be produced, although to a lesser extent, by fibroblasts, keratinocytes, and endothelial cells and plays an important role in the later stages of wound healing. 5 days after injury, T-lymphocytes and B-lymphocytes begin to migrate into the wound site marking the end of the inflammatory phase. T-lymphocytes are believed to play a role in downregulating the inflammatory process and in mediating cell proliferation and the degree of collagen crosslinking. The role of B-lymphocytes still remains largely unknown and more research is needed in order to fully elucidate their role in the wound healing process.

Proliferation Phase The inflammatory phase starts to wane 2–3 days after injury at which point the proliferation phase begins. During the proliferation stage, cells will migrate into the fibrin provisional matrix and begin to produce a new ECM. Fibroblasts are recruited into the wound site by FGF and TGF-b secreted by the macrophages and activated platelets. TGF-b is also responsible for increasing the rate at which new ECM is produced by upregulating the transcription of several genes that promote the synthesis of collagen, proteoglycans, and

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fibronectin. At the same time, TGF-b is also responsible for the downregulation of the matrix metalloproteinase (MMP) family of enzymes that are responsible for the degradation of ECM proteins. Epithelialization also begins as epithelial cells begin to infiltrate into the wound site due to chemotaxis controlled by epidermal growth factor (EGF) and transforming growth factor a (TGF-a) which are produced by platelets and macrophages. The NO produced by macrophages, fibroblasts, keratinocytes, and endothelial cells also plays an important role at this stage of the wound healing process. NO has been shown to control the rate of epithelialization and the rate of collagen syntheses; the inhibition of NO slows the epithelialization process and decreases the rate of collagen synthesis resulting in slower wound contraction and lower clot strength. A process called fibroplasia begins during the proliferation phase of the wound healing process. Fibroplasia starts approximately 5 days after the initial injury and is the process by which fibroblasts begin to produce the collagen matrix backbone found in normal tissue that eventually replaces the provisional fibrin matrix. There are three different classes of collagen that can be deposited during this process. These classes are fibrillar type I, II, or III collagen, nonfibrillar type IV collagen, or fibril-associated collagens with interrupted triple helices (FACIT) type VI and VII collagen. Each class of collagen plays a specific role in the formation of new ECM with fibrillar collagen forming rigid linear networks, nonfibrillar collagen forming complex networks of collagen that are usually found in basement membranes, and FACIT collagen providing matrix stability by integrating with fibrillar collagen and other ECM proteins. Each class of collagen is produced at specific times in the wound healing process. At the onset of fibroplasia, fibroblasts rapidly produce the fibrillar type I and III collagen with peak production occurring between 7 and 14 days following injury. The initial deposition of type I and III collagen is important for increasing the tensile strength of newly formed tissue with the collagen fibers becoming linear strands aligned parallel to the forces felt in the tissue. FACIT type VI collagen peaks one to two weeks after injury and is believed to play an important role in integrating the new collagen matrix with the native neovasculature. The rate of collagen deposition is also self-regulated by the interactions that occur between keratinocytes and fibroblasts and the matrix to which they are attached. For example, keratinocytes will increase the amount of collagenase produced when in direct contact with collagen fibers. The increased amount of collagenase will, in turn, increase the rate of collagen degradation which ensures that the amount of collagen in the ECM is kept at a homeostatic level. With fibroblasts, an ECM made primarily of collagen will result in lower proliferation levels as well as lower collagen deposition rates. The lower proliferation and deposition rates occur due to a negative feedback mechanism initiated by the binding of the a1b1 integrin to collagen. The process of angiogenesis occurs concurrently to fibroplasia. Angiogenesis is the complex process by which new blood vessels form that rely on the proper ECM composition as well as cytokine signaling. FGF, VEGF, and TGF-b are important growth factors involved in angiogenesis; fibrin has also been shown to induce angiogenesis directly. While fibroblasts are beginning to deposit collagen into the fibrin provisional matrix, endothelial cells are stimulated by the FGF and VEGF released by platelets, macrophages, and neutrophils to locally dissolve the basement membrane. PDGF, VEGF, and FGF also increase endothelial cells’ ability to revascularize tissue allowing for the formation of new blood vessels in the basement membrane. NO also plays a role in angiogenesis as NO promotes vasodilation which improves local blood flow to the wound site allowing for the faster recruitment of cells, platelets, neutrophils, and macrophages. The newly formed blood vessels are stabilized by smooth muscle cells which are recruited to the wound site by PDGF that is released by platelets and endothelial cells. Toward the end of angiogenesis, TGF-b secreted by endothelial cells acts as in inhibitory agent impeding vascular proliferation bringing angiogenesis to an end.

Remodeling Phase The final phase of the wound healing process, the remodeling phase, occurs concurrently with granulation tissue formation. The primary purpose of the remodeling phase is the formation of new epithelium and scar tissue and this process can take up to a year or longer to complete. Throughout the remodeling phase, fibroblasts degrade and realign collagen fibers. As fibroblasts realign and deposit collagen fibers, the fibril packing density in each fiber is seen to increase which leads to collagen fibers with larger diameters and higher mechanical properties. The realignment of collagen fibers also transforms the initial, unorganized collagen matrix into a highly organized collagen matrix whose structure closely mimics that of the native tissue. The realigned collagen matrix will have the potential to regain up to 80% of the native tissue’s tensile strength, but the original strength of the tissue will never fully be achieved. How close the new collagen matrix can match the tensile properties of the native tissue depends on the location and size of the wound as well as the duration of the repair process. As the remodeling phase continues, the number of macrophages and fibroblasts decrease due to apoptosis and angiogenesis ends. At the end of the remodeling phase, new tissue with a high tensile strength and a minimal number of cells and vascularization is formed.

Host Responses in Wound Healing There are many different factors that can impact the wound healing process. Systemically, factors such as age, nutrition, and the presence of diseases such as diabetes can alter the wound healing process. Age does not immediately impair the wound healing process, but it has been observed that delayed wound healing is common in older organisms and is believed to be associated

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with an altered inflammatory response. Malnutrition can lead to impaired capillary formation, fibroblast proliferation, and collagen synthesis due to lack in the required proteins or energy needed to perform those specific functions. Diseases that lead to hypoxia, such as diabetes, and diseases that lead to altered immune responses, such as HIV or lupus, have been linked to impaired wound healing and chronic wounds. Locally, factors such as foreign bodies, ischemia, and infection can change wound healing responses. The composition of the foreign body will greatly affect the wound healing response by changing the length and severity of the inflammatory response. Oxygen is extremely important in the wound healing process so ischemia can also drastically change wound healing responses. Ischemia can lead to impaired angiogenesis, fibroblast and keratinocyte proliferation, and reepithelialization all of which will impede wound healing. Infection can lead to a prolonged immune response causing an increased amount of MMPs at the wound site causing excess degradation of ECM.

Abnormal Wound Healing As outlined in previous sections, wound healing is a highly regulated process that requires the fine control of both mechanical and chemical cues in order for normal healing to occur. Any irregularities in the control of the mechanical and/or chemical cues involved can lead to abnormal wound healing in the form of excessive scar tissue formation or deficient matrix production. Excess ECM production is known as fibrosis and has been connected to many life-threatening medical conditions, while deficient ECM is seen in chronic nonhealing wounds which are most readily seen in ulcers. Fig. 2 shows the relationship between the amount of ECM deposition and the subsequent effect it has on fibroblast cells as they try to repopulate the wound site. If left untreated, both fibrosis and chronic nonhealing ulcers can lead to detrimental results, and sometimes death, for the affected organism. Fibrosis is characterized as the overproduction of connective tissue and ECM resulting in excess scar tissue formation that can lead to the disruption of the structure, and in extreme cases the function, of the native tissue or organs. Clinical examples of fibrotic related conditions include keloids, hypertrophic scars, Crohn’s disease, pulmonary fibrosis, and atrial fibrosis. The underlying mechanisms behind many fibrosis-related conditions remain poorly understood and are only just beginning to be elucidated. For example, it has been found that fibroblasts isolated from keloids have increased proliferation rates, produce 2–3 times more ECM, and express higher levels of VEGF, PDGF, and TGF-b when compared to normal fibroblasts. Additionally, a high density of mast cells has been observed at the site of excess scar tissue deposition found in many fibrotic conditions. Mast cells contain enzymes responsible for processing procollagen and it is believed that in fibrotic conditions abnormally high amounts of histamine and renin are produced that upregulate collagen synthesis leading to excess ECM deposition and scar tissue formation. Nonhealing wounds occur when an abnormally low amount of connective tissue or ECM is produced in the wound bed. It has been found that chronic ulcers have reduced levels of PDGF, FGF, EGF, and TGF-b. Chronic ulcers also have increased protease activity which is believed to result from an overexpression of MMPs caused by increased neutrophil infiltration rates and leads to the increased degradation of ECM. Clinical examples of nonhealing wounds include venous ulcers, diabetic ulcers, or pressure ulcers. Venous ulcers occur in the legs and are thought to be caused from venous hypertension leading to ischemia. Diabetic ulcers are believed to stem from neuropathy which inhibits the perception of pain leading to unnoticed injuries that either get infected or are continuously reinjured. Pressure ulcers are normally found in people suffering from conditions such as paralysis. Pressure ulcers arise from the ischemia that occurs when the pressure on the tissue is greater than the pressure in the capillaries.

Amount of ECM Deposition and Cell Morphology in Normal and Abnormal Wound Healing

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Fig. 2 ECM deposition and cell morphology in normal and abnormal wound healing: key characteristics of abnormal wound healing are the amount of ECM deposition and the resulting cell morphologies. In chronic nonhealing wounds, there is an abnormally low amount of ECM found at the injury site. This results in a wound bed with low mechanical properties leading to senescent fibroblasts. Conversely, in fibrosis there is an overproduction of ECM resulting in a wound environment that has much higher mechanical properties than normal tissue and highly activated fibroblasts. Under normal wound healing conditions with normal amounts of ECM deposition, fibroblasts will exhibit appropriate amounts of cell spreading and balanced migration/contraction responses.

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Host Responses to Biomaterials There are many factors that determine how well a biomaterial performs its intended function once implanted in vivo. Such factors include the biomaterial’s composition, mechanical and material properties, surface topography, size, and placement in the body. However, the most important factor that determines the success of the implanted material is the resulting host response that begins immediately following implantation. The host’s response to the biomaterial is a summation of the tissue damage received during implantation and the response caused by the material itself. The response caused by the tissue damage resolves quickly as part of the normal healing process, but the foreign body response caused by the material will last as long as the material is present. The properties of the material implanted dictate the success of the material and whether or not integration or the fibrous encapsulation of the biomaterial will occur. Integration is the ideal response for most implanted biomaterials and involves the incorporation of the biomaterial into the surrounding tissue. In order for integration to occur, a biomaterial should have the proper surface properties to promote the adhesion of the appropriate proteins and cells. Generally, a combination of high porosity and surface roughness or a hydrogel coating is used to increase cell attachment to a biomaterial. An overview of design considerations to promote material integration is shown in Fig. 3. Successful integration of the biomaterial into the surrounding tissue is characterized by cellular ingrowth and ECM deposition along with generating little to no immune response from the host. In practice, complete integration of the biomaterial into the tissue rarely occurs and it is more likely for fibrous encapsulation to occur to some extent. Generally, the end result of the wound healing process in the presence of a biomaterial is fibrous encapsulation. The foreign body response leading to the fibrous encapsulation of the biomaterial is shown in Fig. 4. Fibrous encapsulation is the formation of scar tissue on the surface of the biomaterial produced by myofibroblasts and fibrocytes and occurs as a response to frustrated phagocytosis. Frustrated phagocytosis is the process by which macrophages try to dissolve a foreign body that is too large to internalize. In order to remove the large foreign body, macrophages will join together to form foreign body giant cells (FBGCs) and begin to release superoxides and free radicals that lower the pH of the local environment resulting in damage to the biomaterial and surrounding tissue. Fibrous encapsulation will then occur in order to sequester the biomaterial from the rest of the body and end frustrated phagocytosis. However, if frustrated phagocytosis persists, a chronic nonhealing wound will begin to form at the site of the implant and lead to the ultimate failure of the implant. The scar tissue formed as a result of frustrated phagocytosis acts as a barrier between the biomaterial and the wound site severely limiting the integration of the biomaterial. Fibrous encapsulation also limits the performance of chemical biosensors, electrical leads/ electrodes, therapeutic delivery systems, and orthopedic/cardiovascular prostheses. Research into the acute inflammatory

Properties for Promoting the Integration of a Biomaterial Young’s Modulus

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Fig. 3 Properties for promoting the integration of a biomaterial: the material properties, surface properties, generated immune response, and incorporation of bioactive agents should all be considered when choosing a biomaterial to be used for implantation. The closer the biomaterials properties are to that of the native tissue, the more likely the biomaterial will be incorporated into the native tissue. Properties such as the Young’s modulus, viscoelasticity, fatigue strength, creep, and porosity should all be considered when creating or choosing a biomaterial for a specific application. It is highly unlikely that all of the biomaterial properties will be identical to native tissue so, therefore, it is important to prioritize which properties are the most critical for the specific application. Surface coatings can also be used to increase the rate of cellular attachment to a biomaterial. Plasma coating, ceramic coatings, hydrogel coatings, and ECM proteins can all be used to increase cellular attachment rates. Controlling the immune response generated by the biomaterial is also important to the successful integration of the biomaterial. With a low immune response, the foreign body process will be limited, thereby enhancing the integration of the material. Finally, bioactive agents can be incorporated into the biomaterial in order to direct and control cellular responses. Antibacterial/antimicrobial drugs can be used to prevent infections and lower immune responses, cells can be encapsulated into the material in order to provide naturally derived cues to promote cellular interactions with the biomaterial, and growth factors can also be incorporated in order to direct various cellular responses.

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Biomaterial Fig. 4 Host foreign body response to implanted biomaterials: the host response to a biomaterial begins with protein adsorption onto the surface of that material and the immune response. Due to the large size of biomaterials, neutrophils are unable to phagocytose the implant, begin to undergo frustrated phagocytosis, and recruit monocytes which then differentiate into macrophages at the implant. Due to the proteins adsorbed to the surface, macrophages begin to adhere to the surface of the implant while continuing to perform frustrated phagocytosis. The adhered macrophages will then join together to form foreign body giant cells (FBGCs) which help to recruit fibroblasts to the biomaterial. The recruited fibroblasts will then begin to synthesize new ECM on the surface of the implant resulting in the fibrous encapsulation of the implant.

responses to implanted biomaterials has shown that the adsorption of fibrinogen onto the surface results in integrin-mediated leukocyte recruitment and adhesion. The adhered leukocytes will then secrete growth factors that recruit other cell types leading to fibrous tissue deposition. However, how chronic inflammation mediates fibrous encapsulation and the mechanisms behind it remain poorly understood.

Biomaterials Used in Wound Healing Biomaterials can be classified into two main groups: synthetic or natural biomaterials. Synthetic biomaterials consist of metals, ceramics, and polymers while natural biomaterials consist of protein-based biomaterials, polysaccharide-based biomaterials, and decellularized-based biomaterials. Both types of biomaterials can be characterized by their surface properties, mechanical properties, chemical properties, biological properties, and implantation lifetime resulting in many similarities and differences between both types of materials. A key difference between synthetic and natural biomaterials is seen in their applications. A division of synthetic biomaterials consists of metal alloys and ceramics which have Young’s moduli on a similar order of magnitude as bone. This makes metals and alloys perfect candidates for bone or joint replacements, but not for soft tissues and/or wound healing applications. As such, metals and ceramics are outside of the scope of this book chapter. On the other hand, synthetic polymers have diverse and tunable properties making them ideal for soft tissue and wound healing applications and will be discussed below. Since natural biomaterials are derived from tissues and organs, they are already prime candidates for uses in wound healing applications due to their similar structure and composition to native tissue. However, naturally derived biomaterials can have limitations in certain applications due to lack of mechanical robustness and immune rejection adding additional concerns. The decision of whether to use a synthetic polymer or natural biomaterial for each wound healing application is based on the properties of the specific biomaterial and the properties that are required to elicit the desired biological response. To this end, natural biomaterials are more advantageous to use in applications where the structure of the implant needs to be similar to the environment. Another advantage that natural biomaterials afford is their built-in bioactivity due to being composed of similar macromolecules as native ECM allowing for interactions between the material and cells. Examples of naturally derived biomaterials include silk, chitin, ECM components, and decellularized ECM. Additionally, naturally derived decellularized ECM materials also have growth factors bound to them naturally which can be released to promote specific cellular processes as cells infiltrate and remodel the scaffold. However, a predominant limitation of many naturally derived biomaterials is that they usually have poor

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mechanical properties. Furthermore, naturally derived biomaterials can have a high degree of variability in material properties from batch to batch. Also, since naturally derived biomaterials have such a similar structure and composition to native tissue, these types of biomaterials are more likely to cause an immune response which may or may not be desirable depending on the specific application. The main advantages for using synthetic polymers in wound healing applications stem from the ability to control their mechanical and chemical properties. Since synthetic polymers are fabricated in controlled settings, it is possible to change the materials composition during its synthesis such that the desired mechanical or chemical properties are obtained. This also results in relatively low batch-to-batch variability compared to natural biomaterials. However, the manufactured structures of the synthetic polymer scaffolds often lack the complexity required to accurately mimic native tissue resulting in reduced levels of cellular infiltration and bioactivity. Another disadvantage of synthetic polymers is that their biocompatibility is unknown and has to be tested. Since synthetic polymers are man-made and not normally found in nature, how the host will respond to the material is a serious consideration that needs to be thoroughly tested in order to ensure the success of the material.

Natural Biomaterials As their name suggests, naturally derived biomaterials are commonly found in nature as part of the tissues and organs of various organisms. Naturally derived biomaterials are commonly used to repair or replace damaged tissue or organs since they already contain the required structure and biological composition needed to mimic the native tissue. Because they are naturally derived, these materials have high degrees of biocompatibility, biodegradability, and are subject to high degrees of remodeling by resident cells after implantation; all these features further allow for the integration of the biomaterial into the native tissue. Commonly used examples natural biomaterials that are used in augmenting the wound healing process include silk, chitin, collagen, fibrin, and various decellularized ECM components. Silk is a natural fiber composed of hydrophobic fibroin and hydrophilic sericin. Sericin is composed of 25%–30% of the silk protein and it envelopes the fibroin fibers gluing them together. Most of the sericin needs to be removed before implantation in the body since the combination of the two proteins has been seen to cause allergic responses. Raw silk can be harvested from several natural sources such as silkworms and spiders making it an abundantly available material. In its processed form, silk has a high tensile strength, is biocompatible and biodegradable, and does not generate a significant immune response making it an ideal candidate for scaffolds and wound dressings. Silk can also be electrospun into nanofibers that have a high specific surface area and improved thermal and electrical properties compared to nonelectrospun silk which further increases the number of applications for which it can be used. Clinically, silk is used in sutures for ligament replacement, nonwoven mats for use as wound dressings, and as porous sponge scaffolds for healing bone defects. Chitin is one of the most abundant biopolymers found on Earth. Chitin is a complex polysaccharide of N-acetyl-glucosamine and glucosamine, and is predominately found in the exoskeletons of crustaceans and the scales of fish. The degree of deacetylation and the molecular weight of the polymer can severely impact the material properties and the resulting performance of the chitin polymer. Chitin is used in wound healing dressings due to its antimicrobial properties and the ability for N-acetyl-glucosamine to accelerate the rate of tissue repair and prevent the formation of scars. The original use of chitin was as a powder, but more recently it has been incorporated into films, membranes, and woven/nonwoven dressings. Wound dressing made with chitin was found to adhere well to wound sites as well as allow for the permeability of oxygen into the wound site. One of the most abundant ECM components in the body is collagen and it has been widely applied to many different in vitro and in vivo applications. Clinically, collagen is used because of its nonimmunogenicity and its ability to provide the structural support needed for tissue regeneration. Collagen is capable of being prepared as a crosslinked solid or into un-crosslinked gels allowing for versatility in its application. Collagen’s main use as a biomaterial is in the treatment of burns in the form of a wound dressing or as a bone filling material. Collagen wound dressings have been shown to improve the spreading and growth of chondrocytes, fibroblasts, and keratinocytes while collagen bone fillings have been shown to increase bone regeneration rates. Hyaluronic acid is another naturally derived ECM component that is commonly used in biomaterials. It is a linear polysaccharide of glucuronic acid and N-acetyl glucosamine-glucuronic acid disaccharides. Hyaluronic acid is mainly used as a biodegradable hydrogel used to deliver drugs or growth factors into the wound environment and has been used to increase cell adhesion and proliferation rates onto scaffolds. Finally, fibrinogen, a soluble plasma glycoprotein that plays an important role in blood coagulation, is also commonly used as a biomaterial. The ability for fibrinogen to promote cell adhesion and migration, along with its biocompatibility and nonimmunogenicity make it popular for use in tissue scaffolds and wound dressings. Fibrinogen can be electrospun to form porous, fibrous scaffolds and when combined with thrombin, produce a biodegradable mesh. Fibrin gels are also used clinically with the most popular application being fibrin glues which are formulated by combining fibrinogen and thrombin directly at the surgical site in high concentrations. Fibrin-based glues function by reproducing the fibrin clot found during the normal wound healing process. These types of glues are mainly used as hemostatic agents for bleeding surfaces. The microstructure of a fibrin hydrogel is shown in Fig. 5. Gelatin-based glues are another type of naturally derived tissue glues that are commonly used clinically. Gelatin is obtained by controlling the hydrolysis of collagen extracted from animal tissues resulting in a mixture of polypeptides dispersed based off of size and chemical reactivity. The properties of gelatin are largely based off the tissue and animal it was extracted from and the method of hydrolysis. The collagen content of gelatin glues result in a resorbable material with a greater bonding strength than fibrin glues

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Fig. 5 Network structure of a natural biomaterial gel: a cryo-scanning electron microscopy (SEM) image at 2000 showing the structure of a 3D fibrin gel composed of 2 mg/mL fibrinogen and 0.1 U/mL thrombin. Image courtesy of Ms. Seema Nandi. Authors acknowledge Dr. Elaine Zhou and the Analytical Instrumentation Facility at North Carolina State University for their assistance with sample preparation and acquisition.

making them useful for sealing larger tissues. Tissue glues can be used instead of sutures and have been used to treat burns and adhere skin grafts. Naturally derived ECM can also be obtained from the decellularization of various tissues and organs. Decellularization aims to remove all of the cells from their ECM while keeping the ECM intact. With the removal of the cells, a large number of the antigens that would normally cause an immune response are removed, resulting in a scaffold that has the mechanics and structure of native tissue with a limited immune response. While the decellularization process does remove cells, many of the growth factors bound naturally to the ECM remain intact. The bound growth factors can then be used to direct specific cellular responses as they are released from the ECM by cell-mediated interactions. Common tissues that are decellularized and used as ECM scaffolds commercially include the small intestine submucosa, the urinary bladder matrix, human/porcine/bovine skin, and horse/bovine/porcine pericardium. These matrices have been used to facilitate the regeneration of skin, tendons, and ligaments. Whole organs are also being investigated for use as scaffolds and current research on efficiently decellularizing hearts, lungs, and livers for use as organ replacements is ongoing. Entirely decellularized organs would allow for the exact replication of the organ’s highly complex architecture which is critical for the function of the organ and would normally be impossible to replicate. The newly decellularized organs would then be reseeded with cells from the patient and result in a fully functional organ that is an exact genetic match to the patient. This would then limit the immune response from the patient to the newly implanted organ and increase the safety, efficacy, and lifetime of implanted organs.

Synthetic Biomaterials Unlike natural biomaterials, synthetic biomaterials are not found in nature and can only be manufactured. This makes synthetic materials easy to obtain and available in large quantities. The manufacturing process also allows for the mechanical and chemical properties to be tailored to the specific application. It is also possible to incorporate drugs during the manufacturing process of synthetic materials which adds another degree of functionality of these types of materials. Also, synthetic biomaterials can be coated, functionalized, or conjugated with different proteins or antibodies to limit the immunogenic response and increase the amount of bioactivity. For wound healing applications, synthetic polymers are commonly used since it is possible to make scaffolds with material properties that closely match native tissue. Examples of synthetic polymers used clinically are synthetic skin substitutes, tissue glues, hydrogels, and semisynthetic chitosan. Skin substitutes are artificial skin replacements that provide a protective barrier when placed over burns or other chronic wounds in a similar manner to normal skin. The primary objective of skin substitutes is to facilitate the repair, regeneration, and restoration of the functional properties of skin following a traumatic injury such as second or third degree burns. Properties that make a successful skin substitute include water vapor transmission similar to normal skin, minimal inflammatory response, adherence to the wound site, controlled degradation, impermeable to bacteria, and appropriate mechanical properties. Commercially available skin substitutes are made of ultrathin silicone, nylon, and petrolatum gauze. Skin substitutes can be temporary or permanent with the temporary substitutes acting purely as a biodegradable barrier throughout the wound healing process and the permanent skin substitutes incorporating fibroblasts or epithelial cells along with a treated dermis layer from a human cadaver. Tissue glues are used to replace traditional sutures and staples in order to seal a wound following damage or a surgical procedure. There are three different types of tissue glues commercially available. These are fibrin-based glues, gelatin-based glues, and

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cyanoacrylate-based glues. Fibrin- and gelatin-based glues have been described above, but a synthetic alternative to these naturally derived glues are cyanoacrylate-based glues. These types of glues are limited in their use clinically due to their varying degrees of cytotoxicity. However, they are still used to close dermal wound sites that are still actively bleeding since fibrin- and gelatinbased glues are not as effective on wounds that are actively bleeding. Hydrogels are polymer networks that are insoluble in water and swell to an equilibrium volume while retaining their shape. Hydrogels have structural similarities to the ECM, have the ability to retain a large quantity of water, are biocompatible, have a low interfacial tension, and cause minimal frictional or mechanical irritation to the implantation site making them a popular biomaterial for use in regenerative medicine. The hydrophilicity of the polymer network stems from the presence of chemical residues such as hydroxylic (eOH) and carboxylic (eCOOH) functional groups found along the polymer backbone. The hydrogel remains insoluble to its 3D network and the balance between the dispersive forces acting on the hydrated chain and the cohesive forces that prevent the further penetration of water. Common polymers that are used in the formation of hydrogel are polyethylene oxide, polypropylene oxide, and poly(N-isopropyacrylamide) just to name a few. The properties of hydrogels make them an attractive material to use in the application of implantable biomaterials. Since the structure of hydrogels closely mimics that of the ECM, hydrogels are often used as a coating to promote cell growth and attachment. For example, hyaluronate hydrogels have been used as a coating on an inorganic bone material in order to increase cell attachment. Hydrogels can also be chemically crosslinked in order to tune their properties to fit the specific application. Chemically crosslinked hydrogels offer scaffolds with greater mechanical properties and slower degradation times which make them ideal for scaffolds designed to last in the body for longer periods of times. Chitosan is a semisynthetic biomaterial made from the deacetylated form of chitin, which is found naturally. The deacetylation process makes it a cationic glycosaminoglycan biopolymer with a degree of deacetylation ranging from 30% to 95% and a molecular weight of 250 kDa to greater than 1000 kDa. The solubility, viscosity, and biocompatibility are proportional to the degree of deacetylation while the bioadsorption of chitosan is inversely proportional to molecular weight. Chitosan has been extensively investigated for its potential use in the field of regenerative medicine. One of the most promising applications of chitosan is the development of dressings to control and promote wound healing. It has been demonstrated that simple application of chitosanbased materials with varying degrees of deacetylation to a wound can direct healing and regeneration processes. Accordingly, chitosan and chitosan-derived materials have exhibited promising biocompatibility, mucoadhesive properties, and a broad spectrum of antimicrobial activity. To date, chitosan has been used to bind, deliver, or promote the presence of biomacromolecules and growth factors implicated in wound healing. Chitosan-based materials have been developed to deliver heparin, hyaluronic acid, and other growth factors to the site of the wound. These materials are designed to have a short lifespan, hydrating on contact, and then dissolving slowly while releasing their payload. This has been utilized to create stable basic FGF containing films that resulted in rapid wound healing in diabetic mice. Chitosan has also been covalently modified with the basic FGF and bone morphogenic protein, both of which showed increased wound healing and osteogenesis, respectively. Similarly, chitosan-based materials have been used to deliver whole cells to treat full thickness wounds. Lastly, chitosan and chitosan-based materials are used for their broad spectrum antimicrobial activity. The charge interactions between chitosan and microbial cell membrane components lead to dysregulation of the microbial transport mechanism and ultimately the death of the microbe. Several studies have shown that the bactericidal and bacteriostatic capabilities of chitosan depend on molecular weight and degree of deacetylation.

Engineered Growth Factors/Growth Factor Delivery Systems Aside from biomaterials, growth factors and growth factor delivery systems have recently begun to be engineered for wound healing applications. In a recent study by Martino et al. a method for producing engineered growth factors with a super affinity to the ECM is described. In this article, it was found that a domain in placenta growth factor-2 (PlGF-2) binds to the ECM with a high affinity. This domain was then fused with VEGF-A, PDGF-BB, and BMP-2 in order to create a new, engineered growth factor with a significantly higher binding affinity to the ECM than normal growth factors. This newly engineered growth factor was also seen to augment the wound healing process. When introduced into mouse models of chronic wounds and bone defects, it was seen that wound healing was greatly enhanced when compared to the normal growth factors. Controlling the delivery and release of growth factors is also critically important for augmenting the wound healing process. To that end, there are various strategies that can be employed to make a biomaterial act more like natural ECM. For example, biophysical properties such as the materials density, porosity, charge, and hydrophobicity can all be changed in order to create a material with the desired release kinetics. However, those types of biophysical changes are often not ideal for cellular attachment, growth, and remodeling. Another approach is to slow the release of growth factors by functionalizing the material with growth factor binding sites isolated from ECM molecules. For example, it has been seen that heparin sulfate proteoglycans are capable of binding with several different growth factors. As such, biomaterials have been created with modified heparin or heparin sulfate mimetic molecules to bind and slowly release growth factors such as FGF. Other ECM proteins such a fibronectin, fibrinogen, tenascin C, and vitronectin also have growth factor binding domains that can be isolated and used to bind a wide variety of growth factors to biomaterial matrices. Besides attaching growth factors to the matrix via binding sites, growth factors can also be built into the material. For example, growth factors can be covalently crosslinked into fibrin matrices by creating a recombinant growth factor containing a sequence

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recognized by factor XIIIa. As factor XIIIa naturally polymerizes the fibrin matrix, the recombinant growth factor will also be incorporated. The release of the incorporated growth factor is then determined by the degradation rate of the fibrin matrix. Growth factors can also be incorporated into hyaluronic acid hydrogels by loading the growth factor into the hydrogel prior to creating the scaffold.

Emerging Materials New biomaterials for augmenting the wound repair process are constantly being developed. One of the more recently emerging materials for use as a tissue scaffold are wound-interfacing microgels. Microgels are crosslinked polymeric particles that are composed of water-soluble/swellable polymer chains classifying these types of materials as hydrogels. These types of materials offer unique advantages compared to other types of hydrogels including: tunable size from a range of nanometers to micrometers, a large surface area for bioconjugation, increased reaction time to temperature and pH, and tunable mechanical properties. For example, microgels films created from chemically crosslinked poly(allylamine hydrochloride) were deposited onto surgical sutures in a layerby-later fashion and were shown to have the ability to release ibuprofen into the wound site while also accelerating the wound healing process. Microgels have also been used to create microporous scaffolds that can be injected into wound sites in order to promote cell migration and tissue regeneration. Biomimetic microgels, which mimic the fibrin binding ability and clot retraction features of natural platelets, have also been shown to enhance clotting in vivo and enhance clot stability. Collectively, these recent studies demonstrate that microgels are a useful material strategy for modulation wound responses.

Conclusions Overall, the wound healing process is a highly ordered, highly controlled process that is modulated by cellular interactions with ECM components and cytokines. Fibrin, fibrinogen, and collagen provide the scaffold through which the wound healing process occurs on and modulate cellular adhesion, migration, proliferation, and ECM deposition/remodeling. Cytokines such as TGF-b, PDGF, TNF-a, VEGF, and FGF are also responsible for myriad of important cellular responses including cellular recruitment through chemotaxis, proliferation, migration, and reepithelialization. However, an abnormal amount of ECM components or growth factors can lead to abnormal wound healing as seen as either fibrosis or chronic nonhealing wounds resulting in life-threatening conditions for the organism. The presence of a biomaterial can also cause a change in the wound healing process based off the severity of the immune response of the host. A severe immune response can lead to frustrated phagocytosis resulting in either the start of a chronic nonhealing wound if left uncontrolled or the fibrous encapsulation of the material, both of which are undesirable. Increased rates of the integration of biomaterials can be influenced by controlling the material properties, the surface properties, the resulting immune response, and by the addition of bioactive agents such as growth factors. To this end, both naturally derived biomaterials and synthetically derived biomaterials have advantages and disadvantages when it comes to controlling properties that influence integration. There are also many types of natural biomaterials and synthetic biomaterials that have a myriad of uses clinically. Newly emerging materials focus on bridging the gap between natural and synthetic biomaterials by creating a synthesized material that has highly tunable material properties that accurately mimic biological properties while providing a degree of biomimicry to increase the degree of the resulting material/wound interface. In summation, biomaterials are important clinical tools for the control and augmentation of the wound repair process and with further control on material and biological properties, the better and more substantial the host responses will be.

Further Reading Alrubaiy, L., & Al-Rubaiy, K. K. (2009). Skin substitutes: A brief review of types and clinical applications. Oman Medical Journal, 24, 4–6. Anderson, J. M., Rodriguez, A., & Chang, D. T. (2008). Foreign body reaction to biomaterials. Seminars in Immunology, 20, 86–100. Badylak, S. F. (2015). Host response to biomaterials: The impact of host response on biomaterial selection (1st edn.). Cambridge, MA: Academic Press. Barrientos, S., Stojadinovic, O., Golinko, M. S., Brem, H., & Tomic-Canic, M. (2008). Growth factors and cytokines in wound healing. Wound Repair and Regeneration, 5, 585–601. Boateng, J. S., Matthews, K. H., Stevens, H. N. E., & Eccleston, G. M. (2007). Wound healing dressings and drug delivery systems: A review. Journal of Pharmaceutical Sciences, 97, 2892–2923. Chester, D., & Brown, A. C. (2016). The role of biophysical properties of provisional matrix proteins in wound repair. Matrix Biology, 60–61, 124–140. Clark, R. (2013). The molecular and cellular biology of wound repair (2nd edn.). New York, NY: Springer Science. Crapo, P. M., Gilbert, T. W., & Badylak, S. F. (2011). An overview of tissue and whole organ decellularization processes. Biomaterials, 32, 3233–3243. Elena, I. P., Bazaka, K., & Crawford, R. J. (2013). New functional biomaterials for medicine and healthcare (1st edn.). Cambridge: Elsevier. Gilbert, T. W., Sellaro, T. L., & Badylak, S. F. (2006). Decellularization of tissues and organs. Biomaterials, 27(19), 3675–3683. Griffin, D. R., Weaver, W. M., Scumpia, P. O., Di Carlo, D., & Segura, T. (2015). Accelerated wound healing by injectable microporous gel scaffolds assembled from annealed building blocks. Nature Materials, 14, 737–744. Ha, T. L. B., Quan, T. M., Vu, D. N., & Si, D. M. (2013). Naturally derived biomaterials: Preparation and application. In J. A. Andrades (Ed.), Regenerative Medicine and Tissue Engineering. Croatia: InTech. Halim, A. S., Khoo, T. L., & Mohd. Yussof, S. J. (2010). Biologic and synthetic skin substitutes: An overview. Indian Journal of Plastic Surgery, 43, S23–S28. Hunt, T. K., Hopf, H., & Hussain, Z. (2000). Physiology of wound healing. Advances in Skin & Wound Care, 13, 6–11. Kim, J. K., Kim, H. J., Chung, J., et al. (2014). Natural and synthetic biomaterials for controlled drug delivery. Archives of Pharmacal Research, 37, 60–68.

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Lutolf, M. P., & Hubbell, J. A. (2008). Advances in Tissue Engineering 255–278. London: World Scientific Publishing Co. Martino, M. M., Briquez, P. S., Guc, E., et al. (2014). Growth factors engineered for super-affinity to the extracellular matrix enhance tissue healing. Science, 343, 885–888. Mazza, G., Rombouts, K., Hall, A. R., et al. (2015). Decellularized human liver as a natural 3D-scaffold for liver bioengineering and transplantation. Scientific Reports, 5, 13079. Mogosanu, G. D., & Grumezescu, A. M. (2014). Natural and synthetic polymers for wounds and burns dressing. International Journal of Pharmaceutics, 2, 127–136. Morais, J. M., Papadimitrakopoulos, F., & Burgess, D. J. (2010). Biomaterials/tissue interactions: Possible solutions to overcome foreign body response. The AAPS Journal, 2, 188–196. Mutsaers, S. E., Bishop, J. E., McGrouther, G., & Laurent, G. J. (1997). Mechanisms of tissue repair: From wound healing to fibrosis. The International Journal of Biochemistry & Cell Biology, 29, 5–17. Nguyen, H., Qian, J. J., Bhatnagar, R. S., & Li, S. (2003). Enhanced cell attachment and osteoblastic activity by P-15 peptide-coated matrix in hydrogels. Biochemical and Biophysical Research Communications, 311, 179–186. Roach, P., Eglin, D., Rohde, K., & Perry, C. C. (2007). Modern biomaterials: A reviewdbulk properties and implications of surface modifications. Journal of Materials Science. Materials in Medicine, 7, 1263–1277. Tapias, L. F., & Ott, H. C. (2014). Decellularized scaffolds as a platform for bioengineered organs. Current Opinion in Organ Transplant, 19, 145–152. Velnar, T., Bailey, T., & Smrkolj, V. (2009). The wound healing process: An overview of the cellular and molecular mechanisms. Journal of International Medical Research, 37, 1528–1542.

ENCYCLOPEDIA OF BIOMEDICAL ENGINEERING

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ENCYCLOPEDIA OF BIOMEDICAL ENGINEERING EDITOR IN CHIEF

Roger Narayan University of North Carolina at Chapel Hill, Chapel Hill, NC, United States

VOLUME 2

Section Editors Christian Hellmich TU Wien, Vienna University of Technology, Vienna, Austria

Diego Mantovani Laval University, Quebec City, QC, Canada

Alexander Wong University of Waterloo, Waterloo, ON, Canada

William Z Rymer Rehabilitation Institute of Chicago, Chicago, IL, United States.

Levi Hargrove Rehabilitation Institute of Chicago, Chicago, IL, United States

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Elsevier Radarweg 29, PO Box 211, 1000 AE Amsterdam, Netherlands The Boulevard, Langford Lane, Kidlington, Oxford OX5 1GB, United Kingdom 50 Hampshire Street, 5th Floor, Cambridge, MA 02139, United States Copyright Ó 2019 Elsevier Inc. All rights reserved. No part of this publication may be reproduced or transmitted in any form or by any means, electronic or mechanical, including photocopying, recording, or any information storage and retrieval system, without permission in writing from the publisher. Details on how to seek permission, further information about the Publisher’s permissions policies and our arrangements with organizations such as the Copyright Clearance Center and the Copyright Licensing Agency, can be found at our website: www.elsevier.com/permissions. This book and the individual contributions contained in it are protected under copyright by the Publisher (other than as may be noted herein). Notice Knowledge and best practice in this field are constantly changing. As new research and experience broaden our understanding, changes in research methods, professional practices, or medical treatment may become necessary. Practitioners and researchers may always rely on their own experience and knowledge in evaluating and using any information, methods, compounds, or experiments described herein. In using such information or methods they should be mindful of their own safety and the safety of others, including parties for whom they have a professional responsibility. To the fullest extent of the law, neither the Publisher nor the authors, contributors, or editors, assume any liability for any injury and/or damage to persons or property as a matter of products liability, negligence or otherwise, or from any use or operation of any methods, products, instructions, or ideas contained in the material herein. Library of Congress Cataloging-in-Publication Data A catalog record for this book is available from the Library of Congress British Library Cataloguing-in-Publication Data A catalogue record for this book is available from the British Library ISBN 978-0-12-804829-0

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Publisher: Oliver Walter Acquisition Editor: Blerina Osmanaj Publishing Services Manager: Beckie Brand Associate Content Project Manager: Kshitija Iyer Cover Designer: Greg Harris Printed and bound in the United States

EDITORIAL BOARD

EDITOR IN CHIEF Roger Narayan University of North Carolina at Chapel Hill, Chapel Hill, NC, United States

SECTION EDITORS

Levi Hargrove Rehabilitation Institute of Chicago, Chicago, IL, United States

Christian Hellmich TU Wien, Vienna University of Technology, Vienna, Austria

Sri Krishnan Ryerson University, Toronto, ON, Canada

Cato Laurencin University of Connecticut Health Center, Farmington, CT, United States

Diego Mantovani Laval University, Quebec City, QC, Canada

William Z Rymer Rehabilitation Institute of Chicago, Chicago, IL, United States

Pankaj Vadgama Queen Mary University of London, London, United Kingdom

Min Wang The University of Hong Kong, Pokfulam, Hong Kong

Alexander Wong University of Waterloo, Waterloo, ON, Canada

Xiaojun Yu Stevens Institute of Technology, Hoboken, NJ, United States

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EDITOR IN CHIEF Roger Narayan Dr. Roger Narayan is a professor in the Joint Department of Biomedical Engineering at the University of North Carolina and North Carolina State University. He is an author of over 200 publications as well as several book chapters on processing, characterization, and modeling of biological and biomedical materials. Dr. Narayan has edited several books, including Biomedical Materials, Printed Biomaterials, Computer Aided Biomanufacturing, Diamond-Based Materials for Biomedical Applications, Medical Biosensors for Point of Care (POC) Applications, Monitoring and Evaluation of Biomaterials and their Performance In Vivo, Nanobiomaterials: Nanostructured Materials for Biomedical Applications, and the ASM Handbook on Materials for Medical Devices. He has previously served as chair of the Functional Materials Division of The Minerals, Metals & Materials Society and is currently chair-elect of the Bioceramics Division of American Ceramics Society. Dr. Narayan has received several honors for his research activities, including the North Carolina State University Alcoa Foundation Engineering Research Achievement Award, the North Carolina State University Sigma Xi Faculty Research Award, the University of North Carolina Jefferson-Pilot Fellowship in Academic Medicine, the National Science Faculty Early Career Development Award, the Office of Naval Research Young Investigator Award, the American Ceramic Society Richard M. Fulrath Award, the Royal Academy of Engineering Distinguished Visiting Fellowship, and TMS Brimacombe Medal. He has served as Fulbright Scholar at the University of Otago, the National Polytechnic Institute (Mexico City), and the University of Sao Paulo. He has been elected as Fellow of ASM International, the American Association for the Advancement of Science, the American Ceramic Society, and the American Institute for Medical and Biological Engineering.

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SECTION EDITORS

Levi Hargrove Dr. Hargrove is currently the Director of the Center for Bionic Medicine and of the Neural Engineering for Prosthetic and Orthotics Laboratory at the Shirley Ryan AbilityLab. He is also an Associate Professor in the Departments of Physical Medicine and Rehabilitation and the McCormick School of Engineering at Northwestern University. A major goal of his research is to develop clinically realizable myoelectric control systems that can be made available to persons with limb loss in the near future. His research addresses all levels of amputation and has been published in the Journal of the American Medical Association and the New England Journal of Medicine, and multiple patents. Key projects include the development of advanced and adaptive control systems for prosthetic legs, improving control of robotic hand prostheses, and intramuscular EMG signal processing. In 2012, Dr. Hargrove cofounded Coapt, a company to transition advanced rehabilitation technologies from the research laboratory to patients’ homes.

Christian Hellmich Dr. Christian Hellmich, Full Professor at the Department of Civil Engineering of the Vienna University of Technology (TU Wien), is the director of the Institute for Mechanics of Materials and Structures. At TU Wien, he received his engineering, Ph.D., and habilitation degrees (in 1995, 1999, and 2004, respectively). From 2000 to 2002, he was a Max Kade Postdoctoral Fellow in the Department of Civil and Environmental Engineering at the Massachusetts Institute of Technology. His work is strongly focused on well-validated material and (micro)structural models, in terms of theoretical foundations and applications to concrete, soil, rock, wood, bone, and biomedical implants, up the structural level (tunnels, pipelines, bridges, biological organs such as the skeleton)dwith complementary experimental activities if necessary. He has led several projects for the tunnel, railway, and pipeline industries, as well as international research activities sponsored by the European Commission, including the coordination of the mixed industry-academia consortium “BIO-CT-EXPLOIT” at the crossroads of numerical simulation and computer tomography, or the cross-domain COST action NAMABIO integrating engineers, physicists, (stem) cell biologists, and medical doctors across the European continent and beyond. He has published more than 130 papers in international refereed scientific journals in the fields of engineering mechanics, materials science, and theoretical biology, more than 20 book chapters, and more than 120 papers in refereed conference proceedings. Dr. Hellmich has served as the Chairman of both the Properties of Materials Committee of the Engineering Mechanics Division of the American Society of Civil Engineers (ASCE), and the Poromechanics and Biomechanics Committees of the Engineering Mechanics Institute (EMI), as associate editor of the Journal of Engineering Mechanics (ASCE), and as Coeditor in Chief of the Journal of Nanomechanics and Micromechanics (ASCE). As community service, he has (co)chaired and/or supported more than 50 international conferences (including chairmanship of the 2013 Biot Conference on Poromechanics and the 2015 CONCREEP conference; both EMI-ASCE supported), and he has reviewed for 128 different scientific journals and 15 science foundations. He was awarded the Kardinal Innitzer Science Award of the Archbishopry of Vienna in 2004 (for his habilitation thesis), the Science Award of the State of Lower Austria in 2005 (for his achievements in the micromechanics of hierarchical composites), and he was the recipient of the 2008 Zienkiewicz Award for Young Scientists in Computational Engineering Sciences, sponsored by the European Community on Computational Methods in Applied Sciences (ECCOMAS). For further activities in the multiscale poromicromechanics of bone materials, he received one of the highly prestigious ERC Grants of the

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European Research Council in 2010; and he was elected member of the Young Academy of the Austrian Academy of Sciences in 2011. In 2012, he was rewarded the prestigious Walter L. Huber Research Prize of the ASCE, for his contributions to the microporomechanics of hierarchical geomaterials and biomaterials; he was elected Fellow of EMI in 2014 and was corecipient of the 2017 Kajal Mallick Memorial Award of the Institution of Civil Engineers (United Kingdom).

Sri Krishnan Sridhar (Sri) Krishnan received B.E. degree in Electronics and Communication Engineering from the College of Engineering, Guindy, Anna University, Chennai, India, in 1993, and M.Sc. and Ph.D. degrees (with student fellowship from Alberta Heritage Foundation for Medical Research) in Electrical and Computer Engineering from The University of Calgary, Calgary, Alberta, Canada, in 1996 and 1999, respectively. Sri Krishnan joined Ryerson University in July 1999 and is currently a Professor in the Department of Electrical and Computer Engineering. Since July 2011, he is an Associate Dean (Research, Development and External Partnerships) for the Faculty of Engineering and Architectural Science. He is also the Codirector of the Institute for Biomedical Engineering, Science and Technology (iBEST) and an affiliate scientist at the Keenan Research Centre in St. Michael’s Hospital, Toronto. Since January 2002 Sri Krishnan held various administrative leadership positions in the Department of Electrical and Computer Engineering and the Faculty of Engineering and Architectural Science. In 2010–2011, Sri Krishnan held Visiting Appointments in University of Rennes 1 (France), Grenoble Institute of Technology (France) and Indian Institute of Technology (Madras). Sri Krishnan is a registered professional engineer in the Province of Ontario and is a senior member of IEEE (EMBS and SP societies). He was the Founding Chair (2005–2015) of IEEE Signal Processing Society, Toronto Section and Region 7 (Canada), and a Founding Member of the IEEE Engineering in Medicine and Biology Society, Toronto Section. He currently serves as a Technical Committee Member (Biomedical Signal Processing) of IEEE EMBS. Sri Krishnan held the Canada Research Chair position (2007–2017) in Biomedical Signal Analysis. Sri Krishnan has successfully supervised/trained 10 postdoc fellows, 10 Ph.D., 30 Masters (thesis), 9 Masters (project), 42 RAs, and 20 Visiting RAs. Sri Krishnan’s research interests include adaptive signal representations and analysis and their applications in biomedicine, multimedia (audio), and biometrics. He has published 295 papers in refereed journals and conferences, filed 10 invention disclosures, and has one US patent. He has presented keynote/plenary/invited talks in more than 35 international conferences and workshops. Sri Krishnan also serves as a reviewer, committee member, and chair for many international conferences, journals, and granting bodies. Sri Krishnan’s academic interests include (interdisciplinary) curriculum design, experiential learning, and innovation. Sri Krishnan serves in the advisory boards of research institutes, innovation centers, incubator zones, and business organizations. Sri Krishnan is a recipient/awarded Outstanding Canadian Biomedical Engineer Award 2016; Certificate of Appreciation from PEO York Chapter 2016; Fellow of Canadian Academy of Engineering in 2014; 2014 Exemplary Service Award from IEEE Toronto Section; 2014 Certificate of Merit from IEEE Signal Processing Society; 2013 Achievement in Innovation Award from Innovate Calgary; 2011 Sarwan Sahota Distinguished Scholar Award; 2011 Certificate of Appreciation from IEEE Signal Processing Society; 2010 Shastri Visiting Professorship; 2010 French Embassy Visiting Researcher; 2008 Ontario Research Innovation Award from Biodiscovery Toronto; 2007 Canadian Engineers’ Young Engineer Achievement Award from Engineers’ Canada; 2006 New Pioneers Award in Science and Technology; 2006 South Asian Community Achiever Award; 2006 IEEE Toronto Section Best Chapter Chair Award; 2005 IEEE AESS Best Chapter Chair Award; 2005 IEEE Certificate of Appreciation from Six Societies; Six Best Research Paper Awards coauthored with his graduate students in International Conferences; and 2005 FEAS Research Excellence Award.

Cato Laurencin Cato T. Laurencin, M.D., Ph.D. is the University Professor at UCONN. He is the eighth designated in UCONN’s history. He is Professor of Chemical Engineering, Professor of Materials Science and Engineering, and Professor of Biomedical Engineering, and the Van Dusen Distinguished Endowed Professor of Orthopaedic Surgery. He directs the Institute for Regenerative Engineering and the Raymond and Beverly Sackler Center at the University of Connecticut. Dr. Laurencin earned his B.S.E. degree in Chemical Engineering from Princeton University. He earned his Ph.D. in Biochemical Engineering/Biotechnology from the Massachusetts Institute of Technology where he was named a Hugh Hampton Young Fellow. At the same time, he earned his M.D., Magna Cum Laude from the Harvard Medical School where he received the Robinson Award for Surgery. Dr. Laurencin is an expert in biomaterials, nanotechnology, stem cell science, and, the new field he has pioneered, Regenerative Engineering. He is a fellow of American Institute of Chemical

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Engineers and was named one of the 100 Engineers of the Modern Era by the AICHE. He received the Percy Julian Medal from National Organization of Black Chemists and Chemical Engineers, and the Pierre Galletti Award from the American Institute of Medical and Biological Engineering. He has received the NIH Director’s Pioneer Award and the National Science Foundation Emerging Frontiers in Research and Innovation Award for his research in Regenerative Engineering. Dr. Laurencin is an elected member of the National Academy of Engineering, the National Academy of Medicine, the Indian National Academy of Engineering, the Indian National Academy of Sciences, and the African Academy of Sciences. He is an academician and foreign member of the Chinese Academy of Engineering. Dr. Laurencin has two awards named in his honor. The W. Montague Cobb Institute and the National Medical Association established the Cato T. Laurencin Lifetime Research Achievement Award, while the Society for Biomaterials established The Cato T. Laurencin, M.D., Ph.D. Travel Fellowship Award. Dr. Laurencin received the Presidential Faculty Fellow Award from President Bill Clinton and the Presidential Award for Excellence in Science, Mathematics, and Engineering Mentoring from President Barack Obama. He is the recipient of the National Medal of Technology and Innovation, America’s highest award for technological achievement from President Barack Obama in ceremonies at the White House.

Diego Mantovani, Ph.D., FBSE. Prof. Diego Mantovani is the director of Laboratory for Biomaterials and Bioengineering at Laval University, in Canada, and senior scientist of the Regenerative Medicine Division of the Quebec University Hospital Research Centre. He received his doctoral degree jointly from University of Technology of Compiègne, France, and Laval University in 1999 and his joint Diploma in Engineering from Politecnico di Milano and the University of Technology of Compiegne, France, in 1993. After an industrial postdoc (1999), he becomes professor at Laval University School of Science and Engineering in 2000. Since the beginning he established is Laboratory at the University Hospital Research Center in Quebec City. Within his team, researches focus on surface modifications by plasma, thin polymer functional films, cell–materials interactions, degradable metals, scaffolds, and bioreactors for the replacement and regeneration of cardiovascular tissue. He has authored more than 260 original articles, holds 5 patents, and presented more than 185 keynotes, invited and seminar lectures worldwide. His H-index is 43 (June 2018), and his works were cited more than 7000 times. He was President of the Canadian Society for Biomaterials (2008–2009), and Executive Cochair of the World Biomaterials Congress in 2016 in Montreal, Canada. In 2012, he was elected Fellow of the World Biomaterials Science and Engineering Society. Since 2012, he is the holder of the Canada Research Chair 1 in Biomaterials and Bioengineering for the Innovation in Surgery. He was member of ad hoc panels at FDA, ISO, and Health Canada and member of a number of funding, regulatory and scientific committees worldwide. He is Adjunct Professor at Politecnico di Milano and Universita del Piemonte Orientale in Italy, as well as at the Vellore Institute of Technology, in India. He was invited professor in several universities worldwide, including Campinas, Brasil (2012–2015), Bologna (2015), Bordeaux (2014), Siao Tong West, China (2012), Cergy-Pontoise (2012), ParisTech (2011), Buenos Aires (2010), Namur, Belgium (2008), Tor Vergata, Italy (2007), Ankara, Turkey (2006), and others. He is member of the editorial board of five scientific journals in the field and of the advisory board of three medical devices consortia worldwide.

William Z Rymer Professor William Z Rymer is Professor of Physical Medicine and Rehabilitation and Physiology at the Rehabilitation Institute of Chicago, Chicago, IL, United States. His focus of work includes pathophysiology, stroke, spinal cord injury, spinal circuits, biomedical engineering, and neural signal processing.

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Pankaj Vadgama Pankaj Vadgama qualified in Degree in Medicine and then in Chemistry at the University of Newcastle upon Tyne, United Kingdom, with a First Class Honors BSc. He is a chemical pathologist, becoming a Fellow of the Royal College of Pathology. He completed his Ph.D. on medical biosensors as an MRC Fellow at Newcastle, and while there, he was made Director of the Biosensors Group and later appointed as Professor of Clinical Biochemistry at the University of Manchester, subsequently becoming Research Dean for the Faculty of Medicine. He was appointed Director of the Interdisciplinary Research Centre in Biomedical Materials at Queen Mary, University of London and was, until recently, Head of the Department of Clinical Biochemistry, Barts Health NHS Trust. His main interests are variously biosensors, applied bioelectrochemistry, point-of-care testing, and membrane technology. He has published over 200 papers. He is also Fellow of the Royal Society of Medicine, Institute of Physics, Royal Society of Chemistry, the Institute of Materials Minerals and Mining, and the Royal Society of Biology. He was given the Foundation Award of the Association of Clinical Biochemistry and Laboratory Medicine, has been a Sandoz Lecturer of the British Geriatric Society, and delivered the Latner lecture at the University of Newcastle. He has served on various UK Research Council grants award committees and is at present member of the Institute of Materials Minerals and Mining Smart Materials and Nano Committees and the Biomedical Materials Application Division. He sits on various BSI committees and was Chair of the ISO subpanel on nanomedicine nomenclature. He sits on various editorial boards and is Editor in Chief of Bioelectrochemistry. He is Deputy Chair of the Council for the Frontiers of Science based in Uganda directed at research training in East Africa.

Min Wang Min Wang is a Full Professor at The University of Hong Kong (HKU), and as Programme Director (2013–2018), he has led HKU’s Medical Engineering Programme (which is retitled to “Biomedical Engineering Programme” in 2018). He has worked in universities in the United Kingdom (1991–1997), Singapore (1997–2002), and Hong Kong (2002–Present) and has been a Guest Professor or Adjunct Professor of several universities in mainland China (Shanghai Jiao Tong University, Zhejiang University, Tianjin University, Southwest Jiao tong University, etc.). He was awarded BSc (1985) and Ph.D. (1991), both in Materials Science and Engineering, by Shanghai Jiao Tong University and University of London, respectively. He is a chartered engineer (CEng, 1995; UK) and chartered scientist (CSci, 2005; UK). He is an elected fellow of professional societies in the United Kingdom, Hong Kong, United States, and internationally (FIMMM, 2001; FIMechE, 2007; FHKIE, 2010; FBSE, 2011; FAIMBE, 2012; WAC Academician, 2013). Since 1991, he has been conducting research in biomaterials and tissue engineering and developing new biomaterials using the composite/hybridization approach. He was a founding member of UK’s Interdisciplinary Research Centre (IRC) in Biomedical Materials at the University of London. His biomaterials research has covered metals, polymers, ceramics, and composites and includes surface modification of materials or scaffolds. In recent years, he has focused on nanobiomaterials, electrospinning, and 3D printing. He and his research staff/students have won many awards at international conferences. He has authored a large number of research papers as well as many book chapters. His research has been widely cited by other researchers around the world. He has given many conference presentations, including more than 150 invited talks at international conferences. He has also given more than 110 seminars in universities, research institutes, and hospitals in Europe, North America, Asia, and Australia. He has been Chairman/Organizer of many conferences and has served in committees of more than 70 international conferences. He is the Founding Series Editor of Springer Series in Biomaterials Science and Engineering books and has been Editor, Associate Editor, or member of the Editorial Board of 20 international, printed journals, including International Materials Reviews, Composites Science and Technology, Surface and Coatings Technology, Journal of Materials Science: Materials in Medicine, and Journal of the Royal Society Interface. He has acted as a referee for more than 110 international journals in the fields of materials science and engineering, biomaterials and tissue engineering, physics, chemistry, medicine, dentistry, medical devices, biofabrication, nanoscience, nanotechnology, and 3D printing. He has been active in professional society activities and has served in various roles in these societies. He was Chairman of the Biomedical Division of Hong Kong Institution of Engineers (HKIE). He serves/has served in the Nomination Committee of World Academy of Ceramics (WAC) and the ICF-BSE Steering Committee of the International College of Fellows of the International Union of Societies for Biomaterials Science and Engineering (IUS-BSE). He has been an elected Council Member of Chinese Society for Biomaterials, Hong Kong Institution of Engineers, Asian Biomaterials Federation, World Association for Chinese Biomedical Engineers (WACBE), and Administrative Council of International Federation for Medical and Biological Engineering (IFMBE). (http://web.hku.hk/memwang/).

Section Editors

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Alexander Wong Alexander Wong, P.Eng., is currently the Canada Research Chair in Artificial Intelligence and Medical Imaging, Codirector of the Vision and Image Processing Research Group, and an Associate Professor in the Department of Systems Design Engineering at the University of Waterloo. He had previously received the B.A.Sc. degree in Computer Engineering from the University of Waterloo, Waterloo, ON, Canada, in 2005, the M.A.Sc. degree in Electrical and Computer Engineering from the University of Waterloo, Waterloo, ON, Canada, in 2007, and Ph.D. degree in Systems Design Engineering from the University of Waterloo, ON, Canada, in 2010. He was also an NSERC postdoctoral research fellow at Sunnybrook Health Sciences Centre. He has published over 400 refereed journal and conference papers, as well as patents, in various fields such as computational imaging, artificial intelligence, computer vision, and medical imaging, and has received numerous awards such as 13 paper awards at international conference and an Early Researcher Award from the Ministry of Economic Development and Innovation.

Xiaojun Yu Dr. Yu is Associate Professor, Biomedical Engineering at Stevens Institute of Technology, Hoboken, NJ, United States. Dr. Yu’s primary research interests focus on tissue engineering, polymeric biomaterials and drug delivery. His current research activities include nano- and microscale functionalization of biomimic three-dimensional scaffolds for neural and musculoskeletal tissue repair and regeneration, investigation of cell and material interactions in bioreactors, development of controlled release systems for the delivery of growth factors and drugs, and manipulation of microenvironment for stem cell proliferation and differentiation.

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CONTRIBUTORS TO VOLUME 2 Gursel Alici University of Wollongong, Wollongong, NSW, Australia Robert Amelard University of Waterloo, Waterloo, ON, Canada Orestis G Andriotis TU Wien, Vienna, Austria Michel Assad AccelLAB Inc., a Citoxlab Group Company, Boisbriand, QC, Canada Stéphane Avril Ecole Nationale Supérieure des Mines de Saint-Etienne, CIS-EMSE, SAINBIOSE, Saint Etienne, France; INSERM, Saint Etienne, France; and Université de Lyon, SAINBIOSE, Saint Etienne, France Eliezer Bernart Federal University of Rio Grande do Sul, Porto Alegre, Brazil Thor Besier University of Auckland, Auckland, New Zealand Paolo Bifulco University of Naples Federico II, Naples, Italy Romane Blanchard University of Melbourne, Melbourne, VIC, Australia Francesca Boccafoschi University of Piemonte Orientale, Novara, Italy Michel Boissière Université de Cergy Pontoise, Neuville-sur-Oise, France Stéphane Bolduc Laval University, Québec City, QC, Canada Carolina Catanio Bortolan Laval University, Quebec, QC, Canada Gilbert Bruce Pike University of Calgary, Calgary, AB, Canada Jennifer Shane Williamson Campbell McGill University, Montreal, QC, Canada

Luis Cardoso The Graduate School of The City University of New York, New York, NY, United States Ugo Carraro IRCCS Fondazione Ospedale San Camillo Venezia-Lido, Venezia, Italy Francesco Casella Ospedale Maggiore della Carità, Novara, Italy Marta C Catoira University of Piemonte Orientale, Novara, Italy Mario Cesarelli University of Naples Federico II, Naples, Italy Stéphane Chabaud Laval University, Québec City, QC, Canada Rachel W Chan Sunnybrook Research Institute, Toronto, ON, Canada Frédéric Chaubet Université Paris 13, Paris, France Jean Chen Rotman Research Institute, Baycrest Centre for Geriatric Health, Toronto, ON, Canada; and University of Toronto, Toronto, ON, Canada Pascale Chevallier Laval University, Quebec, QC, Canada Thomas Christian Gasser KTH Royal Institute of Technology, Stockholm, Sweden Paul K Chu City University of Hong Kong, Hong Kong, China Marie-Annick Clavel Laval University, Quebec City, QC, Canada John G Clement University of Melbourne, Melbourne, VIC, Australia Julien Cohen-Adad Polytechnique Montreal, University of Montreal, Montreal, QC, Canada

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Contributors to Volume 2

Caitlyn J Collins TU Wien, Vienna, Austria

Massimiliano Garzaro Eastern Piedmont University, Novara, Italy

David C Cooper University of Saskatchewan, Saskatoon, SK, Canada

Dario Gastaldi Department of Chemistry, Materials and Chemical Engineering Giulio Natta, Politecnico di Milano, Milano, Italy

Nancy Cote Laval University, Quebec City, QC, Canada Tien Tuan Dao Sorbonne Universités, Paris, France; and Université de Technologie de Compiègne, Compiègne, France Valeria Dell’Era Eastern Piedmont University, Novara, Italy Annalisa De Paolis The Graduate School of The City University of New York, New York, NY, United States

Shounak Ghosh Vellore Institute of Technology (VIT) University, Vellore, India Magnus K Gislason Reykjavik University, Reykjavik, Iceland Sara Greenberg University of Waterloo, Waterloo, ON, Canada Martin Guimond University of Montreal, Montreal, QC, Canada

Thomas E Doyle School of Biomedical Engineering, McMaster University, Hamilton, ON, Canada

Ezequiel Guzzetti Laval University, Quebec City, QC, Canada

Bryce T J Dyer Bournemouth University, Bournemouth, United Kingdom

Masoom A Haider University of Toronto, Toronto, ON, Canada; and Ontario Institute for Cancer Research (OICR), Toronto, ON, Canada

Mehran Ebrahimi University of Ontario Institute of Technology (UOIT), Oshawa, ON, Canada Kyle J Edmunds Reykjavik University, Reykjavik, Iceland Luca Esposito University of Naples Federico II, Naples, Italy Justin Fernandez University of Auckland, Auckland, New Zealand Daniel P Ferris University of Florida, Gainesville, FL, United States Eliezer Soares Flores Federal University of Rio Grande do Sul, Porto Alegre, Brazil Massimiliano Fraldi University of Naples Federico II, Naples, Italy Martin Frank TU Wien, Vienna, Austria Luca Fusaro University of Piemonte Orientale, Novara, Italy Ming Gan Purdue University, West Lafayette, IN, United States Paolo Gargiulo Reykjavik University, Reykjavik, Iceland

Matthew G Hanna University of Pittsburgh Medical Center, Pittsburgh, PA, United States Rita Hardiman University of Melbourne, Melbourne, VIC, Australia Christian Hellmich Vienna University of Technology, Vienna, Austria Wendy Hill University of New Brunswick, Fredericton, NB, Canada; and Atlantic Clinic for Upper Limb Prosthetics, Fredericton, NB, Canada Marie-Christine Ho Ba Tho Sorbonne Universités, Paris, France; and Université de Technologie de Compiègne, Compiègne, France Caroline D Hoemann George Mason University, Fairfax, VA, USA Nicolette Jackson AccelLAB Inc., a Citoxlab Group Company, Boisbriand, QC, Canada Weihong Jin City University of Hong Kong, Hong Kong, China Halldór Jónsson University of Iceland, Reykjavík, Iceland; and Landspitali University Hospital, Reykjavík, Iceland

Contributors to Volume 2

Picard Julien AKKA Life Science, Lyon, France

Sergio Loffredo Laval University, Quebec City, QC, Canada

Sumanta Kar North Dakota State University, Fargo, ND, United States

Andrea Malandrino Institute for Bioengineering of Catalonia, Barcelona, Spain; and Massachusetts Institute of Technology, Cambridge, MA, United States

Orestis L Katsamenis University of Southampton, Southampton, United Kingdom Dinesh R Katti North Dakota State University, Fargo, ND, United States Kalpana S Katti North Dakota State University, Fargo, ND, United States

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Geetha Manivasagam VIT University, Vellore, India Diego Mantovani Laval University, Quebec City, QC, Canada Michele Marino Leibniz Universität Hannover, Hannover, Germany Pietro Mascheroni University of Padova, Padova, Italy

Farnoud Kazemzadeh University of Waterloo, Waterloo, ON, Canada; and Elucid Labs Inc., Waterloo, ON, Canada

Carmelo Mastrandrea Jean Monnet University, Saint-Étienne, France; and Lyon University, Lyon, France

Farzad Khalvati University of Toronto, Toronto, ON, Canada; and Lunenfeld-Tanenbaum Research Institute, Toronto, ON, Canada

Mathew T Mathew UIC School of Medicine at Rockford, UIC, Rockford, IL, United States

David A Koff McMaster University, Hamilton, ON, Canada; and Diagnostic Imaging, Hamilton Health Sciences, Hamilton, ON, Canada Abbas Z Kouzani Deakin University, Geelong, VIC, Australia Witold Krasny Ecole Nationale Supérieure des Mines de Saint-Etienne, CIS-EMSE, SAINBIOSE, Saint Etienne, France; INSERM, Saint Etienne, France; Université de Lyon, SAINBIOSE, Saint Etienne, France; and Université de Lyon, Ecole Centrale Lyon, France Hyock Ju Kwon University of Waterloo, Waterloo, ON, Canada Wilfred W Lam Sunnybrook Research Institute, Toronto, ON, Canada Angus Z Lau Sunnybrook Research Institute, Toronto, ON, Canada; and University of Toronto, Toronto, ON, Canada Justin Y C Lau Sunnybrook Research Institute, Toronto, ON, Canada; and University of Toronto, Toronto, ON, Canada Hantao Liu Cardiff University, Cardiff, United Kingdom

John McPhee University of Waterloo, Waterloo, ON, Canada Emad Moeendarbary Massachusetts Institute of Technology, Cambridge, MA, United States; and University College London, London, United Kingdom Shahjahan Molla North Dakota State University, Fargo, ND, United States Vanessa Montaño-Machado Laval University, Quebec, QC, Canada Claire Morin Ecole Nationale Supérieure des Mines de Saint-Etienne, CIS-EMSE, SAINBIOSE, Saint Etienne, France; INSERM, Saint Etienne, France; and Université de Lyon, SAINBIOSE, Saint Etienne, France Ehsan Mostaed Politecnico di Milano, Milan, Italy Roger J Narayan UNC/NCSU Joint Department of Biomedical Engineering, Raleigh, NC, United States Sunita Nayak VIT University, Vellore, India Vedran Nedelkovski TU Wien, Vienna, Austria

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Contributors to Volume 2

Alexander K Nguyen UNC/NCSU Joint Department of Biomedical Engineering, Raleigh, NC, United States Ko Okumura Department of Physics and Soft Matter Center, Ochanomizu University, Tokyo, Japan Hazem Orabi Laval University, Québec City, QC, Canada Miao-Jung Y Ou University of Delaware, Newark, DE, United States Liron Pantanowitz University of Pittsburgh Medical Center, Pittsburgh, PA, United States

Erwan Salaun Laval University, Quebec City, QC, Canada Raffaella Santagiuliana University of Padova, Padova, Italy Antonio Santos-Paulo University of Aveiro, Aveiro, Portugal Jacob Scharcanski Federal University of Rio Grande do Sul, Porto Alegre, Brazil Stefan Scheiner Vienna University of Technology, Vienna, Austria Bryan R Schlink University of Florida, Gainesville, FL, United States

Maria-Ioana Pastrama KU Leuven, Leuven, Belgium

Marco Schneider University of Auckland, Auckland, New Zealand

Carlo Paternoster Laval University, Quebec City, QC, Canada

Bernhard Schrefler Technical University of Munich, Garching bei München, Germany; and Houston Methodist Research Institute, Houston, TX, United States

Philippe Pibarot Laval University, Quebec City, QC, Canada Jonathan Pitocchi Reykjavik University, Reykjavik, Iceland Peter Pivonka Queensland University of Technology, Brisbane, QLD, Australia Dejan B Popovic Serbian Academy of Sciences and Arts, Belgrade, Serbia; and Aalborg University, Aalborg, Denmark Asokami Rajamanikam Tamil Nadu Academy of Sciences, Chennai, India Martina Ramella University of Piemonte Orientale, Novara, Italy Sophie Ramsay Laval University, Québec City, QC, Canada Reza Sharif Razavian University of Waterloo, Waterloo, ON, Canada

Giraudier Sébastien Voisin Consulting, Boulogne, France Dwaipayan Sen Vellore Institute of Technology (VIT) University, Vellore, India Jonathon W Sensinger University of New Brunswick, Fredericton, NB, Canada Ashkan Shafiee Wake Forest School of Medicine, Winston-Salem, NC, United States Bonghun Shin University of Waterloo, Waterloo, ON, Canada Malgorzata Sikora-Jasinska Politecnico di Milano, Milan, Italy; and Laval University, Québec City, Canada

Aakash Reddy VIT University, Vellore, India

Roy V Sillitoe Baylor College of Medicine, Houston, TX, United States; and Texas Children’s Hospital, Houston, TX, United States

Violeta Rodriguez-Ruiz Université de Cergy Pontoise, Neuville-sur-Oise, France

Alex Swee University of Auckland, Auckland, New Zealand

Christian J Roth Technical University of Munich, Munich, Germany

Michelle Sybring University of New Brunswick, Fredericton, NB, Canada

Ingrid Saba Laval University, Québec City, QC, Canada

C David L Thomas University of Melbourne, Melbourne, VIC, Australia

Contributors to Volume 2

Philipp J Thurner TU Wien, Vienna, Austria

Wolfgang A Wall Technical University of Munich, Munich, Germany

Vikas Tomar Purdue University, West Lafayette, IN, United States

Zhou Wang University of Waterloo, Waterloo, ON, Canada

Paolo Aluffi Valletti Eastern Piedmont University, Novara, Italy

Alexander Wong University of Waterloo, Waterloo, ON, Canada; and Elucid Labs Inc., Waterloo, ON, Canada

Hans Van Oosterwyck Biomechanics Section, KU Leuven, Heverlee, Belgium Diego A Vargas Biomechanics Section, KU Leuven, Heverlee, Belgium Maurizio Vedani Politecnico di Milano, Milan, Italy Diego Velasquez Universidad CES, Medellin, Colombia Pasquale Vena Department of Chemistry, Materials and Chemical Engineering Giulio Natta, Politecnico di Milano, Milano, Italy Katari Venkatesh Vellore Institute of Technology (VIT) University, Vellore, India Larreta-Garde Véronique University of Cergy-Pontoise, Pontoise, France Laurence Vico Jean Monnet University, Saint-Étienne, France; and Lyon University, Lyon, France

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Shasha Yeung University of Auckland, Auckland, New Zealand Lena Yoshihara Technical University of Munich, Munich, Germany Aaron J Young Georgia Institute of Technology, Atlanta, GA, United States Anne-Sophie Zenses Laval University, Quebec City, QC, Canada; and AixMarseille University, Marseille, France Ju Zhang University of Auckland, Auckland, New Zealand Yang Zhang Purdue University, West Lafayette, IN, United States Yucheng Zhang University of Toronto, Toronto, ON, Canada; and Lunenfeld-Tanenbaum Research Institute, Toronto, ON, Canada

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CONTENTS OF VOLUME 2 Editorial Board

v

Editor in Chief

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Section Editors

ix

Contents of All Volumes Preface

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Biomechanics Biomechanics of Cells as Potential Biomarkers for Diseases: A New Tool in Mechanobiology Dinesh R Katti, Kalpana S Katti, Shahjahan Molla, and Sumanta Kar

1

Bone Micro- and Nanomechanics Caitlyn J Collins, Orestis G Andriotis, Vedran Nedelkovski, Martin Frank, Orestis L Katsamenis, and Philipp J Thurner

22

Cell Adhesion: Basic Principles and Computational Modeling Diego A Vargas and Hans Van Oosterwyck

45

Centrifugation and Hypergravity in the Bone Carmelo Mastrandrea and Laurence Vico

59

Computational Modeling of Respiratory Biomechanics Christian J Roth, Lena Yoshihara, and Wolfgang A Wall

70

Constitutive Modeling of Soft Tissues Michele Marino

81

Continuum Mechanics and Its Practical Applications at the Level of Scaling Laws Ko Okumura

111

CT-Based Bone and Muscle Assessment in Normal and Pathological Conditions Paolo Gargiulo, Magnus K Gislason, Kyle J Edmunds, Jonathan Pitocchi, Ugo Carraro, Luca Esposito, Massimiliano Fraldi, Paolo Bifulco, Mario Cesarelli, and Halldór Jónsson

119

Knowledge Extraction From Medical Imaging for Advanced Patient-Specific Musculoskeletal Models Marie-Christine Ho Ba Tho and Tien Tuan Dao

135

Mathematical Quantification of the Impact of Microstructure on the Various Effective Properties of Bones Miao-Jung Y Ou, Annalisa De Paolis, and Luis Cardoso

143

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Contents of Volume 2

Multiphase Porous Media Models for Mechanics in Medicine: Applications to Transport Oncophysics and Diabetic Foot Pietro Mascheroni, Raffaella Santagiuliana, and Bernhard Schrefler

155

Multiscale Bone Mechanobiology Stefan Scheiner, Maria-Ioana Pastrama, Peter Pivonka, and Christian Hellmich

167

Multiscale Mechanical Behavior of Large Arteries Claire Morin, Witold Krasny, and Stéphane Avril

180

Nanoindentation-Based Characterization of Hard and Soft Tissues Pasquale Vena and Dario Gastaldi

203

Nanomechanical Raman Spectroscopy in Biological Materials Yang Zhang, Ming Gan, and Vikas Tomar

215

On the Use of Population-Based Statistical Models in Biomechanics Justin Fernandez, Shasha Yeung, Alex Swee, Marco Schneider, Thor Besier, and Ju Zhang

229

Poroelasticity of Living Tissues Andrea Malandrino and Emad Moeendarbary

238

Structural and Material Changes of Human Cortical Bone With Age: Lessons from the Melbourne Femur Research Collection Romane Blanchard, C David L Thomas, Rita Hardiman, John G Clement, David C Cooper, and Peter Pivonka Vascular Tissue Biomechanics: Constitutive Modeling of the Arterial Wall Thomas Christian Gasser

246

265

Medical Devices 3D Printing in the Biomedical Field Alexander K Nguyen, Roger J Narayan, and Ashkan Shafiee Biocompatibility Evaluation of Orthopedic Biomaterials and Medical Devices: A Review of Safety and Efficacy Models Michel Assad and Nicolette Jackson

275

281

Biological Grafts: Surgical Use and Vascular Tissue Engineering Options for Peripheral Vascular Implants 310 Francesca Boccafoschi, Martina Ramella, Luca Fusaro, Marta C Catoira, and Francesco Casella Current Advancements and Challenges in Stent-Mediated Gene Therapy Shounak Ghosh, Katari Venkatesh, and Dwaipayan Sen

322

Dentistry: Restorative and Regenerative Approaches Geetha Manivasagam, Aakash Reddy, Dwaipayan Sen, Sunita Nayak, Mathew T Mathew, and Asokami Rajamanikam

332

Ephemeral Biogels: Potential Applications as Active Dressings and Drug Delivery Devices Larreta-Garde Véronique, Picard Julien, and Giraudier Sébastien

348

Immunological Responses in Orthopedics and Transplantation Caroline D Hoemann and Martin Guimond

359

Iron-Based Degradable Implants Sergio Loffredo, Carlo Paternoster, and Diego Mantovani

374

Contents of Volume 2

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Medical Devices: Coronary Stents Vanessa Montaño-Machado, Malgorzata Sikora-Jasinska, Carolina Catanio Bortolan, Pascale Chevallier, and Diego Mantovani

386

Medical Devices in Otorhinolaryngology Paolo Aluffi Valletti, Massimiliano Garzaro, and Valeria Dell’Era

399

Medical Devices in Neurology Abbas Z Kouzani and Roy V Sillitoe

409

Obstetrics and Gynecology: Hysteroscopy Antonio Santos-Paulo

414

Orthopedic Implants Weihong Jin and Paul K Chu

425

Pharmacology: Drug Delivery Frédéric Chaubet, Violeta Rodriguez-Ruiz, Michel Boissière, and Diego Velasquez

440

Prosthetic Aortic Valves Anne-Sophie Zenses, Philippe Pibarot, Marie-Annick Clavel, Ezequiel Guzzetti, Nancy Cote, and Erwan Salaun

454

Urology and Nephrology: Regenerative Medicine Applications Ingrid Saba, Stéphane Chabaud, Sophie Ramsay, Hazem Orabi, and Stéphane Bolduc

467

Zinc-Based Degradable Implants Ehsan Mostaed, Malgorzata Sikora-Jasinska, and Maurizio Vedani

478

Medical Imaging Biomechanics Imaging and Analysis Reza Sharif Razavian, Sara Greenberg, and John McPhee

488

Breast Imaging: Mammography, Digital Tomosynthesis, Dynamic Contrast Enhancement Mehran Ebrahimi

501

Diffusion Magnetic Resonance Imaging Jennifer Shane Williamson Campbell and Gilbert Bruce Pike

505

Digital Holographic Microscopy Farnoud Kazemzadeh and Alexander Wong

519

Digital Pathology Matthew G Hanna and Liron Pantanowitz

524

Functional Magnetic Resonance Imaging Jean Chen and Julien Cohen-Adad

533

Hemodynamic Imaging Robert Amelard and Alexander Wong

545

Imaging Informatics David A Koff and Thomas E Doyle

551

Macroscopic Pigmented Skin Lesion Prescreening Eliezer Bernart, Eliezer Soares Flores, and Jacob Scharcanski

561

Magnetic Resonance Imaging Rachel W Chan, Justin Y C Lau, Wilfred W Lam, and Angus Z Lau

574

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Contents of Volume 2

Perceptual Quality Assessment of Medical Images Hantao Liu and Zhou Wang

588

Radiomics Farzad Khalvati, Yucheng Zhang, Alexander Wong, and Masoom A Haider

597

Ultrasound Elastography Hyock Ju Kwon and Bonghun Shin

604

Rehabilitation Engineering and Integrative Technologies Functional Electric Stimulation Therapy Dejan B Popovic

614

ProsthesesdAssistive TechnologydSports Bryce T J Dyer

621

ProsthesesdAssistive TechnologydUpper Jonathon W Sensinger, Wendy Hill, and Michelle Sybring

632

Robotics: Exoskeletons Daniel P Ferris, Bryan R Schlink, and Aaron J Young

645

RoboticsdSoft Robotics Gursel Alici

652

CONTENTS OF ALL VOLUMES VOLUME 1 Biomaterials: Science and Engineering Alternative Processing Techniques for CoCr Dental Alloys Lucien Reclaru and Lavinia Cosmina Ardelean

1

Bioceramics Besim Ben-Nissan, Sophie Cazalbou, and Andy H Choi

16

Biomedical Composites Min Wang and Qilong Zhao

34

Bulk Properties of Biomaterials and Testing Techniques Min Wang and Chong Wang

53

Corrosion of Orthopedic Implants Qiong Wang, Felipe Eltit, and Rizhi Wang

65

Decellularized Extracellular Matrix Paul Frank Gratzer

86

Diamond, Carbon Nanotubes and Graphene for Biomedical Applications Aaqil Rifai, Elena Pirogova, and Kate Fox

97

Gold Nanoparticles for Colorimetric Detection of Pathogens Paul Z Chen and Frank X Gu

108

Manufacture of Biomaterials Min Wang, Lin Guo, and Haoran Sun

116

Materials and Their Biomedical Applications Min Wang and Bin Duan

135

Nano-Biomaterials and their Applications Mian Wang and Thomas J Webster

153

Natural Biopolymers for Biomedical Applications Natalia Davidenko, Ruth Cameron, and Serena Best

162

Polymeric Coatings and Their Fabrication for Medical Devices Dimitrios A Lamprou, Nikolaos Scoutaris, Steven A Ross, and Dionysios Douroumis

177

xxv

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Contents of All Volumes

Porous Biomaterials and Scaffolds for Tissue Engineering Liliana Liverani, Vincenzo Guarino, Vincenzo La Carrubba, and Aldo R Boccaccini

188

Preparation and Properties of Coatings and Thin Films on Metal Implants Zhong Li and Khiam Aik Khor

203

Titanium Alloys Mitsuo Niinomi

213

Biomaterials: In Vitro and in Vivo Studies of Biomaterials Anatomy and Physiology for Biomaterials Research and Development Inn Chuan Ng, Pornteera Pawijit, Jordon Tan, and Hanry Yu

225

Animal Models in Biomaterial Development James M Anderson and Sirui Jiang

237

Blood–Biomaterial Interactions Nicholas P Rhodes

242

Interaction Between Mesenchymal Stem Cells and Immune Cells in Tissue Engineering Rong Huang, Yinghong Zhou, and Yin Xiao

249

Osseointegration of Permanent and Temporary Orthopedic Implants J S Hayes and R G Richards

257

Tissue Response to Biomaterials Jiao Jiao Li and Hala Zreiqat

270

Biomaterials: Biomaterial Applications and Advanced Medical Technologies Biomaterials in Dentistry Li Wu Zheng, Jing Yi Wang, and Ru Qing Yu

278

Biomaterials in Ophthalmology 289 Rachel L Williams, Hannah J Levis, Rebecca Lace, Kyle G Doherty, Stephnie M Kennedy, and Victoria R Kearns Biomaterials in Orthopaedics Emmanuel Gibon and Stuart B Goodman

301

Cell Encapsulation and Delivery Stefani Mazzitelli and Claudio Nastruzzi

308

Drug Delivery Systems and Controlled Release Nicholas J Kohrs, Thilanga Liyanage, Nandakumar Venkatesan, Amir Najarzadeh, and David A Puleo

316

Electrospinning and Electrospray for Biomedical Applications Min Wang and Qilong Zhao

330

Gene Delivery and Clinical Applications Mahboob Morshed and Ezharul Hoque Chowdhury

345

Materials for Exoskeletal Orthotic and Prosthetic Systems Man Sang Wong, Babak Hassan Beygi, and Yu Zheng

352

Microfluidics for Biomedical Applications Shiyu Cheng, Jinqi Deng, Wenfu Zheng, and Xingyu Jiang

368

Organs-on-Chips Yunki Lee, Song Ih Ahn, and YongTae Kim

384

Contents of All Volumes

Shape-Memory Polymer Medical Devices Muhammad Y Razzaq, Markus Reinthaler, Mark Schröder, Christian Wischke, and Andreas Lendlein

xxvii

394

Regenerative Engineering Adult Bone Marrow-Derived Stem Cells: Immunomodulation in the Context of Disease and Injury A E Ting and S A Busch

406

Assessment of Cellular Responses of Tissue Constructs in vitro in Regenerative Engineering Margaret A T Freeberg, Jacob G Kallenbach, and Hani A Awad

414

Assessment of Tissue Constructs In Vivo in Regenerative Engineering Anuradha Subramanian and Swaminathan Sethuraman

427

Bioengineered Kidney and Bladder D S Koslov and A Atala

432

Bioengineering Scaffolds for Regenerative Engineering Zichen Qian, Daniel Radke, Wenkai Jia, Mitch Tahtinen, Guifang Wang, and Feng Zhao

444

Biomaterials for Tissue Engineering and Regenerative Medicine Ohan S Manoukian, Naseem Sardashti, Teagen Stedman, Katie Gailiunas, Anurag Ojha, Aura Penalosa, Christopher Mancuso, Michelle Hobert, and Sangamesh G Kumbar

462

Biomimetic Approaches for Regenerative Engineering Nirmalya Tripathy, Rafiq Ahmad, Jeong Eun Song, and Gilson Khang

483

Bioreactors: System Design and Application for Regenerative Engineering Antonio Valdevit

496

Bone Substitute Materials M Bohner

513

Case Studies for Soft Tissue Regenerative Engineering Jorge Luis Escobar Ivirico and Cato T Laurencin

530

Characterizing the Properties of Tissue Constructs for Regenerative Engineering Yusuf Khan

537

Clinical and Laboratory Aspects of Hematopoietic Stem Cell Transplantation S T Avecilla and M M Cushing

546

Dental Stem Cells M Nakashima and Y Hayashi

554

Drug and Gene Delivery for Regenerative Engineering Morgan A Urello, Tianzhi Luo, Bing Fang, Kristi L Kiick, and Millicent O Sullivan

565

Ethics of Issues and Stem Cell Research: the Unresolved Issues Z Master

584

Eye Diseases and Stem Cells H Ouyang, D H Nguyen, and K Zhang

598

Human Parthenogenetic Pluripotent Stem Cells N Turovets and M Csete

608

Human Pluripotent Stem Cells P Rajan

618

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Contents of All Volumes

Introduction to Regenerative Engineering Manisha Jassal, Radoslaw Junka, Cato T Laurencin, and Xiaojun Yu

624

Nanoelectronics for Neuroscience Sahil Kumar Rastogi and Tzahi Cohen-Karni

631

Neural Crest Stem Cells T Hochgreb-Hägele and M E Bronner

650

Osteoarthritis at the Cellular Level: Mechanisms, Clinical Perspectives, and Insights From Development 660 Melanie Fisher, Tyler Ackley, Kelsey Richard, Bridget Oei, and Caroline N Dealy Reproductive Technologies, Assisted D Pergament

677

Tooth Regenerative Therapy: Tooth Tissue Repair and Whole Tooth Replacement M Oshima, K Ishida, R Morita, M Saito, and T Tsuji

686

Vascularized Tissue Regenerative Engineering Using 3D Bioprinting Technology Sungwoo Kim, Arnaud Bruyas, Chi-Chun Pan, Alexander Martin Stahl, and Yunzhi Yang

696

Wound Healing and the Host Response in Regenerative Engineering Daniel Chester, Ethan A Marrow, Michael A Daniele, and Ashley C Brown

707

VOLUME 2 Biomechanics Biomechanics of Cells as Potential Biomarkers for Diseases: A New Tool in Mechanobiology Dinesh R Katti, Kalpana S Katti, Shahjahan Molla, and Sumanta Kar

1

Bone Micro- and Nanomechanics Caitlyn J Collins, Orestis G Andriotis, Vedran Nedelkovski, Martin Frank, Orestis L Katsamenis, and Philipp J Thurner

22

Cell Adhesion: Basic Principles and Computational Modeling Diego A Vargas and Hans Van Oosterwyck

45

Centrifugation and Hypergravity in the Bone Carmelo Mastrandrea and Laurence Vico

59

Computational Modeling of Respiratory Biomechanics Christian J Roth, Lena Yoshihara, and Wolfgang A Wall

70

Constitutive Modeling of Soft Tissues Michele Marino

81

Continuum Mechanics and Its Practical Applications at the Level of Scaling Laws Ko Okumura

111

CT-Based Bone and Muscle Assessment in Normal and Pathological Conditions Paolo Gargiulo, Magnus K Gislason, Kyle J Edmunds, Jonathan Pitocchi, Ugo Carraro, Luca Esposito, Massimiliano Fraldi, Paolo Bifulco, Mario Cesarelli, and Halldór Jónsson

119

Knowledge Extraction From Medical Imaging for Advanced Patient-Specific Musculoskeletal Models Marie-Christine Ho Ba Tho and Tien Tuan Dao

135

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Mathematical Quantification of the Impact of Microstructure on the Various Effective Properties of Bones Miao-Jung Y Ou, Annalisa De Paolis, and Luis Cardoso

143

Multiphase Porous Media Models for Mechanics in Medicine: Applications to Transport Oncophysics and Diabetic Foot Pietro Mascheroni, Raffaella Santagiuliana, and Bernhard Schrefler

155

Multiscale Bone Mechanobiology Stefan Scheiner, Maria-Ioana Pastrama, Peter Pivonka, and Christian Hellmich

167

Multiscale Mechanical Behavior of Large Arteries Claire Morin, Witold Krasny, and Stéphane Avril

180

Nanoindentation-Based Characterization of Hard and Soft Tissues Pasquale Vena and Dario Gastaldi

203

Nanomechanical Raman Spectroscopy in Biological Materials Yang Zhang, Ming Gan, and Vikas Tomar

215

On the Use of Population-Based Statistical Models in Biomechanics Justin Fernandez, Shasha Yeung, Alex Swee, Marco Schneider, Thor Besier, and Ju Zhang

229

Poroelasticity of Living Tissues Andrea Malandrino and Emad Moeendarbary

238

Structural and Material Changes of Human Cortical Bone With Age: Lessons from the Melbourne Femur Research Collection Romane Blanchard, C David L Thomas, Rita Hardiman, John G Clement, David C Cooper, and Peter Pivonka Vascular Tissue Biomechanics: Constitutive Modeling of the Arterial Wall Thomas Christian Gasser

246

265

Medical Devices 3D Printing in the Biomedical Field Alexander K Nguyen, Roger J Narayan, and Ashkan Shafiee Biocompatibility Evaluation of Orthopedic Biomaterials and Medical Devices: A Review of Safety and Efficacy Models Michel Assad and Nicolette Jackson

275

281

Biological Grafts: Surgical Use and Vascular Tissue Engineering Options for Peripheral Vascular Implants 310 Francesca Boccafoschi, Martina Ramella, Luca Fusaro, Marta C Catoira, and Francesco Casella Current Advancements and Challenges in Stent-Mediated Gene Therapy Shounak Ghosh, Katari Venkatesh, and Dwaipayan Sen

322

Dentistry: Restorative and Regenerative Approaches Geetha Manivasagam, Aakash Reddy, Dwaipayan Sen, Sunita Nayak, Mathew T Mathew, and Asokami Rajamanikam

332

Ephemeral Biogels: Potential Applications as Active Dressings and Drug Delivery Devices Larreta-Garde Véronique, Picard Julien, and Giraudier Sébastien

348

Immunological Responses in Orthopedics and Transplantation Caroline D Hoemann and Martin Guimond

359

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Iron-Based Degradable Implants Sergio Loffredo, Carlo Paternoster, and Diego Mantovani

374

Medical Devices: Coronary Stents Vanessa Montaño-Machado, Malgorzata Sikora-Jasinska, Carolina Catanio Bortolan, Pascale Chevallier, and Diego Mantovani

386

Medical Devices in Otorhinolaryngology Paolo Aluffi Valletti, Massimiliano Garzaro, and Valeria Dell’Era

399

Medical Devices in Neurology Abbas Z Kouzani and Roy V Sillitoe

409

Obstetrics and Gynecology: Hysteroscopy Antonio Santos-Paulo

414

Orthopedic Implants Weihong Jin and Paul K Chu

425

Pharmacology: Drug Delivery Frédéric Chaubet, Violeta Rodriguez-Ruiz, Michel Boissière, and Diego Velasquez

440

Prosthetic Aortic Valves Anne-Sophie Zenses, Philippe Pibarot, Marie-Annick Clavel, Ezequiel Guzzetti, Nancy Cote, and Erwan Salaun

454

Urology and Nephrology: Regenerative Medicine Applications Ingrid Saba, Stéphane Chabaud, Sophie Ramsay, Hazem Orabi, and Stéphane Bolduc

467

Zinc-Based Degradable Implants Ehsan Mostaed, Malgorzata Sikora-Jasinska, and Maurizio Vedani

478

Medical Imaging Biomechanics Imaging and Analysis Reza Sharif Razavian, Sara Greenberg, and John McPhee

488

Breast Imaging: Mammography, Digital Tomosynthesis, Dynamic Contrast Enhancement Mehran Ebrahimi

501

Diffusion Magnetic Resonance Imaging Jennifer Shane Williamson Campbell and Gilbert Bruce Pike

505

Digital Holographic Microscopy Farnoud Kazemzadeh and Alexander Wong

519

Digital Pathology Matthew G Hanna and Liron Pantanowitz

524

Functional Magnetic Resonance Imaging Jean Chen and Julien Cohen-Adad

533

Hemodynamic Imaging Robert Amelard and Alexander Wong

545

Imaging Informatics David A Koff and Thomas E Doyle

551

Macroscopic Pigmented Skin Lesion Prescreening Eliezer Bernart, Eliezer Soares Flores, and Jacob Scharcanski

561

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Magnetic Resonance Imaging Rachel W Chan, Justin Y C Lau, Wilfred W Lam, and Angus Z Lau

574

Perceptual Quality Assessment of Medical Images Hantao Liu and Zhou Wang

588

Radiomics Farzad Khalvati, Yucheng Zhang, Alexander Wong, and Masoom A Haider

597

Ultrasound Elastography Hyock Ju Kwon and Bonghun Shin

604

Rehabilitation Engineering and Integrative Technologies Functional Electric Stimulation Therapy Dejan B Popovic

614

ProsthesesdAssistive TechnologydSports Bryce T J Dyer

621

ProsthesesdAssistive TechnologydUpper Jonathon W Sensinger, Wendy Hill, and Michelle Sybring

632

Robotics: Exoskeletons Daniel P Ferris, Bryan R Schlink, and Aaron J Young

645

RoboticsdSoft Robotics Gursel Alici

652

VOLUME 3 Mathematical Techniques in Biomedical Engineering Cardiac Modeling Edward Vigmond and Gernot Plank

1

Mathematical Approaches for Medical Image Registration Barbara Zitova

21

Mathematical Modeling of Gene Networks Lakshmi Sugavaneswaran

33

Mathematical Modeling Tools and Software for BME Applications Fred J Vermolen and Amit Gefen

56

Mathematical Techniques for Biomedical Image Segmentation Roberto Rodríguez and Juan H Sossa

64

Mathematical Techniques for Circulatory Systems Jason Carson, Raoul Van Loon, and Perumal Nithiarasu

79

Mathematical Techniques for Noninvasive Muscle Signal Analysis and Interpretation Roberto Merletti, Ales Holobar, and Dario Farina

95

Optimization Techniques in BME Jeevan Kumar Pant

112

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Single-Cell-Based In Silico Models: A Tool for Dissecting Tumor Heterogeneity Aleksandra Karolak, Saharsh Agrawal, Samantha Lee, and Katarzyna A Rejniak

130

Spectrotemporal Modeling of Biomedical Signals: Theoretical Foundation and Applications Raymundo Cassani and Tiago H Falk

144

Statistical Modeling in Biomedical Engineering Yunfeng Wu and Pinnan Chen

164

Time–Frequency Distributions in Biomedical Signal Processing Yashodhan Athavale and Sridhar Krishnan

177

Wavelets in Biomedical Signal Processing and Analysis Babak Azmoudeh and Dean Cvetkovic

193

Bioinstrumentation and Bioinformatics A Systematic Workflow for Design and Computational Analysis of Protein Microarrays Jonatan Fernández-García, Rodrigo García-Valiente, Javier Carabias-Sánchez, Alicia Landeira-Viñuela, Rafael Góngora, María Gonzalez-Gonzalez, and Manuel Fuentes

213

Ambulatory EEG Monitoring Bernard Grundlehner and Vojkan Mihajlovic

223

Automated EEG Analysis for Neonatal Intensive Care Nathan Stevenson and Anton Tokariev

240

Big Data Calls for Machine Learning Andreas Holzinger

258

Bioimpedance Spectroscopy Processing and Applications Herschel Caytak, Alistair Boyle, Andy Adler, and Miodrag Bolic

265

Bioinformatics in Design of Antiviral Vaccines Ashesh Nandy and Subhash C Basak

280

Bioinformatics in Disease Classification Miguel Ángel Medina

291

Biopotential Monitoring Julián Oreggioni, Angel A Caputi, and Fernando Silveira

296

Blood Gas Analysis and Instrumentation Rebecca Symons, Robindro Chatterji, Kirsty Whenan, Rita Horvath, and Paul S Thomas

305

Computational Approaches in microRNA Biology Ulf Schmitz, Shailendra K Gupta, Julio Vera, and Olaf Wolkenhauer

317

Detection and Classification of Breast Lesions Using Ultrasound-Based Imaging Modalities Md Kamrul Hasan and Sharmin R Ara

331

DNA Microarrays: Fundamentals, Data Integration and Applications Eduardo Valente and Miguel Rocha

349

ECG Monitoring: Present Status and Future Trend Saurabh Pal

363

Genetic Algorithms for Breast Cancer Diagnostics Florin Gorunescu and Smaranda Belciug

380

Contents of All Volumes

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Machine Learning in Biomedical Informatics Carlos Fernandez-Lozano, Adrián Carballal, Cristian R Munteanu, Marcos Gestal, Víctor Maojo, and Alejandro Pazos

389

Matrix Assisted Laser Desorption/Ionization as a New Cancer Diagnostic Tool Bozena Hosnedlova, Marta Kepinska, Branislav Ruttkay-Nedecky, Carlos Fernandez, Tomas Parak, Halina Milnerowicz, Jiri Sochor, Geir Bjørklund, and Rene Kizek

400

Medical Utility of NIR Monitoring Zuzana Kovacsova, Gemma Bale, and Ilias Tachtsidis

415

Metabolomics in Biomaterial Research Ana M Gil, Maria H Fernandes, and Iola F Duarte

432

Nucleic-Acid Sequencing G Dorado, S Gálvez, H Budak, T Unver, and P Hernández

443

Optical Techniques for Blood and Tissue Oxygenation Panayiotis Kyriacou, Karthik Budidha, and Tomas Y Abay

461

Polymerase Chain Reaction (PCR) G Dorado, G Besnard, T Unver, and P Hernández

473

Single-Photon Emission Computed Tomography: Principles and Applications Yong Du and Habib Zaidi

493

Translational Bioinformatics: Informatics, Medicine, and -Omics Sergio Paraiso-Medina, David Perez-Rey, Raul Alonso-Calvo, Cristian R Munteanu, Alejandro Pazos, Casimir A Kulikowski, and Victor Maojo

507

Ultrasonic Imaging in Biomedical Applications Roman Gr Maev and Fedar M Severin

515

Index

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PREFACE The use by man of available technology to treat damaged or diseased tissue is older than the written historical record. For example, the Mayan people created artificial teeth out of nacre, which were shown to be fused to the bone (Bobbio, 1972; Westbroek & Marin, 1998). Giovanni Borelli’s studies of the cardiovascular system (e.g., the elasticity of arteries), which were published in De Motu Animalium (On Animal Motion), can be considered as one of the foundations of the field of biomechanics (Parker, 2009). The hypothesis of an intrinsic ’animal electricity’ by Luigi Galvani in the 18th century led to the development of the field of electrophysiology (Piccolino, 1997). In the 19th century, the development of the antiseptic approach to surgical procedures by Joseph Lister made implantation of medical devices without certain postoperative infection possible; for example, Lister described the use of silver wire for treatment of a fractured patella (Worboys, 2013). The discovery of Xrays by Roentgen at the end of the 19th century was rapidly translated for medical imaging (Rowland, 1896; Schuster, 1896). In our lifetime, the work by W. T. Green on generating new cartilage by seeding of chondrocytes as well as by John Burke and Ioannis Yannas on generating skin substitutes is recognized as the birth of the field of regenerative engineering (Vacanti, 2006). At the beginning of the 21st century, the American Institute for Medical and Biological Engineering identified several research areas for the field of biomedical engineering. These include: (a) functional genomics and proteomics, (b) nanotechnology, (c) targeted drug delivery, (d) tissue engineering, and (e) the development of new types of medical instrumentation (Hendee, Chien, Maynard, & Dean, 2002). As some of these research areas have matured, others have become more prominent. Over the past few years, the use of 3D printing and bioprinting technologies to create medical devices has become more prominent. One benefit of utilizing 3D printing and bioprinting for patient care is that medical imaging data (e.g., magnetic resonance imaging and computed tomography data) may be employed to fashion prostheses or artificial tissues with submicroscale features that conform with the requirements of the patient (Narayan, Doraiswamy, Chrisey, & Chichkov, 2010; Skoog & Narayan, 2013). Another technology that will likely transform the field of biomedical engineering over the coming decades involves the use of clustered regularly interspaced short palindromic repeats (CRISPR)/Cas9 for engineering of the human genome. The interface between biomedical engineering and the new field of genome engineering has already spawned research into new technologies for delivery of genome editing tools into the body; the synergy between these fields will only grow over time (Wright, Nuñez, & Doudna, 2016). The goal of the Encyclopedia of Biomedical Engineering is to consider the principles and technologies that underlie the field of biomedical engineering. The encyclopedia is divided into three volumes. The first volume contains a section on biomaterials, which was edited by Min Wang at the University of Hong Kong, and a section on regenerative engineering, which was edited by Cato Laurencin at the University of Connecticut and Xiaojun Yu at the Stevens Institute of Technology. The second volume contains a section on rehabilitation engineering and integrative technologies, which was edited by William Rymer and Levi Hargrove at Northwestern University, a section on biomechanics, which was edited by Christian Hellmich at the Vienna University of Technology, a section on medical devices, which was edited by Diego Mantovani at the University of Laval, and a section on medical imaging, which was edited by Alexander Wong at the University of Waterloo. The third volume contains a section on mathematical techniques in biomedical engineering, which was edited by Sri Krishnan at Ryerson University, and a section on bioinstrumentation and bioinformatics, which was edited by Pankaj Vadgama at Queen Mary University of London. I would like express my sincere appreciation to the section editors and authors for all of their efforts on the encyclopedia. I would also like thank Beckie Brand, Susan Dennis, Becky Gelson, Ginny Mills, Blerina Osmanaj,

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Laura Escalante Santos, and Will Smaldon at Elsevier for their outstanding efforts to bring the encyclopedia to publication. I hope that this work serves the biomedical engineering community by providing a resource that considers topics at the interface of the biological sciences and engineering. Roger J Narayan, M.D., Ph.D. UNC/NCSU Joint Department of Biomedical Engineering. References Bobbio, A. (1972). The first endosseous alloplastic implant in the history of man. Bull. Hist. Dent, 20, 1–6. Hendee, W. R., Chien, S., Maynard, C. D., & Dean, D. J. (2002). The National Institute of biomedical imaging and Bioengineering: history, status, and potential impact. Annals of Biomedical Engineering, 30, 2–10. Narayan, R. J., Doraiswamy, A., Chrisey, D. B., & Chichkov, B. N. (2010). Medical prototyping using two photon polymerization. Materials Today, 13, 44–50. Parker, K. H. (February 2009). A brief history of arterial wave mechanics. Medical & Biological Engineering & Computing, 47(2), 111–118. Piccolino, M. (October 1997). Luigi Galvani and animal electricity: two centuries after the foundation of electrophysiology. Trends in Neurosciences, 20(10), 443–448. Rowland, S. (March 7, 1896). Report on the Application of the new Photography to medicine and surgery. Br Med J, 1(1836), 620–622. Schuster, A. (January 18, 1896). On the new Kind of Radiation. Br Med J, 1(1829), 172–173. Skoog, S. A., & Narayan, R. J. (2013). Stereolithography in medical device fabrication. Advanced Materials & Processes, 171, 32–36. Vacanti, C. A. (July 2006). The history of tissue engineering. J Cell Mol Med, 10(3), 569–576. Westbroek, P., & Marin, F. (1998). A marriage of bone and nacre. Nature, 392, 861–862. Worboys, M. (September 20, 2013). Joseph Lister and the performance of antiseptic surgery. Notes and Records of the Royal Society of London, 67(3), 199–209. Wright, A. V., Nuñez, J. K., & Doudna, J. A. (January 14, 2016). Biology and Applications of CRISPR systems: Harnessing Nature’s Toolbox for genome engineering. Cell, 164(1–2), 29–44.

PERMISSION ACKNOWLEDGMENTS The following material is reproduced with kind permission of Taylor & Francis Figure 1 On the Use of Population-Based Statistical Models in Biomechanics Figure 2 On the Use of Population-Based Statistical Models in Biomechanics Figure 3 On the Use of Population-Based Statistical Models in Biomechanics www.taylorandfrancisgroup.com The following material is reproduced with kind permission of American Association for the Advancement of Science Figure 8a Nanotechnology for Regenerative Engineering www.aaas.org The following material is reproduced with kind permission of Oxford University Press Figure 1 Translational Bioinformatics for Personalised Medicine www.oup.com The following material is reproduced with kind permission of Nature Publishing Group Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure Figure

5 Continuum Mechanics and Its Practical Applications at the Level of Scaling Laws 6 Continuum Mechanics and Its Practical Applications at the Level of Scaling Laws 7 Continuum Mechanics and Its Practical Applications at the Level of Scaling Laws 8 Continuum Mechanics and Its Practical Applications at the Level of Scaling Laws 9 Continuum Mechanics and Its Practical Applications at the Level of Scaling Laws 2 Multi-scale Bone Mechanobiology 3 Multi-scale Bone Mechanobiology 5 Functional Magnetic Resonance Imaging 6 Functional Magnetic Resonance Imaging 4b Cell Mechanics and Cell Adhesion - Basic Principles and Computational Modeling 9 Diffusion Tensor Imaging 8 Corrosion of Biomaterials 8 Biomaterials for Tissue Engineering and Regenerative Medicine 9 Biomaterials for Tissue Engineering and Regenerative Medicine 12 Biomaterials for Tissue Engineering and Regenerative Medicine 3c Nanotechnology for Regenerative Engineering 4a Nanotechnology for Regenerative Engineering 4b Nanotechnology for Regenerative Engineering 5b Nanotechnology for Regenerative Engineering 5c Nanotechnology for Regenerative Engineering 6 Nanotechnology for Regenerative Engineering 7b Nanotechnology for Regenerative Engineering

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Figure 7c Nanotechnology for Regenerative Engineering Figure 8b Nanotechnology for Regenerative Engineering Figure 5 Drug and Gene Delivery for Regenerative Engineering Figure 3 Holographic Microscopy Figure 4 Radiomics http://www.nature.com

BIOMECHANICS Biomechanics of Cells as Potential Biomarkers for Diseases: A New Tool in Mechanobiology Dinesh R Katti, Kalpana S Katti, Shahjahan Molla, and Sumanta Kar, North Dakota State University, Fargo, ND, United States © 2019 Elsevier Inc. All rights reserved.

Introduction Theory Theory of AFM Force Curves Operational modes of AFM Theory of Nanoindentation Modes of nanoindentation Examples of Nanomechanical Testing of Cells AFM Nanoindentation Perspectives Some Current Challenges in Use of AFM Current Challenges in Nanoindentation of Live Cells Future Perspectives Acknowledgements References

1 2 2 4 4 6 7 7 12 16 16 16 17 17 17

Abbreviations AFM Atomic force microscopy nanoDMA Nano dynamic mechanical analysis

Introduction Living cells in the human body, as physical entities, possess structural and physical properties and are constantly subjected to various mechanical stimulations. Studies in the evaluation of mechanics of single cells, and also various cellular components have indeed grown in importance due to their potential impact on medicine and human health (Suresh, 2007a,b; Suresh et al., 2005; Bao and Suresh, 2003). It has been shown that cellular functions such as growth, differentiation, migration, apoptosis, motility, gene expression, signal transduction, etc. are influenced by cellular biophysics and biomechanics (Chen et al., 1997; Boudreau and Bissell, 1998; Huang and Ingber, 1999; Schwartz and Ginsberg, 2002; Suresh et al., 2015). Any changes to the biophysical or biomechanical properties of human cells may interrupt regular physiological functions of cells and may lead to diseases. For example, when red blood cells are affected by malaria-causing bacteria, plasmodium falciparum, the molecular and structural properties of red blood cells are significantly compromised (Fedosov et al., 2011; Suresh et al., 2015; Cooke et al., 2001; Bannister and Mitchell, 2003). Biomechanical properties of individual cells can influence the physio-structural properties of the whole tissue by interacting with the extracellular matrix. Also, mechanical stress on the tissue is transmitted to the cellular level, compromising their physiological functions significantly (Guilak and Mow, 2000). Other than disease causing virus and bacteria, many chemicals are also known to have an effect on mechanical properties of human cells. For instance, latrunculin B cytochalasin D have an adverse effect on the cytoskeletal structure of the cells (Nagayama et al., 2006; Wakatsuki et al., 2001; Sato et al., 1990). Chemotactic agent fMLP significantly increases the stiffness of neutrophils (Zahalak et al., 1990; Worthen et al., 1989). The primary structural element of the cytoskeleton, the actin filament and microtubules are affected by colchicine (Tsai et al., 1998; Borisy and Taylor, 1967; Imazio et al., 2014). The past few decades has seen a substantial growth of the research in the field of biomechanical properties of cells. Various techniques have been applied to determine the biophysical properties of cells (Stamenovic et al., 1996; Lim et al., 2004; Stamenovic and Ingber, 2002). However, it is still challenging to evaluate the nanomechanics of a single cell, considering the dynamic nature of cells during progression of different diseases (Bao and Suresh, 2003; Zhu et al., 2000; Fung and Liu, 1993).

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With the advent of nanotechnology, it has now become feasible to probe cellular mechanics at single cell level. For instance, biophysical techniques such as micropipette aspiration (Evans and Yeung, 1989; Guo et al., 2012; Sato et al., 1987, 1990, 1996; Thoumine et al., 1999; Tsai et al., 1998), optical tweezers (Balland et al., 2006; Coceano et al., 2016; Henon et al., 1999), cell stretchers (Harris et al., 2013; Katsumi et al., 2002; Krishnan et al., 2009; Steward et al., 2011), flow rheometry (Gossett et al., 2012; Guo et al., 2012; Lu et al., 2004; Steward et al., 2011), magnetic bead cytometry (Deng et al., 2006; Fabry et al., 2001), traction force microscopy (TFM) (Munevar et al., 2001; Legant et al., 2010), atomic force microscopy (AFM) (Ketene et al., 2012; Kuznetsova et al., 2007; Lekka, 2016; Nematbakhsh and Lim, 2015), micropillar arrays (Fu et al., 2010), and nanoindentation (Chen et al., 2010; Duan and Chen, 2015; Khanna et al., 2011, 2012; McDaniel et al., 2007; Yang et al., 2010; Miyamoto et al., 2015), have been used extensively in the past decade to probe mechanical properties of various cell types. Of these above-mentioned techniques, a majority have been employed to apply controlled force to the cells (micropipette aspiration, optical tweezers, cell stretchers, flow rheometry, magnetic bead cytometry, AFM, and nanoindentation), while the rest are used to monitor the cells’ ability to deform itself by intracellular forces (TFM, and micropillar arrays) to obtain their mechanical properties. The connection between human disease and biophysics of cells has become a subject of intense scientific research recently. Cellular nanomechanics is strongly connected to the molecular and physiological changes introduced by the progression of certain disease and invasion by various external organisms such as a virus, bacteria or other parasites (Chaffey et al., 2003; Boal and Boal, 2012; Miller et al., 2002; Cooke et al., 2001). Cellular pathology and pathophysiology are heavily influenced by these changes in elastic and viscoelastic properties of cells (Miller et al., 2002; Cooke et al., 2001; Bao and Suresh, 2003). In the past two decades, advancement in the field of tissue engineering and bioengineering has enabled us to investigate the real-time biomechanical changes during the progression of certain diseases. The nanomechanical properties of living cells such as elastic modulus or stiffness can either increase or decrease in a pathogenic progression depending on the biomolecular and biochemical restructure (Suwanarusk et al., 2004). It has been demonstrated that the mechanics of cells play a critical role in cellular growth and disease progression such as cancer, cardiovascular diseases, liver diseases, renal glomerular diseases etc. (Engler et al., 2007; Yallapu et al., 2015; Chaturvedi et al., 2010; Georges et al., 2007; Wells, 2008; Wyss et al., 2011). Several studies showed that the stiffness of the substrate in the in vitro studies hugely affects the growth and differentiation of cells. For instance, it is shown that when cortical brain cells are cultured in a softer substrate (0.15–0.30 kPa), neurons grow selectively. On the other hand, when same cells were cultured in a stiffer substrate (2 kPa), the neurons attach to glial cells, and astrocytes proliferate (Lu et al., 2006; Georges et al., 2006). The elastic modulus of some of the tissues such as liver, lung, breast, kidney has been reported in the range of 0.2–4.0 kPa (Levental et al., 2007). The variance in elastic modulus of cells within the same tissue has been reported to be within a 10%–15% range, and under normal physiological conditions, this controlled mechanical property of cells helps them to maintain homeostatic cell-cycle progression (Georges et al., 2007). Changes in elastic properties > 12 kPa leads to abnormal cell cycle progression which may cause irregular cell growth and differentiation (Kumar and Weaver, 2009; Assoian and Klein, 2008; Klein et al., 2009; Levental et al., 2009). It has been reported that live breast, liver and pancreatic cancer cells obtained from a body fluid of patients possess very different nanomechanical properties than their normal counterparts although cancerous and normal cells exhibit similar morphologies (Cross et al., 2008). Baker et al. (2009) demonstrated that cancerous breast cancer cells are stiffer (4 kPa) than the normal breast tissue (0.2 kPa) (Baker et al., 2009; Levental et al., 2007). It has been reported that dense and stiffer breast tissue is more likely to develop cancer. In vitro studies suggested that stiffer substrate increases the invasive nature of cancer cells by increasing ERK and Rho activity of breast cancer cells (Paszek et al., 2005). Further, normal cells (Hu609 and HCV29) have higher Young’s modulus than their cancerous counterpart (Hu456, T24, BC3726) (Lekka et al., 1999). Nanomechanical properties of cells also play a significant role in various cardiac diseases (Engler et al., 2008; Chaturvedi et al., 2010). It has been reported that in some cardiac-related diseases, heart ventricle is stiffer than the normal which causes heart failure. It was demonstrated that the muscle tissues of these hearts are stiffer than the regular heart muscle cells (Chaturvedi et al., 2010). Arterial stiffness is another risk factor for cardiac diseases such as stroke and heart failure (Cecelja and Chowienczyk, 2009; Mitchell et al., 2010). Arterial stiffness is usually developed by numerous factors such as aging, genetic factors, blood pressure, etc. (DeLoach and Townsend, 2008; Lacolley et al., 2009). High aortic stiffness leads to irregular mechanical properties of the arterial system, impaired cardiac performance, and difficulty in supplying blood to the heart and cardiac hypertrophy. Cell mechanics is also associated with liver diseases (Wells, 2008). The elastic modulus of healthy liver cells (0.4–0.6 kPa) could increase to as much as15 kPa following an injury (Georges et al., 2007). Different liver cells (stellate cells, hepatocytes, portal fibroblasts) develop changes in differentiation and proliferative characteristics when subjected to the stiffer matrix (Li et al., 2007). It has been reported that renal diseases are also associated with reduced stiffness renal glomerular podocyte cells. The glomerular podocytes from a mouse model of HIV-associated nephropathy are significantly softer than normal podocytes (Tandon et al., 2007). In this article, we will overview two advanced techniques, namely AFM and nanoindentation, as relevant methodologies for evaluation of mechanical properties of cells that will possibly play an important role in medical diagnostics as well as contribute towards understanding mechanisms of disease progression.

Theory Theory of AFM Force Curves In the AFM, the elasticity of living cells is determined from interactions of AFM probes with sample surfaces that are interpreted as a measure of Young’s modulus. AFM consists of three main parts, namely a cantilever, a system that detects its deflection and a system that enables scanning and positioning (Fig. 1). The most important part of an AFM is the cantilever. AFM cantilevers

Biomechanics j Biomechanics of Cells as Potential Biomarkers for Diseases: A New Tool in Mechanobiology

Fig. 1

3

Schematic showing different components of atomic force microscopy (AFM).

are made of silicon or silicon nitride with a sharp tip at the end. When the tip comes into proximity of a sample surface, forces between the sample and the tip cause a deflection of the cantilever. The deflection can easily be detected using an optical system comprised of a laser and a photodetector. The laser is focused on the free end of the cantilever, and when it is reflected, it goes directly towards the center of the photodiode. Deflection of the cantilever by interaction with features on the sample surface is monitored during scanning and is translated into a three-dimensional image of the surface. A force–displacement curve is also obtained from the plot of force (tip/sample) and deflection of the cantilever which can further be used to calculate mechanical properties of the sample. A typical force–displacement curve has three parts: approach, contact with the sample, and retraction. The approaching curve provides the information repulsive or attractive forces between tip and sample, whereas the second part sheds light on mechanical properties of the sample, such as Young’s modulus, relaxation time, etc. Finally, the retracting curve describes the adhesion forces that exist between tip and sample (Benitez and Toca-Herrera, 2014; Vanlandingham et al., 1997) (Fig. 2). Young’s modulus (E) of cells from the force–displacement (force-indentation) curve is estimated using Hertz model from contact mechanics which describes the indentation of two purely elastic spheres (Hertz, 1881). The model has been widely used to determine the apparent cell membrane elasticity. Two important assumptions of the model are the following: the indentation depth is less than  10% of the sample thickness, and the indentation depth is greater than 200 nm (Pelling et al., 2007; Rico et al., 2005; Stolz et al., 2004). The thickness of adherent cells varies from  200 nm at the cell edge to  5–10 mm over the center (nucleus) of the cell. Consequently, elastic modulus at the cell edge is substantially influenced by substrate’s (tissue culture polystyrene (TCPS) or glass) modulus and usually appears to be higher than that of the nucleus (Pelling et al., 2007). If the indenter is spherical the relation between the load force and the indentation depth is given by (Hertz, 1881): 4 1 3 F ¼ E0 R2 d2 3

(1)

where F is the load force, d is the indentation depth, and R is the radius of the spherical indenter, and E0 is the reduced Young’s modulus of the sample. The E0 is the reduced Young’s modulus of a sample, expressed by the following equation,

Fig. 2

@Schematic representation of a force–displacement curve.

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Biomechanics j Biomechanics of Cells as Potential Biomarkers for Diseases: A New Tool in Mechanobiology

 1  m2tip 1  m2sample 1 ¼ þ E0 Etip Esample

(2)

where mtip and msample are the Poisson ratios of the tip and a sample. It ranges from 0 to 0.5. For living cells, the elastic modulus is much smaller than Young’s modulus of the tip, that is, Ecell  Etip; therefore, the reduced Young’s modulus can be expressed as, E0 z

Ecell  1  m2cell

(3)

The value of mcell is assumed as 0.5 for the ease of calculation as cells can be treated as an incompressible material. Hertz model has been modified by Sneddon (1965), introducing axisymmetric shapes of the indenter (i.e., conical and pyramidal) into the relation between the load force and the indentation depth: Cone: 2 F ¼ E0 tanad2 p

(4)

E0 tanad2 O2

(5)

Pyramid: F¼

where a is the opening angle of the cone. The Hertz model does not take into account the attractive forces (van der Waals) within contact. The JKR theory considers the influence of finite surface energy (Johnson et al., 1971). The JKR model is ideal for tips with large curvature radius and small stiffness: F¼

Ka3  O6psKa3 R

where a is contact area radius, s is work of adhesion and K is effective Young’s modulus of the sample and is given by, 
# "
2 1  m2sample 1 3 1  mtip ¼ þ K 4 Etip Esample

(6)

(7)

The model assumes that adhesion forces (van der Waals) acting along the contact area perimeter results in an additional probe sample attraction which weakens forces of elastic repulsion (DMT) (Derjaguin et al., 1975). The DMT model is applied to tips with small curvature radius and high stiffness: F¼

Ka3  2pRs R

(8)

Operational modes of AFM AFM measures the repulsive (hard sphere) or attractive (van der Waals) interaction forces between the atoms at the proximity of a fine tip and the atoms at the sample surface (Binnig et al., 1986). Different operational modes of AFM are described below. In contact mode, the repulsive force between tip and sample is kept constant during the raster scan of the sample, providing a topographical image of the sample. In tapping mode, the cantilever oscillates at its resonant frequency. When the tip is brought into the proximity of the surface, the tip-sample interactions (involves atomic repulsions) reduce the amplitude of the oscillations (Putman et al., 1992) In non-contact mode, the cantilever either oscillates at either its resonant frequency (frequency modulation) or just above (amplitude modulation) where the amplitude of oscillation is typically a few nanometers down to a few picometers (Gross et al., 2009).

Theory of Nanoindentation Nanoindentation is a sophisticated technique which involves the application of a controlled load to penetrate an indenter into the surface of a specimen and uses recorded load–displacement curve to measure the hardness, elastic modulus, and, some other mechanical properties of the specimen. It is considered as a nondestructive technique as it imprints the sample at very shallow depths. Besides the elastic modulus, other mechanical properties such as fracture toughness, yield strength, hardening index, and residual stress, etc. can be measured using nanoindentation technique. It is essential to understand the basic principle of nanoindentation for an accurate interpretation of the experimental data. There are several review papers and book chapters on the basic principle of nanoindentation (Bhushan, 2010; Cheng and Cheng, 2004; Fischer-Cripps, 2006, 2011; Sharpe, 2008). At the later part of the 19th century, Hertz developed elastic equations of contact for contacting spherical surfaces (Hertz, 1881). To relate contact radius a, to the combined radius R, where R [ a, Hertz equation is: a3 ¼

3 PR 4 E

(9)

Biomechanics j Biomechanics of Cells as Potential Biomarkers for Diseases: A New Tool in Mechanobiology

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where, P is the applied force and E* is the elastic modulus. Love (1939) formulated an equation to relate the force P, implemented by the indenter and the resulting penetration depth or displacement h, as follows:
a P 2E tana 2 and P ¼ h (10) Pm ¼ 2 cos h1 pa r p Where a denotes the semi-angle of the conical indenter and the combined elastic modulus E*, defined regarding Young modulus E, and the Poisson ratio n, as follows: E ¼

E 1  n2

(11)

Sneddon who employed the integral transformation method to derive the same formula as (10) concluded the slope of a load– displacement curve for a linearly isotropic half space indented by an axisymmetric arbitrary profile as follows (Sneddon, 1965): dP 4E tana ¼ h dh p

(12)

where penetration depth h is related to contact depth hc as follows: p h ¼ hc 2

(13)

The relationship between h and hc, the height of the part of the indenter which is in contact with elastic half-space is illustrated in Fig. 3. Now for a conical indenter of half-apical angle a, the contact area is given by A ¼ pa2 ¼ ph2c tan 2 a ¼

4 2 h tan 2 a p

(14)

Now if we introduce Eq. (14) into Eqs. (10) and (12), P E cot a ¼ A 2

(15)

pffiffiffi. dP ¼ 2E A p dh

(16)

Although the tangent stiffness regarding projected contact area in Eq. (16) is derived for a conical indenter, it is also applicable for all asymmetric indenters such as cylindrical punches and spherical indenters (Bulychev et al., 1975; Pharr et al., 1992). However, in the case of the circular flat punch of radius a (Maugis, 2013), the relation between applied load P, and depth of penetration h, is as follows:  0:5 P r2 Pm ¼ 1  and P ¼ 2aE h (17) 2pa2 a2 where, A ¼ pa2. Now if we consider a rigid spherical indenter of radius R that is pressed into a linearly elastic specimen, then Hertz formula can be expressed (Johnson, 1985):  0:5 3P r2 4 pffiffiffi Pm ¼ 1  and P ¼ E Rh3 (18) 2pa2 3 a2 Now if a is expressed as the radius circular contact surface then a2 ¼ Rh and Eq. (18) certainly satisfies Eq. (16).

Fig. 3

Schematic of the nanoindentation of an elastoplastic solid by a conical cone at full load and unload.

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Biomechanics j Biomechanics of Cells as Potential Biomarkers for Diseases: A New Tool in Mechanobiology

Using Eqs. (10) and (12), reduced modulus of a linearly isotropic solid could be easily determined theoretically. However, in practice, it is very complex to determine elastic modulus using this simple equation from a load–displacement curve. The indenter with the three-sided geometry, known as Berkovich indenter, is the most popular indenter used to perform nanoindentation experiments. For a Berkovich indenter the contact area is calculated as follows: pffiffiffi A ¼ 3 3h2c tan 2 q (19) Where hc is the contact depth, and semi-angle q for a Berkovich indenter is 65.27. So from Eq. (14), we can obtain the relationship between contact area A and contact depth hc for a Berkovich indenter as follows: A ¼ 24:5 h2c

(20)

In most nanoindentation experiments, the indenter is in contact with a semi-infinite half space which makes R as the radius of the indenter alone but E* as the elastic modulus of the combined contacting bodies. Usually, the indenter is made of diamond which is strongly resistant to the deformation, and so for most materials, E* is dominated by the elastic properties of the specimen. This relationship is described in the Eq. (2): Further the experimental P–h curve is quite different from the parabolic P–h curve that is predicted by the Eq. (10). Fig. 4 shows the schematic diagram of a P–h curve for an elastoplastic specimen indented by a conical indenter. Here, maximum penetration (hmax) by the indenter is the combination of elastic penetration (he) and plastic penetration (hp) or hmax ¼ he þ hp. The equations presented above have some limitations to analyze a P–h curve obtained by a nanoindentation experiments as these equations do not consider several factors such as the tip has a finite radius, and indenter leaves a residual imprint on the specimen due to plastic deformations. So, Eq. (16) is modified as follows: rffiffiffi dP A rhmax ¼ 2bE (21) dh p where b is a dimensionless constant which accounts for the factors in a nanoindentation experiments which are ignored in the above equations and the value of which, is a debatable issue? Considering the factors of a nanoindentation experiments, Oliver and Pharr (1992) proposed an equation which is widely used in nanoindentation experiments (Oliver and Pharr, 1992): dP rhmax (22) hc ¼ hmax  bxPmax dh where ¼ 1 is the correction above factor, x z 0.72 for a conical indenter. Although the Oliver–Pharr method is quite popular in nanoindentation experiments, it has its limitations.

Modes of nanoindentation Dynamic nanoindentation: Dynamic nanoindentation is a mode of nanoindentation to determine the viscoelastic properties (Odegard et al., 2005; Loubet et al., 2000). It involves applying an oscillatory force to the indenter during static indentation. Dynamic properties can be determined by monitoring phase lag between applied displacement and measured force. The dynamic

Fig. 4

Schematic of a load–displacement curve corresponding to the nanoindentation.

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nanoindentation technique was reported by Loubet et al. for the very first time, who used this technique to determine the storage and loss modulus of the rubber polyisoprene (Loubet et al., 2000). Modulus mapping: Modulus mapping is another technique to investigate the loss modulus and storage modulus characteristics of materials by combining the aspects of nanoDMA and in situ SPM. In this technique, the indenter tip oscillates with a small force to monitor the phase lag due to material response as a result of displacement. The advantage of this technique is in situ SPM imaging enables the indenter tip to scan across the specimen to obtain a topographic image of the sample surface. In this way, modulus mapping allows performing dynamic indentation at each point of the surface. A large number of indentations can be carried out within a short period to obtain overall material characteristics like loss and storage modulus, loss and storage stiffness, tan delta maps, etc. Load control: In this mode of operation, one can specify the force applied to the indenter as a function of time, the displacement is achieved with that applied force, and the resultant load–displacement curve is used to determine the stiffness, elastic modulus, hardness and other mechanical properties. In a load control test, the load is increased at a particular rate. In load controlled operation, force path that includes loading and unloading can be applied. Load controlled nanoindentation operation is preferred for conducting experiments as one can apply constant or variable load rate and also the load can be held constant for a period of time. The forces mimicking real conditions can be more closely represented in this operational mode. Displacement control: In this mode of operation, displacement path is prescribed. At any time, displacement/position of the tip is known, and the force is determined to plot the force–displacement curve. In this mode, the softening behavior of the material (beyond the peak load) can be evaluated. This mode of operation is preferred when indentations to certain depths are to be performed. For example, if the thickness of the sample is known and the depth of penetration is to be limited to 10% of sample thickness. Alternatively, in the case where the interface between two materials is to be probed.

Examples of Nanomechanical Testing of Cells AFM Table 1 summarizes the results of Young’s modulus measurement for various types of living cells. The magnitudes indicate that the elastic modulus value of living cells varies over a wide range. It has long been established that external conditions influence the elasticity of cell or cell membrane. Hence, the first factor that brings about changes in the elasticity of cells is AFM sample handling. Erythrocytes treated with 5% formalin showed a 10-fold (119.5  15 kPa) increase in their elastic modulus compared to native erythrocytes (16.05  2.3 kPa) (Mozhanova et al., 2003). The transverse stiffness of cardiomyocytes is also increased by a factor of 16 after fixing with formalin (Shroff et al., 1995). A study involving AFM to probe leukemia cells placed in microwells demonstrated the influence of microwells on the elastic modulus of the cell (Rosenbluth et al., 2006). The cell deformation was described using Hertzian contact mechanics. The second factor is related to the heterogeneity of mechanical properties of cells. There are significant variations in the values of elastic modulus at different cell regions. The elastic modulus value of human umbilical vein endothelial cells was 7.22  0.46 kPa over the nucleus; 2.97  0.79 kPa over the area away from the nucleus towards cell edge, and 1.27  0.36 kPa at the cell edge (Mathur et al., 2000). A similar study on cardiomyocytes revealed that cells are softer at the nuclear region, and become stiffer towards the periphery (Shroff et al., 1995). Elastic modulus values of chicken cardiomyocytes’ soft region between 5 and 30 kPa, with stress fibers favored region showing stiffness value of 100–200 kPa are reported (Hofmann et al., 1997). The elastic modulus of the surface layer of living astrocytes ranged from 1 to 40 kPa, being influenced by the inner structure of cells such as nucleus and F-actin. On the other hand, the elastic modulus of fixed cells was relatively uniform (200–700 kPa), irrespective of the inner structures of cells (Yamane et al., 2000). The third factor influencing elastic modulus measurement relates to cell thickness. Substrate contributions can be neglected if AFM tip does not indent > 10% of the cell thickness (Mathur et al., 2001). The elastic modulus of cells has also been found to play an important role in delineating functions of various cell types, some examples of which are presented below. Some of the earliest AFM studies on living aortic endothelial cells were carried out to study the effects of shear stress on cellular organization and other factors which may influence the mechanical response of cells to flow (Barbee et al., 1994; Sato et al., 2000). It was reported that the average elastic modulus of bovine pulmonary artery endothelial cells (BPAECs) in the range 0.2–2.0 kPa (Pesen and Hoh, 2005). A similar study showed the difference in mechanical properties of rabbit endothelium, that is, cells were found to be stiffer in the medial wall of aortic bifurcation than in the lateral wall (Miyazaki and Hayashi, 1999). The elastic modulus of activated platelets was found to be in the range of 1–50 kPa using force mapping techniques (Radmacher et al., 1996). A few AFM studies have also been carried out to delineate the mechanistic mechanism of cell migration using fibroblasts (Haga et al., 2000; Nagayama et al., 2001). A relation between the local stiffness distribution on cell and cell migration was reported, that is, stiffness distribution of cell surface was constant until cells started to move, a sharp decrease in the stiffness in the nuclear region of the cells was observed upon movement (Nagayama et al., 2001). It has been reported in previous studies that mechanical properties of the lamellipodium are important for a deeper understanding of cell migration mechanism. It was demonstrated that the rigidity of fibroblast cell body was due to the over-condensation of the actin network, and hardening of the cell cortex (Haga et al., 2000). An identical study on fish epidermal keratocytes showed rigidity of the cell body as a function of distance from cell edge, that is, rigidity was the highest at the cell edge, and it gradually decreased towards the cell center. Also, the change in rigidity was not perturbed by the vertical lamellipodia thickness (Laurent et al., 2005). AFM has also been found to be effective to study cell mechanical properties during cell adhesion. There are some AFM studies which characterized osteoblast elastic properties during adhesion (Domke et al., 2000; Simon et al., 2003). The effect of substrate on

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Biomechanics j Biomechanics of Cells as Potential Biomarkers for Diseases: A New Tool in Mechanobiology Table 1

Summary of the AFM based experiments carried out to evaluate elastic modulus for various cell types

Cell type

Young’s modulus, E (kPa)

References

Rat astrocytes Endothelial cells HUVEC –

2–20

Yamane et al. (2000)

10–11 1.3–7.2

Sato et al. (2004) Mathur et al. (2000)

0.2–2.0 0.3–20 2.1–8.8 4–5 4–5 1.86  1.13 10–55

Pesen and Hoh (2005) Simon et al. (2003) Domke et al. (2000) Bushell et al. (1999) Wu et al. (1998) Nikkhah et al. (2010) Laurent et al. (2005)

0.2–1.4 0.02–0.08 0.2–0.07 1.24  0.09 0.51  0.06 1–50 19–33 0.1–0.2 0.2–0.4 1–1.4 0.6–0.7

Rosenbluth et al. (2006)

11–45 28–21 40–45 95–100 5–200

Collinsworth et al. (2002) Mathur et al. (2001) Yoshikawa et al. (1999) Mathur et al. (2001) Hofmann et al. (1997)

2.797  0.491 1.401  0.162 0.287  0.052 1.139  0.320 3.09  0.84 0.45  0.21 1.95  0.47 1.36  0.42 0.6  0.4

Faria et al. (2008)

1.13  0.44 0.63  0.22 2.26  0.56 1.24  0.46 1.20  0.28 1.75  0.12 1.3  0.15 0.8  0.05 1.13  0.84 0.51  0.35

Li et al. (2008)

9.7  3.6 7.5  3.6 0.3  0.2 0.8  0.4 27.57 2.46 5.7  0.45 1.9  0.26 2.7  0.24

Lekka et al. (1999)

BPAEC Osteoblasts SaOS2 Fibroblasts L929 HS68 Epidermal keratocytes Blood Leukemia myeloid cells (HL60) Leukemia lymphoid (Jurkat) cells Neutrophils Lymphocytes Leukemia lymphoid Jurkat cells Platelets Erythrocytes Red blood cells (RBCs) Raji Hut K-562 Skeletal muscle cells Murine C2C12 myoblasts – Myofibrils Cardiocytes Chicken Prostate cells BPH LNCaP PC-3 PNT2-C2 PZHPV-7 LNCaP PC-3 DU145 DU145 Breast cells MCF10A MCF7 184A1 MCF7 T47D MCF10A MCF7 MDA-MB-231 MCF10A MDA-MB-231 Bladder cells HU609 HCV29 HU456 T24 SV-HUC-U1 MGH-U1 RT112 T24 J82

Cai et al. (2010) Radmacher et al. (1996) Dulinska et al. (2006) Li et al. (2012)

Lekka et al. (2012)

Efremov et al. (2015)

Lee et al. (2012) Nikkhah et al. (2010)

Canetta et al. (2014) Abidine et al. (2015)

Biomechanics j Biomechanics of Cells as Potential Biomarkers for Diseases: A New Tool in Mechanobiology Table 1

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Summary of the AFM based experiments carried out to evaluate elastic modulus for various cell typesdcont'd

Cell type Mesothelial cells Normal Cancerous Cervical cells CRL2614 CaSki End1/E6E7 HeLa Thyroid cells S748 S277 Ovarian cells MOSE early MOSE-intermediate MOSE-late IOSE OVCAR4 HEY OVCAR3 HEYA8 Keratinocytes HaCaT HaCaT

Young’s modulus, E (kPa)

References

1.97  0.70 0.53  0.10

Cross et al. (2008)

2.58  0.08 2.03  0.14 4.8  0.8 2.46  0.6

Zhao et al. (2015)

2.2–6.8 1.2–1.4

Prabhune et al. (2012)

1.097  0.682 0.796  0.441 0.549  0.281 2.472  2.048 1.120  0.865 0.884  0.529 0.576  0.236 0.494  0.222

Prabhune et al. (2012)

84.5 77–109

Fung et al. (2011) Heu et al. (2012)

Hayashi and Iwata (2015)

Xu et al. (2012)

cell adhesion using osteoblast cell was studied (Domke et al., 2000). Osteoblasts cultured on different substrates (CoCr, Ti, TiV, glass, and TCPS) exhibited the elastic modulus values ranging from 2 kPa (CoCr and TiV) to 9 kPa (Ti surface). The latter modulus was found to be in good agreement with the modulus obtained from osteoblasts cultured on TCPS. Another study on the osteoblasts demonstrated the influence of cytoskeletal reorganization and state of the cell on the mechanical properties of the cell. The elastic modulus of osteoblasts was found to be in the range of 0.3–200 kPa depending upon the type of substrate the cells were cultured on (Simon et al., 2003). The changes in mechanical properties and cytoskeleton reorganization have been found to correlate well with cell cycle stages in the following studies (Berdyyeva et al., 2005; Collinsworth et al., 2002; Sato et al., 2004). These results have paved the way for understanding the mechanism behind cell differentiation and aging. A study on human umbilical vein endothelial cells (HUVEC) revealed that the culture period influences the mechanical properties of cells (Sato et al., 2004). It was reported that epithelial cells cultured on type IV collagen for longer than 4 days showed average elasticity values > 10 kPa (Berdyyeva et al., 2005). A similar study on mouse skeletal myocyte cells hypothesized cell differentiation as a key regulator of transverse elastic properties of the cells. They observed a significant increase (from 11.5  1.3 to 45.3  4.0 kPa) in elastic modulus of cells upon cell differentiation over the course of 8 days (Collinsworth et al., 2002). Actin and myosin were found to be major contributors to changes in the transverse elastic modulus. AFM studies on single cells have shown that depending on the cell type, either actin filaments or microtubules have a strong influence on the mechanics of cells. There are a few studies where living cells are subjected to incubation with cytoskeletondisrupting drugs to identify the type of cytoskeletal elements that dominate the mechanical response (Cai et al., 2010; Heu et al., 2012; Hofmann et al., 1997; Lulevich et al., 2010; Rotsch et al., 1997; Rotsch and Radmacher, 2000; Spedden et al., 2012; Wu et al., 1998). The chemical disassembly of the rat liver macrophages’ actin filaments by applying cytochalasin B and latrunculin was studied (Rotsch et al., 1997). Treating cells with cytochalasin B resulted in a sevenfold decrease in elastic modulus after 40 min. On the other hand, latrunculin treatment induced a twofold decrease in the elastic modulus of the perinuclear region after 40 min, whereas other parts of cell remained intact. The latrunculin-induced disruption of cytoskeleton network of various fibroblast cell lines was reported by two different groups (Braet et al., 2001; Rotsch and Radmacher, 2000). A decrease in the elastic modulus of chicken cardiocytes (Hofmann et al., 1997) and L929 cells (Wu et al., 1998) was observed upon treating them with cytochalasin B and cytochalasin D, respectively. In contrast, treatment of fibroblasts with nocodazole or colcemid resulted in an increase in their elastic modulus. Another study used cytochalasins B and D, latrunculin A and Jasplakinolide to show disruption of actin filaments results in a decrease in the elastic modulus of fibroblasts, while reorganization of microtubules does not affect cell elastic modulus (Rotsch and Radmacher, 2000). Cai et al. (2010) showed that cytochalasin B induced 70% and 60% decrease in the elastic modulus of lymphocyte and Jurkat cells, respectively. A similar study on keratinocytes reported that latrunculin and nocodazole treatment not only altered the cell elastic modulus but also treated cells appeared softer than control (Fig. 5). They observed around 50% and 27% decrease in elastic modulus of cells treated with latrunculin and nocodazole, respectively (Lulevich et al., 2010). In a recent study, glyphosate has been shown to induce cell membrane stiffening and the emergence of cytoskeleton structures at the

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Biomechanics j Biomechanics of Cells as Potential Biomarkers for Diseases: A New Tool in Mechanobiology

Fig. 5 Effect of latrunculin A on keratinocytes, (A) characteristic force versus relative deformation curve for keratinocyte cells treated with latrunculin A (blue, dashed) as compared to control (red, solid), immunocytochemistry for actin (green) in (B) control and (C) latrunculin A-treated keratinocyte, (D) and (E) are optical micrographs of a latrunculin A-treated keratinocyte before and after compression, respectively. Reprinted with permission from Lulevich, V., Yang, H. Y., Isseroff, R. R. and Liu, G. Y. (2010). Single cell mechanics of keratinocyte cells. Ultramicroscopy 110, 1435–1442.

subcellular scale at low concentration (15 mM) (Fig. 6). However, HaCaT cells showed a flattened external cell membrane along with reduced number of protrusions. Also, quercetin was found to reverse the effects of glyphosate (Heu et al., 2012). Spedden et al. (2012) showed an increase by > 30% in the elastic modulus of neuronal cell bodies upon Taxol (10 mM) treatment, while nocodazole (10 nM) did not induce any substantial change in cell body elastic modulus. AFM estimation of cell mechanical properties has shown the potential for the diagnostics of different pathologies in the past decade as it facilitates the measurement of the

Fig. 6 Effect of glyphosate on HaCaT keratinocytes, (i) topographical changes, (ii) change in mechanical properties, and (iii) induced deformation for control cells (A) and treated cells (B) at 15 mM glyphosate for a 6 h–incubation. The color scales are (from bottom to top): 0–1 nN, 0–500 kPa, and 0–250 nm for the peak force error, Young’s modulus and deformation images, respectively. Reprinted with permission from Heu, C., Berquand, A., Elie-Caille, C. and Nicod, L. (2012). Glyphosate-induced stiffening of HaCaT keratinocytes, a peak force tapping study on living cells. Journal of Structural Biology 178, 1–7.

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influence of various factors on the mechanical properties of the living cells. A comparative study of normal and cancerous human bladder cells showed that normal cells’ elastic modulus was found to be one order of magnitude higher than cancerous counterparts (Lekka et al., 1999). In a later study, (Lekka et al., 2001) showed a strong correlation between the reduction of glycolytic activity and the increase in Young’s modulus values obtained for the bladder cancer cells treated with chitosan. An AFM-based cytomechanical analysis showed that cancer cell stiffness is 70%–80% less than that of normal cells (Cross et al., 2008) Using AFM and confocal fluorescence microscope, (Li et al., 2008) demonstrated that MCF-7 (malignant breast) cells had an apparent Young’s modulus significantly lower (1.4–1.8 times) than that of MCF-10A (normal breast epithelial) cells and their apparent Young’s modulus increased with loading rate. A similar study on various prostate cells (BPH, LNCaP, and PC-3) revealed that BPH had higher Young’s modulus among three cell lines examined. Interestingly, the highly invasive PC-3 cells were found to have a higher Young’s modulus than the non-invasive LNCaP clone FGC cells, showing that there are variations between different types of prostate cancer cells, a feature which is reflected in their clinical behavior (Faria et al., 2008). The measurement of cell’s elastic modulus using AFM indentation revealed that HS68 (normal fibroblast cells) cells are significantly stiffer than MCF-10A and MDA-MB-231 (breast cancer) cells. Upon microtubule disruption with nocodazole, fibroblasts no longer stretched, but adhesion of MCF-10A and MDA-MB-231 within the etched features remained unaltered (Nikkhah et al., 2010). Recent studies on ovarian cancer cells have demonstrated cancerous ovarian cells are softer than their normal counterparts but possess higher motility whereas ovarian cancer cells with lower migratory or invasive potential are almost five times stiffer than the highly invasive ones (Swaminathan et al., 2011; Xu et al., 2012) Analogous to other cancers, thyroid cancer cells were found to be three to fivefold more deformable than normal thyroid cells (Prabhune et al., 2012). Lekka et al. (2012) studied both breast and prostate cancer cells using AFM indentation. The calculated elastic modulus values for the presented results of prostate cancers were 0.45  0.21, 1.36  0.42, and 1.95  0.47 kPa for LNCaP, Du145, and PC-3, respectively. However, the elastic modulus calculated for PZHPV-7 cells was 3.09  0.84 kPa. For breast cancers, the results were 1.20  0.28 kPa, 1.24  0.46 kPa, and 2.26  0.56 kPa for T47D, MCF7, and 184A cell lines, respectively. These results showed that the normal cells were characterized by larger Young’s modulus values, irrespective of the cell types (Fig. 7A). It also indicates their lower ability to deform compared with the normal cells coming from later stages of cancer progression. Since indentation depth governs cells’ elastic modulus, hence, the dependence of elastic modulus on indentation depth was studied. For instance, the largest indentation depth resulted in the lowest modulus value for 184A and MCF-7 cells (Fig. 7B). The loading rate has been known to influence elastic modulus. Therefore, the influence of displacement rate was examined for the human breast cells (Fig. 7C). It demonstrates an increase in modulus of about 48% and 157% for cancerous MCF7 and normal 184A cells, respectively, for data recorded at 0.5 mm/s as compared with those recorded with a velocity of 10 mm/s. The influence of prolonged indentation on a single cell was also determined using the modulus measurements obtained in T47D cells (Fig. 7D; black dots). 232 curves were recorded over an area of 10  10 mm. However, no influence of time over modulus measurements was observed. When indentation is carried out at the constant position (Fig. 7D; open squares), the continuous indentation influences the measurements. For instance, for the same T47D breast cancer cells, such experiments ended quickly at the breakdown at the curve No. 145. AFM and modulated Raman spectroscopy (MRS) have been used to discriminate between normal human urothelial cells (SV-HUC-1) and bladder cancer cells (MGH-U1). The results demonstrated that MGH-U1 cells were 1.5-fold smaller, 1.7-fold thicker and 1.4-fold rougher than normal SV-HUC-1 cell and the adhesion energy was 2.6-fold higher in the MGH-U1 cells compared to normal SV-HUC-1 cells, indicating higher deformability of bladder cancer cells. Also, the elastic modulus of MGHU1 cells was found to be 12-fold lower than that of SV-HUC-1 cells (Canetta et al., 2014). In another study, AFM-based nanomechanics and fluorescence based intracellular calcium dynamics studies were performed on poorly (LNCaP) and highly (CL-1, CL-2) metastatic human prostate cancer cells. The elastic moduli and calcium dynamics were found to be greater in CL-1 and CL-2 than LNCaP. The authors suggested that the enhanced elastic moduli and calcium dynamics were observed due to the intensified tensile stress generated by actin cytoskeletons anchored at more focal adhesion sites (Bastatas et al., 2012). In a recent study, the effects of actin on the mechanics of normal Vero and prostate cancer cell line DU145 was determined using AFM before and after treating the cells with specific nucleation inhibitors (SMIFH2, CK-666), cytochalasin D, Y-27632 and trypsin. The elastic modulus of Vero cells was found to be around 1.3  0.9 kPa, while the prostate cancer cell DU145 exhibited significantly lower modulus (0.6  0.4 kPa). Interestingly, cancer cells exhibited diverse viscoelastic behavior and different responses to actin cytoskeleton reorganization upon drug treatment (Efremov et al., 2015). It is shown that the elastic modulus values of late-stage mouse ovarian surface epithelial (MOSE) cells (0.549  0.281 kPa) were significantly less than that of their early-stage ones (1.097  0.632 kPa) (Ketene et al., 2012). Apparent cell viscosity was also found to be reduced significantly from early (144.7  102.4 Pa s) to late stage (50.74  29.72 Pa s), indicating higher stiffness and viscosity of normal ovarian cells. The increase in cell deformability was found to correlate directly with the transition of benign phenotype to malignant one. The decrease in the level of actin in the cytoskeleton and its organization facilitated the changes in cell biomechanical property. Using spherical probes, (Li et al., 2012) investigated the mechanical properties of RBCs, Raji, HuT, and K562 cells. The elastic modulus was found to be 0.1  0.2, 0.2  0.4, 1  1.4, and 0.6  0.7 kPa, for RBCs, Raji, HuT, and K562 cells, respectively, indicating aggressive cancer cells are softer than normal cells. A similar study on HeLa cells (human cervical cancer cell) and End1/E6E7 cells (squamous epithelial cell from normal human cervix) showed that cancer cells were softer than normal cells, and there were no significant locational differences in the stiffness of cancer cells between the central and the peripheral regions (Hayashi and Iwata, 2015). A mechanical and adhesive mismatch between transformed and non-transformed cells in a cell monolayer could trigger enhanced pulsating migration using various breast cell lines (Lee et al., 2012). They also examined the influence of neighboring stiff epithelial cells on cancer cells in the early steps of cancer progression. A recent study has shown that the elastic modulus and the transition frequency can be used as markers of invasiveness for cancer cells (Abidine et al., 2015). The longitudinal elastic modulus

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Fig. 7 (A) Comparison of the stiffness of prostate and breast cancer cells, (B) influence of indentation depth, (C) force load on the elastic modulus of breast cells (184A and MCF7), and (D) effect of prolonged indentation on breast cancer cell (T74D). Reprinted with permission from Lekka, M., Gil, D., Pogoda, K., Dulinska-Litewka, J., Jach, R., Gostek, J., Klymenko, O., Prauzner-Bechcicki, S., Stachura, Z. and Wiltowska-Zuber, J. (2012). Cancer cell detection in tissue sections using AFM. Archives of Biochemistry and Biophysics 518, 151–156.

distribution of human cervical squamous carcinoma cells (CaSki) and normal cervical epithelial cells (CRL2614) was found using AFM. The results indicated that CaSki cells were less stiff than their normal counterparts, and the nuclear region appeared to be softer. Also, authors have shown a correlation between heterogenicity in longitudinal elastic modulus and indentation depth. For instance, CaSki cells, with a modulus of 0.35–0.47 kPa, was found to be located at 237–225 nm; while normal cells showed an elastic modulus of 1.20–1.32 kPa at 113–128 nm (Zhao et al., 2015).

Nanoindentation In this section, we discuss some examples of measuring cellular nanomechanics using nanoindentation technique. For the very first time, mechanical properties of soft tissue using nanoindentation technique was measured by Ebenstein et al. in 2004 (Ebenstein and Pruitt, 2004). The purpose of the study was to develop a methodology to keep the sample hydrated, select an appropriate indenter for soft tissue like material indentation, determine an appropriate control material for the development of future indentation protocols and identify a substrate to be used for blunt tip alignment. They developed a hydration system to keep the sample hydrated for > 8 h without completely submerging the sample. To identify appropriate indenter for the hydrated, soft tissue, they performed indentation with different types of indenters. Three sided sharp Berkovich indenter with an average radius of curvature of 100–200 nm and conosperical diamond indenters with radius of 10, 50, 100, and 800 mm radius of curvature were used. They found that conosperical tip with a 100-mm radius of curvature is most suitable for the indentation of the vascular tissue sample and agarose gel is most effective for the tip alignment. They used Oliver and Pharr method to calculate reduced elastic modulus and found that artery tissue has an elastic modulus of z 0.73 MPa. In orthopedic research, cartilage repair is one of the main challenges. To address this challenge they studied the mechanical properties of cartilage repair tissue using nanoindentation experiments (Ebenstein et al., 2004). They reported that the mechanical properties of cartilage inferior tissue are less than the control cartilage

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tissue. In another study, nanoindentation technique to study the mechanical properties of fibrous tissue, blood clots, partially calcified fibrous tissue, and bulk calcifications from human atherosclerotic plaque tissue (Ebenstein et al., 2009). The goal of that study was to investigate the mechanical changes in plaques tissues during atherosclerotic disease progression as clinical events such as stroke, and heart attack can be caused by the rupture of atherosclerotic plaques in blood vessel walls. They demonstrated that the stiffness of the plaque tissue increases with the increase in the mineral content in the tissue. In bone tissue engineering, the interaction between bone cells and engineered biomaterial is an important area to explore. We studied the nanomechanical properties of bone cells and bone cell-substrate construct using nanoindentation technique under physiological condition. Cells were seeded on tissue culture treated polystyrene (TCPS) (Khanna et al., 2011). Once cells are attached, cell seeded TCPS was placed in a fluid cell submerged in cell culture media and then nanoindentation experiment carried out by placing the sample on the temperature controlled stage of nanoindentation device. In displacement control mode operation, cells demonstrated completely elastic load-deformation response (Fig. 8) and cell elastic modulus was found 1.3–12.4 MPa using Oliver and Pharr method. In another study, we reported a novel in situ nanoindentation technique developed to evaluate the composite mechanical behavior of cell-biomaterial construct under physiological conditions over the time scale of bone nodule generation (Khanna et al., 2012). The schematic illustration of cell-substrate indentations on a single cell attached to Chi-PgAHAP film deposited onto TCPS substrate is shown in Fig. 9. Nanomechanical properties of cells over the time of adhesion, proliferation, development and bone nodule formation were evaluated (Fig. 10). The study implied that unique interactions between cells and nanocomposite films provided a favorable mechanical environment for the formation of bone nodules. In situ quasi-static and dynamic nanoindentation tests on calcified nodules formed by mouse osteoblasts to investigate the effects of glucocorticoid hormones such as dexamethasone and hydrocortisone on the nanomechanical properties have been conducted (Miyamoto et al., 2015). How dexamethasone affects the hardness, elastic modulus and storage modulus of the calcified nodules formed by mouse osteoblasts is illustrated in Fig. 11. The proliferation or calcium deposition of the cells were not affected by dexamethasone, but nodules formed in the presence of dexamethasone were significantly stiffer than the nodules formed in the absence of dexamethasone. Their result suggested that hormones like dexamethasone are essential for in vitro formation of more mature calcified nodules by bone cells.

Fig. 8 Representative load–displacement curves obtained on live osteoblast. (A) Cells exhibit complete elastic recovery upon indentation as indicated by linear loading and unloading curves (B). The first portion which is flat shows the cell indentation response and steep loading slope beyond 1250 nm displacement show the stiffer response due to TCPS substrate lying underneath the cell. The flat portion of the loading and unloading curve in (B) are replotted separately in (C) and (D) respectively. The figure is adapted from Khanna, R., Katti, K. S. and Katti, D. R. (2011). Experiments in nanomechanical properties of live osteoblast cells and cell–biomaterial interface. Journal of Nanotechnology in Engineering and Medicine 2, 041005 and reprinted with permission.

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Fig. 9 Schematic showing cell-substrate indentations on a single cell attached to Chi-PgA-HAP film deposited onto TCPS substrate and sectional view of an indenter.

Fig. 10 In situ nanomechanical responses of osteoblast cells and cell-substrates obtained by nanomechanical experiment (A–F). Mechanical responses of cells and cell-CPH composites are completely reversible as shown by load–displacement curves (A and B).

In an attempt to map the elastic properties of soft tissues, nanoindentation experiments are performed to investigate the micromechanical properties of 5-mm-thick sections of ferret aorta and vena cava and to relate these mechanical properties to the histological distribution of fluorescent elastic fibers (Akhtar et al., 2009). They reported the elastic modulus of the blood vessel tissues varies from 8 MPa to 35 MPa depending on a different layer of the tissue using extended Oliver and Pharr method and 10 mm conosperical indenter. They demonstrated that it is possible to distinguish the differences in the micromechanical properties of blood

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Fig. 11 Nanoindentation technique is used to determine the hardness and elastic modulus of calcified nodules formed by MC3T3-E1 cells in the presence and absence of dexamethasone. Representative hardness (upper panels) and storage modulus (lower panels) data versus contact depth of calcified nodules formed in dexamethasone negative medium (A) and dexamethasone positive medium (B). Hardness (C) and elastic modulus (D) obtained in quasistatic nanoindentation tests and hardness (E) and storage modulus (F) obtained in dynamic nanoindentation tests for calcified nodules formed in cultures with and without dexamethasone. The figure is adapted from Miyamoto, S., Miyamoto, Y., Shibata, Y., Yoshimura, K., Izumida, E., Suzuki, H., Miyazaki, T., Maki, K. and Kamijo, R. (2015). In situ quasi-static and dynamic nanoindentation tests on calcified nodules formed by osteoblasts: Implication of glucocorticoids responsible for osteoblast calcification. Acta Biomaterialia 12, 216–226 and reprinted with permission.

vessels, which are related to the tissue microstructure. Continuous depth-sensing nanomechanical characterization of living, fixed and dehydrated cells attached on a glass substrate using a dynamic contact module are conducted (Yang et al., 2010). The nanomechanical characteristics of the cells were reported as the continuous harmonic contact stiffness (HSC). In this study, they tried to understand how the underlying substrate influences the interpretation of the nanomechanical property of thin soft matter on a hard substrate (Table 2).

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Table 2

Summary of the nanoindentation based nanomechanical experiments carried out to evaluate elastic modulus for different cells and tissues

Cell/tissue

Indenter

Method

Elastic modulus

Reference

Mineralized matrix formed on glass coverslips by immortalized cell line Y201 from human mesenchymal stem cells Calcospherulite crystals produced by rat osteoblasts Cartilage repair tissue Porcine aorta tissue Human carotid bifurcation plaque tissue Osteoblast Osteoblast Nodule formed by mouse osteoblastic MC3T3-E1 cells Embedded cross section of aorta and vena Cava tissue 3T3 mouse embryonic fibroblast cells

Berkovich

Oliver and Pharr

10–45 GPa

Duan and Chen (2015)

Berkovich 100 mm conosperical 100 mm conosperical 100 mm conosperical Berkovich Berkovich Berkovich 10 mm conosperical

Oliver and Pharr Oliver and Pharr Oliver and Pharr Compliance method Oliver and Pharr Oliver and Pharr Oliver and Pharr Oliver and Pharr

0.41  0.15 GPa Not reported 0.7–0.8 MPa 0.05–43.4 MPa 1.1–12.0 MPa 1.3–12.4 MPa 10 GPa 8–35 MPa

Chen et al. (2010) Ebenstein et al. (2004) Ebenstein and Pruitt (2004) Ebenstein et al. (2009) Khanna et al. (2012) Khanna et al. (2011) Miyamoto et al. (2015) Akhtar et al. (2009)

Berkovich

Reported as continuous harmonic contact stiffness (HCS)

Yang et al. (2010)

Perspectives Some Current Challenges in Use of AFM Here we discuss the fundamental limitations of AFM such as scan range, scan speed, piezoelectric transducer nonlinearity, probe size and shape, and sample size. An AFM often scans a maximum area of about 150  150 mm and a maximum height on the order of 10–20 mm. On the other hand, scanning electron microscope can image an area on the order of square 1  1 mm with a depth of field on the order of millimeters. The scanning speed of an AFM is also a limitation. Usually, an AFM requires several minutes for a typical scan, and the relatively slow rate of scanning often leads to the formation of thermal drift in the image. It makes AFM less suited for measuring accurate distances between topographical features on the image (Lapshin, 1998, 2004) AFM images can also get adulterated by nonlinearity, hysteresis, and creep of the piezoelectric material and cross-talk between the x, y, z-axes. Theoretically, piezoelectric transducers allow convenient positioning at or below the nanometer scale. However, the position is not a single-valued, linear function of the applied voltage, so, if a certain point on the sample is in the center of the image as the AFM scans in the þ x direction, it may not be at the image center when the AFM scans in the  x direction (Lapshin, 1995). Moreover, scanner response is also subjected to change depending on scan frequency, temperature, and age of the scanner. AFM imaging often induces the formation of image artifacts by either of the following factors; an incompatible tip, substandard operating environment, or even by the sample itself (Velegol et al., 2003) Image artifacts are unavoidable. However, their effect on results can be lessened through various methods. Small sample size seems to be another limitation that AFM suffers from. The overall dimensions of the samples to be measured by AFM are small so that they can be easily mounted on the sample holder. The sample holder containing sample is moved in x, y, and z-axes by a piezoelectric transducer. Consequently, even small deflections of a stationary cantilever are counted as probe signal. Another limitation of AFM is the size and shape of the probe itself. It comes into play when the features to be imaged have high aspect ratios. For instance, a probe that is large or blunt will easily measure very flat surfaces without much loss of information, but will not be able to trace the true profile of a surface that includes high aspect ratio features that are sharper, or of smaller dimensions than the probe.

Current Challenges in Nanoindentation of Live Cells Challenges of nanoindentation experiments on biological soft samples include problem associated with instrumentation, optimization of indenter geometry and load–displacement range, an appropriate model for analyzing experimental data, maintaining hydration state of the biological samples, etc. Commercially available instruments are usually designed for time independent materials, but the mechanical response of biological materials such as cells is mostly time-dependent. Due to the low elastic moduli of several biological samples such as cells the evaluation of force–displacement curves for soft samples is challenging. Indentation tip force is directly related to the tip geometry and material properties of the specimen. If a sharp indenter is used for indentation on soft tissues, the contact area and the corresponding force is very small for a shallow indentation. It makes it very complicated to determine when the indenter touches the sample which leads to a zero-point error (Kaufman and Klapperich, 2009). Another nanoindentation instrumentation challenge for biological materials like cells or tissue is maintaining hydration state at the physiological condition. For this purpose, the Katti research group has used cell culture media in a fluid cell on the temperature controlled stage, but it is a still a challenge to maintain physiological pH, as no CO2 was supplied to the cell culture media, thus limiting the time for conducting the tests. (Khanna et al., 2011, 2012). Using fluid to hydrate the sample and temperature controlled stage might introduce further experimental complexities which have not been sufficiently addressed in the literature. Most popular indenter for nanoindentation experiments is three sided Berkovich tip. The Berkovich indenter has been used for soft tissues and glossy polymers (VanLandingham, 2003; Habelitz et al., 2001; Rho et al., 1999; Larsson and Carlsson, 1998). Sharp indenters such as Berkovich

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can introduce plastic deformation and stress concentration as well as damage the specimen. In an attempt to minimize this problem, blunt tips such as spherical, conosperical and cylindrical punch have been used (Larsson and Carlsson, 1998; Gerberich et al., 1998; Ebenstein et al., 2004, 2009; Balooch et al., 1998; Lundkvist et al., 1997). The disadvantage of using flat tips is that it has a high-stress concentration at the contact perimeter. In addition to the indenter geometry, the size of the indenter needs to be considered as well. Indenters with small radius can provide nanomechanical properties of a single cell meanwhile indenter with a larger diameter will provide information on nanomechanical characteristics of the whole tissue including extracellular matrix. Most of the commercially available nanoindentation instruments are designed with software to analyze experimental data using Oliver and Pharr method (Oliver and Pharr, 1992). This approach works extremely well for the specimen with a time-independent mechanical response. However, this method introduces significant errors to calculate elastic modulus for samples with a viscoelastic time-dependent response like soft tissues as the assumption of elastic unloading is invalidated by the continuing time-dependent deformation (Oyen and Cook, 2003; Oyen, 2013). Creep or sinking of tip under constant force into the specimen is the most common effect of viscoelasticity. For viscoelastic specimen, a “nose” is observed on the unloading curve when there is no hold time which results in a negative slope at the beginning of the unloading curve and makes it very difficult to measure the elastic modulus (Briscoe et al., 1998). Some correction has been proposed to minimize the creep effect in calculating elastic modulus of the viscoelastic materials (Tang and Ngan, 2003; Feng and Ngan, 2002; Ngan et al., 2005). Some other models have been developed to analyze the nanoindentation experimental data obtained from soft tissue like viscoelastic materials (VanLandingham, 2003; Oyen, 2005; Lu et al., 2003; Oyen and Cook, 2003). Another problem associated with indenting soft tissue samples is the adhesion between the indenter and the sample (Carrillo et al., 2005, 2006). It has been reported that modulus calculation is significantly affected due to adhesion between the tip and the soft tissues (Carrillo et al., 2005, 2006). JKR (Johnson et al., 1971) model has been reported to be more accurate to minimize the contact effect between the soft samples and the indenters while measuring elastic modulus (Ebenstein and Wahl, 2006; Carrillo et al., 2005, 2006). More sophisticated approaches have also been reported combining JKR model and viscoelastic analysis (Ebenstein and Wahl, 2006; Greenwood and Johnson, 2006; Wahl et al., 2006).

Future Perspectives Overall, as instrumentation capabilities improve, including improved and reliable feedback mechanisms, the capability to conduct AFM force curve experiments and nanoindentation is likely to become routine and thus present itself as a new and novel mechanistic approach to understanding disease progression as well as diagnostics. The combination of experiments and modeling will also play an important role in improving the accuracy and interpretation of experimental data. Next generation medical treatments and research will see a large increase in the evaluation of mechanics using these techniques. The critical advances will arise from making implicit connections between gene expression and other traditional biochemical assays evaluated for disease diagnostics and mechanisms and mechanics of the cell during the time evolution of the disease state.

Acknowledgements The authors would like to acknowledge partial support from the NDSU Grand Challenge program for the Center for Engineered Cancer Test-Beds.

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Bone Micro- and Nanomechanics Caitlyn J Collins, Orestis G Andriotis, Vedran Nedelkovski, and Martin Frank, TU Wien, Vienna, Austria Orestis L Katsamenis, University of Southampton, Southampton, United Kingdom Philipp J Thurner, TU Wien, Vienna, Austria © 2019 Elsevier Inc. All rights reserved.

Bone Structure and Hierarchy Trabecular Bone Mechanical Testing of Trabecular Bone at the Tissue-Level Nanoindentation Micromechanical tests Scanning acoustic microscopy (SAM) Effects of Age and Disease on Micromechanical Properties of Trabecular Bone Cortical Bone Mechanical Testing of Cortical Bone at the Tissue and Osteonal Levels Micromechanical tests Reference point indentation (RPI) Effects of Age and Disease on the Micromechanical Properties of Cortical Bone Individual Lamellae, Interlamellar Areas, and Cement Lines Mechanical Testing of Cortical Bone at the Lamellar Length Scale Fracture toughness Micromechanical tests Collagen Fibrils and Nanocrystals: Individual Components of Bone Collagens and Collagen Fibrils Collagen Fibril Structure Collagen Fibril Hydration and Mechanics Mineralized Collagen Fibrils (MCFs) Mineralization of Collagen Fibrils Mechanics of Mineralized Collagen Fibrils (MCFs) Noncollagenous Proteins References Further Reading

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Glossary Carbon-substituted hydroxyapatite Calcium phosphate mineral crystals that make up most of the inorganic part of bone. Collagen Protein family that makes up most of the organic part of bone. Collagen fibril Sub-mm diameter fibers that are formed from certain collagen types and are one basic building block of bone and most other skeletal tissues. Cortical/compact bone A type of bone tissue consisting of closely packed osteons and forms the dense, stiff exterior of bones. Fracture toughness Measure of susceptibility of a material to resist fracture. Hierarchical structure Exhibiting structural elements on more than one length scale. Lamellae A thin layer of tissue made up of a network of mineralized collagen fibers. Osteon Cylindrical structures made up of concentrically arranged lamellae surrounding individual Haversian canals. Trabecular/cancellous bone A highly porous type of bone tissue that pervades the interior or cavity of bones with a lattice of mineralized struts and plates (trabeculae) and intervening spaces filled with marrow or fat. Transverse isotropic A material with the same physical properties in one plane (e.g., x–y plane) and different physical properties in the direction normal to this plane (e.g., z-axis). These materials can be described by five independent elastic constants.

Nomenclature εkl A dimensionless component of the second-order strain tensor, representing the strain in the l-direction acting on a surface that is oriented perpendicular to the k-direction Da Crack extension, measured in mm

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sij A component of the second-order stress tensor, representing the stress in the j-direction acting on a surface that is oriented perpendicular to the i-direction with dimensions of force per unit area AFM Atomic force microscopy AGEs Advanced glycation end products, that is, proteins or lipids that become glycated as a result of exposure to sugars) Cijkl The fourth-order stiffness tensor that is a property of the material and is often dependent on temperature, pressure, and microstructure. Cijkl is given in units of force per unit area. FE Finite element IDI Indentation distance increase, measured in mm K Stress intensity factor, measured in units of stress * length1/2 Kc Critical stress intensity factor, measured in units of stress * length1/2 NCPs Noncollagenous proteins NEG Nonenzymatic glycation OPM Oliver–Pharr method qBEI Quantitative backscattered electron imaging R-curves Crack resistance curves generated from plots of fracture energy versus crack length RPI Reference point indentation SAM Scanning acoustic microscopy SEM Scanning electric microscopy TID Total indentation distance, measured in mm

Bone Structure and Hierarchy When considering bone, the first structures that come to mind are likely whole bones. At this macroscale or organ level, bones have complex shapes, which cater to the needs of the human body and its musculoskeletal system, that is, shielding internal organs, acting as lever arms for load transfer, providing structural stability, and enabling movement and locomotion via muscle attachment sites. Whole bones also serve as mineral reservoirs and are responsible for the majority of blood cell formation. While bone tissue at the macroscale generally appears white in color and is considered to be structurally stiff, it is unlike any engineering material. In fact, bones are complicated composite materials with a distinct hierarchical organization; at the lowest organizational level, bone tissue consists of collagen in the form of fibrils, carbon-substituted hydroxyapatite mineral in the form of nanocrystals as well as water and a small amount of noncollagenous proteins (NCPs). A more detailed description of these basic building blocks can be found toward the end of this article as we will start our description at the macroscale. Here, bone tissue can be examined using a low power light microscope. Sequentially increasing the microscope magnification and changing imaging modalities, as required, reveals the whole structural hierarchy of bone. This hierarchy has been well described in the works of Weiner and Wagner as well as Rho et al. (Weiner and Wagner, 1998; Rho et al., 1998). In fact, seven different structural levels can be isolated (Fig. 1). Earlier works by Katz et al. and Lakes, focused on the hierarchical organization of cortical bone, differentiated between four and five structural levels (Katz et al., 1984; Lakes, 1993). Due to this complex structural organization, micro- and nanomechanics of bone are of particular interest as they characterize specific levels within the hierarchy. Starting at the macroscale, the bone structural hierarchy begins with the whole bone level, with individual bones being classified as flat, short, long (femur or tibia) or irregular (vertebrae) (Standring, 2016). Venturing to the next smaller length scale, for example, cutting open a long bone, the two major types of bone tissue are encountered, namely cortical (or compact) bone and trabecular (or cancellous) bone (tissue level). These two bone tissue types mainly differ in apparent density as a result of porosity. Cortical bone has low porosity (several %), while trabecular bone is highly porous and consists of small, sub-mm size plates and rods. Below the length scale of these two bone tissue types, further structure is observed. The classification of these structures is dependent on which type of bone tissue is investigated. Osteonal bone is encountered within cortical bone (in humans). So-called osteons are cylindrical features with a central Haversian canal oriented parallel to the main loading direction, that is, along the long bone axis in long bones. Typical osteon diameters range from 100 to 400 mm. The equivalent structural feature in trabecular bone at this length scale is the so-called trabecular bone packet. Both features originate due to the remodeling process in bone and are interfaced into the surrounding bone via a so-called cement line. Both the bone packet and the osteon consist of bone lamellae with typical thicknesses of 5–10 mm. At this structural level in osteons, lamellae are arranged in a concentric fashion around the Haversian canal with interspersed boundary layers, so-called interlamellar areas (sometimes also called thin lamellae). Within individual lamellae, mineralized collagen fibrils, the basic building block of bone, are organized into aligned bundles or arrays. Orientation of these fibril arrays varies across individual lamellae, and there are several models which describe the fibril arrangement within the lamellae of osteons: the rotated plywood, the twisted plywood model, or spiral twisting model (Giraud-Guille et al., 2003; Wagermaier et al., 2006). More recently, three-dimensional reconstruction of human lamellar bone has revealed the presence of two distinct interpenetrating materials, which contribute to the hierarchical organization at this length scale (Reznikov et al., 2014). Occasionally,

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Fig. 1 Hierarchical structural organization of bone from macro- to sub-nanostructure: whole bone divided into cortical and trabecular (cancellous) bone; osteons with Haversian systems; lamellae; collagen fiber assemblies of collagen fibrils; individual collagen fibrils; and bone mineral crystals, collagen molecules, and noncollagenous proteins. Adapted from Rho, J.Y., Kuhn-Spearing, L., and Zioupos, P. (1998). Mechanical properties and the hierarchical structure of bone. Medical, Engineering and Physics 20, 92–102., Copyright (2017) Elsevier, and Thurner, P.J. (2009). Atomic force microscopy and indentation force measurement of bone. Wiley Interdisciplinary Reviews: Nanomedicine and Nanobiotechnology. 1, 624–649. https://doi. org/10.1002/wnan.56., Copyright (2017) Wiley.

a local arrangement or order of several fibrils may be called a fiber. As mentioned, fibril arrays consist of a network of mineralized collagen fibrils. In their pristine form, these fibrils have diameters between 100 and 300 nm. The fibrils themselves consist primarily of type I collagen, mineral crystals, and water. The mineral crystals are located in the gap regions within the fibrils (intrafibrillar mineral) (Knapp et al., 2002; Sasaki et al., 2002) as well as decorating the entire fibril surface (extrafibrillar mineral) (Lees et al., 1984; Prostak and Lees, 1996; Kindt et al., 2007). The attachment of minerals to the collagen is most likely facilitated via NCPs. The interface between adjacent mineralized fibrils is also thought to consist of NCPs (Thurner, 2009). Here, NCPs have been found to play both a structural role in terms of controlling crystal size and shape as well as a mechanical role (see Noncollagenous proteins). The main body of this article will address the measurement or investigation of mechanical properties of bone at different length scales ranging from the sub-mm size to the individual components (e.g., NCPs). Evaluation of bone at each length scale is of interest as potential changes in whole bone tissue mechanics could arise from changes at any hierarchical level.

Trabecular Bone Trabecular bone is a highly porous variant of bone tissue and is mainly present in the terminal regions of long bones and in the middle regions of short, flat, and irregular bones such as vertebrae. Trabecular bone is composed of a complex network of interconnected rods and plates, called trabeculae (Lucchinetti et al., 2000). On the microscopic scale, trabeculae are built up of trabecular bone packets (Jee, 2001), with an average wall thickness of around 50 mm (Lips et al., 1978). These trabecular bone packets consist of layers of lamellae oriented in slightly different directions, and the packets are separated by cement lines and interstitial lamellae (Jee, 2001). Individual lamellae consist of mineralized collagen fibrils that are preferentially aligned with the long axis of each lamella. Within the lamellae and packets, ellipsoid cavities, called lacunae, typically 5–15 mm in cross-section and 25 mm in length, house osteocytes (Hamed et al., 2012). These cells, which have a higher activity level in trabecular than cortical bone, sense deformation and damage of the bone matrix and trigger bone remodeling and adaptation processes (Oftadeh et al., 2015). At the component level, trabecular and cortical bone consists mainly of carbon-substituted hydroxyapatite, collagen, and water. However, trabecular bone has a higher water content (vol. 27%) compared to cortical bone (vol. 23%) (Oftadeh et al., 2015; Gong et al., 1964). The relative composition of the main bone tissue constituents is highly dependent on age, anatomical site, gender, and potential pathology (Boskey, 2013). Since water is a major constituent, it may play a significant role in the overall mechanical behavior of bone. Dehydration has been shown to cause changes in structure as well as mechanical properties of trabecular bone (Lievers et al., 2010). Further, dehydration results in a transition from ductile buckling to a brittle behavior in compression loading of individual bone trabeculae (Townsend et al., 2017).

Mechanical Testing of Trabecular Bone at the Tissue-Level The porous structure of trabecular bone makes the determination of tissue-level mechanical or material properties difficult. At the trabecular bone tissue-level, measurements at the length scale of individual trabeculae or smaller are of interest. Tissue-level

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properties of trabecular bone can be determined in three different ways: with direct measurement via micromechanical tests or nanoindentation, with scanning acoustic microscopy (SAM), or with indirect assessment from large scale finite element (FE) models. To remain within the scope of this article, indirect assessment methods are not reviewed.

Nanoindentation Nanoindentation of bone is usually performed in force-controlled mode (Thurner, 2009). In principle, a tip with defined geometry is driven into the sample at a specified loading rate until a certain maximum load is reached. Then, this maximum load is usually held constant for some time (on the order of tens of seconds) to wait for any transient processes to cede. Subsequently, the sample is unloaded. The elastic modulus and the hardness of the sample are then calculated from the unloading curve using the Oliver–Pharr method (OPM) (Oliver and Pharr, 1992, 2004), which assumes elastic behavior during the first part of the unloading process. This method can be seen as the gold standard for data analysis of nanoindentation measurements and is based on several assumptions, as reviewed in Thurner (2009). While some assumptions are not met for bone, for example, isotropy, the OPM is still commonly used to extract material properties of cortical and trabecular bone at the microscale. However, the majority of studies reported in the literature use a variety of test protocols making it difficult to cross compare quantitative results. The elastic modulus of trabecular bone, determined over the course of several studies, was reported to vary from 6.9 to 23.5 GPa (Thurner, 2009). Comparatively, the elastic modulus of human cortical bone, excised from the femur midshaft, was reported to vary between 17 and 27 GPa, depending on the type of lamellae (Thurner, 2009; Zysset et al., 1999; Rho et al., 1999). As yet, no clear distinction has been drawn between the mechanical properties of cortical and trabecular bone measured via nanoindentation; however, some studies have reported differences in hardness (Hodgskinson et al., 1989; Weaver, 1966). Exhaustive reviews of bone nanoindentation results can be found in Thurner (2009) and Lewis and Nyman (2008). Perhaps the largest limitation of this technique is that nanoindentation is typically conducted on dried samples. Sample holders and indentation tip assemblies that would allow testing in a liquid environment do exist; however, testing under these conditions generally renders the operator blind for the choice of indentation location. Elastic modulus has been shown to be correlated to the mineral content of (calcified) cartilage (Gupta et al., 2005); however, similar correlations within bone of a single species or across an anatomical location are not clearly present (Spiesz et al., 2013). This lack of clear correlation between elastic modulus and mineral content is most likely due to the fact that bone is not only heterogeneous but also heterogeneously anisotropic. As an exception to this observation, tissue stiffness and mineralization were found to be highly correlated across transverse sections of individual trabeculae (see Fig. 2). Both variables were found to be higher within the core compared to the outer surface (Mulder et al., 2007). The distribution of mineralization between the core and the surface of individual trabeculae stems from the remodeling process in trabecular bone. Remodeling initiates at the surface of the trabeculae leading to a distinct difference between shell and core regions. In addition to heterogeneity over the trabeculae cross-section, a significant difference has been reported between the transverse and the longitudinal elastic modulus within individual trabeculae (Rho et al., 1999). For the purpose of nanoindentation measurements, bone tissue is often assumed to be isotropic with a Poisson’s ratio of 0.3 (Zysset, 2009). In a study using nanoindentation to evaluate the elastic properties of cortical and trabecular bone lamellae, varying the Poisson’s ratio from 0.2 to 0.4 was found to only generate relative errors between þ 9.9% to  8.2% (Zysset et al., 1999). However, bone is inherently anisotropic, and the indentation modulus for anisotropic materials is a function of the axis of indentation and the full elasticity tensor (Zysset, 2009). As such, measurements carried out under the assumption of an isotropic material model either over- or underestimate the “true” elastic modulus, that is, the elastic constants of the stiffness matrix from the generalized Hooke’s Law (Swadener et al., 2001). Stiffness and hardness have also been shown to be significantly dependent on lamellar type, anatomical site, and the individual (Zysset et al., 1999). Such dependence might be explained by local differences in the bone mineral density distribution (Roschger et al., 2008) in addition to differences in mineralization and collagen organization. A separate study reported only a weak correlation between the mean degree of mineralization and the indentation modulus and hardness (Hengsberger et al., 2002). However, the authors mention that other factors such as hypermineralization or a reduction in collagen cross-links might contribute to the behavior of the extracellular matrix. The literature is in disagreement with regards to the correlation of nanoindentation results with macroscale mechanical tests. Hengsberger et al. reported good agreement between nanoindentation and macroscopic tests (Hengsberger et al., 2003), whereas Silva et al. reported no significant correlation between nanoindentation and three-point-bending tests of macroscopic samples (Silva et al., 2004).

Micromechanical tests Due to the aforementioned limitations of nanoindentation (generally dry samples, information only on elastic modulus and hardness), the determination of mechanical or material properties from micromechanical tests of individual trabeculae is still a matter of ongoing research. However, this faces several challenges, as reviewed by Lucchinetti et al. (2000). Most studies have been conducted on rodlike trabeculae focusing on the elastic region and only a few studies have investigated postyield behavior (Hodgskinson et al., 1989; Weaver, 1966; Lewis and Nyman, 2008). Based on the generalized Hooke’s law, the elastic properties of trabecular bone tissue might be described by sij ¼ Cijklekl, where Cijkl is a fourth-order tensor (the stiffness tensor) that relates stress (s) to strain (ε). This law is only valid up to the proportionality limit, that is, the point after which stress is no longer proportional to strain, and ceases to apply past the elastic limit of a material. As such, the generalized Hooke’s law should only be applied in cases where small forces or displacements are applied. Within this tensor context, trabecular bone is assumed to behave as an orthotropic material, that is, nine

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Fig. 2 (A) The tissue stiffness profile from nanoindentation and (B) the degree of mineralization profile across transverse sections of individual trabeculae from a new born pig. Measurement areas are indicated in black in both images on the right, taken using light microscopy (top) and microcomputed tomography (mCT) (bottom). Tissue stiffness and degree of mineralization increase toward the central section of the trabeculae. Note that the two outermost indents are located in the embedding medium. (C) Image of a trabecular cross-section captured using quantitative backscattered electron imaging (qBEI); tissue lamellae, canaliculi, and the spatial distribution of mineralization across the cross-section can be seen in the image, where higher brightness indicates higher mineral concentration. (A and B) Reprinted with permission from Mulder, L., Koolstra, J.H., den Toonder, J.M.J., and van Eijden, T.M.G.J. (2007). Intratrabecular distribution of tissue stiffness and mineralization in developing trabecular bone. Bone 41, 256–265. https://doi.org/10.1016/j.bone.2007.04.188. Copyright (2017) Elsevier. (C) Adapted from Brennan, M.A., Gleeson, J.P., Browne, M., O’Brien, F.J., Thurner, P.J., and McNamara, L.M. (2011). Site specific increase in heterogeneity of trabecular bone tissue mineral during estrogen deficiency. European Cells & Materials 21, 396–406.

independent components fully describe the material behavior. Since single trabeculae tend to have a preferred lamellar orientation along the longitudinal trabecular axis, they can be considered as transverse isotropic (Lucchinetti et al., 2000). Trabeculae are primarily subjected to bending, tensile and compressive loads. Thus, the elastic constant/modulus in the longitudinal direction might be sufficient to describe the elastic behavior. Several approaches have been used to determine the tissue-level elastic modulus via micromechanical testing, these include: buckling (Townsend et al., 2017; Runkle and Pugh, 1975), three-point bending (Carretta et al., 2013a, b; Szabó et al., 2011; Busse et al., 2009; Kuhn et al., 1989; Jungmann et al., 2011; Ridha and Thurner, 2013; Szabó and Thurner, 2013), four-point bending (Choi and Goldstein, 1992) and tensile tests (Carretta et al., 2013a, b; Rho et al., 1993; Bini et al., 2002; Hernandez et al., 2005; Jirousek et al., 2011; Ryan and Williams, 1989; Yamada et al., 2014). Tissue-level elastic moduli range from 0.8 GPa (Ryan and Williams, 1989) up to 16.2 GPa (Carretta et al., 2013b). Traditionally, this large variation is attributed to the use of different test setups and their associated boundary conditions. For example, Carretta et al. tested trabeculae in tension as well as three-point bending and reported significantly higher values for the elastic modulus of the trabeculae subjected to tensile loading (Carretta et al., 2013a). However, by definition, elasticity is a measure of the energy, which can be fully transformed into efficient mechanical work without dissipation, and must be independent of both strain distributions and strain rates (Rajagopal and Srinivasa, 2009). Hence, the elastic modulus, which describes material elasticity, should not differ between experiments. There is a discussion in recent literature on whether bone is an elastic material or if it exhibits elastoplastic hardening, thereby necessitating loading–unloading experiments to measure the elastic modulus (Schwiedrzik et al., 2014; Luczynski et al., 2015). Most of the studies utilizing micromechanical tests to evaluate individual trabeculae treated them as simple rods with a constant elliptical cross-section. Carretta et al. argued that true tissue material parameters can only be obtained when combining an experimental approach with FE analysis (Carretta et al., 2013a). Indeed, Frank et al. showed that the curvature of individual trabeculae

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Fig. 3 The determined tissue stress–strain curves of trabeculae tested in uniaxial tension, based on an elliptical cross-section (light blue). The bold line represents the mean tissue stress–strain curve, calculated using the formulas shown in the graph and the mean values of each parameter. The red crosses and circles illustrate the point of yield and the point of failure, respectively. Reprinted with permission from Frank, M., Marx, D., Pahr, D.H., Thurner, P.J. (2017) Mechanical properties of individual trabeculae in a physiological environment. Proc. IASTED Int. Conf. Biomed. Eng. (BioMed 2017). https://doi.org/10.2316/P.2017.852-023. Copyright (2017) IEEE.

results in a structural influence on the tissue-level elastic modulus, but an upper and lower boundary can be estimated by simple geometric assumptions (Frank et al., 2017). Determination of yield and postyield properties of trabecular bone has been mainly carried out on larger samples (Chang et al., 1999; Turner, 1989; Kopperdahl and Keaveny, 1998), thereby involving a structural component. Only a few studies have determined the yield properties of trabecular bone at the tissue-level (Carretta et al., 2013a, b; Frank et al., 2017; Busse et al., 2009; Hernandez et al., 2005). Carretta et al. reported on the postyield behavior of bovine (Carretta et al., 2013a) and human (Carretta et al., 2013b) trabeculae under both tension and three-point bending test conditions. In both studies, yield strain, ultimate strain, and postyield work were reported to be significantly higher in the three-point bending group. Busse et al. subjected osteoporotic and skeletally intact trabeculae to three-point bending test (Busse et al., 2009); significant differences in yield strength, ultimate stress, and work to failure were found, highlighting the impact of disease on the postyield behavior of bone. Hernandez et al. demonstrated that ultimate tensile strain was weakly influenced by nonenzymatic glycation (NEG), also called advanced glycation end products (AGEs) of type I collagen, a factor associated with aging and diabetes (Hernandez et al., 2005). In this study, individual trabeculae were hydrated in physiologic salt solution (pH ¼ 7.4) prior to tension testing. The average ultimate tensile strain was reported to be 8.8% (Hernandez et al., 2005). Carretta et al., in contrast, reported an ultimate tensile strain of 5.1% for individual trabeculae (Hodgskinson et al., 1989; Weaver, 1966); however, trabeculae in these studies were prepared and tested in a dry environment. Frank et al. performed tensile tests on individual trabeculae in a hydrated environment (physiologic salt solution, pH ¼ 7.4) and reported an average ultimate strain of 9.8% (Frank et al., 2017), thus confirming that hydration leads to significant postyield deformation (as shown in Fig. 3).

Scanning acoustic microscopy (SAM) SAM enables nondestructive determination of the elastic mechanical properties of relatively small, stiff samples, such as bone at the tissue level (Laugier et al., 2013). SAM can be used to determine changes in the longitudinal wave velocity within bone in response to aging or disease, for example, in individuals before and after menopause (Hasegawa et al., 1995). Bumrerraj and Katz used SAM to determine the correlation between acoustic microscope brightness and the elastic modulus of known materials. This correlation was then used to predict the elastic modulus of bone samples subjected to SAM (Bumrerraj and Katz, 2001). The predicted elastic modulus values were found to have a good correlation with elastic modulus values of cortical and trabecular bone measured using nanoindentation testing. Turner et al. also used acoustic microscopy and nanoindentation to determine the elastic modulus of trabecular and cortical bone tissue (Turner et al., 1999). The measured elastic modulus for trabecular bone was reported as 17.5  1.1 GPa using acoustic microscopy and as 18.1  1.7 GPa using nanoindentaion, both were slightly higher than the measured elastic moduli of cortical bone in the transverse direction (14.9  0.5 and 16.6  0.32, respectively) (Turner et al., 1999). In a new approach, Litniewski measured the velocity of surface waves to determine elastic modulus. Such an approach enables evaluation of samples that are only accessible from one side (Litniewski, 2005). A good agreement was found between the reported elastic modulus and that of a previous study (Turner et al., 1999).

Effects of Age and Disease on Micromechanical Properties of Trabecular Bone With respect to age and disease, the influence of strain rate on tissue level material properties is of particular interest. Most experiments are carried out at low strain rates in a monotonic fashion. However, the physiological loading experienced by patients is either cyclical (fatigue) or occurs at high strain rates (falls and fracture). Therefore, a discussion of the micromechanics of aging

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trabecular bone under these conditions is important. Unfortunately, very few studies exist that shine light on these questions and, as such, more research is required to expand the current knowledge and understanding of trabecular bone behavior. Whether the mechanical properties of individual trabeculae are in fact affected by strain rate is not clear as several conflicting studies exist (Szabó et al., 2011; Hansen et al., 2008; Ferreira et al., 2006; Currey, 1975). Upon aging, the collagen phase within bone (the main organic component) can suffer from increased cross-linking via sugars (e.g., present due to diet or disease), known as AGEs. This results in a tissue which is structurally very different from that generated by the naturally occurring enzymatic cross-linking process. Two studies that have investigated the effects of AGEs and are briefly discussed below. Hernandez et al. determined that there was only a weak correlation between AGEs and ductility in individual trabeculae, although ductility varied tremendously in their study (Hernandez et al., 2005). Ductility was characterized by the strain at maximum load-carrying capacity and was found to range between roughly 3% and 18%. In a separate study, Tang et al. determined that AGEs cause stiffness loss and a reduction in damage accumulation within individual trabeculae, thus resulting in a loss in the ability to dissipate energy (Tang et al., 2007). These studies suggest that the accumulation of sugars and the resultant cross-linking, that is, AGEs within collagen, are a potential issue, especially in individuals with a high sugar diet, obesity, or diabetes. Osteoporosis, one of the most prominent bone diseases, is associated with a significant increase in the incidence of bone fracture and may be paired with a significant loss of bone mass. One line of thought considers that the properties of bone at the tissue level might be altered over the course of this disease. In this context, individual trabeculae excised from osteoporotic bone have been reported to have a lower ultimate strain and postyield work under tensile testing conditions when compared to trabeculae excised from healthy donors (Carretta et al., 2013b). However, no significant difference was found between the elastic behavior of either osteoporotic or healthy donors. Similarly, studies using nanoindentation reported no significant difference between osteoporotic and healthy trabecular bone (Wang et al., 2008; Hu et al., 2015). Wang et al. reported no significant difference between the elastic properties of trabecular bone measured from patients that had experienced vertebral osteoporotic fracture and those of a control group (Wang et al., 2008). Hu et al. reported that the elastic modulus and hardness of trabecular bone tissue were unaffected by an ovariectomy in a murine animal model (rat) (Hu et al., 2015). Studies investigating effects of bisphosphonate treatment on the microstructure and mechanical strength of bone have reported a reduced capacity to resist fracture in the treated bone compared to untreated bone (Ma et al., 2017; Acevedo et al., 2015). Although bisphosphonates have been shown to impede the loss of bone density in osteoporosis patients, the effect of long-term bisphosphonate treatment on preyield (elastic) properties is still subject to debate. Microscopic damage is also thought to be an important determinant of bone fragility (Seref-Ferlengez et al., 2015; Fazzalari et al., 1998). Two different types of microscopic damage have been identified: linear microcracks, which may run parallel or cross-hatch to one another, and diffuse damage (Fig. 4). Linear microcracks are sharply defined cracks of roughly 50–100 mm in length that primarily occur within interstitial bone (Seref-Ferlengez et al., 2015). Diffuse damage is characterized by an accumulation of short submicroscopic cracks and is preferentially induced inside trabecular bone packets (Vashishth et al., 2000). Interestingly, the presence of microcracks in trabecular bone of the femoral head have been reported to be significantly higher in old bone from both healthy and fractured cohorts compared to young bone (Mori et al., 1997). The impact of age and disease on the type of damage mechanisms found in trabecular bone is a matter of ongoing and future research.

Fig. 4 Micrograph of basic fuchsin stained trabeculae highlighting the morphological differences between microcracks and diffuse damage. The arrows identify microcracks in linear (left) and cross-hatch (center) configurations as well as diffuse damage with no discernible microcracks (right). Reprinted with permission from Fazzalari, N.L., Forwood, M.R., Smith, K., Manthey, B.A., Herreen, P. (1998). Assessment of cancellous bone quality in severe osteoarthrosis: Bone mineral density, mechanics, and microdamage, Bone 22, 381–388. https://doi.org/10.1016/S8756-3282(97)00298-6. Copyright (2017) Elsevier.

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Cortical Bone Cortical bone is the denser variant of bone tissue and makes up the shells and shafts of long bones as well as the external shells of short, flat, and irregular bones. The internal microstructure of cortical bone is organized into concentrically arranged cylindrical structures called osteons (or Haversian systems). This functional unit of human cortical bone is comprised of a central so-called Haversian canal surrounded by concentric rings of lamellae roughly 5–7 mm in thickness (Gupta et al., 2006a). Haversian canals house both blood vessels and nerve cells and interconnect via Volkmann’s canals that run transverse to the bone diaphysis. The average osteon diameter is reported to range from 50 to 500 mm in diameter (Black et al., 1974). Interstitial lamellae, the remnants of osteons partially resorbed during the remodeling process, occupy the space between individual osteons, that is, interstitial bone. Osteons and interstitial lamellae are separated by cement lines, zones primarily made up of calcified mucopolysaccharides, lacking collagen fibrils (Burr et al., 1988), and rich in NCPs (Sodek et al., 2000). Although the degree of cement line mineralization has been historically controversial, recent work supports the conclusion that cement lines are not poorly mineralized when compared to surrounding bone (Skedros et al., 2005). Within cortical bone lamellae, osteocyte housing lacunae and canaliculi are found. Canaliculi, small channels roughly 0.5 mm in diameter, provide paths for the long dendritic extensions of osteocytes to interconnect with cells in adjacent lacunae as well as to cells within the Haversian canal. The spatial organization of osteonal lamellae remains a topic of debate. Under polarized light microscopy, lamellae appear either light or dark. Models attempting to explain the observed differences in adjacent lamellae can be grouped into either fibril orientation or fibril density models. In the first, orientation of the mineralized collagen fibril differentiates adjacent layers. In the second, the collagen is assumed to be homogeneously distributed in adjacent layers and only the relative collagen fibril density is responsible for the observed differences. A comprehensive review of the existing models of lamellae organization can be found in Mitchell and van Heteren (2016).

Mechanical Testing of Cortical Bone at the Tissue and Osteonal Levels At the tissue level, measurements at the length scale of individual osteon or smaller are of interest. A primary argument to attempt such measurements is that osteons are in fact the product of the bone remodeling process. Therefore, changes in bone mechanics due to age and disease will be most pronounced at this level. Sub-millimeter specimens of cortical bone allow researchers to isolate and focus on the desired microstructural features, for example, individual osteons and interstitial lamellae, and access their mechanical behavior under well-controlled conditions. Tissue-level properties of cortical bone can be determined with direct measurement via micromechanical tests, nanoindentation, or reference point indentation (RPI). Imaging modalities are commonly combined with mechanical testing protocols to aid in the assessment of postyield and fracture behavior, particularly when evaluating mechanisms for energy dissipation at smaller length scales. This section will focus on micromechanical tests and RPI. For detailed reviews on nanoindentation refer to Thurner (2009) and Zysset (2009).

Micromechanical tests Rho et al. used microtensile experiments to measure the elastic modulus of cortical bone obtained from the diaphyseal region of a human tibia as well as individual trabeculae (Rho et al., 1993). The tensile specimens had approximate dimensions of an individual trabecula and the measured elastic modulus (18.6  3.5 GPa) was larger than that of single trabeculae (10.4  3.5 GPa). This study was one of the first to evaluate mineral density versus elastic modulus relationship in cortical and trabecular bone. By miniaturizing the cortical specimens to sizes similar to single trabeculae, the authors effectively demonstrated that cortical bone and trabecular bone have differing material properties. Ever since, a number of studies have used similar approaches to measure the elastic modulus of miniaturized cortical bone specimens (Reilly and Burstein, 1974, 1975). Other micromechanical testing setups have included cantilever beam bending, three- or four-point bending, torsion, and compression. The elastic modulus of cortical bone varies depending on the testing method, level of hydration, porosity, anatomical location as well as orientation of the tissue sample (Currey, 2002). Dry cortical bone was found to have a 20% higher elastic modulus than hydrated cortical bone (Currey, 2002). Compression tests also revealed that the elastic modulus of human cortical bone is higher in the longitudinal direction (16–23 GPa) compared to transverse (6–13 GPa) (Rho et al., 1998). This anisotropy mirrors the primary physiological loading direction. As early as the 1970s, researchers have sought to understand the mechanical properties of single osteons. The pioneering studies of Ascenzi and Bonucci which evaluated the tensile (Ascenzi and Bonucci, 1967), compressive (Ascenzi and Bonucci, 1968), shearing (Ascenzi and Bonucci, 1972), bending (Ascenzi et al., 1990), and torsional (Ascenzi et al., 1994) properties of individual osteons paved the way for a whole new era of quantitative assessment of bone at and below the microscale. Using basic tools, such as stereoscopes, microscopes, drills, surgical blades, and impressive manual handling skills, they were able to isolate individual osteons from bulk cortical bone and subject them to controlled mechanical tests using custom-made apparatuses (Fig. 5). Elastic moduli of single osteons were reported to range from 2.6 to 11.7 in tension (Ascenzi and Bonucci, 1967), 1.6–9.3 GPa in compression (Ascenzi and Bonucci, 1968), and 0.9–2.7 GPa in three-point bending (Ascenzi et al., 1990). All moduli were found to be sensitive to changes in mineral content and the predominant collagen fibril orientation in neighboring lamellae. Osteons at an initial stage of calcification exhibited greater ductility and reduced stiffness compared to more mature osteons. Further, osteons with primarily longitudinal collagen fiber alignment were able to support greater stresses in tension and torsion, whereas osteons with alternating collagen fiber alignment in adjacent lamellae were able to support greater stresses in compression. Notably, the

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Fig. 5 (A) Images of (a) an osteon prepared for three-point bending; (b) an “alternate” osteon, with fiber bundles changing orientation in successive lamellae though an angle of 90 degree, at ultimate bending strength; (c) a “longitudinal” osteon, with longitudinally aligned fiber bundles in adjacent lamellae, at ultimate bending strength; and microradiographs of fractured (a) “longitudinal,” (e) “alternate,” and (d) irregular “alternate” osteon samples. (B) Polarized light microscopy images of (a) an isolated osteon ready to be tested in tension (50 magnification); (b) a “longitudinal” osteon (100 magnification); and (c) an “alternate” osteon (100 magnification). (C) From left to right: isolated osteon sample; isolated osteon sample as seen in transmitted light; an end surface of the sample. (A) Reprinted with permission from Ascenzi, A., Baschieri, P., and Benvenuti, A. (1990). The bending properties of single osteons. Journal of Biomechanics 23, 763–771. https://doi.org/10.1016/0021-9290(90)90023-V. Copyright (2017) Elsevier. (B) Reprinted with permission from Ascenzi, A. and Bonucci, E. (1967). The tensile properties of single osteons. The Anatomical Record 158, 375–386. https://doi.org/10.1002/ar.1091580403. Copyright (2017) John Wiley and Sons. (C) Reprinted with permission from Ascenzi, A. and Bonucci, E. (1968). The compressive properties of single osteons. The Anatomical Record 161, 377–391. Copyright (2017) John Wiley and Sons.

values reported by Ascenzi and Bonucci are well below elastic modulus values from nanoindentation or other micromechanical tests. Since there have been no studies reproducing these experiments, it remains unclear why this is the case.

Reference point indentation (RPI) As noted previously, indentation techniques at various length scales have been used to measure hardness and stiffness of both trabecular and cortical bone (Thurner, 2009; Zysset, 2009). Given the search for complementary diagnosis of osteoporosis and age-related bone fracture risk, microindentation has been established as a micromechanical test that can even be conducted in patients (Diez-Perez et al., 2010). Microindentation via Reference Point Indentation (RPI), which does not require miniaturized specimens, utilizes a similar approach as nanoindentation (described above), driving an indenter into the bone tissue; although this technique utilizes larger forces and penetration distances, it can still be considered as a micromechanical technique. Clinically oriented RPI evaluates the resistance of bone to indenter penetration. Currently, two approaches are used. One is cyclic microindentation, where the indenter makes repeated loading and unloading cycles (Hansma et al., 2006). This approach is predominantly used in preclinical studies. The other is single load microindentation, where the indenter impacts into bone with a given force (Bridges et al., 2012). This approach is predominantly used in clinical studies. Both approaches are implemented in RPI devices. RPI typically uses either a test probe with a spherical tip contained within a hypodermic needle (cyclic RPI) (Fig. 6), or a single probe and a preload (single load RPI) to establish a reference point. Once a reference point has been established the indentation process begins. In case of a single load cycle (clinical device), the indentation distance is compared to a PMMA reference material giving the “bone material strength index” (BMSi). For multiple loading cycles (laboratory device), the indentation distance increase (IDI) and the total indentation distance (TID) are measured. IDI is the indentation distance between the first and last cycle and TID is the indentation depth of the last indentation relative to the initial reference point. Since RPI causes microscopic damage to open and may relate to the separation of mineralized collagen (Diez-Perez et al., 2010), IDI and TID parameters have been assumed

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Fig. 6 Reference Point Indentation. (A) Force–distance curves for first and last indentation cycle. (B) Diagram of indentation probe at the first and last cycle. Adapted from Diez-Perez, A., Güerri, R., Nogues, X., Cáceres, E., Peña, M.J., Mellibovsky, L., Randall, C., Bridges, D., Weaver, J.C., Proctor, A., Brimer, D., Koester, K.J., Ritchie, R.O., Hansma, P.K. (2010). Microindentation for in vivo measurement of bone tissue mechanical properties in humans. Journal of Bone and Mineral Research 25, 1877–1885. https://doi.org/10.1002/jbmr.73. Copyright (2017) John Wiley and Sons.

to be related to fracture toughness. However, the measured parameters in both approaches have been found to have little to no correlation with a single material property. Instead, they likely relate to multiple material and structural properties of bone such as fracture mechanics, elastoplastic behavior, and structural properties (porosity and pore proximity) of bone (Jenkins et al., 2017). While RPI has been used in a number of preclinical and clinical studies (Diez-Perez et al., 2010; Güerri-Fernández et al., 2013; Poundarik et al., 2015; Sundh et al., 2016; Uppuganti et al., 2017; Rozental et al., 2017; Abraham et al., 2015; Jenkins et al., 2016), whether or not such approaches can indeed improve diagnosis of bone fracture risk in individuals remains to be shown (Allen et al., 2015).

Effects of Age and Disease on the Micromechanical Properties of Cortical Bone Age, anatomical location, sex, and pathology impact the morphology of cortical bone at the macro- and microscale. Cortical thinning and an increase in porosity are associated with aging and various pathologies (Zebaze et al., 2005; Granke et al., 2016; Mirzaali et al., 2015). Human osteons exhibit distinctive morphological heterogeneity depending on the age of the individual, skeletal site, and the presence or absence of both local and systemic factors, that is, hormones, cytokines, chemokines, etc. (Ascenzi, 2012). With age, osteon cross-sectional area and diameter have been shown to decrease (Currey, 2002; Bernhard et al., 2013) while cortical bone osteon density has been shown to increase (Currey, 2002). One of the important functions of an osteon is to serve as a barrier to crack propagation, similar to grain boundaries in metals, in order to increase bone toughness. Studies have shown that the ability of an osteon to inhibit crack propagation significantly decreases with age (Diab and Vashishth, 2005, 2007). Cortical bone specimens (4  4  48 mm in dimension) from younger donors (38  9 years) were reported to have a significantly longer bending fatigue life than bone from older donors (82  5 years); moreover, histological analysis revealed that younger bone predominantly formed diffuse damage as opposed to linear microcracks (Diab and Vashishth, 2005, 2007). Although osteon damage morphology appears to change with age, the underlying mechanisms for this shift from diffuse damage to microcrack accumulation is still a matter of ongoing research (Katsamenis et al., 2015; Yeni and Norman, 2000a).

Individual Lamellae, Interlamellar Areas, and Cement Lines At the several-micron length scale, the structural features of cortical bone observable by light or electron microscopy are the so-called lamellae, which can be found with varying thicknesses in the range 2–10 mm. Although lamellae are generally characterized as packed layers of mineralized collagen fibrils with slightly alternating orientation, details regarding composition, arrangement, and collagen fiber orientation are still poorly understood (Reznikov et al., 2014; Mitchell and van Heteren, 2016). Lamellae are interfaced with so-called interlamellar areas (also labeled thin lamellae), distinct features that differ from lamellae in mineralization degree and collagen orientation (Katsamenis et al., 2013a). Lamellae and interlamellar areas are arranged concentrically around Haversian canals in osteons and are built-up layer by layer during bone remodeling to form secondary osteons. Other types of lamellae include interstitial lamellae, found in the space between osteons, and those formed in fibrolamellar bone (or plexiform bone). Fibrolamellar bone is found in fast growing animals where bone tissue needs to be laid down before tissue organization via bone remodeling can take place (Weiner and Wagner, 1998). At the length scale of the lamella, only osteocyte lacunae and canaliculi contribute to cortical bone porosity.

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Mechanical Testing of Cortical Bone at the Lamellar Length Scale The determination of mechanical or material properties at the lamellar level is a challenge since the relevant structural features are on the order of a few microns. Imaging techniques such as scanning electron microscopy (SEM) and atomic force microscopy (AFM) have been used as indirect methods for assessing fracture surfaces, that is, the micro-level fracture mechanisms of larger, millimeter sized cortical bone specimens. In recent years, bone biomechanics researchers have utilized focused ion beam (FIB) milling techniques to isolate well defined micron-sized volumes on the order of a single lamella. Direct measurement of lamellar mechanical properties can then be achieved by subjecting these isolated volumes to micromechanical tests using AFM or nanoindenters.

Fracture toughness Fracture toughness is a mechanical property used to describe the ability of a material to resist fracture, typically measured in terms of stress intensity at the crack tip. A review of how to measure fracture toughness of bone can be found in Ritchie et al. (2008). Within bone, distinct toughening mechanisms have been identified at each hierarchical level (Fig. 7), each contributing to whole bone fracture resistance (Launey et al., 2010). Intrinsic toughening mechanisms work to inhibit crack initiation, while extrinsic toughening mechanisms primarily inhibit crack propagation and, to some extent, rupture. SEM is an imaging technique capable of generating high-resolution images with detailed topographical, morphological, and compositional information. As such, SEM is commonly used to evaluate the extrinsic toughening mechanisms of bone, which are observable at length scales from a single micron to hundreds of micrometers. In this context, studies assessing bone crack propagation have been conducted using environmental SEM (Ritchie et al., 2005; Koester et al., 2008, 2011), which does not require bone samples to be completely dehydrated and coated with a conductive material layer. However, only small scale samples with sample thicknesses ranging between 1 and 4 mm have been assessed using environmental SEM (Koester et al., 2008; Nalla et al., 2005a). Despite this limitation, SEM has been instrumental in the determination of cortical bone fracture toughness, particularly in the development of full crack resistance curves (R-curves). Fracture toughness testing is typically conducted on small, sub-mm compact tension (Norman et al., 1995; Yeni and Norman, 2000b) or single-edge notched bending specimens (Katsamenis et al., 2015; Ritchie et al., 2008) machined from cortical bone in accordance to ASTM standards E399 (ASTM, 1997) and E1820 (ASTM, 2001), respectively. R-curves can then be constructed by tracking the crack propagation, or crack extension, on each loading-step and plotting it against the stress intensity factor, yielding information on crack initiation and propagation behavior. Crack extension can be measured directly if tests are conducted within an environmental SEM (Ritchie et al., 2005; Koester et al., 2008, 2011), inferred by measuring the crack-tip opening displacement, or through the use of high definition videography (Katsamenis et al., 2013b). Crack extension can also be indirectly measured using standardized load-line compliance calibrations. Since fracture toughness is a material property, changes in composition and structure due to aging or tissue ultrastructure will have an effect. Nalla et al. reported a 40% reduction in crack initiation toughness with age from experiments on compact tension specimens of human cortical bone (donor age: 34–99 years). Cortical bone propagation toughness was also reported to be nearly eliminated in the older donors (Nalla et al., 2006). Single-edge notched bending tests revealed a link between the energy release rate (crack extension energy per unit area) and the collagen fiber orientation (Fig. 8); specifically, the energy release rate perpendicular to the collagen fibrils was reported to be nearly two orders of magnitude higher than in the direction of the collagen fibers (Peterlik et al., 2006). Osteon orientation further contributes to this observed anisotropy in cortical bone fracture toughness. Crack propagation around osteons has been shown to require significantly less energy than cracking through osteons (Katsamenis et al., 2015, 2013b; Nalla et al., 2005b). Similar to SEM, crack propagation at the micrometer scale can be captured using AFM, a surface characterization technique that can be used both for imaging and mechanical assessment via indentation or pulling. For example, AFM imaging of bovine trabecular bone fracture surfaces revealed that exposed collagen fibrils, as previously hypothesised from SEM images (Braidotti et al., 1997, 2000), are densely coated with mineral platelets, implying that the nature and the mechanical behavior of the interface between neighboring mineralized fibers is of significant importance (Kindt et al., 2007). Similarly, AFM analysis of fractured cortical bone revealed that cement lines and interlamellar areas, apart from providing a crack propagation path of least resistance (Katsamenis et al., 2013a; Peterlik et al., 2006; Fratzl, 2008), exhibit reduced modulus of elasticity compared to lamellae (Fig. 8) (Katsamenis et al., 2013a). As such, cement lines and interlamellar areas are thought to positively contribute to bone toughness.

Micromechanical tests As mentioned above, FIB milling can be used to prepare microscale specimens for mechanical testing. During FIB milling, a focused beam of heavy ions (typically Gallium) accelerated by high voltage (in the range of 30 kV) is directed toward a sample. The kinetic energy of the impinging ions knocks out target atoms, microscopically eroding and milling out material in a controlled manner. One of the earliest studies using FIB milling on biological mineralized tissue (Chan et al., 2009) employed a dual beam FIB/ SEM to fabricate microscopic cantilever beams with a triangular cross-section from a human primary molar. These beams were then bent using a conventional nanoindenter and Berkovich tip. Further studies were conducted using micron-sized cantilevers and micropillars machined to the length scale of a cortical bone lamella. The FIB only operates under high vacuum conditions, which can lead to substantial dehydration of the material and alteration of the mechanical behavior of the tissue. Jimenez-Palomar et al. investigated the effect of high vacuum exposure on micron-sized

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Fig. 7 The toughness of bone results from a mutual competition between extrinsic (crack-tip shielding) toughening mechanisms and intrinsic (plastic deformation) toughening mechanisms. Distinct toughening mechanisms occur at each level of bone hierarchy. Molecular uncoiling and intermolecular sliding of molecules are observed at the smallest level (see Noncollagenous proteins). Microcracking and fibrillar sliding are observed at the level of fibril arrays. At higher levels, crack bridging by collagen fibrils combines with the breaking of sacrificial bonds to increase the energy dissipation capacity of bone at the interface of fibril arrays. At the highest length scales (10–100 mm range), the primary sources of toughening result from extensive crack deflection and crack bridging by uncracked ligaments, both motivated by the occurrence of microcracking. Reprinted with permission from Launey, M.E., Buehler, M.J., Ritchie, R.O. (2010). On the mechanistic origins of toughness in bone. Annual Review of Materials Research 40, 25–53. https://doi.org/10.1146/annurev-matsci-070909-104427. Copyright (2017) Annual Reviews.

cantilever beams (10  2  2 mm in dimension) FIB-milled from a rat femur. Displacement was applied near the end of the beams using an in situ AFM with a FIB flattened AFM-cantilever tip (Jimenez-Palomar et al., 2012). Microbeams were exposed to and measured in three different environments, high vacuum, low vacuum and air, and were rehydrated in a vapor chamber between experiments. Environmental conditions were reported to have no significant effect on the elastic moduli ( 5 GPa) for all microbeams.

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Fig. 8 (Left) The energy required for crack extension in cortical bone is strongly correlated with the collagen fibril orientation angle (g). A significant jump in the crack extension energy is observed at an orientation angle of approximately 50 degrees. (Right) Cracks in cortical bone preferentially propagate through cement lines and interlamellar areas. Inserts i, ii and iii are time-lapsed AFM images of stable crack propagation (scale bar: 20 mm). (Left) Adapted from Peterlik, H., Roschger, P., Klaushofer, K., and Fratzl, P. (2006). From brittle to ductile fracture of bone. Nature Materials 5, 52–55. https://doi.org/10.1038/nmat1545. Copyright (2017) Nature Publishing Group. (Right) Reprinted with permission from Katsamenis, O.L., Chong, H.M.H., Andriotis, O.G., and Thurner, P.J. (2013). Load-bearing in cortical bone microstructure: Selective stiffening and heterogeneous strain distribution at the lamellar level. Journal of the Mechanical Behavior of Biomedical Materials 17, 152–165. https://doi.org/10.1016/j.jmbbm.2012.08.016. Copyright (2017) Elsevier.

Similar to osteons, collagen fibril orientation and tissue pathology may have an impact on the mechanical response of lamellae. These effects were investigated in successive studies using the same microbeam bending technique. Here, a range of elastic moduli from 3.7 to 11.2 GPa was reported, dependent on the alternating collagen fibril orientation within individual microbeams (Fig. 9) (Jimenez-Palomar et al., 2015a). In a further study, microbeams FIB-milled from femurs of ovariectomized (OVH) rats, an animal model for osteoporosis, were reported to have lower elastic modulus (1.59  1.26 GPa) but higher strain at failure (10  4.04%)

Fig. 9 (Left) High-resolution SEM micrographs taken after testing and FIB cross-sections showing failure modes encountered in compression tests of micropillars milled in the axial direction. Micropillars mostly deformed homogeneously and failed in shear by development of a single slip plane (Left, top). A minority of the axial pillars failed by mushrooming (Left, middle) or axial splitting (Left, bottom), which is a brittle failure mode. (Right) SEM micrographs showing (a) in situ cantilever beam testing in bending provided by the AFM tip pushing into the free end of the bone microbeam until (b) failure of the microbeam occurs. (Left) Adapted from Schwiedrzik, J., Raghavan, R., Bürki, A., LeNader, V., Wolfram, U., Michler, J., and Zysset, P. (2014). In situ micropillar compression reveals superior strength and ductility but an absence of damage in lamellar bone. Nature Materials 13, 740–747. https://doi.org/10.1038/nmat3959. Copyright 2017, Nature Publishing Group. (Right) Reprinted with permission from Jimenez-Palomar, I., Shipov, A., Shahar, R., Barber, A.H. (2015). Structural orientation dependent sub-lamellar bone mechanics. Journal of the Mechanical Behavior of Biomedical Materials 52, 63–71. https://doi.org/10.1016/j.jmbbm.2015.02.031. Copyright (2017) Elsevier.

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than control mice (2.9  1.45 GPa and 6.3  1.89%) (Jimenez-Palomar et al., 2015b). No difference in strength was reported between the control and OVH mice. Note that these beams were tested in compression, with the AFM tip applying a displacement parallel to the long access of the beam. Bending experiments characterize the behavior of a structural element, that is, beam, subjected to an external load applied perpendicular to its longitudinal axis. This results in a complex stress state within the beam, dependent on both beam geometry in a nonlinear fashion as well as material properties. As a result, bending experiments require the use of a beam theory such as Euler–Bernoulli or Timoshenko, both of which have inherent assumptions, to derive an elastic modulus. Due to these complications, uniaxial compression tests on FIB-milled samples, offering a much simpler loading case, were also investigated: Studies utilizing micropillar compression on FIB-milled ovine osteonal lamellae have been conducted in both dry (Schwiedrzik et al., 2014) and rehydrated (Schwiedrzik et al., 2017) environments (Fig. 9). Micropillars were machined from ovine tibiae in both axial and transverse directions and compression was performed both monotonically and cyclically beyond the elastic limit. Both dry and rehydrated micropillars were reported to exhibit anisotropic behavior in the elastic and postyield regime. Axial micropillars were reported to yield at higher stresses compared to transverse ones and were less likely to exhibit strain hardening postyield behavior. Rehydration reduced the yield stress anisotropy ratio between the axial and transverse micropillars. High-resolution SEM micrographs taken after testing revealed that shear via a single slip plane was the most predominant mode of failure; however, a minority of the axial micropillars exhibited brittle failure modes (Fig. 9) (Schwiedrzik et al., 2014). Note that the fabrication technique used in these studies results in a tapered cylindrical micropillar that remains attached to the underlying bone. The tapered geometry results in an inhomogeneous stress distribution over the pillar height and the substrate acts as an elastic half-space impacting the overall mechanical response of the micropillar. Luczynski et al. developed a protocol to mill out, extract, and transfer untapered micropillars with square cross sections from the lamellae of bovine tibia onto a silicon wafer, a much more rigid substrate. Here, once fixed, the micropillars were subjected to loading and unloading cycles with a flat punch nanoindenter tip (Luczynski et al., 2015). Elastic moduli, derived from unloading force–deflection curves, were reported to range from 24.1 to 32.2 GPa. Note that nearly all of the micropillars in the aforementioned studies were FIB-milled from sections of lamellae lacking canaliculi or lacunae.

Collagen Fibrils and Nanocrystals: Individual Components of Bone Beyond the individual lamellae, one approaches the individual nanoscale building blocks of bone tissue. Here, collagen, carbonsubstituted calcium phosphate nanocrystals, water, and NCPs are the most important structural elements. At this nanoscale level, these components are organized into mineralized collagen fibrils (MCFs), and the individual MCF can be regarded as the building block of bone.

Collagens and Collagen Fibrils Collagen is a superfamily of chemically distinct but closely related proteins found in different quantities in the body. In total, > 25 different types of collagen molecules exist. The most abundant of these collagen types (type I, II, III, V and XI) assemble into cylindrical-like structures, that is, the collagen fibril, with diameters varying from 20 to 500 nm and lengths up to 1 mm (Starborg et al., 2013). Collagen type I is the most abundant of the collagen molecules and is the main constituent of collagen fibrils within bone; small amounts of type III and V are also present. A characteristic of all collagen molecules is the close packing of three alpha polypeptide chains (not to be confused with a-helices) into a right-handed twisted triple helix. The triple helix is mediated by the high content of glycine amino acids (Gly). Glycines occupy every third position in the amino acid sequence of each a-chain and are always positioned toward the core of the triple helix (Brodsky and Persikov, 2005). The location and high abundance of Gly allows for the formation of a large number of hydrogen bonds (Launey et al., 2010; Ramachandran et al., 1973) as well as hydroxylation of proline and lysine residues (Kivirikko et al., 1967) between the a-chains, both of which stabilize the triple helix into a collagen molecule characterized by enhanced structural integrity.

Collagen Fibril Structure Collagen fibrils are characterized by the  67 nm D-periodicity (Kadler et al., 1996), which results from the presence of overlap and gap regions between self-assembled collagen molecules; this D-periodicity was first described by the Hodge–Petruska twodimensional model of collagen molecule packing. Orgel et al. showed that collagen fibrils are formed of quasihexagonally packed collagen molecules, which are supertwisted into a right-handed structure across the longitudinal direction of the collagen fibril (Orgel et al., 2006). Moreover, Orgel et al. proposed a microfibril model to describe the collagen fibril ultrastructure derived from X-ray diffraction patterns. The microfibril model has a triclinic unit cell (a z 40.0 Å, b z 27.0 Å, c z 678 Å, a z 89.28, b z 94.68 and g z 105.68) comprised of parts from five different collagen molecules (Fig. 10). As a result, a microfibril cannot be considered as a functional structural unit but rather as a structural model. Currently, the Orgel-model of collagen fibril structure is widely used in molecular dynamics simulations in the field of collagen biochemistry, structure, and mechanics (Orgel et al., 2006; Perumal et al., 2008; Gautieri et al., 2011; Vesentini et al., 2013).

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Fig. 10 Collagen structure. Panel (A) displays the front and side view of a collagen molecule showing the right-handed twist of the triple helix (Jmol, collagen structure from Protein Data Bank: PDB 1CAG). Panel (B) shows the Hodge–Petruska model of the collagen molecules self-assembly with overlap and gap regions. Panel (C) shows the a–b plane (top) of the Orgel microfibril model (bottom) with the quasihexagonal packing of collagen molecules. The colors in the microfibril model represent different parts or portions of the collagen molecules and the c-axis is 67.8 nm. Panel (D) shows an atomic force microscopy height topography image (top) of a collagen fibril and the corresponding height profile showing the 67 nm D-periodicity (bottom). Figure adapted from Orgel, J.P.R.O., Irving, T.C., Miller, A., and Wess, T.J. (2006). Microfibrillar structure of type I collagen in situ. Proceedings of the National Academy of Sciences of the United States of America 103, 9001–9005. https://doi.org/10.1073/pnas. 0502718103. Andriotis, O.G., Chang, S.W., Vanleene, M., Howarth, P.H., Davies, D.E., Shefelbine, S.J., Buehler, M.J., and Thurner, P.J. (2015). Structure–mechanics relationships of collagen fibrils in the osteogenesis imperfecta mouse model. Journal of The Royal Society Interface 12, http://rsif.royalsocietypublishing.org/content/12/111/20150701.abstract with permissions. Copyright (2006) National Academy of Sciences, USA.

Collagen Fibril Hydration and Mechanics Hydration water exists in two forms in collagen: bound water and unbound or freely moving water. Bound water is found in close proximity to the collagen molecule, forms an organized structure surrounding the triple helix, and helps stabilize the collagen molecule’s tight conformation (Bella et al., 1995). The unbound water is less organized and can freely move in the intrafibrillar spaces. The intrafibrillar hydration level strongly influences the intermolecular distance. Early experiments using X-ray diffraction to evaluate tendons from both murine and human donors showed that the intermolecular distance increases with increasing hydration (Price et al., 1997). More recently, AFM experiments showed that native collagen fibrils swell upon hydration (Heim et al., 2007), resulting in a three orders of magnitude decrease in the measured indentation modulus (Grant et al., 2008). Collagen fibril hydration could also be influenced by the chemistry of the aqueous environment. Transverse elasticity of individual collagen fibrils was shown to be tunable up to a sevenfold increase by merely changing the pH and ionic strength (Grant et al., 2009). The importance of hydration in collagen fibril mechanics was further revealed in a recent study on native collagen fibrils excised from the osteogenesis imperfecta mouse (OIM) model. OIM collagen fibrils were reported to have reduced hydration levels compared to wild-type counterparts. This reduced hydration was accompanied by a fivefold change in the transverse elasticity, measured using AFM cantilever-based nanoindentation experiments (Andriotis et al., 2015). Although collagen fibril transverse elasticity is influenced by ionic strength, the tensile elastic modulus was shown to be less sensitive to environmental salts in a separate study (Svensson et al., 2010). Svensson et al. proposed that this discrepancy was due to the displacement of water molecules within the fibril upon exposure to environmental salts, which primarily affect the radial properties, that is, swelling and elasticity, rather than the axial (Svensson et al., 2010). Molecular dynamic simulations on collagen microfibril models show that a loss of water results in molecular unfolding of the collagen triple helix (Gautieri et al., 2011); however, such unfolding only results in small changes in the longitudinal swelling of collagen (Masic et al., 2015) compared to swelling in the radial direction. Although small changes occur in the length of collagen molecule during dehydration, Masic et al. showed that osmotic pressure induced by partial dehydration generates tensile forces in collagen (106 pN per collagen molecule) (Masic et al., 2015).

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Mineralized Collagen Fibrils (MCFs) At the nanoscale in bone, MCFs are the structural and functional elements that contain carbon-substituted hydroxyapatite crystals. The crystals are plate-like structures with their c-axis oriented along the collagen longitudinal axis (Landis et al., 1996). MCFs are formed through a multistage mineralization process, during which the collagen and noncollagenous proteins are believed to play an active role, differing from the classical view of crystal formation (Nudelman et al., 2010). Generally, calcium-based minerals in biological materials are believed to be formed through prenucleation clusters (Gebauer et al., 2008) before they transform into a crystal (Pouget et al., 2009).

Mineralization of Collagen Fibrils As noted, the mineralization of a collagen fibril differs from the classical view of crystal formation. Collagen fibrils have been shown to mineralize in the presence of acidic proteins, for example, poly-l-aspartic acid, fetuin or NCPs (Nudelman et al., 2010; Deshpande and Beniash, 2008; Price et al., 2009). Nudelman et al. proposed a mechanism for the mineralization of a collagen fibril in the presence of polyAsp, a nucleation inhibitor, based on experiments combining cryogenic transmission electron microscopy (cryoTEM) and low-dose selected-area electron diffraction (LDSAED) (Fig. 11) (Nudelman et al., 2010). In the first stage of this mineralization process, calcium and phosphate ions form prenucleation sites with the polymer, that is, polyAsp. Then, the prenucleation site stabilizes into a negatively charged polymer–amorphous calcium phosphate (ACP) complex. This polymer–ACP complex then binds to positively charged sites in the collagen fibril located at the gap-overlap borders. Infiltration of minerals occurs at these positively charged regions, resulting in a dense network of prenucleation clusters. Subsequently, these transform into ACP clusters which finally form calcium phosphate crystals within the collagen fibril with the c-axis aligned with the collagen longitudinal axis. In addition to this intrafibrillar mineralization, collagen fibrils have extrafibrillar mineral, which is thought to be mediated via NCPs (detailed in Noncollagenous proteins). Extrafibrillar mineral was first detected in neutron diffraction studies of mineralized tendon and bone (Lees et al., 1984; Bonar et al., 1985); later, transmission electron micrographs of mineralized tendon, dentin, and bone revealed differences in orientation and position of intra- and extrafibrillar mineral (Prostak and Lees, 1996; Lees and Prostak, 1988). As a result of AFM surface maps, Sasaki et al. estimated that up to 77% of the mineral in bovine bone is extrafibrillar (Sasaki et al., 2002). Micromechanical models, designed to characterize the anisotropic behavior of heterogeneous materials such as bone, are commonly employed to investigate the impact of individual constituents on the overall mechanical behavior of bone. Early models of mineralized tissues supported the presence of extrafibrillar mineral and stressed their impact on the ultrastructural stiffness of bone (Prostak and Lees, 1996; Hellmich and Ulm, 2002a, b). More recently, micromechanical models have been used to link the apparent hardening effect in bone to energy dissipation in the extrafibrillar mineral and have been used to predict the strength characteristics of bone from different species at various hierarchical levels (Fritsch et al., 2009; Morin et al., 2017). Foundational work on quantitative validation of such models can be found in Hellmich and Ulm (2003) and Hellmich et al. (2004).

Mechanics of Mineralized Collagen Fibrils (MCFs) In situ mechanical testing combined with X-ray diffraction reveals that tensile loading on micro-sized bone samples induces a cooperative deformation mechanism between mineral crystals and collagen fibrils (Jäger and Fratzl, 2000; Gupta et al., 2006b). This results in load transfer between the inorganic (mineral crystals) and organic matrix (collagen fibril) (Jäger and Fratzl, 2000; Gupta et al., 2006b). Gupta et al. reported differences in the load transfer mechanisms between wet and dry samples. Drying results in a stiffer organic matrix (higher elastic modulus), promoting more effective load transfer from the organic to the mineral crystals. Therefore, crystals bear a higher strain fraction leading to a stiffer overall sample. Beyond elasticity, the presence of intrafibrillar mineral crystals affects yield and postyield behavior of bone at the nanoscale. Molecular dynamics simulations have revealed that MCFs have a higher yield point and greater fracture toughness compared to nonmineralized collagen fibrils (Buehler, 2007). Experimentally, the mechanical properties of native MCFs have been assessed by in situ AFM with SEM imaging (Hang and Barber, 2011). In this work, exposed MCFs on the fractured surface of bone were first imaged using SEM and then an epoxy droplet on the tip apex of an AFM cantilever was used to mechanically fix the exposed MCFs. After fixation, the AFM cantilever was used to pull the fixed MCFs while measuring the applied force. The complex composition and deformation mechanisms of MCFs resulted in bilinear stress–strain behavior, which was heavily dependent on the mineral density. All MCFs were found to exhibit an initial linear response, which the authors attributed to loads transferring from the organic to the inorganic matrix via uncoiling of the collagen molecules. Beyond the initial linear response, two distinct mechanical behaviors were observed. High mineral density MCFs exhibited strain hardening whereas low mineral density MCFs plastically deformed. Strain hardening was attributed to elevated stress transfer between collagen molecules while plastic deformation was attributed to intermolecular sliding. Both the stiffness of the initial linear-region and the behavior within the heterogeneous deformation zone were dependent on the total mineral content. This observed nano-mechanical heterogeneity aids in energy dissipation, likely contributing to the fracture toughness of bone at larger length scales. Note that the samples in this study were exposed to vacuum, resulting in considerable dehydration, thereby influencing the resulting deformation and mechanical behavior (Hang and Barber, 2011). Herein, the structure and mechanics of native and mineralized collagen fibrils have been discussed. Regarding the process of mineralization, there is evidence suggesting that NCPs could promote extrafibrillar mineralization by serving as points for

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Fig. 11 Intrafibrillar collagen mineralization. (A) Prenucleation calcium clusters (in green) form into complexes with the polymer (orange), resulting in more stable mineral droplets depicted as a red cluster (Left). Binding of the mineral droplets to a region on the collagen fibril (Right). (B) Depending on the size of the polymer, either the full droplet or only the calcium clusters enter the intrafibrillar space of the collagen fibril (Left). The intrafibrillar mineral droplets then diffuse within the fibril space (Right). (C) The mineral droplets gradually solidify into amorphous calcium phosphate (ACP; black shown Top). The last step of intrafibrillar mineralization process includes the transformation of the ACP into apatite crystals (yellow) oriented along the long fibril axis (Bottom). Figure adapted from Colfen, H. (2010). Biomineralization: A crystal-clear view. Nature Materials 9, 960– 961. https://doi.org/10.1038/nmat2911 with permissions. Copyright (2017) Nature Publishing Group. (D) A fibril mineralized for 48 h, where the deformation caused by the presence of mineral can be observed (Top). Reconstructed cryo-electron tomography images reveal plate-shaped apatite crystals (pink) embedded within the collagen matrix (Bottom). Adapted from Nudelman, F., Pieterse, K., George, A., Bomans, P.H.H., Friedrich, H., Brylka, L.J., Hilbers, P.A.J., de With, G., and Sommerdijk, N.A.J.M. (2010). The role of collagen in bone apatite formation in the presence of hydroxyapatite nucleation inhibitors. Nature Materials 9, 1004–1009. https://doi.org/10.1038/nmat2875. Copyright (2017) Nature Publishing Group.

prenucleation clusters to form. In addition to this role in the mineralization process, there is further evidence pointing to a nanomechanial function of the NCPs within bone.

Noncollagenous Proteins In addition to collagen, carbon-substituted hydroxyapatite, and water, bone consists of a small fraction (< 10 wt%) of NCPs (many of these are unstructured). These NCPs include osteopontin, osteocalcin, and bone sialoprotein, among others, and are mostly from the small integrin-binding ligand N-linked glycoprotein (SIBLING) family (Fischer et al., 2001). Initially, the scientific focus on such

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proteins was concentrated on their ability to steer and direct biomineralization. They offer attachment sites for collagen type I (Tye et al., n.d.) and nucleation sites for carbon-substituted hydroxyapatite (Hunter and Goldberg, 1993) and cells (Fischer et al., 2001) as well as being thought to control crystal size and shape (Qiu et al., 2004; Habelitz et al., 2004). In the early 2000s, high-resolution imaging via AFM revealed unstructured material phases on fracture surfaces and between mineralized collagen fibrils; these phases were attributed to the NCP fraction contained in bone and suggested a possible mechanical role of NCPs (Fantner et al., 2005). The potential mechanical role of NCPs was further supported by in vivo experiments in which patches of purified proteins exhibited an ability to repeatedly dissipate large amounts of energy during pull apart tests via sacrificial bonds and a so-called “hidden length mechanism” (Fantner et al., 2007; Zappone et al., 2008; Adams et al., 2008). These sacrificial bonds are weak, reformable bonds between neighboring mineralized collagen fibrils that break prior to the structural bonds of the mineralized collagen. The opening of these sacrificial bonds reveals hidden length and the stretching of these bonds increases the total energy needed to fracture the material, increasing overall material toughness. Studies involving animal models deficient of NCPs have also reported significant reductions in bone strength as well as fracture toughness (Duvall et al., 2007; Thurner et al., 2010; Poundarik et al., 2012), neither of which could be explained via structural or other compositional alterations (Thurner et al., 2010; Poundarik et al., 2012). More recently, comparisons between interstitial and osteonal human bone showed changes in NCP content (Sroga et al., 2011). However, whether NCP fractions change due to age and disease leading to similar significant changes in human bone material properties has not been investigated to date. This is despite reports that the protein matrix in bone does indeed change with age and disease (Grynpas et al., 1994). It may well be that the reported shift from diffuse damage to microcrack accumulation associated with aging bone is related to an increase in bone brittleness and fracture risk with age (Diab and Vashishth, 2007). The role of NCPs in this phenomenon is not yet known. An additional finding of interest in this context is that diffuse damage in bone heals without remodeling (Seref-Ferlengez et al., 2014). One may speculate that this is an effect mediated by NCPs; however, to this point there exists no data to prove or disprove this hypothesis.

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Wang, X., Sudhaker Rao, D., Ajdelsztajn, L., Ciarelli, T. E., Lavernia, E. J., & Fyhrie, D. P. (2008). Human iliac crest cancellous bone elastic modulus and hardness differ with bone formation rate per bone surface but not by existence of prevalent vertebral fracture. Journal of Biomedical Materials Research Part B: Applied Biomaterials, 85B, 68–77. https:// doi.org/10.1002/jbm.b.30918. Weaver, J. K. (1966). The microscopic hardness of bone. The Journal of Bone & Joint Surgery, 48, 273–288. Weiner, S., & Wagner, H. D. (1998). THE MATERIAL BONE: Structure-mechanical function relations. Annual Review of Materials Science, 28, 271–298. Yamada, S., Tadano, S., & Fukuda, S. (2014). Nanostructure and elastic modulus of single trabecula in bovine cancellous bone. Journal of Biomechanics, 47, 3482–3487. https:// doi.org/10.1016/j.jbiomech.2014.09.009. Yeni, Y. N., & Norman, T. L. (2000a). Calculation of porosity and osteonal cement line effects on the effective fracture toughness of cortical bone in longitudinal crack growth. Journal of Biomedical Materials Research, 51, 504–509. https://doi.org/10.1002/1097-4636(20000905)51:33.0.CO;2-I. Yeni, Y. N., & Norman, T. L. (2000b). Fracture toughness of human femoral neck: Effect of microstructure, composition, and age. Bone, 26, 499–504. https://doi.org/10.1016/ S8756-3282(00)00258-1. Zappone, B., Thurner, P. J., Adams, J., Fantner, G. E., & Hansma, P. K. (2008). Effect of Ca(2 þ) ions on the adhesion and mechanical properties of adsorbed layers of human osteopontin. Biophysical Journal, 95, 2939–2950. https://doi.org/10.1529/biophysj.108.135889. Zebaze, R. M. D., Jones, A., Welsh, F., Knackstedt, M., & Seeman, E. (2005). Femoral neck shape and the spatial distribution of its mineral mass varies with its size: Clinical and biomechanical implications. Bone, 37, 243–252. https://doi.org/10.1016/j.bone.2005.03.019. Zysset, P. K. (2009). Indentation of bone tissue: A short review. Osteoporosis International. https://doi.org/10.1007/s00198-009-0854-9. Zysset, P. K., Edward Guo, X., Edward Hoffler, C., Moore, K. E., & Goldstein, S. A. (1999). Elastic modulus and hardness of cortical and trabecular bone lamellae measured by nanoindentation in the human femur. Journal of Biomechanics, 32, 1005–1012. https://doi.org/10.1016/S0021-9290(99)00111-6.

Further Reading Currey, J. D. (2002). Bones: Struture and mechanics. Princeton: Princeton University Press. ISBN: 9781400849505. Fratzl, P. (2008b). Collagen: Structure and mechanics. New York: Springer. https://doi.org/10.1007/978-0-387-73906-9. Reznikov, N., Shahar, R., & Weiner, S. (2014). Bone hierarchical structure in three dimensions. Acta Biomaterialia, 10, 3815–3826. https://doi.org/10.1016/j.actbio.2014.05.024. Rosen, C. J., Bouillon, R., Compston, J. E., & Rosen, V. (2013). Primer on the metabolic bone diseases and disorders of mineral metabolism. Wiley. https://doi.org/10.1016/ S0021-9290(00)00074-9.

Cell Adhesion: Basic Principles and Computational Modeling Diego A Vargas and Hans Van Oosterwyck, Biomechanics Section, KU Leuven, Heverlee, Belgium © 2019 Elsevier Inc. All rights reserved.

Introduction Basics of Cell Adhesion Mechanotransduction The Bell Model: Motivation Behind a First Model Transcending Scales: The Subcellular and Multicellular Directions The Integrin–Ligand Bond in More Detail The Impact of Force Spectroscopy Zooming In: Other Players at the Molecular Scale Zooming Out: Cell Aggregates, Cell Sheets, Tissues Refining Understanding Through Experiments Imparting Forces at a Cellular Scale Measuring Forces and Imaging at a Cellular Scale Challenges and Conclusions Further Reading

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Glossary Actomyosin contractility An actin network’s ability to contract due to pulling of myosin on actin filaments in the network through a chemical-energy spending mechanism. Agent-based model Phenomenological model that treats cells as individual units interacting according to a set of rules. Apical domain Refers to cellular region opposite to the basal lamina in an epithelial layer. Physiologically it is exposed to a lumen. It is rich in junctional proteins, making it distinct compositionally, structurally, and functionally from the lateral and basal domains. Cellularized materials Cell aggregate connected by intercellular junctions that interconnect internal cytoskeletons. Includes cell sheets, cysts, amorphous aggregates, cancerous tumors, etc. Continuum model Model in which matter is represented as a continuous homogeneous substance. Continuum models are often deterministic in nature and defined by a set of analytic equations. Force spectroscopy Category of techniques that measure forces in biological or molecular systems. Glycocalyx Carbohydrate-rich layer connected to the cell membrane through proteoglycans and glycoproteins. Ground-state configuration Configuration of a system characterized by the lowest possible energy value. Mechanosensitivity Ability to sense a mechanical stimulus. Mechanotransduction Process of converting a mechanical signal into a chemical signal. Multiscale model Model that integrates processes that occur at different temporal or spatial scales.

Introduction As a physical system, a cell’s interaction with its environment is dictated by electrical, chemical, and mechanical interactions of the different components in the cell. Its response in turn affects interaction of the cell with the outside through feedback mechanisms. The interface of the cell with its environment is its plasma membrane, rich in proteins that mediate this signal transduction. Those involved in transduction of signals are transmembrane proteins that create a bridge between the extracellular and intracellular worlds; however, these constitute only one piece of the puzzle. In the case of mechanical signals, these proteins will transmit forces inside the cell to the cytoskeleton, the structure responsible for the shape and deformation of the cell. On the outside these proteins bind the structural elements of the environment or other cells. This is not a straightforward process though. On the outside of the cell, these proteins are modified changing how they interact with the environment. While inside, they do not bind the cytoskeleton directly; multiple proteins make up this connection, many of them with their own complex and dynamic structures. A study of cell adhesion must account for chemical as well as mechanical interactions of the cell and cellular components. Chemical interactions have been studied for cellular components extensively using traditional binding kinetics assays. These experiments, however, looked at proteins in isolation. Mechanical studies on the other hand are relatively recent and initially suffered from a paradoxical limitation, they could only quantify the response of a cell as a whole. A theoretical approach through mathematical

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modeling was able to bridge this gap. Since the development of the first theoretical model, the different pieces that had been uncovered by biochemists (e.g., diffusion rates along the plasma membrane and binding kinetics of protein pairs) were put together to explain adhesion at the cell level. These initial theoretical studies were able to predict the existence of adhesion dynamics later discovered, such as the catch bond in rolling leukocytes in our immune system. Computational modeling has only expanded the use of theoretical models to study cell adhesion. The ability to use numerical analysis as well as increased computational power has permitted the use of increasingly comprehensive models. This has allowed computational modeling to keep up with the advancement of experimental techniques, playing the role of an equal partner in development of model-driven hypotheses. Modeling techniques from different fields ranging from N-body simulations to continuum mechanics have been used to mechanically model from a single adhesion receptor, through an entire cell, to cell aggregates.

Basics of Cell Adhesion In an organism, most cells are found in a scaffold of proteins and polysaccharides known as the extracellular matrix (ECM). The exact composition of the ECM is tissue-dependent: it is heterogeneous in composition, and its function ranges from providing structural integrity and aiding in tissue organization to signaling, guiding cell survival, migration, proliferation, and differentiation, and survival. The primary structural components of the ECM are fibrous proteins, which include collagen, elastin, fibronectin, and laminin. Collagen, found in high abundance in many tissues, is a protein that self-assembles triple-helical molecules into thicker fibrils. It is an adhesive component of the ECM, meaning cells can attach to it. The space in between the multiple proteins is occupied by hydrophilic glycosaminoglycans and water. Based on its composition, mechanical properties of the ECM change, providing a series of binding sites and a barrier to free diffusion of signaling molecules. Integrins constitute the principal transmembrane adhesion molecule. An integrin receptor is active as a heterodimer, formed by selective pairing of 1 of 18 known a-subunits and 1 of 8 known b-subunits, producing 24 distinct receptors. Named aptly, these receptors integrate the intracellular and extracellular environments by linking the cytoskeleton to the ECM and provide a bidirectional signaling path. Integrins bind specific amino acid sequences in some of the ECM fibrous proteins, notably the RGD motif found in fibronectin, laminin, and, under some conditions, collagen. Integrins are the molecule modeled the most in computational studies. Nonetheless, it is not alone in forming an adhesion complex. The complete transmembrane macromolecular complex is known as a focal adhesion (FA); and it is formed upon maturation of the nascent complex formed when integrins bind a ligand in the ECM. Similarly, intercellular adhesions consist of protein complexes with a transmembrane protein at their core. Based on their composition, different types of intercellular adhesions exist in different tissues; these have differing degrees of strength and are involved differently in tissue organization and signaling. These include gap junctions, tight junctions, adherens junctions (AJs), and desmosomes. Gap junctions provide communication for chemical and osmotic regulation of the intracellular environment. In a complementary role, tight junctions have a barrier function, preventing lateral diffusion of molecules in the extracellular space between adhered cells, allowing for distinct chemical environments to be maintained. AJs and desmosomes anchor cells and provide mechanical communication between cells. Further, desmosomes form patches that link intercellular junctions to intermediate filaments, while AJs exist along the cell–cell boundary and bind to the actin cytoskeleton to form adhesion belts. In both AJs and desmosomes, members of the cadherin protein family bind to each other on opposing cells, forming homotypic bonds. Based on the tissue type, different members of the cadherin family are found. This family includes N-cadherin, common in neural tissues, P-cadherin, originally found in the placenta but increasingly in other tissues, VE-cadherin in endothelial cells, and E-cadherin in epithelium. Cadherins consist of a cytoplasmic domain, a transmembrane domain, and an extracellular region comprising multiple ectodomains. Along with cadherin, AJs also include catenin molecules. Catenin molecules are the link between AJs with the cytoskeleton. The individual or joint presence of the different catenin molecules in AJs indicates stability and maturity of junctions. b-catenin, gcatenin, and p120-catenin can bind cadherin, with g-catenin substituting b-catenin as a junction matures. a-Catenin can bind either b-catenin or g-catenin and the membrane–cytoskeletal protein vinculin: vinculin in turn binds actin, as it occurs in FAs, demonstrating a similarity between mechanisms of cell–cell and cell–ECM adhesion.

Mechanotransduction The structural commonalities between FAs and AJs are evident: they both share the structural components, bind the actin cytoskeleton, carry mechanical signals from the outside to the inside of the cell, and transform a mechanical stimulus into a chemical one. The latter process is known as mechanotransduction. Thus the role of mechanical stimulation of cells has implications not only in the shape and structural conformation of the cells, but also in biological processes. The targets of mechanotransduction include regulatory proteins involved in determining cell growth, proliferation, differentiation, apoptosis, and organogenesis. There is extensive evidence of the role of FAs and AJs in mechanotransduction. For example, FAs only mature once the cytoskeleton is bound, at which point integrins cluster increasing the size of the FA. Sufficient force is necessary for the FA to mature and become stable; otherwise, the initial complex between integrins and the ECM will disintegrate. Proteins with mechanosensitive functions include talin, vinculin, p130Cas, zyxin, and filamin A. The cytoskeleton plays the biggest role. For example, tension on a FA could alter the conformation of scaffolding proteins; upon being stretched, talin reveals

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a binding site for vinculin which leads to the interaction of these proteins. In the case of p130Cas, tension increases its phosphorylation status, which in turn activate multiple Rho GTPases (including Rac) involved in proliferation, differentiation, cell–cell adhesion, and migration. Zyxin accumulates in areas of the actin fibers with large strains and recruits proteins involved in fiber repair. Once a FA has reached maturity, it will activate additional proteins such as focal adhesion kinase (FAK), which is part of intracellular signaling pathways involved in differentiation and migration. Tension on AJs also activates FAK, thus creating an interplay between FAs and AJs. These observations make it clear that a study of cell adhesion must incorporate chemical kinetics as well as mechanics. This was observed in the first theoretical models and maintained through computational models.

The Bell Model: Motivation Behind a First Model George Bell introduced in 1978 the first theoretical framework for the study of cell adhesion by considering spatial limitations (e.g., molecular orientation, membrane diffusion) in the reversible binding of a receptor and a ligand as well as the potential diversity of molecular mechanisms that could exist. Although cells were known to have an electric charge, observations had made it evident that nonspecific binding due to electrical forces are less relevant than specific-binding forces. The model is built from the assumption of specific binding. The first steps in building a model were to define three aspects of the adhesion molecules: 1. Binding affinity: Described by rates of bond formation and breaking from elementary rate constants 2. Number of binding molecules: The number of receptors per unit surface area of the plasma membrane 3. Mobility along the membrane: Described by the rates of diffusion of the receptors along the membrane Although Bell applied the same principles to cell–cell and cell–substrate adhesion, here the example of two adjacent cells having complementary receptors is presented. The reaction is conceptually separated into two processes: The two adhesion molecules in two neighboring cells encounter each other by coming into proximity (i.e., diffusion along the corresponding cell membranes) and then create the adhesion complex. This initial encounter is modeled through conception of an intermediate “encounter” complex, as in enzyme kinetics models (i.e., Michaelis–Menten). This reaction was written by Bell as: N1f þ N2f

dm rþ þ % N1 N2 % Nb dm r 

(1)

N1 and N2 are the total number of receptors per unit area (i.e., density) of membrane of cells 1 and 2, respectively. Nif is the density of free receptors in cell i (i ¼ 1,2), N1N2 the encounter complex, and Nb the density of bound receptors. Given our current knowledge of adhesion complexes, N1f and N2f can be taken to represent cadherins, and consequently, Nb represents an immature AJ. The rates d þ m and d  m represent the rates of formation and dissolution of the encounter complex in a membrane, respectively, and are thus dependent on the translational diffusion constants for free receptor motion in the membrane (Dm(Nif)) and an encounter distance at which the molecules can form an encounter complex (R1,2):
    dm (2) þ ¼ 2p Dm N1f þ Dm N2f
    2 dm  ¼ 2 Dm N1f þ Dm N2f R1;2

(3)

The rates rþ and r represent the forward and reverse rate constants for formation of the adhesion complex (Nb) from the encounter complex (N1N2). Under assumption that the encounter complex is transient and exists in a very small concentration relative to the unbound concentrations of the receptors (N1f, N2f) and the adhesion complex (Nb), then by setting dN1N2/dt ¼ 0, Eq. (1) can be simplified to: kþ N1f þ N2f % Nb k

(4)

The rate constants for this overall reaction of cell–cell adhesion formation (kþ, k) can be described in terms of the rate constants of the intermediate steps: kþ ¼

dm þ rþ dm  þ rþ

(5)

k ¼

dm  r dm  þ rþ

(6)

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Since the encounter complex is assumed to exist in small concentration, then Ni ¼ Nif þ Nb (i ¼ 1,2). The rate of adhesion complex formation is shown to be: dNb ¼ kþ ðN1  Nb ÞðN2  Nb Þ  k Nb dt

(7)

This equation provided a theoretical description of the temporal evolution of adhesion between cells with a homogeneous distribution and a constant number of adhesion molecules. It also provided an equilibrium concentration of cell–cell adhesions depending on binding dynamics. Even though up to that point the proposed model was novel and revealing, it was solely based on diffusion and chemical kinetics. At the time Bell wrote this, there were already experiments that demonstrates the relevance of mechanics on cell adhesion. Experiments would track how many cells (lymphocytes or fibroblasts) would remain attached to lectin-coated fibers when shaken (with specified amplitude and frequency) in a fluid. Bell went on to consider adhesion force and provided a way to couple forces and chemical kinetics. Bell theorized that a force must be present to separate two cells: the probability of all adhesion complexes, which keep two cells together, dissociating simultaneously is extremely small. To consider the effect of force on the rate of bond formation Bell postulated that kinetic theory of the strength of solids could be applied to receptor–ligand bonds. In this theory, the lifetime of a bond (between atoms in a solid) as a function of the force applied on the bond (f) is given by: sð f Þ ¼ s0 exp½  ðgf  E0 Þ=kT

(8)

The lifetime of a bond (once formed) is dependent on a natural lifetime of the bond (s0) dependent on the frequency of oscillation of atoms, the energy of a single bond (E0), and is scaled by the Boltzmann constant and temperature (kT). Additionally, g is an empirical parameter (with units of length) accounting for structure of the solid. The negative exponent shows how increased force reduces the lifetime of the bond (this behavior is used to classify this bond as a slip bond). This equation can be integrated into Eq. (7) to be used in the context of cell adhesion by identifying that the inverse of the lifetime describes the rate of dissociation (i.e., s(f ¼ 0) ¼ (k) 1). Also, the force separating two cells (F) is taken to be distributed equally between all adhesion complexes, therefore (f ¼ F/Nb). The resulting equation is:   dNb gF ¼ kþ ðN1  Nb ÞðN2  Nb Þ  k Nb exp (9) kT Nb dt This is the first equation to take into account the effect of force applied on the cell in adhesion. Beyond this Bell mentioned other more temporary factors that may affect binding, he termed these transients. In a display of premonitory abilities, Bell considered transients such as the adsorption properties of ligand molecules for the case of cell–substrate adhesion, orientation of adhesion molecules with respect to each other when they first encounter each other, and glycosylation of membrane proteins. All of these transients would be explored in future studies by other researchers. In this way Bell’s work could be considered visionary. As will be presented, today most models of adhesion are very specific to a cell type under specific conditions; nonetheless, there is Bell to thank for the underlying theory.

Transcending Scales: The Subcellular and Multicellular Directions In a fully formed organ, it is not difficult to conceive that the stroma and the parenchyma (structural and functional tissue, respectively) would have different mechanical properties and that this fact plays a functional role. Given that many cells are sensitive to force, do the mechanical properties of the distinct tissue determine the functionality of cells composing it, or alternatively, are mechanical properties of the tissue a consequence of the identity or role of cells present? Given what we know about mechanotransduction, both possibilities are true. Observations in developmental processes such as delamination of the neural crest or metastasis in cancerous tumors, have given credit to the theory that differential adhesion of cells in tissue can be attributed the emergence of supracellular organization (differential adhesion hypothesis or DAH). During these processes of cellular rearrangement, the same cell can go from stiff and strongly attached to soft and detached. Further analysis of these processes revealed that there are distinct molecular signatures of cells associated with the different cellular behaviors and ability to adhere. Moving forward, it will become evident how after Bell’s seminal study theorizing about a single cell binding a surface or two cells binding each other (referred to as cellular scale), researchers started thinking about the effects of mechanical properties at smaller and larger spatial scales. At the subcellular scale, the mechanical properties of the plasma membrane or the response to force of single molecules in the adhesion complex were explored and incorporated in cell adhesion models. The mechanics of subcellular elements are considered all the way down to the atomic scale: although not considered here, steered molecular dynamics simulations look at the effect of forces on single adhesion molecules using force fields to simulate the change in protein configuration with force application. In the opposite direction, there is the multicellular scale. Models of multiple cells, either forming a monolayer or a 3D mass, have been developed explicitly incorporating cellular adhesion. These models can be used to study the mechanical (and functional) properties of tissue.

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The Integrin–Ligand Bond in More Detail Dembo and colleagues set up a model to describe adhesion dynamics in a particular theoretical experiment, peeling a cell (or more precisely a fragment of membrane) from a substrate, just as a strip of tape is peeled off a surface (Fig. 1). This “peel test” is an empirical test used to assess the effectiveness of adhesives in industrial applications. Their purpose was to use the peel test as a means of scientific exploration of biological adhesion and develop an analytical expression for steady-state peeling velocity for a biological system (i.e., cell membrane). To do this, the equations for deformation of an elastic membrane are coupled with the equations for the chemical kinetics of the adhesion molecules. This contrasts the Bell approach by including mechanical properties of the membrane and describing the change in binding energy not just in terms of potential difference of bond breaking but mechanical parameters of the adhesion molecule. Fig. 1 shows the setup for the theoretical experiment. Every point in the membrane is tracked by using an arc length coordinate s: the clamped end is designated by (s / -N) and the end being pulled by (s / þN). The point at which attachment begins is designated by (s ¼ 0), and the peeling velocity (Vpl) is defined in terms of the change in position of this point of detachment. As peeling is a time-dependent process, the shape of the membrane at a time t is taken relative to the point at which the membrane detaches (s ¼ 0). Thus the position of the membrane is described in Cartesian coordinates by x(s,t) and y(s,t); the x-axis is the surface of attachment, and the y-axis describes the distance of the membrane to the surface. Tension applied on the free end of the membrane is designated Tfx, and it is applied at an angle qfx with respect to the surface. At each point along the length of the membrane, the density of adhesion molecules is considered: Nb(s,t) (in keeping with the notation used to present the Bell model). Consequently, it can be stated that for the every point in the membrane: N1 ¼ N1f(s,t) þ Nb(s,t). Because here the cell is attached to a surface, and not to a second cell, adhesion is assumed to be dependent solely on the presence of the adhesion molecule on the membrane (i.e., no ligand molecule considered): kþ ðyÞ N1f ðs; t Þ % Nb ðs; t Þ k ðyÞ

(10)

vNb ðs; t Þ ¼ kþ ðyÞN1  ½kþ ðyÞ þ k ðyÞNb ðs; t Þ vt

(11)

Eq. (7) is modified accordingly:

In contrast to Eq. (7), Eq. (11) shows the rate constants for the overall reaction as a function of the distance between the membrane and the surface (given by coordinate y(s,t)). Because adhesions are modeled as Hookean springs, Dembo and colleagues

Fig. 1 Schematic describing hypothetical setup modeled by Dembo et al. in which a fragment of cell membrane is peeled from a surface by applying a tension Tfx on one end of the membrane while the opposite end is clamped. The membrane is attached to a surface via adhesion proteins that have the mechanical characteristics of a Hookean spring and adhesion rates described by chemical kinetics (Bell model). The surface is taken to be in the x-axis, while the y-axis describes the distance of the membrane from the surface. Tension is applied at a particular angle to this axis (qfx). Reproduced from Dembo, M. et al. (1988). The reaction-limited kinetics of membrane-to-surface adhesion and detachment. Proceedings of the Royal Society B: Biological Sciences 234, 55–83.

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state that for thermodynamic consistency one must consider the change in a bond’s Gibbs free energy with the bond stress. The relation between energy and stress is derived from the dependence on the bond’s free energy on bond strain: Eh ¼ E0 þ 0:5kðy  l0 Þ2

(12)

E0 represents the Gibbs free energy of an unstretched spring (same quantity as an Eq. 8), k is the Hookean spring constant, and l0 represents the resting length. In determining the rate constants, Dembo and colleagues consider the existence of a transition state in bond formation. They make two assumptions: First, the free energy of a molecule that is not attached is constant and independent of the distance to its ligand, and second, the transition state is also described by a Hookean spring and thus has the same form of free energy (Eq. 12). In this way they apply the transition state theory and arrive at the following rate constants:   0:5kts ðy  l0 Þ2 (13) kþ ðyÞ ¼ kþ ðl0 Þexp  kT k ðyÞ ¼ k ðl0 Þexp

  0:5ðk  kts Þðy  l0 Þ2 kT

(14)

kts is the spring constant of the transition state, in contrast to that of the formed bond (k). Theoretically, this raised the possibility of the existence of a bond for which the rate of disruption under tension is lower than when unstressed (this behavior is used to classify this bond as a catch bond). This occurs when k < kts. These bonds would be directly observed much later in leukocytes under various degrees of force in microfluidic chambers. In fulfilling their objective of analytically describing peeling velocity, Eq. (11) is added a convection term describing the loss of adhesion molecules due to peeling: vNb ðs; t Þ vN ðs; t Þ ¼ Vpl b þ kþ ðyÞN1  ½kþ ðyÞ þ k ðyÞNb ðs; t Þ vt vs

(15)

The researchers go on to assume steady-state (v Nb(s,t)/vt ¼ 0), expanding Nb(s,t) as a power series in Vpl, and substitute with Eqs. (13) and (14). This provides the following relation:   (16a) Nb ðs; t Þ ¼ ab0 þ ab1 Vpl þ ab1 Vpl 2 . ¼ ab0 þ O Vpl

ab0

h 2i 0Þ N1 Keq exp  0:5kðyl kT i h ¼ 0:5kðyl0 Þ2 1 þ Keq exp  kT

; Keq ¼

kþ ðl0 Þ k ðl0 Þ

(16b)

Dembo and colleagues separate the tangential and normal components to stress for the Hookean springs along the membrane. The tangential component, scaled by the density of adhesions at the specific location along the membrane, is given by: stan ¼ Nb ðs; t Þ k ðy  l0 Þ

vy vs

(17)

They then proceed to take advantage of another definition of tangential stress that uses the mechanical properties of the membrane itself, specifically the modulus of bending (Mb):   v T þ 0:5Mb C2 stan ¼  (18) vs C represents the curvature of the membrane and T the tension. Finally, setting Eqs. (17) and (18) equal to each other and substituting Nb(s,t) with the relation (16a and b), Dembo and colleagues arrive at the analytical description of the peeling velocity in terms of both mechanical properties of the membrane and adhesion molecules:     vy v T þ 0:5Mb C2 O Vpl ¼ ab0 kðy  l0 Þ  (19) vs vs As a premise to their approach, the researchers reject the concept of using a continuous adhesive because of the relatively low number of adhesion molecules found in a cell. They thus provide a way to spatially and temporally describe adhesion while accounting for properties of the different components, in their case membrane and adhesion molecules. However, as discussed previously, bonds between cells or between a cell and a substrate are not formed by a single pair of molecules, but they are dynamic molecular complexes. Their composition may change spatially along the cellular surface, and in turn change the mechanical properties of the membrane. They can also not only bind to similar complexes on other cells or surfaces (heterotypic binding), but can bind to each other at different sites (homotypic binding) while still being able to form heterotypic bonds. The diversity of binding mechanisms and complexity of force response became more evident when experimental techniques were developed exclusively to study the binding forces between individual molecules. This category of techniques is known as force spectroscopy.

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The Impact of Force Spectroscopy Force spectroscopy seeks to quantify the strength of individual bonds by looking at a spectrum of bond rupture force under tension. The specifics of force spectroscopy will be explored in the next section (“Refining Understanding through Experiments” section), where methods that measure response at the single cell level or even single molecule level (single molecule force spectroscopy (SMFS)) will be presented. In the context of modeling, these techniques led for the first time to a less hypothetical consideration of the effects of experimental factors on cell binding dynamics and the effect of individual bonds on one another in multiple bond attachments. Evans presented some of the pioneering work in SMFS beginning in 1997, with quantitative work on the bond lifetime under tension. At that time it was possible to make measurements of single bonds. He developed a theoretical explanation for the force response observed to increasing loading rates (force applied in time). Detachment force (or bond strength) had been shown to be dependent on loading rate. What Evans noted was that the spectrum of bond rupture force as a function of loading rate can be interpreted as a landscape displaying prominent energy barriers. This energy landscape reveals barriers that arise from molecular interactions that are “difficult or impossible to detect in assays of near equilibrium dissociation,” in Evans’ words. These barriers are what determine bond lifetime. Experiments quickly showed that the response of a bond to tension was not linear; therefore, a Hookean spring is not an appropriate model. Taking the adhesion molecules to be highly flexible polymers, a better description of the dependence of force (f) on adhesion molecule total length (x) is provided: kT

f   cb a 1  Lxp

(20)

Lp and b are the contour length and persistence length of the adhesion molecule, respectively. Meanwhile, a and c are constants used to describe the specific polymer and distinguish between different behaviors (e.g., a ¼ 1 and c ¼ 1 describe the polymer as a freely jointed chain). Once again kT represents the Boltzmann constant multiplied by the temperature. In an experiment where the separation speed (vs) is made to be constant, then the loading rate will be time dependent (df/dt ¼ rf(t)):  a kT c Lp b vs (21) rf ðt Þ ¼  aþ1 1  vLspt This is postulated for a single bond, and as such the transition and binding equilibrium is not presented in terms of adhesion density but rather adhesion probability for a single bond. At each point in time, an adhesion molecules is said to have a probability of being unbound Sf(t) or bound Sb(t): Sf(t) ¼ 1  Sb(t) (as there is no other possible state). The forward and backward transitions between these states are described by the rate constants kþ(t) and k(t). Therefore, reminiscent of Eq. (10), binding and unbinding is described by kþ ðt Þ Sf ðt Þ % Sb ðt Þ k ðt Þ

(22)

Since Bell postulated that the kinetic theory of the strength of solids could be applied to receptor–ligand bonds (Eq. 8), a relation between force and unbinding rate had been proposed. In seeking to make a theoretical basis for bond behavior observed in force spectroscopy experiments, in which the force applied is not constant, Evans was confronted with a system where force evolves in time. This evolution is described by the loading rate, which implies that both the state probabilities and the rate constants can be considered functions of force. Since the loading rate describes how time and force are related (df ¼ rf(t)dt), an equation describing the change in binding probability with force can be postulated: dSb ½k ð f Þ þ kþ ð f ÞSb ð f Þ kþ ð f Þ ¼ þ rf ð f Þ rf ð f Þ df

(23)

Analogous to Eqs. (9) (from Bell) and (11) (from Dembo and colleagues), this equation provides a quantitative relation between binding dynamics and mechanical properties of the adhesion molecule. Evans’ equation is much more complete in that it accounts for the different response to force by the molecule (i.e., rf(t)), the rate constants (i.e., kþ(f),k(f)), and the system as a whole (i.e., Sb(f)). Even when considering disruption of a single cellular bond, Evans believed that it is naïve to model the shortening of a bond’s lifetime under external force as the lowering of a single activation (energy) barrier. There are in reality many widely distributed atomic-scale interactions that create a complex energetic landscape. Evan provides the theoretical tools to model the response of different proteins given their conformation, but also brings to our attention the large set of factors that come into play. There is no general description for binding kinetics of proteins. Taking all of these factors into account, it is evident that binding kinetics can only be known for the specific condition measured. Extrapolations to different applied forces or force-rates cannot be assumed correct.

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Zooming In: Other Players at the Molecular Scale As much as Evans explores in detail the effect of adhesion molecule architecture on binding dynamics, he did not consider a biologically active system where multiple molecules come into play. His work can be applied to an adhesion complex as a whole, yet he did not address this explicitly as his work is very general to molecular adhesion in general. Subsequent models focusing precisely on cell adhesion would address this; In one of the most refined models to date, Paszek and colleagues simulate an area of the cell membrane explicitly considering individual integrin molecules, the glycocalyx (a carbohydrate-rich layer connected to the cell membrane through proteoglycans and glycoproteins), and ligand molecules in the adjacent fraction of ECM. The model is set up under the hypothesis that when an adhesive bond is formed, it induces a mechanical deformation of the membrane and ECM, such that it modifies the distance between them. This change in distance can be cooperative to cell adhesion due to mechanical coupling. The glycocalyx is found in some bacterial cells as well as some mammalian tissues, including epithelial and endothelial. Despite its function not being fully characterized, the large nature of carbohydrates has prompted theoretical studies on steric effects on cellular systems and its role in disease. Rather than limiting their inquiry to the force felt at the receptor ligand complex, Paszek and colleagues look at the source of the force itself by incorporating the compliance of the materials where ligand and receptor molecules are embedded, cell membrane and ECM, respectively. The model considers the possibility that integrin clustering can occur both from factors that control cell adhesion (e.g., ligand binding dynamics, matrix stiffness) and factors that mediate repulsion (e.g., glyxocalyx). It also seeks to determine how integrin clustering affects adhesion between the cell and ECM. The model describes integrin–ligand bonds as Hookean springs with distance-dependent kinetic rates of bond association and dissociation, just as Dembo and colleagues (Eqs. 13 and 14); however, it incorporates a lattice spring model (LSM) to represent both the cell membrane and ECM. This consists in describing a body as a series of lattice points connected by harmonic springs; it is a computationally efficient mesoscopic approach commonly used to study fracture mechanics. A schematic of the model used by Paszek and colleagues is shown in Fig. 2. A section, with an area of 1.4  1.4 mm2, of the interface between cell membrane and ECM is modeled as a single spring network. Two separate LSM networks are used to represent the cell membrane (including the cortex) and the ECM, with a thicknesses of 40 and 400 nm, respectively. This corresponds to 70  70  3 lattice points for the former and 70  70  21 for the latter: The distance between nodes (Dx) and the stiffness of the component springs (s) were selected such that the Young’s modulus (E) could be adjusted according to the following equation: E¼

5s 2 Dx

(24)

Integrin molecules, ECM ligand molecules (i.e., integrin binding sites), and the glycocalyx are located in between the two LSM networks. Integrin molecules are bound to the top network (membrane and cortex); however, these can move off lattice when not bound to the ECM. Integrin diffusion along the cellular membrane is modeled by a set of hop reactions in which the time evolution of the system is given by a Gillespie algorithm: a dynamic Monte Carlo method allowing for discrete stochastic simulation where both the time step duration and actual diffusion reactions taking place are variable at each step and calculated from random numbers. The ligand molecules are fixed in certain nodes of the top of the spring network representing the ECM. Finally, the glycocalyx is also represented as a series of Hookean springs always at node positions; these connect the top and bottom networks into a unified mesh, for which forces are solved for each node based on minimization of the system’s energy. With this model, mechanical properties of different cellular components (i.e., glycocalyx stiffness, glycocalyx thickness, and cell cortex rigidity) are be modified. The effects of these changes are quantified in terms of cell adhesion, specifically the degree of integrin clustering at the membrane and the cooperativity in binding. The authors found that cooperativity in integrin binding increased with glycocalyx thickness, and that integrin clustering is primarily driven by bond formation. Similarly, an increased glycocalyx stiffness caused enhanced cooperativity in integrin binding to the ECM and integrin clustering, as well as the formation of denser integrin clusters. Finally, the study found that a minimum value of cortex rigidity is necessary to support cooperative binding of integrins.

Zooming Out: Cell Aggregates, Cell Sheets, Tissues In the field of developmental biology, large homogeneous cell aggregates are studied for their composition, organization, and mechanical properties. These systems can be grouped into a category of cellularized materials, which consist of a large number of cells connected by intercellular junctions, linking neighboring cytoskeletons, providing an avenue for force transduction across multiple cells. Forces applied on the ECM also become relevant to collective displacement, particularly in the case of epithelial sheets. To capture their mechanical properties, these materials were initially studied as bulk materials. They were modeled through continuum approaches, as was the case of tumor modeling. However, discretization of the bulk material is important when the material is active, as is the case of living tissue. This led to the application of agent-based models for the study of cellular systems with tumor growth models dominating the field. In a popular modeling approach, individual cells are reduced to single points representing the center of mass. This is useful when studying very large systems with hundreds or thousands of cells; however, mechanical interactions are simplified to pairwise interactions described in the form of a force potential or Hertz contact mechanics (in which cellular interactions are modeled after that of elastic spheres). These approaches have provided insight into phenomena such as collective cell dynamics and the emergence of

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Fig. 2 Schematics showing different scales of cell adhesion models. Images A and B show a model at the subcellular level, representing a section of membrane and ECM of surface area 1.4  1.4 mm2. Image C shows three isolated cells that form part of a larger hexagonal lattice used to represent an epithelial sheet. (A) Depiction of cell–ECM interface. Integrin molecules are on the bottom face of the membrane, while ligand molecules are on the top face of the ECM. (B) The lattice spring model (LSM) representation of a 2D section of the cell–ECM section. Integrin molecules and the glycocalyx are represented by springs as well. Integrin molecules are not spatially constrained to lattice nodes, while ligand molecules and the glycocalyx are. (C) Schematic showing apicolateral enrichment of cytoskeletal and junctional components in cells packed in epithelia (top). This packing is modeled in 2D by using polygonal cells (numbered a) described by vertices (i,j) connected by cellular edges (lij) (bottom). Configuration of the polygonal network is described by an energy function. (A and B) Reprinted from Paszek, M. J. et al. (2009). Integrin clustering is driven by mechanical resistance from the glycocalyx and the substrate. PLoS Computational Biology 5(12), e1000604; (C) reprinted from Farhadifar, R., Röper, J., Aigouy, B., Eaton, S. and Jülicher, F. (2007). The influence of cell mechanics, cell–cell interactions, and proliferation on epithelial packing, Current Biology 17(24), 2095–2104, Copyright (2007), with permission from Elsevier.

homeostatic stress distribution in tightly packed cellular systems. These approaches, nonetheless, fail to account for cell morphology and are insufficient to describe cell adhesion in these contexts. To remedy this and look more carefully at the evolution of cellular boundaries, vertex models and cellular Potts models (CPMs) were adopted from statistical mechanics. Both types of models discretize the cell, and the cell geometry at each time step is determined through minimization of an energy function calculated by summing up contributions of each discretized section. While CPMs discretize the cellular area in 2D (or volume in 3D), vertex models discretize the perimeter in the form of vertices and line segments connecting them (Fig. 2C). CPMs have proven very efficient for simulations looking into cell sorting of distinct cellular populations; they fail though to recreate mechanical properties of the cellularized material as a whole. Meanwhile, vertex models focus entirely on cellular boundaries and were developed with cell mechanics in the spotlight. Farhadifar and colleagues developed a widely used vertex model to characterize the role of mechanical factors in the evolution of epithelial tissue during wing formation in Drosophila melanogaster (fruit fly). The schematic presented in Fig. 2C shows three cells represented by vertices and the line segments connecting them, representing the cellular boundaries. The epithelial sheet is represented by a 2D network under the assumptions that all cells are equal in height and that the epithelium is highly polarized (i.e., cortical proteins are only located in the apical domain of the cells): force application is restricted to the vertices. The researchers seek to quantify the contributions to epithelial cellular configuration of cell–cell adhesion, actomyosin (cortical) contractility, and cell elasticity. Rather than defining forces due to each factor explicitly, an energy-based approach is used to determine the configuration of the cellular system given vertex position. The energy function expresses the constraints related to the ability of each cell to undergo a deformation due to these factors:

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Biomechanics j Cell Adhesion: Basic Principles and Computational Modeling X

UðRi Þ ¼

Lij lij þ

X Ga a

2

L2a þ

X Ka  a

2

Aa  Aað0Þ

2

(25)

U is the energy function for a node (labeled i, with position Ri) in a cell population. The first summation term describes line tensions (Lij) summed over cell boundaries between cells that contain vertex i. lij represents the length of the boundaries. Whether Lij is positive or negative will determine if surface tension or cell–cell adhesion dominate the behavior of the system, respectively. The second summation term describes the contribution of actomyosin contractility for each of cell containing vertex i in its perimeter (labeled a ¼ 1,.,Nc); the cells contract their cell perimeters (La) according to a coefficient (Ga). The third summation term describes elasticity based on deviation of the cellular area (Aa) for each cell containing vertex i in its perimeter (labeled a ¼ 1,.,Nc), from a preferred area (Aa(0)). Given the energy function, the force acting on each vertex to reach a local minimum is obtained by calculating the derivative with respect to the vertex coordinates: Fi ¼ 

vU vRi

(26)

By substituting Eq. (25) into (26) and computing the derivative, the equation of force can be rewritten for clarity: Fi ¼ 

X

Lij

 vAa vlij X vLa X   Ga  Ka Aa  Aað0Þ vRi vR vRi i a a

(27)

By varying the parameters describing linear tension (Lij) and cortex contractility (Ga) while keeping the target cellular area (Aa(0)) constant, Farhadifar and colleagues find the possible ground-states (lowest energy or most relaxed network configuration) of the system for the different cellular properties. The advantage of vertex models is that with them it is possible to look at large-scale cellular systems while still allowing parameters to have a physiological meaning that can be spatially defined and potentially measured experimentally. This is not usually the case with continuum models. In the context of Drosophila wing development, Farhadifar and colleagues compared the simulated system configurations with experimental observations. They counted the number of neighbors and how this value is distributed in the population at the steady state. They also compared the evolution of the system upon removal of tension at specific cell– cell boundaries through ablation of intercellular boundaries. The researchers were thus able to use the model to estimate the relative magnitude of coefficients that describe line tension and cell contractility in epithelial tissue.

Refining Understanding Through Experiments In the experiments cited by Bell, lymphocytes and fibroblasts were put in the same container as lectin-coated fibers, the containers were shaken to apply shear forces to the cells, and then the number of cells that remained attached was quantified. This principle of exposing cells to mechanical stimulation and quantifying adhesion has been maintained; however, the ways cells are stimulated and visualized and measurements made have been refined. The need to probe deeper comes from the fact that mechanosensitivity of cell adhesions is attributable to the complexity and modular nature of the adhesion complexes. Mechanosensitive elements are found in every functional module of the adhesion complex and in the ECM itself. Fibronectin, for example, stretches three to four times its rest length under cell imposed stress and has been shown to have multiple hidden binding sites. Similar statements can be made for integrins in the membrane and the actin polymerization machinery inside the cell. This fact is reflected in modeling in the trend to move to the subcellular scale, as discussed earlier. Both movements of experimental and modeling approaches to the subcellular scale are not coincidental, as these are complementary and insights gathered on each front spark new inquiries in the other. The refinement of experiments occurs due to advancements in molecular probes, sensors, and imaging technology. These categories sometimes blend as ingenious scientists can modify sensors to act as actuators, or imaging methods can be used to probe the cell or its environment. These advances in the field of cell biology have changed our understanding of cell adhesion.

Imparting Forces at a Cellular Scale Methods to probe a whole cell or its internal or external structures have been applied to study biological system since the late 1980s, a few years after their development. Among the more popular ones are atomic force microscopy (AFM), laser optical tweezers (LOT), magnetic tweezers, and biomembrane force probe (BFP) or micro-needle manipulation. The first two are widely available commercially, making them more popular. In all methods, the probe can be functionalized, making them very versatile and thus used to study a wide array of protein, molecular, and cellular interactions. Through this functionalization, it is possible to study response of a single molecule; for this reason, these methods have been grouped under the umbrella term single molecule force spectroscopy (SMFS). They can impart forces that range from 10 14 to 10 8 N and permit manipulation at a length scale with a range from 10 10 to 10 14 m.

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Fig. 3 Experimental techniques used to impart forces at cell adhesions. (A) Schematic of atomic force microscopy (AFM). (B) Schematic of biomembrane force probe (BFP) or micro-needle manipulation. CHO cells expressing N-cadherin bind to RBC functionalized with mutant N-glycosylated N-cadherin. Abbreviations: CHO, Chinese hamster ovary; ckN-Cad-Fc, Fc-tagged chicken N-cadherin; hN, wild-type human N-cadherin; RBC, red blood cell. (A) Reproduced in part from Dufrêne Y.F. and Pelling A.E. (2013). Force nanoscopy of cell mechanics and cell adhesion. Nanoscale 5, 40944104, with permission of The Royal Society of Chemistry. (B) Reproduced from Langer M., et al. (2012). N-glycosylation alters cadherin-mediated intercellular binding kinetics. Journal of Cell Science 125(10), 2478–2485.

Atomic Force Microscopy (AFM): This technology consists of a small tip positioned at the end of a small cantilever which is then used to probe the surface a sample. The deflection of the cantilever upon contact with the sample is measured from the deflection of a laser beam reflected of it; this contact occurs over an area in the nanometer scale, allowing for single cell probing. Although conceived as an imaging tool, it is now equally used to measure mechanical properties by functionalizing the tip and applying forces at the molecular level (Fig. 3A). This system can measure forces at the scale of piconewtons. Studies targeting cell adhesion molecules through AFM have uncovered the adhesion strength in living cells: Examples include measurements of cell–cell adhesion through the glycoprotein contact site A (csA) in dictyostelium discoideum (slime mold) and measurements of cell–ECM adhesion (i.e., binding between collagen-I and a2b1 integrins) in Chinese hamster ovary (CHO) cells. Laser optical tweezer (LOT): Also known as optical traps, this method consists on trapping a bead in a focused laser beam due to the attraction of the bead to dielectric particles in the vicinity of the focus. Under small movements (approximately 150 nm) from the focus, the force experienced by the bead is linear with the displacement, thus providing a simple relationship that can be used for probing the cellular systems. Optical tweezers can exert forces ranging from 0.1 to 1 nN. Magnetic tweezers are similar to LOT, but a magnetic field is used instead of a laser to trap the bead. It has the advantage over the former method that the bead can be easily rotated. For this reason, it is also known as magnetic twisting cytometry. Because cells are mostly unaffected by the magnetic field (unlike in LOT where damage by bleaching or heating can occur), this option is particularly useful for probing studies inside the cell. Biomembrane force probe (BFP): Also known as micro-needle manipulation. A cell is captured in a droplet using a micropipette. Unlike the methods described earlier, the probe is not a rigid tip or sphere, but rather a cell itself. In many cases a red blood cell or a lipid vesicle, which have no nucleus and whose mechanical properties have been characterized, is functionalized and used as a probe by capturing it with a second micropipette as brought into proximity of the probed cell (Fig. 3B). Using a cell as the probe provides the advantage that functionalization by chimeric proteins is easy, which in turn can be genetically modified upon synthesis to see how specific amino acid residues affect the force response. This method has been used to study the mechanical response of Ecadherin. Researchers have gone as far as using differently N-glycosylated (i.e., asparagine linked saccharide) E-cadherin molecules in these experiments.

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Of the methods presented earlier, AFM is arguably the most popular in the study of cell adhesion. This is the case, because cell adhesion involves forces below the sensitivity of LOT or magnetic tweezers. And although it presents the disadvantage that it cannot be used both outside and inside the cell, studies with LOT or magnetic tweezer that probe inside the cell remain scarce.

Measuring Forces and Imaging at a Cellular Scale In addition to using probes, microfabrication techniques have been used to impart forces on cells. Controlled environments can be created where cells can be exposed to specific forces, even at the site of FAs. For example, microfluidic channels can be made to control the fluid flow and thus shear stress cells are exposed to. This presents a refinement over the original studies quantifying the effect of shear stress on adhesion by shaking the container where cells resided. Also, rather than having cells sit on collagen gels or synthetic gels entirely coated with ECM proteins (e.g., fibronectin, collagen), lithographic methods (chemical or photoactivated) can be used to pattern the surface of a gel (usually polydimethylsiloxane (PDMS)) with protein islands of a desired geometry. By using fluorescently tagged proteins, deformation of the gel by cellular tractions (force per unit area) can be quantified and the spatial source of these tractions identified; this technique is known as traction force microscopy (TFM). TFM has been used extensively due to its reproducibility and low cost. By using markers inside the substrate, whether fluorescent beads or fibrillar components of the ECM, traction components in 3D can be calculated. In a modality of TFM, cells are made to rest on PDMS pillars (or microposts) of a few micrometers in diameter, rather than on a flat surface. The tractions can be easily known by taking into account mechanical and geometrical properties of the pillar. However, the pillars can be used as more than sensors by introducing magnetic nanorods that can be controlled, providing a way to exert force at the sites of cell adhesion by inducing torque in the pillars. This approach was developed by Sniadecki and colleagues and was used to study the response of FAs in 3T3 fibroblasts on functionalized pillars with fibronectin. FA area was monitored as a function of a force applied on the pillars containing the nanorods (Fig. 4A). A similar novel approach uses optomechanical actuators (OMAs) to functionalize the substrate. The actuator consists of RGD-binding sites attached at the end of a thermosensitive polymer with a nanorod core. Upon heating of the nanorod core with near-infrared light, the OMA contracts exerting 13–50 pN forces on FAs. This method presents the advantage that it is up to two orders of magnitude faster than activation of magnetic pillars. These methods work for cells on a 2D configuration (i.e., on a surface). For exploration of cellular adhesion in physiological settings this presents a limitation, as most cells are in a 3D environment in vivo. The trend in the field shows continuing specialization in specific molecules involved in mechanotransduction at the site of the cell adhesions. In contrast to AFM and other SMFS techniques, which do not require imaging as the methods themselves provide a reading of force, TFM does. This requires better imaging methods, the integration of novel microscopy techniques into the study of cell adhesion. Laser scanning confocal microscopy (LSCM) has provided better resolution in 3D imaging by using a pinhole to exclude fluorescent information from regions outside the plane of focus. AFM has been used simultaneously with LSCM to visualize cell indentation simultaneously with measurements of elastic moduli of a cell. Similarly, cytoskeletal rearrangements have been observed as the cell is mechanically stimulated. Imaging techniques with resolution beyond the diffraction limit, dubbed superresolution imaging (e.g., photo-activated localization microscopy (PALM)), have recently started to be implemented in the study of cell adhesion. Though higher in resolution, most of these techniques are relatively slow, creating an image much slower than molecular events occur at the adhesions. A possible solution for this limitation is the use of sensors that produce a signal in response to force. This is the case of Förster resonance energy transfer (FRET) sensors. These are sensors consisting of two fluorophores, a donor and an acceptor, whose proximity determines whether energy transfer between donor and acceptor occurs and which is detected by a shift in emission wavelength. By connecting the donor and acceptor by means of a calibrated molecular spring, FRET can act as a force sensor, measuring the force acting on the molecular spring. A FRET tension sensor has been developed for vinculin, by replacing the flexible linker connecting the head domain and tail domain of vinculin with an elastic domain flanked by fluorophores. This sensor is used to measure tension at cell–cell and cell–ECM adhesions. These sensors have been used to correlate cellular traction with size and duration of FAs (Fig. 4B); this revealed that FAs are responsive to mechanical simulation, undergoing remodeling upon increased strain. Similarly, tension at AJs has also been quantified using VE-cadherin FRET sensors.

Challenges and Conclusions The models presented in this work demonstrate two major trends in the study of cell adhesion mechanics: first, capturing the mechanical properties of the cell–cell or cell–ECM bond itself, and second, accounting for the contribution of additional cellular components beyond the adhesion molecules. Despite these common trends, the models clearly differ in the spatial scale of the system they recreate. All the models, with the exception of the vertex model, start by focusing on the rates of protein–protein association (or dissociation) and attempt to describe how different parameters affect this rate. This provides a clue into what are some of the major challenges in the study of cell adhesion. Regarding the first trend, findings have demonstrated that a mechanical description of the adhesion molecule is not enough to describe the evolution of cellular bonds. This is true, in part, due to the extensive number of molecules adhesion complexes comprise, as well as the occurrence of extensive mechanical and chemical signaling. This is further complicated by mechanotransduction, for which unlike in chemical signaling, there is no single ligand that elicits a response by a specific mechano-receptor.

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Fig. 4 Novel experimental techniques used to measure forces at cell adhesions. (A) 3T3 fibroblast on microfabricated pillars in a magnetic field of magnitude B ¼ 2T. Inset: Some of the pillars have a magnetic nanorod inside. This set up is used to measure and exert forces on cells at site of adhesion (fluorescent labeling). (B) Schematic of action of vinculin FRET sensor in randomly migrating cells. Under high tension FRET efficiency drops. Abbreviations: FN, fibronectin; FRET, Förster resonance energy transfer; mTFP, monomeric teal fluorescent protein; PM, plasma membrane; Vh, vinculin head domain; Vt, vinculin tail domain. (A) Adapted from Sniadecki, N. J. et al. (2007). Magnetic microposts as an approach to apply forces to living cells. PNAS 104(37), 14553-14558, Copyright (2007) National Academy of Sciences. (B) Reprinted from Hernández-Varas P. (2015). A plastic relationship between vinculin-mediated tension and adhesions complex area defines adhesion size and lifetime. Nature Communications 6, 7524.

Continuing research in this direction entails further development of molecular sensors, such as FRET force sensors, and creative experiments that can isolate roles of single molecules in the adhesion complexes. Application of SMFS techniques inside cells to the level in which AFM or TFM is applied still needs to occur. In addition to the difficulty in understanding the molecular mechanisms, it is important to recognize that there is also cell to cell variability. The mechanisms do not work in the same manner in all cells and this should be acknowledged in modeling. This challenge is inherent to biological research. Even if this is accounted for by including stochasticity, cells are programmed differently (i.e., express different genes) in different tissues when they have different purposes. This means models are made specific to the cell modeled, or more relevantly to the biological question probed. The second trend can be seen as a tendency toward integrating models at different spatial scales (multiscale model): how many cellular components are required to faithfully capture cellular behavior? Associated to this trend is the challenge common to modeling, making the model complex enough to capture behavior of a complex system but simple enough that it can be efficiently solved and studied. This tradeoff is most obvious in the models that try to capture large cellular systems or cellularized materials. Tumor growth models were some of the first models of cellularized materials. Their evolution from continuum models to agentbased models demonstrate the need to account for individual cell behavior to capture tissue behavior, and inherently a need to model adhesion. These models have yet to integrate mechanics at the protein level. Yet a tumor growth model by Ramis-Conde and colleagues is the first example of integrating cell mechanics and dynamics of adhesion in a flexible manner that could be applied to many different cellular systems in 2D and 3D.

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Increased computational power and development of numerous numerical methods have allowed for multiscaling. Additionally, the use of modeling repositories and simulation environments exclusively for cell-based modeling (such as CompuCell3D developed at Indiana University, Chaste developed at the University of Oxford, etc.) have also provided a way to share models between researchers in a way that these can be easily modified to overcome the challenge of the variability observed in biological systems. The integration of mechanical description of cellular adhesions in these more widely used modeling platforms is not far down the road.

Further Reading Basic Principles Bell, G. I. (1978). Models for the specific adhesion of cells to cells. Science, 200, 618. Dembo, M., et al. (1988). The reaction-limited kinetics of membrane-to-surface adhesion and detachment. Proceedings of the Royal Society B, 234, 55–83. Evans, E. (2001). Probing the relation between forcedLifetimedAnd chemistry in single molecular bonds. Annual Review of Biophysics and Biomolecular Structure, 30(1), 105–128. Thomas Parsons, J., et al. (2010). Cell adhesion: Integrating cytoskeletal dynamics and cellular tension. Nature Reviews Molecular Cell Biology, 11, 633–643. Computational Modeling DiMilla, P. A., et al. (1991). Mathematical model for the effects of adhesion and mechanics on cell migration speed. Biophysical Journal, 60(1), 15–37. Farhadifar, R., et al. (2007). The influence of cell mechanics, cell-cell interactions, and proliferation on epithelial packing. Current Biology, 17(24), 2095–2104. Kim, M.-C., et al. (2012). Integrating focal adhesion dynamics, cytoskeleton remodeling, and actin motor activity for predicting cell migration on 3D curved surfaces of the extracellular matrix. Integrative Biology, 4, 1386–1397. Paszek, M. J., et al. (2009). Integrin clustering is driven by mechanical resistance from the glycocalyx and the substrate. PLoS Computational Biology, 5(12). Ramis-Conde, I., et al. (2008). Modeling the influence of the E-cadherin-b-catenin pathway in cancer cell invasion: A multiscale approach. Biophysical Journal, 95, 155–165. Experimental Techniques Grashoff, C., et al. (2010). Measuring mechanical tension across vinculin reveals regulation of focal adhesion dynamics. Nature Letters, 466, 263–266. Neuman, K. C., et al. (2008). Single-molecule force spectroscopy: Optical tweezers, magnetic tweezers and atomic force microscopy. Nature Methods, 5(6), 491–505. Sniadecki, N. J., et al. (2007). Magnetic microposts as an approach to apply forces to living cells. PNAS, 104(37), 14553–14558.

Centrifugation and Hypergravity in the Bone Carmelo Mastrandrea and Laurence Vico, Jean Monnet University, Saint-Étienne, France; and Lyon University, Lyon, France © 2019 Elsevier Inc. All rights reserved.

Introduction Disuse, Hypogravity, Microgravity, and Their Effects on the Bone Bed-Rest Studies Centrifugation Principles, Difficulties, and Side Effects Hypergravity, Centrifugation, and Skeletal Changes in Humans Centrifugation and Skeletal Changes in Animals Centrifuge Acclimatization, Animal Size, and Animal Age as Research Considerations Discussion and Conclusions References Further Reading

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Glossary Bone mineral content (BMC) A measurement of the amount of bone mineral within a specified area, usually measured in grams (g). Bone mineral density (BMD) An average measurement of the amount of bone mineral per unit area, usually measured in g/ cm2. Dual emission X-ray absorption (DXA) A method for calculating the BMD of a tissue using X-ray imaging. Hypergravity Conditions in which a body is exposed to an acceleration greater than that of its surrounding environment. At the Earth’s surface this is an acceleration >9.81 m s
2. Hypogravity In this chapter hypogravity will be defined as an environment in which a body is exposed to acceleration lower than that experienced at the Earth’s surface, or 9.81 m s
2, but which is still >0 m s
2. ISS International Space Station. Microgravity An environment in which a body experiences a negligible difference in acceleration relative to its environment, such that the body is perceived to be weightless within this environment. Strictly speaking, this term is wrong since objects are still in a strong gravitational field (the ISS altitude being 350 km, 92% of Earth’s gravitational force still acts on all objects inside), but this term is commonly used and will therefore be applied in this article.

Introduction Gravity, the force of attraction between two particles with mass, is the weakest of the four fundamental forces of the universe. With a theoretically unlimited range, any particle in the universe, at any given time, experiences gravitational attraction to every other particle in existence. However, as any two objects’ gravitational attraction diminishes exponentially as they move away from each other, the majority of these forces produce a negligible, cumulative local effect. On Earth, the culmination of all the individual gravitational forces in the universe produces a sea-level, groundward accelerative force of 9.81 m s
2 (1g), with the Earth producing the majority of this; it is this continuous gravitational acceleration that has arguably provided the most significant evolutionary stimulus for all nonaquatic life. Humans, as erect ambulatory creatures, have evolved complex physiological mechanisms to counteract gravity and enable mobilization in our environment. For example, the cardiovascular system produces a powerful, continuous pressure network through which blood travels to the whole body, supplying oxygen and nutrients to our tissues and organs; the musculoskeletal system produces forces that overcome gravity and allow us to walk, run, and climb. Despite many important adaptations to the Earth’s gravitational field, it is the human skeleton and the associated gravitational effects on the bone that will be discussed further in this article. It may seem obvious that the human skeleton has evolved to provide a rigid framework for muscular attachment in order to facilitate an individual’s movement. However, far from being an inert, passive framework, healthy adult bone is under continual remodeling and repair. Three main cell types control this continuous process: osteoclasts, osteoblasts, and osteocytes. Osteoclast cells, found throughout the skeleton, digest bone and release calcium and phosphate from mineralized mature bone into the blood. This produces areas in which the bone has lost both its extracellular matrix and mineral content. Osteoblasts, cells that deposit new extracellular bone matrix, subsequently replenish these areas with new naive bone that in time mineralizes to maturity. Following the replenishment of an area of the bone, some osteoblasts transform into osteocytes, producing a connected cellular network

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within the mineralized matrix of all bones. The remaining osteoblasts remain at the periosteum as lining cells. Osteocytes can stimulate bone resorption in their immediate vicinity, possibly secondary to local bone damage or due to their own injury or death. Osteocytes also act as mechanosensors. Sufficient mechanoactivation of the three distinct osteocyte mechanoreceptor classes – ion channel, G-protein-coupled, and cytoskeletal-integrin complexes, either through direct cellular strain or due to changes in interstitial fluid flow, leads to a signaling pathway that ultimately results in bone growth (Govey et al., 2015; Robling, 2012; Iwaniec and Turner, 2016). This process of removal and replenishment, known as bone remodeling, is essential to the ongoing health of the skeleton and creates a plasticity that is invaluable. Should a bone need to increase in strength, for example, due to heavy workload, deposition of new bone exceeds removal, and the bone density increases or the bone hypertrophies. The opposite occurs in disuse, where reabsorption of existing bone may exceed new bone deposition, resulting in an overall loss of bone mass. In young animals including humans, new bone growth, known as bone modeling, occurs during childhood and adolescence. Bone modeling produces the normal age-related growth in the size of the skeleton, and it is important to understand that bone deposition during modeling is mainly under genetic control. It is this genetic control and mechanical stimulation in adolescence that drives bone deposition at a significantly higher rate than would be typical of normal adult bone metabolism. Consequently, research that aims to identify skeletal remodeling changes in young animals must consider the possibility of genetic factors driving bone modeling that may otherwise confound results. In adults, bone modeling can occur at the periosteal margin, where load simulation may cause a bone’s diameter to increase, but its length remains static. Unilateral increases in bone mineral content, bone mineral density (BMD), and cortical thickness and an increase in bone diameter were found in the playing arm of tennis players (Haapasalo et al., 1996, 2000). This offers an initial example of higher than normal use and load leading to anabolic bone changes and bone growth in adults. Bone loading or force directly imparted on the bone is now understood to be a major stimulus in its normal physiological maintenance and remodeling. However, it was not until the 1960s that Harold Frost, furthering theories on bone repair published by Julius Wolff in 1892, incorporated strain and force as direct stimuli for bone deposition, maintenance, and loss (Frost, 1987). Frost postulated that the growth and remodeling of any one particular area of the bone was a direct consequence of the forces experienced in that particular area. Frost coined the term “mechanostat” to describe this phenomenon. His theory provides an initial framework by which one can understand the effects of force and therefore gravity on bone homeostasis. As shown in Fig. 1, the magnitude of regular forces exerted on the bone, measured in microstrains, directly corresponds to subsequent bone turnover; disuse at one extreme causes bone loss, and overstress at the other risks fracture. The major sources of these regular forces are routine contact with the ground created during activities such as walking, running, and jumping and direct muscle contraction that imparts forces on the bone. As the force needed to lift an object such as the leg is greater in hypergravity due to increased weight, one can comprehend that higher g-forces result in greater muscle strain on the bone and that the forces of ground impacts when walking or running are also increased. This in turn would cause increased bone deposition and in theory produce a higher bone mass in any one individual as the gravitational force increased. The opposite would be true in a reduced gravitational environment. In the absence of gravity, the resistance to movement is minimal, as weight is abolished and muscles no longer have to fight downward gravitational force. In addition, weightlessness removes the compressive force of an individual’s body weight on weight-bearing bones within the skeleton, especially in the hips, legs, and feet. These two fundamental differences almost completely eliminate the bone’s anabolic stimuli as according to Frost’s mechanostat theory. This understanding of the relationship between force, gravity, and bone mass is of particular interest not only to the study and research of bone diseases such as osteoporosis, but also when considering the detrimental effects of microgravity and hypogravity in astronauts, and the possible uses of hypergravity and centrifugation as a countermeasure to microgravity-induced bone loss.

Fig. 1 Intepretation of Frost’s mechanostat relationship. The potential for bone mass change, represented by the thick curved line, varies for differing magnitudes of forces imparted on bone mechanostat units (Frost, 1987).

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The mechanostat theory therefore provides a theoretical understanding of how the bone will react to different magnitudes of mechanical force. However, one must consider that the forces required to maintain bone mass in weight-bearing areas, such as the hip, may not be the same as those required to maintain bone mass in the non-weight-bearing areas such as the arms or skull. Subsequent research has therefore attempted to provide further classification for the stimuli and thresholds that lead to bone density increases and, although still incomplete, provides a set of intriguing categories for consideration: (1) Different threshold strains must exist in different types of bone; (2) peak magnitude, duration, and frequency of strains all contribute to the mechanostat response; (3) the anabolic responses of the mechanostat diminish with age; (4) systemic interactions, such as circulating hormones, cytokines, and nutrition, all alter responses to a given mechanical stimulus. (Skerry, 2006; Iwaniec and Turner, 2016). Skeletal homeostasis therefore involves a complex set of interactions that results in an overall maintenance of bone mass, or an anabolic or catabolic bone state, with routine muscle and ground impact forces seemingly playing an important role.

Disuse, Hypogravity, Microgravity, and Their Effects on the Bone Spaceflight is often thought to take place in the absence of gravity; however, this is not true; even aboard the International Space Station (ISS), astronauts still experience around 90% of the pull of the Earth’s gravitational field. As the space station and therefore astronauts within it are orbiting the Earth at 27,600 km/h, their linear momentum prevents them from falling to the ground. The ISS is orbiting in a state of equilibrium, with both crew and ship continually moving at the same speed and maintaining a constant altitude. Therefore, astronauts experience microgravity, as they do not experience an accelerative vector in relation to their environment. Should the ISS accelerate, perhaps due to a rocket firing to elevate its altitude, those astronauts inside would experience an accelerative force that may be a small percentage of the 1g experienced on Earth. For the duration of the rocket firing, astronauts would experience hypogravity. Other times in which astronauts may experience hypogravity could be during extravehicular activity on the moon or Mars, where the gravitational field is much weaker than that of Earth. Bone loss in astronauts is common, and pre- and postflight measurements often show significant reductions in BMD. These reductions differ between bones in the body, following a pattern of greater losses in those parts of the skeleton that bear more weight. Fig. 2 shows collated, measured BMD changes in astronauts and in cosmonauts. This loss in bone mass causes a significant increase in calcium liberation from the skeleton, with subsequent increases in calcium excretion in the urine and feces. Measured calcium losses of 200–300 mg/day were typical during the Skylab missions (Rambaut and Johnston, 1979), and large negative calcium balances have been recorded during numerous other missions (Whedon et al., 1976; Leach and Rambaut, 1977; Smith et al., 1999). Not only is this vast volume of excreted calcium a worrying marker of bone loss, but also it increases the risk of the development of urinary calculi. Should an astronaut develop such a complication, it could result in significant pain and risk of serious infection, possibly requiring the termination of the mission. In order to prevent such dramatic losses of bone mass, space agencies have employed microgravity countermeasures that aim to increase the forces imparted on astronauts’ bodies and in turn reduce the loss of bone mass during flight. These typically include exercise regimes involving simulated gravity and ground impact forces using, as an example, bungee cords to allow treadmill running. Additionally resistance exercise is also possible using the advanced resistance exercise device aboard the ISS. Although

Fig. 2 Bone mineral density loss during spaceflight. (A) The percentage change in BMD following 6 month spaceflight at six different body locations. (B) Similar data collected from cosmonauts spending, on average, 188 days in space. As indicated, BMD loss appears to be greater in those areas that typically bear more weight. (A) From Vico L. et al. (2010). Adaptation du squelette humain dans l’espace. 10.106/S0246-0521(10)52493-8. (B) Data from LeBlanc A. et al. (2000). Bone mineral and lean tissue loss after long duration space flight. Musculoskeletal Neuronal Interaction 1(2), 157–160.

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somewhat successful, these methods typically reproduce approximately 25% of an astronaut’s body weight intermittently during exercise, therefore not replicating those forces experienced in the 1g environment on Earth, and bone loss still occurs. Alternatively, continual compressive force through the use of a weight-loading g-suit may act to reduce the detrimental effects of unloading of the spine, although the definitive effects of such garments are still being investigated. Disuse on Earth can cause similar, although less dramatic, reductions in bone mass, with typical examples including immobilized patients, bed-rest studies, and water immersion. Human research in space and on the ground can therefore provide invaluable information as to the processes that cause bone changes in differing gravitational environments. New countermeasure research may achieve increases in the anabolic stimuli for bone maintenance and repair in space, with one such possibility being the use of centrifugation, either continual or intermittent, to provide an artificial gravity.

Bed-Rest Studies Bed-rest studies aim to reproduce, as closely as possible, the effects of microgravity on the human body. They are performed in order to provide a significantly cheaper and more accessible method for microgravity research, which would otherwise only be possible during spaceflight. Six degrees of head-down tilt is now most commonly used as the model of microgravity in bed-rest studies, and this figure was chosen following initial investigations concerning resultant physiological parameters and the subjects’ tolerance of different degrees of head-down tilt. However, the physiological parameters recorded were cardiovascular, not skeletal, and the exact angle and duration of head-down tilt should depend on the physiological system under investigation (Prisk et al., 2002; HinghoferSzalkay et al., 2004). Typically, subjects in these studies are continually confined to their beds, unable to stand or sit in order to eat and wash. Such research is therefore very difficult and time-consuming as a level of care akin to 24 h nursing is required for each participant.

Centrifugation Principles, Difficulties, and Side Effects Centrifugation as a spaceflight countermeasure has been historically proposed multiple times, with different sizes and speeds of centrifugation suggested to achieve some degree of Earth-equivalent gravitational acceleration. However, despite the likely abolishment of the majority of microgravity-induced deconditioning that would occur should this countermeasure be deployed, no permanent centrifuge has ever been launched into or constructed in space. In the history of human spaceflight, centrifuges have only been launched twice, aboard two Skylab missions, in which short-arm centrifuges were used for human neurovestibular study rather than as a microgravity countermeasure. In order to produce artificial gravitational acceleration, centrifuges spin a subject at high speed around a circular enclosure. The size and rotational speed of a centrifuge are dependent on the desired artificial gravitational acceleration and in space can be calculated from the following equation: Acceleration ðGravitational equivalentÞ ¼ Velocity2 =Radius To give some examples, in order to achieve a gravitational acceleration equivalent to that at the Earth’s surface of 9.81 m s
2, a theoretical space-based centrifuge would have the following approximate attributes: A radius of 5 m and rotational speed of 13 rpm A radius of 10 m and rotational speed of 9 rpm A radius of 15 m and rotational speed of 8 rpm On Earth or any other planet or the moon, one must also factor in the continuous surrounding gravitational field. As an additional consideration, one must remain conscious of the Coriolis forces experienced by the subject. The Coriolis force defines the abnormal motion of moving objects within a centrifuge from the viewpoint of an individual within that same centrifuge. As the rotational speed of a centrifuge increases, it results in increasing disorientation, motion sickness, and difficulty performing tasks due to constant rotation. This limits rotational speeds for long-duration centrifugation to approximately 3 rpm (Musgrave and Larsen, 2009). Therefore, in order to achieve a rotational speed no greater than 3 rpm and a continual gravitational acceleration of 1g, the minimum required radius of a space-based centrifuge would have to be approximately 75 m or 150 m in diameter. Given that the current modules that form the ISS are approximately 4.5 m in diameter and that the ISS is 108.5 m along its longest axis, it becomes apparent that such a centrifuge would be impractical. It may be possible to rotate the spacecraft itself, much like the space hotel in 2001: A Space Odyssey, thereby incorporating the centrifuge into the structure of the station. However, such a station would come with immense cost and engineering difficulties. With the aforementioned problems generating continual 1g acceleration, future proposed centrifuge countermeasures compromise on certain parameters. Either the duration of centrifugation falls, thereby allowing high rotational speeds for short periods of time, or the exposed gravitational acceleration falls, allowing for slower rotational speeds. Lower target gravitational accelerations and intermittent rather than continual centrifugation also allow for smaller centrifuge sizes, an important consideration for space countermeasure designs. Despite the theoretical benefits of hypogravity over microgravity (25%–50% of 1g rather than 0g), further

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research must be performed in space to validate hypogravity as a countermeasure and to ensure adequate bone maintenance in astronauts. Terrestrial human and animal centrifugation takes place relatively regularly, either in the routine training of pilots who undergo hypergravity stress training or in scientific research such as countermeasure trails in microgravity analogue bed-rest studies or hypergravity animal studies. In human scientific research, centrifugation devices typically have a radius no larger than a few meters, and centrifugation is intermittent; subjects do not perform complex tasks, and head movement can cause significant motion sickness due to the Coriolis effect. It should be noted that the Russian Samara State Medical University (SSMU) has developed a humanrated centrifuge for clinical use (SSMU; see Fig. 3). They state that it provides benefits for multiple physiological conditions, advocating its use in bone fracture repair and secondary posttraumatic osteoporosis. We are not aware of any other hypergravity therapy options currently available in other countries. Animal research typically involves longer durations of centrifugation over many days to weeks, with continuous hypergravity throughout the trial. Examples of human and animal centrifuges are shown in Fig. 3. Larger centrifuges have been proposed in order to test the effects of continual hypergravity on human subjects, but none have ever been constructed. Van Loon (2008) proposed the design and construction of a large-diameter centrifuge for continual hypergravity experimentation. His design utilizes high-speed trains on a 942 m circular track, with carriages or modules to support long-duration

Fig. 3

Examples of human and animal rated centrifuges.

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Fig. 4 Mean food consumption in centrifuged rats. Food consumption as a mean of six rats kept in continual 4.7g centrifugation. The marked reduction in intake during the initial 2–3 weeks may lead to catabolic metabolism in these animals. Food consumption increased beyond that of noncentrifuged rats if the hypergravity exposure continued for a sufficient period of time. From Oyama et al. (1965). Effects of prolonged centrifugation on growth and organ development of rats. American Journal of Physiology 209(3), 611–615.

living. With rotational speeds under 3 rpm, human and animal subjects would experience long-duration, constant gravitational accelerations greater than 1g, and it offers a potentially exciting opportunity for both space and terrestrial medical research. Animal centrifugation, both in space and on the Earth’s surface, has been performed multiple times in the past as a method to study the effects of hypergravity. Given the smaller size of animals such as mice, smaller apparatus is required, and so cost and engineering difficulty are reduced. Studies have been performed with gravitational accelerations in excess of 3g; however, one must be careful when interpreting results from these experiments. The stress of animals in high-velocity centrifuge environments is known to be elevated, with corticosterone levels rising during acclimatization to 2g and for the duration of a 21-day 3g centrifugation (Guéguinou et al., 2012). Elevations in corticosterone also occur at 4g, as well as activation of the sympathetic nervous system (Petrak et al., 2008). Additional evidence for animal stress in these environments includes the animal’s weight post centrifugation. Wunder et al. (1966) investigated if the weight loss in centrifuged animals was due to nausea or motion sickness, a direct consequence of the high rotational speeds. They found that in labyrinthectomized hamsters, in which motion could not be sensed by the inner ear, a hypergravity-induced reduction in weight persisted, proving that reduced appetite as a direct consequence of nausea is not the cause for weight loss in centrifugation studies. Wade et al. (2002) evaluated the metabolic rate of rats in hypergravity, providing evidence that resting metabolism in these animals was elevated, a further mechanism that would lead to weight loss. Whatever the reason, food consumption falls, and this plays a significant role in the bone formation ability of centrifuged animals (see Fig. 4). Even at 1g centrifugation, rats spun in space at 53.3 rpm for 18.5 days showed significant suppression of weight gain compared with control 1g animals kept in an identical, noncentrifuged environment (Spengler et al., 1983). As an elevation in stress, circulating corticosterone and sympathetic activity and a reduction in food intake can all contribute to a reduction in body weight and subsequent bone deposition; interpretation of results in experimental centrifugation research must be considered carefully. In summary, despite a clear understanding of the principles and design requirements for continual human hypergravity centrifugation either in space or on Earth, engineering and financial constraints prevent their use. Therefore, intermittent terrestrial centrifugation and continual animal centrifugation provide the majority of our understanding of the physiological changes to the skeletal system in hypergravity.

Hypergravity, Centrifugation, and Skeletal Changes in Humans Unfortunately, human centrifuge studies investigating skeletal changes in hypergravity are typically limited. The nature of such experimentation and the need to provide exposure for long durations in order to achieve identifiable and reproducible results prohibit their routine undertaking. In addition, it has historically been impossible to visualize microarchitectural bone changes, as one could not dissect and section the femur of a living human subject. Recently, however, the development of highresolution peripheral quantitative computed tomography (HR pQCT) now allows for these investigations, although at this time these experiments are yet to be performed. Kos et al. (2013) and Rittweger et al. (2015) performed analyses of 1g centrifugation at the center of mass, 2.5G at the feet, as a countermeasure to bed-rest-induced bone loss. Comparing short-duration, bed-rest-induced elevations in markers of bone resorption, they concluded that 30 min centrifugation per day, either as a single 30 min exposure or as six 5 min exposures, resulted in a minimal reduction in markers of bone loss. This trial was short, lasting only 5 days for each condition. Smith et al. (2009) examined the effects of 1 h of centrifugation (1 g center of mass and 2.5 g at the feet) per day for 21 days, again in bed-rest subjects. Measuring serum and urine bone resorption markers and performing dual-energy X-ray absorption (DXA) scans, they did not find any differences in either markers of bone resorption, BMD, or bone mineral content (BMC) between control and experimental groups. Finally, Iwase et al. (2004) identified a suppression of the bone resorption marker deoxypyridinoline in subjects undertaking 14 or 20 days of bed rest, with 30 min per day 1.4g centrifugation with concurrent 60 W exercise. The full results of this

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experiment were not, however, published, and the exercise intervention may have contributed to the protective effect seen in experimental subjects. The possible reason for a lack of statistical difference in these studies is due to the fact that bone mass changes take from many weeks to months; therefore, identifiable changes are likely to be minimal after only a few days’ intervention. In addition, it would be ethically unacceptable to allow experimental subjects in bed-rest groups to develop significant bone loss that may lead to complication in later life. Also, due to the huge cost and technical difficulties of caring for subjects who are confined to bed 24 h a day, sample sizes are low, and this may impede the ability of studies to find statistically significant differences. Therefore, current human trials are of limited value in this respect, although they do offer an interesting platform for the investigation of more rapid changes such as cardiovascular fluid shifts and muscular degeneration that occur in microgravity. One very interesting research paper focuses on the hypergravity-induced bone changes in fast-jet pilots. Naumann et al. (2001) followed fighter pilot trainees over the course of a 1-year flight training course, in which they regularly experienced positive gravitational accelerations between 2g and 6g through their spines and pelvis. It was found that compared to age- and height-matched control subjects, and following correction for changes in total body weight and fat mass, the trainee pilots showed an increased BMD and BMC of 11.0% and 4.9% in the thoracic spine and pelvis, respectively. So far as we are aware, this is the only method by which a comparison of bone parameters between 1g and accelerative forces greater than 1g has been performed in humans. These results strongly suggest that regular hypergravity exposure has an anabolic effect on the bone; however, this exposure method is not suitable for more standard research.

Centrifugation and Skeletal Changes in Animals The centrifugation of many different animal species has been performed in the past, including chickens, dogs, rats, and mice. Both on Earth and in space, animal centrifugation offers a method for the study of long-duration, continual hypergravity exposure on bone parameters. Another advantage in using animals is that the effects of hypergravity during early life and adolescence can be studied in order to identify any significant differences in primary bone growth. However, the number of studies that have investigated these changes remains limited. In 1977, Riggins et al. assessed chickens’ tibial strength following a gradual increase in gravitational acceleration to 3g over the course of 18 weeks. They found that although the cortical bone thickness increased in the centrifuged animals, midshaft diameter was significantly reduced, and bone density was unchanged. Despite little change in fracture patterns on stress testing, it appeared that hypergravity induced some degree of bone growth and remodeling. Additional findings in hypergravity were also identified in another study, where rats exposed to 2.8g centrifugation for 810 days showed an increase in femur density when compared with similar control animals (Jaekel et al., 1977). Finally, Doden et al. (1978) exposed young beagle dogs to 3 months of continual centrifugation at 2g, finding that the density of thinner bones was higher and thicker bones lower, in the centrifuged animals when compared with controls. More recent research has begun to identify the microscopic changes that occur in hypergravity. In 2015, Gnyubkin et al. published their findings regarding the changes in bone deposition and resorption surfaces in the femur and vertebrae of C57BL/6 mice. Following 21 days of continual centrifugation at 2g, it was found that trabecular fraction and density increased in the femur by 18% and 32% and in the vertebrae by 13% and 9%, respectively. Within the bone, osteoclast surfaces were reduced and mineralized surfaces increased, suggesting an overall reduction in bone resorption and an increase in bone formation. These differences were accompanied by an increase in the number and volume of blood vessels in the distal femur metaphysis. As an additional comparison, the same protocols and investigations were undertaken on another set of mice with one difference; in the second group of animals, the gravitational acceleration was increased to 3g. In these animals, it was found that the bones showed cortical thinning, osteoclast surface area increase, and reduction in the rate of bone formation. The author states that the changes in 3g may possibly be due to increased stress rather than the true effects of hypergravity on the bone. Nonetheless, it was concluded that continual 2g acceleration does cause an increase in bone deposition, with a possible detrimental effect of continual gravitational accelerations around and above 3g in magnitude due to significant stress responses. In another study, Ikawa et al. (2010) identified a reduction in the amount of bone resorption in ovariectomized rats over 28 days at 2.9g. Normally, estrogen, produced by the ovaries, inhibits bone resorption and protects the female animal from osteoporosis; following ovariectomy, levels of estrogen decline, and bone density falls. By comparing ovariectomized rats kept at 2.9g with those at 1g, it was found that hypergravity prevented much of this bone loss, and the conclusion that 2.9g hypergravity inhibits bone resorption was therefore made. These findings correlate with the decreased bone resorption surfaces in 2g mice found by Gnyubkin et al.; however, they directly conflict with the findings of increased bone resorption surfaces in 3g mice. Finally, Kawao et al. (2016) investigated the effects of the vestibular system on bone metabolism in hypergravity. They hypothesized that skeletal responses to gravitational accelerations are influenced by the vestibular apparatus and that subsequent sympathetic signaling affects gravitational bone and muscle metabolism. Bilateral vestibular lesions were created in young mice, leading to reductions in BMD and BMC. Beta-blockade in nonlesioned mice seemed to inhibit the increase in body-weight-corrected BMC in hypergravity but did not achieve statistical significance. However, the authors state that following vestibular surgery, mice did not eat properly for 2 weeks and immediately after this they were introduced into a 3g environment, which is known to cause a reduction in food intake for another 2–3 weeks (see Fig. 4). Therefore, the mice may not have eaten adequately for 4–5 weeks preceding sacrifice at 12 weeks of age, causing significant weight loss in lesioned mice, a factor that can independently affect bone deposition. It should be noted that Wunder et al. (1966) induced vestibular lesions

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Biomechanics j Centrifugation and Hypergravity in the Bone Table 1

Animal and bone parameters following 18.5 days of spaceflight

Animal group

Animal number

Animal weight (g)

Femur failure (torque, N cm)

Femur density (g/cm3)

Ground control Ground simulated flight 0g flight 1g centrifugation flight

5 5 5 5

289  9 294  16 295  29 264  8

32.2  9.6* 26.1  6.9 17.5  4.4 29.2  1.5*

1.60  0.02 1.60  0.01 1.56  0.05 1.61  0.01

* Significantly different from the corresponding value for 0g flight group. Adapted from Spengler, D., Morey, E., Carter, D., Turner, R., Baylink, D. (1983). Effects of spaceflight on structural and material strength of growing bone. Proceedings of the Society for Experimental Biology and Medicine 174, 224–228.

in hamsters with no significant differences in weight at 1g and minimal differences in weight at 5g, findings that contradict Kawao et al. Finally, by comparing bone parameters when corrected for body weight, values may be confounded by the significant reduction in body weight in hypergravity over such a short period of time. The actual skeletal density changes in differing gravitational environments were relatively static, but due to a large loss of animal weight, the correction for body weight created a significant difference. These three studies provide evidence for alterations to the skeleton in hypergravity, suggesting potential causes for these changes. Animal age, size, duration of centrifugation, and nutrition were not consistent, and as discussed later, this may adversely affect skeletal changes during centrifugation. Centrifugation studies of animals in space have been previously achieved, aboard the Biocosmos 4 (Cosmos 936) satellite. Two groups of male Wistar rats spent 18.5 days aboard the satellite before its return to Earth, with two additional control groups spending 18.5 days in similar conditions on the ground. The pertinent results are shown in Table 1(Spengler et al., 1983). Those rats in the simulated flight ground group experienced vibration and noise equivalent to those that flew aboard the Biocosmos 4. From the data, one can see that induced femur failure occurred at much lower torque stresses in those animals having spent 18 days in a 0g environment, and this is to be expected given the previous evidence presented in this article. Those rats that were centrifuged in space (1g centrifugation flight group) had femur failure rates equivalent to those rats kept in similar environments on the ground. Despite the relatively short duration of the flight, the centrifuged animals were exposed to continual Earthequivalent gravitational acceleration, and this prevented the significant loss of femur strength and density seen in the 0g cohort. However, mirroring the findings of ground-based centrifugation studies, there is a noticeable difference in the weight of these animals, and this is likely to be secondary to the fast revolution speed of the centrifuge (53.3  3 rpm). Unfortunately, no video telemetry of the rats exists, but it may be that they did not eat and as would have been expected. This would be a consequence of both stress (as previously discussed) and due to the fast revolution speed and significant Coriolis forces causing disorientation and nausea. Nonsacrificed rats did regain weight following landing (Spengler et al., 1983). As previously alluded to, stress due to high centrifugation speeds must always be considered in experiments of this nature, with careful interpretation of subsequent results. In addition, the acclimatization to centrifugation may play an extremely important role in the inconsistent results seen in relatively short-duration, continual centrifugation research.

Centrifuge Acclimatization, Animal Size, and Animal Age as Research Considerations Smith (1975) described in great detail the effect of alterations in gravitational acceleration on different organs in a variety of species. He concluded that the tolerance of animals to both acute and chronic hypergravity was inversely proportional to the animal’s size and weight. In addition, an organism’s skeleton must contribute to a larger percentage of overall body mass as the animal’s size increases. This correlates with the animal findings in hypergravity, where the skeletal BMD increases as a proportion of total body weight. Summarized by Clément and Slenzka (2006), in addition to the skeletal changes and the intrinsic ability of smaller animals to resist hypergravity, posture and hydrostatic pressure greatly contribute to g-tolerance (see Fig. 5). Therefore, one would expect smaller animals to better tolerate higher gravitational accelerations, with the subsequent production of greater anabolic skeletal changes. However, using two studies discussed previously in this article (Gnyubkin et al., 2015; Ikawa et al., 2010), it appears that mice had more detrimental skeletal changes, showing increased femur resorption areas, when compared with the larger rats, which showed increased femur density, which is inconsistent with this theory. These discrepancies may be due to centrifuge acclimatization, a direct consequence of the duration of centrifugation. Both centrifuged animal populations experienced weight loss compared with their control counterparts (12% and 20%, respectively), suggesting that the hypergravity cohort did not eat as much as the control cohort. This is a common finding in all centrifuge studies. What is different however is the total duration of centrifugation. Gnyubkin et al. exposed mice to 21 days of centrifugation, whereas Ikawa et al. provided 28 days. Given the initial stress of centrifugation, 21 days may have not been enough time for the mice to adapt and subsequently adequately respond to the hypergravity stimulus on the skeleton. Conversely, 28 days may have allowed the rats to resume normal nourishment and for stress level to lower to a point at which the skeleton could anabolically respond to the hypergravity stimulation. It would be expected, given the theories forwarded by Wolff, Frost, and Skerry, that the expected response of the skeleton would be to increase in volume and density given the higher mechanical strain produced by this level of centrifugation. In

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Fig. 5 The skeleton, hypergravity, and species size. (A) Relationship between nonaquatic organisms’ weights and their relative proportional skeletal masses. (B) Peak survivable g-forces in relation to an animal’s size. Animal size is also inversely related to chronic hypergravity tolerance. From Clément and Slenzka (2006). Gravitational biology. In: Fundamentals of space biology. Chapter 2. ISBN: 0-387-33113-1.

Table 2

Femur mass following long-duration hypergravity

Weight (g) Femur mass (mg/100 g) Femur mass (g)

Age matched

2.5g

3.5g

4.7g

261  11 376  11 0.98

216  9 409  21 0.88

211  10 383  19 0.81

205  14 393  85 0.81

All animals’ final weights are lower than their age-matched controls, due to the stresses inherent in centrifugation. Femur mass, expressed as a proportion of overall animal mass, is increased in all animals; however, the absolute mass added on the bottom row (g) is decreased, accounted for by the smaller animal size. Adapted from Oyama, J., Platt, W. T. (1965). Effects of prolonged centrifugation on growth and organ development of rats. American Journal of Physiology 209(3), 611–615.

longer-duration studies, hypergravity, even to levels of 4.7g, produces an anabolic effect on weight-bearing bones (Oyama and Platt, 1965); after 4.5 months of continual centrifugation, rats exposed to hypergravity developed increased femur masses, and when expressed as a proportion of body weight, these findings show the relative anabolic effect of high gravitational acceleration (see Table 2). These mice also experienced an initial weight loss due to decreased calorific intake (Fig. 4) (Oyama and Platt, 1965); however, following acclimatization, the relative femur mass increased proportionally to the animal’s overall weight. As an additional consideration, total activity and gait may be altered in hypergravity. In the developing mouse, continual hypergravity is known to cause alterations to gait parameters including stride frequency and length, with walking causing extension of the hind limbs beyond that normally shown in mice at 1g (Bojados et al., 2013). This may lead to differences in imparted strain on joints and bones, in turn altering the skeleton’s response to load. Hypergravity may even cause a certain degree of immobility in animals, with high accelerations leading to a reduction in walking and running distance. Therefore, further investigations into voluntary movement, including running distances, must be made during hypergravity exposure in animals. This will ensure that the skeleton is exposed to a similar frequency of ground impact forces and muscular contractions in both control and hypergravity animals. Consideration of animal age when investigating the consequences of the gravitational environment on skeletal parameters must be addressed. Young animals and primary bone growth may be impacted negatively due to the lack of normal mechanical stimulation during skeletal development. This is an important consideration for future physiological gravitational research. Martinez et al. (1998) performed 14 days of continual centrifugation on rats at 2g. No significant differences in bone formation or cortical bone composition were identified; however, it was noted that in these relatively young animals, hypergravity resulted in the stunting of femur growth by approximately 3% and possible evidence for an earlier onset of bone maturation. Vico et al. (1999) compared the effects of 4 days of continual 2g centrifugation with 1g controls in young Sprague Dawley rats. Within the primary spongiosa of the tibia and femur, trabecular thickness and density were significantly elevated with a reduction in the overall bone length. These differences occurred despite a noted weight loss in the 2g animals. Therefore, the changes identified in these two papers indicate that hypergravity may reduce the mature length of weight-bearing bones in juvenile animals but still indicate that overall bone density is increased in a similar fashion to the mature skeleton.

Discussion and Conclusions Currently, very little data exist with respect to the effects of hypergravity on the bone. In addition, the human research that has been performed in this field typically involves a comparison between unloaded or microgravity analogue models and proposed, intermittent countermeasures. Therefore, the majority of bone-related continual hypergravity experimentation has been performed on

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animals, and even these are few in number. Despite this limitation, it can be concluded that in comparatively higher than normal gravitational environments, the metabolic state of the bone is stable or anabolic relative to the mass of the organism. Given the historical findings that show significant loss of appetite, growth retardation, hormonal alterations, and autonomic changes in animal centrifuge studies, careful consideration must be employed in the interpretation of the results of these studies. True skeletal alterations to hypergravity in research studies may only occur following the acclimatization of animals to centrifugation stresses, so such experiments may require exposures of many months rather than a few weeks. In humans, centrifugation remains as an untested microgravity countermeasure. It is the potential reduction in bone resorption in hypergravity that is of greatest interest for space physiologists, as the loss of bone in microgravity causes considerable risk both during and following flight, with current countermeasures that include resistance exercise and tethered treadmill running unable to provide adequate long-duration skeletal protection. Therefore, the use of artificial gravity in space may provide a method to reduce skeletal resorption with relatively physiologically normal bone stresses and without pharmacological intervention. However, one must remain aware that with high continual gravitational acceleration, stresses on both the bone and organism as a whole may be detrimental. Significant adverse physiological changes in humans who are exposed to continual or intermittent 1g centrifugation in space may not occur, but given the findings in animal trials during continual gravitational acceleration in excess of 1g, future ground-based human research is required. The increased bone densities in fighter pilots exposed to hypergravity, as previously discussed, suggests intermittent hyper gravity exposure, with adequate recovery for normal nourishment and a reduction in stress responses, may provide the optimal circumstances for anabolic bone metabolism. There is therefore a significant need, in both space and terrestrial medical research, for the future investigation of centrifugation and the hypergravity consequences for the skeleton. Both as a countermeasure to microgravity-induced bone loss and as a possible preventative measure for osteoporosis, hypergravity research offers exciting potential avenues for future therapeutic therapies. However, given the significant costs involved in the construction of continual human hypergravity centrifuges, either in space or on the ground, there is currently a significant fiscal barrier to this type of experimentation. Additionally, the duration of exposure required to see significant beneficial changes in the bone density of humans during hypergravity centrifugation remains unknown. Therefore, this field is open to both new physiological discoveries and development of possible future centrifugation treatments and space countermeasures.

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Rambaut, P. C., & Johnston, R. S. (1979). Prolonged weightlessness and calcium loss in man. Acta Astronautica, 6(9), 1113–1122. Rittweger, J., Bareille, M. P., Clément, G., Linnarsson, D., Paloski, W., Wuyts, F., Zange, J., & Angerer, O. (2015). Short-arm centrifugation as a partially effective musculoskeletal countermeasure during 5-day head-down tilt bed rest-results from the BRAG1 study. Journal of Applied Physiology, 115(6), 1233–1244. Robling, A. G. (2012). The interaction of biological factors with mechanical signals in bone adaptation: Recent developments. Current Osteoporosis Reports, 10(2), 126–131. Skerry, T. M. (2006). One mechanostat or many? Modifications of the site-specific response of bone to mechanical loading by nature and nurture. Musculoskeletal and Neuronal Interaction, 6(2), 122–127. Smith, A. H. (1975). Principles of gravitational biology. In M. Calvin, & O. Gazenko (Eds.), Foundations of space biology and medicine (Vol II Book 1, pp. 129–162). Washington DC: NASA. Smith, S. M., Wastney, M. E., & Morukov, B. V. (1999). Calcium metabolism before, during, and after a 3-month space flight: Kinetic and biochemical changes. American Journal of Physiology, 277, R1. Smith, S. M., Zwart, S. R., Heer, M. A., Baecker, N., Evans, H. J., Feiveson, A. H., Shackelford, L. C., & LeBlanc, A. D. (2009). Effects of artificial gravity during bed rest on bone metabolism in humans. Journal of Applied Physiology, 107(1), 47–53. Spengler, D., Morey, E., Carter, D., Turner, R., & Baylink, D. (1983). Effects of spaceflight on structural and material strength of growing bone. Proceedings of the Society for Experimental Biology and Medicine, 174, 224–228. Van Loon, J. (2008). The human centrifuge. Microgravity Science Technology, 21, 203–207. Vico, L., Barou, O., Laroche, N., Alexandre, C., & Lafage-Proust, M. H. (1999). Effects of centrifuging at 2G on rat long bone metaphyses. European Journal of Applied Physiology and Occupational Physiology, 80(4), 360–366. Wade, C., Moran, M., & Oyaya, J. (2002). Resting energy expenditure of rats acclimated to hypergravity. Aviation Space and Environmental Medicine, 73(9), 859–864. Whedon, G. D., Lutwak, L., Rambaut, P. C., Whittle, M. W., Reid, J., Smith, M. C., Leach, C., Stadler, C. R., & Sanford, D. D. (1976). Mineral and nitrogen balance study observations: The second manned Skylab mission. Aviation, Space and Environmental Medicine, 47(4), 391–396. Wunder, C., Milojevic, B., & Eberly, L. (1966). Growth and Food consumption of labyrinthectomized hamsters during chronic centrifugation at 5G and 6G. Nature, 210(5032), 177–179.

Further Reading LeBlanc, A., Schneider, V., Shackelford, L., West, S., Oganov, V., Bakulin, A., & Voronin, L. (2000). Bone mineral and lean tissue loss after long duration space flight. Musculoskeletal Neuronal Interaction, 1(2), 157–160. Riggins, R. S., & Chacko, K. A. (1977). The effect of increased gravitational stress on bone. Life Science and Space Research, 15, 263–265. Samara Sate Medical University Website. (November 2016). Synergy project. Accessed on, 03. http://www.smuit.ru/projects/medicinskoe-priborostroenie/lechebnye-ehlektronnyepribory-medicinskogo-naznacheniya/gravitacionnyj-stend-sinergiya/. Vico L, Pavy-Le Traon A (2010) Adaptation du squelette humain dans l’espace. https://doi.org/10.1016/S0246-0521(10)52493-8 Wolff, J. (1982). Das Gesetz der Transformation des Knochen. Berlin: Hirschwald.

Computational Modeling of Respiratory Biomechanics Christian J Roth, Lena Yoshihara, and Wolfgang A Wall, Technical University of Munich, Munich, Germany © 2019 Elsevier Inc. All rights reserved.

The Medical Reason Structure of the Respiratory System Reduced Models of the Conducting Zone Continuum Models of the Conducting Zone Reduced Models of the Respiratory Zone Continuum Models of the Respiratory Zone Reduced-Dimensional Coupled Lung Models Single-Compartment Models Multi-Compartment Models Hybrid Representations of the Respiratory System Hybrid Model (Conducting) Hybrid Model (Respiratory) Continuum Coupled Lung Models Concluding Remarks Further Reading Relevant Websites

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Glossary Acinus Composition of several air spaces (so-called alveolar ducts) distal to a terminal bronchiole. The acinus is the largest unit of lung tissue in which all parts participate in gas exchange. Compliance Measure for lung stiffness defined as quotient between volume change DV and associated pressure change DP. Interdependence Interaction between neighboring air spaces competing for lung volume. Parenchyma Lung tissue on a macroscopic point of view including air spaces, alveolar walls, small airways, and blood vessels. Recruitment/Derecruitment Reopening (recruitment) and closure (derecruitment) of airways and/or lung tissue during the respiratory process. Resistance Measure for viscous and convective flow resistance in the conducting zone of the lung. The resistance is defined as the quotient between pressure drop DP and associated flow DQ through an airway. Surfactant Liquid lining of surface active agents on the inside of the alveoli. Surfactant reduces the surface tension on the gas exchange interface and consequently facilitates breathing.

The Medical Reason Respiratory biomechanics covers a variety of different complex and interacting phenomena including tissue mechanical, fluid dynamical, gas and particle transport processes. Their underlying physics has been directly linked to observed immunological reactions and to the current health state of a patient, leading to the fact that breathing is actually a mechanical problem. However, single phenomena in respiration are difficult to measure in vivo due to both ethical and technical reasons. Therefore, advanced modeling techniques have been developed to adequately represent important effects and to provide new insights into respiratory biomechanics in general as well as to promote suitable medical treatment in case of acute and chronic respiratory disease. Especially approaches with the ability to represent patient-specific and regional information show the potential to deliver valuable information on patient-tailored treatment in respiratory care. The current goal of developing predictive models based on the underlying physics of the lung is driven by the awareness that it is more powerful to understand underlying processes in respiratory biomechanics than to only see their effects in medical imaging and clinical monitoring of a patient.

Structure of the Respiratory System To allow for a better understanding of the governing physics in respiratory biomechanics, a deep knowledge about the anatomy and physiology of the lung is indispensable. From the anatomical point of view (see Fig. 1), the lung as an organ is subdivided into the

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Fig. 1 Human anatomy showing the thoracic cage with all respiratory components highlighted. The conducting zone is visualized in blue and comprises the first 16 generations of the airway tree. The respiratory zone is marked in pink and denoted as parenchyma on the macroscopic scale. On the microscopic scale, lung tissue is composed of single alveoli arranged in a sponge-like structure. The alveolar walls contain the capillary network and constitute the blood-gas barrier.

conducting zone (airway tree generations 1-16) and the respiratory zone (airway tree generations 17-23). The conducting zone is characterized by a dichotomously branching tree of large to medium-size airway segments and its main task is to efficiently distribute air towards the smaller structures of the lung. The respiratory zone starts with the terminal bronchioles leading to the alveolar ducts which consist of single clustered air spaces, the so-called alveoli. The respiratory zone is ultimately responsible for gas exchange with the blood. To allow for a maximum of exchange surface area, alveoli are embedded within a tissue scaffold following the concept of a dense spherical packing (see Fig. 1). For many applications, it is not possible to resolve the complex microstructure containing alveoli, alveolar walls and blood vessels in full detail. Therefore, the respiratory zone is often considered from a macroscopic point of view. This macro-scale is referred to as lung parenchyma (see Fig. 1). From the biomechanical point of view, the respiratory system can be seen as a complex multi-physics and multi-scale problem in nature. Its full physical description is governed by airflow dynamics in the conducting airways, tissue mechanics of lung parenchyma covered by a film of surfactant, gas transport and mixing processes as well as particle and aerosol transport on scales ranging from a few centimeters to a few nanometres. Further, the single components visualized in Fig. 1 highly interact with each other. The deformation of lung tissue introduced by contraction of the respiratory muscles and the diaphragm creates a negative pressure in the pleural cavity and respiratory air spaces and thus induces flow in the airways. Airflow vice versa can lead to the movement of accumulated liquid in occluded airways and triggers the regional inflation of lung tissue in healthy and diseased regions of the entire organ.

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Fig. 2 Classification of respiratory biomechanics models. Approaches for the respiratory zone and for the conducting zone can be found in the red and blue parts of the table, respectively, whereas coupled models combining both zones are given in the white part of the table. Reduced models are highlighted in gray, continuum models in yellow and hybrid (i.e., combined continuum and reduced) models in green.

From this complexity it becomes clear that a simultaneous and full investigation of all aspects in respiratory biomechanics is virtually impossible and in many cases not even reasonable. Different clinical questions are related to effects on different scales that are governed by different physical fields. Only relevant phenomena have to be resolved in detail, while other effects can be modeled appropriately. Finding suitable settings in mechanical ventilation of a patient for example, requires other effects and scales to be resolved than particle and aerosol transport on the alveolar scale. These considerations confirm that there cannot be a one-sizefits-all respiratory model but that each question in respiratory biomechanics requires a suitable approach for its investigation. In this work, an overview of the current state of the art in modeling respiratory biomechanics is provided. In general, models for the single zones as well as coupled models combining both zones are available. Each zone can be realized at different levels of resolution, either as fully resolved three-dimensional continuum or as dimensionally reduced model. Dimensional reduction in this context means any reduction of complexity either by elimination of spatial or governing physical details. Specific examples will be presented in the following subsections. A classification with these two criteria (i) conducting zone/respiratory zone and (ii) reduced/continuum based description is visualized in Fig. 2. In the following sections the visualized approaches will be discussed in detail, with model complexity increasing successively.

Reduced Models of the Conducting Zone Several decades ago, modeling approaches in respiratory biomechanics were restricted to reduced-dimensional observations of the conducting zone only. These models were motivated by the fact that the anatomy of the branching airway tree has been well known from lung casts which contained statistically relevant geometrical data up to the 16th generation of the airway tree. Specific effects observed in lung physiology were attributed to this complex branching airway tree structure and the flow patterns within. At that time, computational methods were not powerful enough to resolve the flow field in the entire conducting airway tree. Therefore, simplified approaches computing a flow resistance for each airway segment based on its geometrical dimensions have been used. These resistance models were mainly based on observations of laminar and turbulent flow in rigid pipes and the corresponding analytical solutions for example, Poiseuille’s resistance or modifications respecting turbulence. The assumption required for such reduced-dimensional models of the conducting zone was the applicability of the derived reduced-dimensional pressure-flow relationships in the airways under physiological conditions. Hence, all flow properties from fully laminar to turbulent flow conditions have to be represented correctly. In particular, it is assumed that airflow within a single airway segment is axisymmetric, fully developed, and that curvature of the single airway segments can be neglected. These prerequisites allow a spatial reduction of the three-dimensional Navier–Stokes equations (potentially including fluid– structure interaction effects) towards a one-dimensional (1D) formulation via integration within an airway cross section. It is further assumed that the pulsatile character of this 1D formulation can be omitted in respiratory biomechanics as the pressure wave travels throughout the entire airway tree in less than 0.1 ms and thus in a much smaller timescale than those relevant in respiration. Consequently, an integration in axial direction can be performed leading to the zero-dimensional (0D, or lumped) description of fluid flow in (compliant) airways. Another possible mathematical approach to build a formulation for airflow is the assumption that the

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Womersley solution holds for each airway segment. Then, an equivalent impedance can be computed to account for the viscous resistance, compliance, and inertance of airflow in each airway segment. The resistances or impedances for single airway segments can then been combined to a network specified by the anatomy of the airway tree and have been used for computations of airflow throughout the entire conducting zone. This basic idea of dimensional reduction of the conducting zone has been an important first step towards building up the discipline of airflow modeling in respiratory biomechanics. The applied resistance models were derived from basic physics of pipe flow and perfectly valid in the physiological range of airflow in the lungs. One major limitation of these models has, however, been the lack of respiratory tissue with all its important effects in lung physiology. Nevertheless, reduced-dimensional models of the conducting zone have successfully been used to investigate the resistance distribution across different generations of the airway tree. Further, the effect of expiratory flow limitation has been investigated using the extension of flexible airway walls. Expiratory flow limitation describes the maximum flow rate at which a subject is able to exhale. This flow rate cannot be increased with a higher forced positive pressure generated by the respiratory muscles due to the involved compression of the flexible airways and the resulting increase in flow resistance. Today, these models of the conducting zone have been extended by adequate representations of the respiratory zone towards reduced-dimensional multi-compartment models presented later in this article (see “Multi-Compartment Models” section).

Continuum Models of the Conducting Zone If a detailed spatial resolution of the airflow field is crucial, the use of continuum mechanics based conducting zone models is a prerequisite. In this case, airflow is modeled and simulated in two-dimensional or three-dimensional geometries of the conducting airways. Depending on the specific application, different regions of interest are considered in existing models, starting with individual airway models via single bifurcation models to multiple bifurcation models including different numbers of airway generations between the upper respiratory tract and the terminal bronchioles. Previously, idealized geometric representationsdthat is, individual tubes or a system of bifurcating tubesdhave been used predominantly. Well-known examples are the Weibel and Horsfield lung models, which are based on general morphological data about branching angles and generation-dependent airway diameters and lengths. Recently, the importance of patientspecific geometric features for the development of airflow patterns has gained more attention. Consequently, many models are meanwhile based on imaging-based geometries, that is, geometric reconstructions from, for example, bronchoscopic, X-ray, MR, or CT data. However, due to the limited resolution of imaging techniques and the high computational cost related to patientspecific (i.e., extremely complex) geometric models, it is not reasonable to resolve all airways in the conducting zone with such a level of detail. Airflow in the conducting zone is in general modeled by the incompressible Navier–Stokes equations. The assumption of incompressibility is appropriate since typical Mach numbers under physiological conditions are below 0.2. It is known that a high-speed jet is formed as air passes through the larynx. Resulting turbulence effects can affect flow patterns in the trachea and bronchi but are believed to decay already after few generations of airways. Depending on the specific region of interest, some models (particularly of the upper and central airways) consider the influence of potential turbulence effects either by resolving all relevant spatial and time scales (i.e., direct numerical simulation) or by applying turbulence models such as large-eddy simulation (LES) or Reynoldsaveraged Navier–Stokes (RANS) models. For some particular investigations also multiphase or non-isothermal models can be used. Most models of the larger airways consider the airway walls as being rigid. This assumption is often deemed suitable because the walls of the trachea and the first airway generations mainly consist of cartilage and are consequently rather stiff. Hence, cross sectional deformations in this region are usually small. However, in certain diseases (e.g., chronic obstructive pulmonary disease) airway cross sections can change significantly even during normal breathing. Furthermore, the composition of the airway wall changes over the generations and smaller airways are considerably more compliant. To account for the mutual interaction of airflow and airway wall deformation in these cases, so-called fluid–structure interaction (FSI) models have been proposed as an alternative to classical computational fluid dynamics (CFD) models. Apart from enabling the prediction of realistic airflow patterns in the deforming airways for above-mentioned scenarios, FSI models also allow for the quantification of airway wall strains and stresses. The latter are assumed to play an essential role, for example, for the alteration of cell shapes, biological signaling, and liquid secretion. It has to be noted, though, that a detailed modeling of airway wall properties is indispensable for realistic FSI models. So far, however, only few studies providing corresponding experimental data can be found in literature. The airways are embedded in the surrounding lung tissue, which exerts stresses on the airway walls during breathing or mechanical ventilation. Some models account for this so-called parenchymal tethering effect by introducing non-linear springs attached to the outside of the airway walls or considering an additional, empirically derived tethering pressure to the airway model. This way, the changes in flow characteristics as a consequence of disease-related alterations in the tethering forces (e.g., in asthma or fibrosis) can be investigated theoretically. Most existing continuum models of the conducting zone consider only a small part of the conducting zone, either since only local effects are of interest or because the inherent complexity of the lung inhibits a detailed modeling of all airways down to the respiratory zone. At the truncated airways in the models, suitable inflow and outflow boundary conditions have to be specified. At the inlet of the modeled airway tree, usually a time-dependent parabolic or constant velocity profile or a (time-dependent)

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pressure is prescribed. The definition of appropriate outflow conditions is more intricate. Boundary conditions at the outlets of the modeled airway tree have to be chosen such that (i) flow separates realistically between the different branches; (ii) during expiration (i.e., in case of a reversal of the flow direction), the problem remains well-posed and the volumes of previously inspired and expired air at each outlet are equal. These requirements are often difficult to fulfill in practice since they actually necessitate a detailed consideration of the unresolved downstream region. In most continuum models of the conducting zone, however, this influence is either disregarded or taken into account in an often unphysiological way. At the truncated outflows, rather simple boundary conditions are frequently prescribed. For instance, a priori defined outlet pressures, velocity profiles, or mass flow percentages are commonly prescribed. Alternatively, a so-called individual path model can be used. In this case, some pathways are resolved down to the peripheral airways, whereas most other pathways are truncated at low generations. The pressure or mass flow rate at equivalent interior locations of the resolved pathways is then mapped randomly to the outlets of the truncated branches. Other approaches have been based on subject specific boundary conditions to model the outflow from specific regions of the lung based on imaging data. These imaging-based outflow rates prescribed at single outlets are only valid for the specific flow scenario under image acquisition. However, they are of great value to validate a computed flow and radio marker distribution against available clinical measurements (e.g., of hyperpolarised noble gas distribution) in this specific scenario. The resulting numerical models can be applied for improving the general understanding of airflow in the conducting zone in health and disease. For instance, the influence of the airway tree geometry on airflow dynamics has been studied extensively. Severe differences with regard to airflow characteristics have been reported between patient-specific and idealized simplified geometries. Airway curvature and complex shapes at junctions have been found to be responsible for more complex flow patterns (including significantly more off-axis flow) in patient-specific geometries. Furthermore, the impact of pathological geometry changes (e.g., airway obstructions) on simulated airflow patterns has been surveyed. In case of airway constructions, for example, disturbed flows (including flow recirculation zones) as well as an increase in both pressure drop and work of breathing have been shown. Besides, the consequences of rapid inhalation (e.g., during sniffing) on airflow dynamics have been studied. While spontaneous fluctuations associated with turbulence have been reported in the supraglottic region, these fluctuations have been found to decay rapidly. Apart from the analysis of airflow patterns in the conducting zone, several studies have also been concerned with calculating airway wall (shear) stresses, for example, during coughing. For instance, substantial shear stress levels have been detected at expiratory flow-rates equivalent to cough, which can exacerbate damage to the epithelial layer of airway walls. Another important application field of continuum models of the conducting zone is the investigation of airway stability and reopening. Small airways are prone to fluid-elastic instabilities that can lead to their collapse and occlusion by a liquid bridge formed by the airway liquid lining. Since a persistent occlusion of the airways can lead to a severe impairment of gas exchange, medical treatment aims at a quick reopening of airways. At the same time, however, tissue forces resulting from the propagation of the “air finger” into the liquid-filled airway have to be minimized to avoid cellular injury and inflammation. Continuum models of the conducting zone can be used to determine (i) the propagation speed of the air finger as a function of the applied pressure and (ii) the stresses in the airway wall. Continuum models of the conducting zone are also the basis for the prediction of particle transport and deposition in the lung. Example scenarios are targeted drug delivery and the inhalation of toxic pollutants from the environment. For this type of application, the continuum models of the conducting zone have to be extended by a particle transport formulation and appropriate absorption boundary conditions. Although in general a mutual interaction is possible, the effect of the particle transport on the airflow is commonly deemed insignificant and, consequently, often neglected. In several studies, the influence of physical quantities (e.g., flow rate and particle size) as well as model parameters (e.g., time-dependent vs. steady flow conditions, idealized vs. imagingbased geometry) on particle deposition patterns and deposition efficiency have been investigated. It has been found that geometric features of airways, particle distributions, and the history of airflow fields can affect particle deposition significantly.

Reduced Models of the Respiratory Zone Similar to the conducting zone, the first models of the respiratory zone were restricted to reduced-dimensional observations. These approaches were motivated by the fact that the respiratory zone is composed of a huge number of small clustered units (i.e., the alveolar ducts) similar in their architecture. If the properties of a single alveolar duct can be expressed in a reduced-dimensional sense, the entire respiratory zone can then be adequately described. Models based on this idea have replicated alveolar tissue as a network of pin-jointed line elements (i.e., springs and dashpots) representing bundles of collagen and elastin fibers. These line elements are arranged at the edges and across the surfaces of regular polyhedra representing individual alveoli. Different quasi-static stress-strain behaviors are commonly attributed to the fiber bundles (i.e., linear for elastin and highly nonlinear for collagen), whereas identical stress relaxation functions are assumed to define the history of the stress response. Additionally, surface tension effects at the air-wall interface caused by a thin liquid lining on the inside of the single alveoli, the so-called surfactant, can be respected using membrane elements across the surfaces of these polyhedra. The main assumption that is required for a reduced-dimensional model of the respiratory zone is, that the chosen alveolar geometry and arrangement as well as constitutive model for the line elements are representative for human lung tissue. Inflation/deflation of such alveolar clusters then allows to compute quasi-static or dynamic relationships between pressure and volume of an alveolar duct. These relationships provide a computationally very efficient representation of the respiratory zone based on the underlying physics of collagen/elastin fibers and surfactant. They are used, for example, as terminal units in coupled lung

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models to mimic the behaviour of the respiratory zone and to quantify regional inflation of lung tissue (see “Reduced-Dimensional Coupled Lung Models” section). Further, reduced-dimensional models of the respiratory zone have also been used to study the effect of different diseases on the behaviour of lung tissue. For instance, by eliminating single pin-jointed line elements within a cluster, an emphysemic condition has been modeled. By this means, the failure of single alveolar walls on the microscopic scale has been related to a disease-related tissue softening on the macroscopic scale. Finally, spring models of the respiratory zone have been used to investigate the volume-competition between different lung regions during inflation in a simplified way and to quantify interactions between neighboring air spaces in health and disease. Still, reduced models of the respiratory zone are only describing one part of the lung, namely the respiratory zone. Therefore, they are limited to isolated observations of this region or have to be coupled to an adequate model of the conducting zone, as presented in the “Reduced-Dimensional Coupled Lung Models” section.

Continuum Models of the Respiratory Zone For a more detailed investigation of phenomena in the respiratory zone, continuum mechanics based models of single alveoli, alveolar ducts, or entire acini have been developed in the past. Many of these models are based on artificial geometric representations of individual alveoli ranging from simple spherical or polyhedral shapes to more realistic irregular cells. Some approaches also use imaging-based reconstructions, for example, obtained from synchrotron-based X-ray microscopic tomography. However, due to the small size of the air spaces, in vivo imaging of alveolar structures remains difficult. Continuum models of the respiratory zone have been used to study acinar flow phenomena in detail. Despite the low Reynolds numbers, airflow in this region is still complex due to the rhythmic expansion and contraction of the tissue during respiration. For instance, recirculating structures and radial flows induced by the septal wall movement can be found in the alveolar cavities. In addition to acinar flow patterns, the transport, mixing, and deposition of particles have been studied extensively. Applications include the prediction of acinar aerosol deposition for therapeutic delivery and the assessment of health risks associated with inhaled hazardous particles. Furthermore, the transport of oxygen molecules through the acinus and the exchange of oxygen at the blood-gas barrier have been simulated. Finally, continuum mechanics based models of the respiratory zone have also been applied to investigate alveolar stresses and strains. This way, local tissue strain “hotspots” have been identified which are at risk of overdistension during mechanical ventilation. Since stresses and strains cannot be measured experimentally, continuum models of the respiratory zone can make an important contribution to a better understanding of involved stress raising phenomena and associated risks at tissue overdistension. As with all isolated models of parts of the respiratory system, the formulation of physiologically reasonable boundary conditions (i.e., flow boundary conditions, boundary loads, and deformations) is difficult. In reality, the modeled alveolar or acinar structure is connected to the conducting zone and embedded in the surrounding respiratory zone. However, most existing models presented in this section are not capable of considering these effects adequately. Therefore, continuum models of the respiratory zone are primarily used for isolated, qualitative investigations.

Reduced-Dimensional Coupled Lung Models In the majority of applications, both a representation of the conducting zone as well as the respiratory zone are necessary to cover all aspects that are relevant for an accurate description of lung physiology. If, however, no detailed resolution of the flow field in the conducting zone and no fully resolved continuum mechanical description of the respiratory zone is required, the lung can be represented using reduced-dimensional coupled approaches. These simplified reduced-dimensional representations of respiratory biomechanics comprise approaches ranging from pure phenomenological single-compartment models to imaging-based, physiologically realistic approaches grounded on the physics of previously introduced airflow dynamics and tissue mechanics.

Single-Compartment Models In analogy to the basic anatomy of the lung, single-compartment models consist of a single pipe representing the conducting airways, which is connected to a single elastic compartment representing lung parenchyma. The motivation behind these models is that the resistance of the entire conducting airway tree can be combined into a single equivalent (pipe) resistance. Further, the compliance of all tissue components in the respiratory zone and the elastic properties of the chest wall are united into a single equivalent (compartment) compliance. The mechanical behaviour of the lung can then be expressed via fitting these two parameters (i.e., resistance and compliance) to patient measurements using the method of least squares. Since this approach is no physically based model in a strict sense, no modeling assumptions in a strict sense apply. The only two necessary assumptions for single-compartment models are that the chosen model equation is able to replicate lung behaviour and that a sufficient number of clinical measurements of pressure, flow and volume exists to fit the required parameters. A further assumption that is implicit in the representation via a single compartment is that the fit mechanical behaviour is averaged over the entire organ. This means that any regional differences in lung function are omitted in these models. Depending on the complexity of the model, different realizations for resistance and compliance are conceivable. If both the resistance and the compliance are constant the so-called linear single-compartment model is obtained. Several investigations have,

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however, shown that both the conducting zone and the respiratory zone contain significant sources of non-linearity. To begin with the respiratory zone, clinically measured transpulmonary pressures at different lung volumes recorded as quasi-static pressure– volume (P–V) measurements have revealed that the lung shows a sigmoidal-like inflation/deflation behaviour. This behaviour can be successfully described using exponential mathematical functions of transpulmonary pressure depending on lung volume. Especially the lower and the upper sigmoidal part of such P–V curves are important in modeling respiratory biomechanics. The lower part (i.e., at small pressures and lung volumes) reflects recruitment/derecruitment of air spaces from the initially degassed state and the upper part (i.e., high pressures and lung volumes) characterizes tissue overdistension. These two phenomena cause elevated stresses in the tissue and therefore require accurate quantification in clinically relevant modeling. Observing the conducting zone, airflow through the larynx and in the larger airways can become turbulent especially at high flow rates or in cases of highly pulsatile flows which leads to a flow-dependent non-linear resistance for example, described by Rohrer’s equation. Further, the pressure difference between the inside of a conducting airway and its surrounding causes diameter changes of the compliant airway, which in turn affect the airway’s resistance. This effect leads to resistance differences between inspiration and expiration due to the varying pleural pressure that is propagated to the airway’s surrounding and ultimately defines expiratory flow limitation explained previously. In general, single-compartment models are a phenomenological description of lung mechanics based on mathematical equations. The required parameters, namely equivalent resistance and equivalent compliance, are fit from clinical pressure and flow measurements allowing a quick and easy assessment of basic respiratory function. The required fitting algorithms for linear and for non-linear models are easy to implement, they operate in real-time and deliver values that are easy to interpret in a clinical setting. Still, these models do not have a strict physical background and rather have to be seen as a fitting technique. With more effects included for example, non-linearity or a relationship for recruitment/derecruitment, single-compartment models become more realistic and allow a better adaption to the physiology of the lung. On the other hand, each additional parameter requires more reliable data for fitting and thus the model becomes less predictive, or is even not predictive at all. Finally, one major drawback of the single compartment models is that no specific anatomy of the patient can be respected and no regional information on mechanical overstraining or recruited/derecruited regions can be given. This means that if equivalent resistance and compliance values lead to the conclusion that the lung of a specific patient is only working at 50%, no specification can be made whether the entire lung is working at only 50% or if one part is working completely normal whereas the other one is not working at all. Despite these limitations, single-compartment models are up to now the widest used modeling approach for respiratory biomechanics and still under development. They have successfully served as starting point for investigations of gas transport and transfer into the blood and for general coupling with the cardiovascular system. Further, they have been used in diagnosis of respiratory diseases as an easy to apply bedside tool in respiratory care. Recently, single-compartment models have been extended by a functionality to quantify ventilator-associated lung injury resulting from recruitment/derecruitment and overstraining. Finally, single-compartment models have successfully been applied in optimizing ventilatory settings in several randomized clinical trials. Nevertheless as it has been proven that the extent of lung heterogeneity is directly linked to disease severity and mortality, it would be desirable to use more precise regional models for this optimisation task.

Multi-Compartment Models Multi-compartment models are the next step towards more realistic modeling of respiratory biomechanics. They comprise all approaches that are characterized by multiple reduced-dimensional components for both the conducting and the respiratory zone of the lung and mark the transition from pure phenomenological approaches towards physically motivated models in respiratory biomechanics. In general, multi-compartment models are motivated by the idea that a reduced-dimensional description is the most efficient way to describe respiratory biomechanics on the organ-level and the awareness that the lack of regional information has to be overcome to allow precise conclusions in a clinical setting. Pure phenomenological multi-compartment models are characterized by a parallel arrangement of single-compartment models with distributed parameter values for equivalent resistance and compliance extended by models governing recruitment/derecruitment dynamics. The same assumptions hold as for single-compartment models except for the assumption that the behaviour is averaged over the entire organ (see “Single-Compartment Models” section). Required model parameters are still identified via fitting to patient measurements. Physically motivated multi-compartment models on the other hand are built upon the underlying physics. Specific assumptions are made to enable the reduced-dimensional description of both the conducting and the respiratory zone. The one-dimensional, zero-dimensional, or impedance-based representations of single airway segments of the conducting zone (see “Reduced Models of the Conducting Zone” section) are then combined to a morphologically realistic tree structure using either data from lung casts or tree-growing algorithms that generate a space-filling airway tree within a patient-specific imaging-based lung hull geometry. Additionally, each airway segment can be equipped with a representation of recruitment/derecruitment dynamics based on an additional variable that describes the opening state and its progression. The respiratory zone at the terminal ends of the airway tree or in the parallel arrangements of single-compartment models can also either be fit to pure phenomenological equations of lung tissue using for example, the previously mentioned exponential compliance equations, or be derived from physically motivated descriptions of lung tissue for example, based on alveolar duct models (see “Reduced Models of the Respiratory Zone” section). An important recent extension related to the conducting zone in multi-compartment models is the consideration of the interplay between single neighboring compartments, also known as lung interdependence, adding realistic stability to single inflating/deflating air spaces.

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Essentially all multi-compartment models are a functional relationship between pressure and flow in the conducting and the respiratory zone and allow for a spatial resolution of computed quantities in different regions of the lung. With the possibility to respect spatially distributed material properties and regionally varying threshold reopening pressures as well as gravitational effects, they allow a more realistic examination of lung function. Simple parallel arrangements of single-compartment models are still phenomenological representations of lung mechanics that have to be fit to measurements and thus are easy to adapt to a specific patient. For a satisfactory fit, the quality of the available measurements is decisive. The predictive character of these models suffers from the fact that it is not understood what happens in scenarios beyond those where fitting data are available. Conclusions on higher pressures than those measured are then only a more sophisticated mathematical extrapolation without deeper knowledge about potential critical points in system behaviour and thus dangerous for prediction in a clinical application. Physically based multi-compartment models allow a deeper insight into airflow throughout a network of compliant airway segments and inflation of (visco-)elastic lung tissue. In these models, the descriptions of the conducting and the respiratory zone are derived from physically sound airflow dynamics and tissue mechanics and extended by all capabilities that are necessary to describe the behaviour of the lung. They can include interdependence as well as the dynamics of recruitment/derecruitment. Verification against continuum mechanical representations of the conducting zone show that results from the reduced-dimensional models is in good agreement and even able to adequately take into account turbulence effects. Yet, the reduced-dimensional models are fast in their computation and deliver pressure and flow data that are easy to interpret in a clinical setting. These models allow a closer look into the black box of lung modeling and thus are more powerful than pure fitting approaches in terms of predicting critical or extremely beneficial states of lung function. They require only few data for patient-specific calibration, which means that they can deliver reliable data in the entire physiological pressure range in respiration. Furthermore, it is possible to integrate patientspecific information from medical imaging in form of the lung contours that serve as a limitation of the artificially grown airway tree. So far, several questions in respiratory biomechanics have successfully been investigated using multi-compartment models. Most importantly, the reopening dynamics of collapsed lung regions in acute respiratory distress syndrome have been assessed as a function of reopening pressure and time of the maneuver. In this context the optimal moments, pressures and duration of deep inflations during mechanical ventilation could be determined. Further, it has been possible to predict flow limitations in a healthy airway tree as well as the effect of heterogeneous bronchoconstriction and regional tissue heterogeneity on regional ventilation in diseased lungs. Besides, the propagation of a liquid plug in a complex network of reduced-dimensional airways could be studied and the associated frequency dependency of conducting airway and lung tissue behaviour could be determined. The aforementioned investigations address the basic concepts of cyclic closure/reopening and overstraining during mechanical ventilation of critically ill patients. The multi-compartment models have successfully enabled the identification of minimally injurious modes of ventilation in this context.

Hybrid Representations of the Respiratory System Hybrid models are the next level of detail in modeling respiratory biomechanics. They are used if only one zone requires specific attention and has to be modeled as a continuum while the other one can be respected using a reduced-dimensional approach. By coupling reduced and resolved representations, the interaction between the conducting and the respiratory zone of the lung is automatically assured. In practice, there are two main possibilities how such hybrid models can be composed. First, a continuum mechanical description of the conducting zone can be linked to a reduced-dimensional model of the respiratory zone. According to the zone that is resolved as continuum, this variant is denoted “hybrid (conducting)” in the overview (see Fig. 2). Second, a reduced-dimensional model of the conducting zone can be coupled to a continuum mechanical description of the respiratory zone, denoted as “hybrid (respiratory).” The two variants will be discussed in detail in the following.

Hybrid Model (Conducting) For some investigations, only the conducting zone is required to be resolved in full detail as far as possible from available imaging data, while phenomena related to the terminal airways and the respiratory tissue are of no specific interest. Still, the influence of the airways beyond full resolution and the respiratory zone cannot be neglected for a realistic description of airflow in the conducting zone. Therefore, the effects of flow resistance in smaller non-resolved airways, and the inflation of lung tissue have to be modeled via appropriate boundary conditions at the terminal ends of the fully resolved airway tree. The idea is to respect all downstream effects that generate dynamic pressures during inflation/deflation of the lung in a reduced-dimensional fashion. Additionally, a storage capacity for the air that leaves the fully resolved tree during inspiration, and can consistently re-enter the domain during expiration, is provided. Hence, the model denoted as “hybrid (conducting)” is essentially a continuum description of the conducting zone extended by appropriate methods to take into account all effects that occur downstream the fully resolved region. One assumption that is required for this kind of hybrid models is that there are no recirculation zones present at the terminal ends of the fully resolved conducting zone. Further, the elongation of airways in axial direction and the distortion of the airway tree with regards to branching angles between different generations, introduced by the inflation of lung tissue, is usually not included in this type of hybrid models. If elongation/distortion occurs, recirculation zones in the resolved tree could be influenced and investigations of, for example, particle transport and deposition as well as gas transport and mixing have to be carefully evaluated. For

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such observations, and investigations in flow limitation or bronchoconstriction, fluid–structure interaction and parenchymal tethering effects have to be considered (see “Continuum Models of the Conducting Zone” section). The governing equations for airflow in the continuum part of the conducting zone of such hybrid models are the incompressible Navier–Stokes equations. The non-resolved respiratory zone downstream of each resolved tree outlet is governed by a non-linear relationship between pressure at the outlet cross section and flow into as well as volume of the region of the respiratory zone attached. Different possibilities exist to derive such a relationship. It is for example, possible to couple the continuum conducting zone to a reduced-dimensional multi-compartment tree-like model to account for the resistance and compliance of the smaller nonresolved airways. Thereby, physically-based information can be included into the hybrid model such as the dimensions and connection of the non-resolved airways coming from a tree-growing algorithm. At the terminal ends of that tree, a reduced-dimensional respiratory zone model is attached as known from physically based multi-compartment models. Alternatively, regional functional diagnostics of a patient can be included to represent varying material properties of the different regions attached to the single outlets. One popular approach in this context is to derive an equivalent regional resistance and compliance of the respiratory tissue to be modeled and to prescribe a pressure-flow-volume relationship similar to the single-compartment equation at each outlet. In this case, additional measurements are required to fit the involved parameters such as equivalent regional resistance and compliance. Both approaches then lead to a realistic pressure level over time at the outlets of the continuum mechanical conducting zone and consequently to a correct pressure distribution in the entire domain. Neglecting the pressures resulting from non-resolved downstream passages can significantly corrupt flow patterns. Imagine for example, a case where airflow splits up between two almost symmetrical daughter airways of a resolved domain. Without the representation of the downstream regions, airflow would distribute almost equally between the two daughter airways. If however, one of these daughters supplies a stiffer lung region, in reality airflow to this region would be much lower while the other region would be overdistended. Also, the transport of curative aerosols to the stiffer region would be overrated and with it the potential for healing the increased stiffness. This brief example already shows that it is extremely important to model the non-resolvable regions of the respiratory system, especially in heterogeneous, that is, diseased lungs. A further important aspect is the storage possibility for air during inspiration. In some modeling studies in respiratory biomechanics expiration is modeled via a prescribed negative inflow at the trachea inlet. However, it is not clear from which of the single airway branches the outflowing air comes fromdnot to speak about the associated dynamics. This may lead to the fact that at some outlets more air is expired than actually inspired before, a behaviour which is clearly unphysical. In general, a storage capacity is an important aspect of the respiratory zone and has to be included in realistic models of respiratory biomechanics. This is why an impedance modeling of the reduced-dimensional respiratory zone is in general not suitable, as the impedance does not include a storage capacity for air volume. Important applications of the hybrid (conducting) models are in general all problems where a detailed flow field in resolvable airways is required and where induced deformation and distortion of the bronchial geometry is small. These requirements usually hold in large airways which are stiff enough and for small volume changes during respiration. Possible examples of such scenarios have already been mentioned in the “Continuum Models of the Conducting Zone” section and virtually all of these investigations would profit from a hybrid representation of the respiratory zone. One specific application in mechanical ventilation that is particularly suitable for this hybrid model is the ventilation with high frequencies and low tidal volumes known as high frequency oscillatory ventilation. In that case introduced deformation is low due to low tidal volumes and a resolved fluid field is required to fully understand the complex mixing processes in the conducting system during this efficient technique of mechanical ventilation.

Hybrid Model (Respiratory) A second approach of hybrid modeling is used if processes on the macroscopic tissue scale require specific attention (e.g., to quantify tissue overdistension) whereas airflow in the conducting system is only of interest inasmuch as it is responsible for the distribution of air in the respiratory zone. In this case, a suitable hybrid model can be realized by combining a physically based reduced-dimensional model of the conducting zone with a continuum mechanical imaging-based description of the respiratory zone. The geometry of the airway tree can be obtained using tree growing algorithms within the hull geometry of the lung segmented from a CT or MRI scan. The governing equations for the reduced-dimensional compliant airway tree in the conducting zone have already been introduced along with multi-compartment models in the “Multi-Compartment Models” section. The respiratory zone can be modeled both as a solid continuum and a porous medium. While the former model only accounts for solid mechanical phenomena, the latter approach additionally considers airflow through the sponge-like parenchyma (see Fig. 1) in a volume averaged sense, that is, without explicitly resolving the detailed microstructure. As regards the coupling of the conducting and respiratory zone models, pressure and flow have to be in equilibrium at the ends of the reduced-dimensional airway tree and the associated tissue regions. Hence, inflation of a tissue region can only happen if the conducting system supplies air to it and vice versa a prescribed deformation of a tissue region induces an airflow in the associated reduced-dimensional airway. The derivation of this hybrid model with a reduced conducting and resolved respiratory zone requires some assumptions. For the conducting zone, they are equal to those of multi-compartment models (see “Multi-Compartment Models” section). For the respiratory zone, the assumptions are less restrictive and only require sufficiently resolved imaging data to extract the geometry of the respiratory zone as well as a suitable continuum mechanically based constitutive law for lung tissue.

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In brief, the “hybrid model (respiratory)” is a continuum model for inflation/deflation of the respiratory zone governed not only by tissue/porous media mechanics but also by correct airflow dynamics to inflate/deflate a specific lung region. The benefit of the hybrid approach comes from the fact that the conducting zone is represented in a dimensional reduced fashion. Consequently, pressure and flow in the conducting airway network can be computed very efficiently. Further, more tree generations than obtainable from imaging data can be integrated in the model using well-known tree-growing algorithms. In theory it is possible to create an entire airway tree up to the 16th generation within the lung hull geometry segmented from medical imaging data. Then each terminal end of the airway tree can be coupled to a continuum model of the associated respiratory region. In this context, equilibrium of pressure and flow through the airway tree and at the specific region to be modeled has to be ensured. Since the structure of the lung parenchyma is modeled as a continuum, the interplay between neighboring air spaces during inflation/deflation is automatically included and thus lung interdependence is inherently considered in the model. Further, due to the full representation of the tissue, prescribing additional boundary traction forces resulting from kinematic constraints (e.g., from the thoracic cage or the diaphragm) is straightforward. Moreover, not only traction forces but also deformation paths for example, obtained from a temporal series of imaging data can be prescribed to the continuum representation of lung tissue. Equipped with these properties, the “hybrid model (respiratory)” is applicable for detailed investigations on regional overstraining of lung tissue and resulting stresses. Therefore it is perfectly suited for investigations in ventilator-associated lung injury which has been directly linked to structural stresses and strains. Further, the availability of a fully resolved structure is beneficial for the identification of suspicious tissue regions in tumor tracking and associated radiation therapy applications from an inverse fit of measured tissue movements. Recently, the effect of airway constriction and heterogeneous tissue elasticity on the distribution of tissue stress and alveolar pressure have been evaluated based on this type of hybrid models. These are important information for treatment planning in chronic obstructive pulmonary disease, emphysema, or fibrosis. Further hybrid couplings are conceivable such as, for example, a continuum mechanical description of the large airways combined with a reduced-dimensional description of the small airways in the conducting zone and a fully resolved representation of the respiratory zone. The decision which part has to be resolved as continuum and which one is sufficiently modeled with reduced approaches can be tailored to the current question to be investigated. In this way, very efficient and modular hybrid models can be generated to exactly fulfill the specific requirements for the specific application.

Continuum Coupled Lung Models Continuum mechanics based overall lung models offer the possibility to simulate coupled phenomena in two- or threedimensional representations of both the conducting and the respiratory zone at a high level of detail. For instance, a recently proposed continuum overall lung model is based on the coupling of a three-dimensional model of the airway tree with a threedimensional model of lung parenchyma. Similar to the model denoted as “hybrid (respiratory),” the parenchyma model is divided into subdomains associated with the outlets of airway tree now fully resolved as continuum. The volume of air passing through each outlet is then constrained to equal the change in volume of the corresponding parenchyma subdomain. This mutual coupling of airway and parenchyma models enables the simulation of reasonable deformation states and pressure levels since it links flow quantities, for example, pressure and resolved velocity profiles in the conducting zone, to local stresses and strains in the respiratory zone of the lung.

Concluding Remarks In current literature a variety of approaches exist for modeling respiratory biomechanics and their sensible usage is largely dependent on the problem to be investigateddsomething which is true for all meaningful modeling. For clinical monitoring of a patient with plenty of measurement data available and a fast and only global classification required at bedside, phenomenological fitting approaches are a quite reasonable choice. For more advanced investigations with single effects to be isolated or even to be predicted, more realistic respiratory models grounded on the underlying physics of airflow dynamics and soft tissue mechanics are more powerful and allow a deeper insight into physiological effects beyond classical black-box parameterisations of lung function. In very specific cases, realistic predictions are possible based on models comprising the conducting or respiratory zone only. The majority of phenomena in respiratory biomechanics, however, live from the interplay between the conducting and the respiratory zone and how alterations in one zone ultimately affect the behaviour of the other. Therefore, many models consider both the conducting and the respiratory zone as well as their interplay. Different levels of resolution for the two functional components have been introduced, namely fully resolved continuum and reduced-dimensional representations. Starting from phenomenological approaches that are purely based on parameter fitting of patient measurements, physically based respiratory models have been defined with increasing complexity throughout this work. At first, single-compartment models have been introduced which are often used in clinical monitoring to date. They have been extended towards physically based reduced-dimensional multi-compartment approaches retaining the fast and efficient solution and offering reliable quality in the prediction of regional pressures and ventilation. In case either the conducting or respiratory zone requires specific attention, hybrid models can be defined where one zone can be realized by a continuum description while the remaining zone can be respected by

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a reduced dimensional model. If a maximum level of detail is required either for extremely complex investigations or for validation of reduced-dimensional models, both zones can be resolved as three-dimensional continuum. Specific examples for potential applications of the single approaches have been presented in this article. They outline the fact that there is no “one-size-fits-all” approach in respiratory biomechanics, but rather a toolbox of methods available to solve the current problem at hand. An introduction towards this toolbox including model assumptions and a fair discussion of potentials and limitations for each approach is provided in this work and guides the reader towards the best model for current and future investigations in a clinical or scientific practice. We hope that this categorization sheds light on the current state of the art in modeling respiratory biomechanics and further promotes the development of methods that ultimately improve diagnosis and advances in individual patient treatment.

Further Reading Amini, R., & Kaczka, D. W. (2013). Impact of ventilation frequency and parenchymal stiffness on flow and pressure distribution in a canine lung model. Annals of Biomedical Engineering, 41, 2699–2711. Bates, J. H. T. (2009). Lung mechanics: An inverse modeling approach (1st edn.). Cambridge: Cambridge University Press. Bertram, C., & Gaver, D. P., III (2005). Biofluid mechanics of the pulmonary system. Annals of Biomedical Engineering, 33, 1681–1688. Denny, E., & Schroter, R. C. (2000). Viscoelastic behavior of a lung alveolar duct model. Journal of Biomechanical Engineering, 122, 143–151. Kleinstreuer, C., & Zhang, Z. (2010). Airflow and particle transport in the human respiratory system. Annual Review of Fluid Mechanics, 42, 301–334. Ma, B., & Bates, J. H. T. (2010). Modeling the complex dynamics of derecruitment in the lung. Annals of Biomedical Engineering, 38, 3466–3477. Maury, B. (2013). The respiratory system in equations. In A. Quateroni (Ed.), Modeling, simulation and applications (7th edn., pp. 1–278). Italia: Springer-Verlag. Rausch, S. M., Haberthür, D., Stampanoni, M., Schitty, J. C., & Wall, W. A. (2011). Local strain distribution in three-dimensional alveolar geometries. Annals of Biomedical Engineering, 39, 2835–2843. Roth, C. J., Yoshihara, L., Ismail, M., & Wall, W. A. (2016). Computational modeling of the respiratory system: Discussion of coupled modeling approaches and two recent extensions. Computer Methods in Applied Mechanics and Engineering, 314, 473–493. Smith, B. J., & Bates, J. H. T. (2015). Variable ventilation as a diagnostic tool for the injured lung. IEEE Transactions on Biomedical Engineering, 62, 2106–2113. Sznitman, J. (2013). Respiratory microflows in the pulmonary Acinus. Journal of Biomechanics, 46, 284–298. Tawhai, M. H., & Bates, J. H. T. (2011). Multi-scale Lung Modeling. Journal of Applied Physiology, 110, 1466–1472. West, J. B. (2012). Respiratory physiology–the essentials (9th edn.). Philadelphia, PA: Lippincott Williams & Wilkins. Yoshihara, L., Roth, C. J., & Wall, W. A. (2016). Fluid–structure interaction including volumetric coupling with homogenized subdomains for modeling respiratory mechanics. International Journal of Numerical Methods in Biomedical Engineering, 33, e2812.

Relevant Websites Auckland, n.d.dhttp://www.abi.auckland.ac.nz/en/about/our-research/lungs-respiratory-system.html, Auckland Bioengineering InstitutedLungs and Respiratory System. INRIA, n.d.dhttps://team.inria.fr/reo/respiration_modeling/, INRIA Paris-RocquencourtdRespiratory Modeling. Technical University of Munich, n.d.dhttps://www.lnm.mw.tum.de/research/applications/biomedical-respiratory-system/, Insititute for Computational Mechanics, Technical University of MunichdRespiratory Modeling. Tulane University, n.d.dhttp://www.tulane.edu/dpg/index.html, Biofluid Mechanics Laboratory, Tulane University.

Constitutive Modeling of Soft Tissues Michele Marino, Leibniz Universität Hannover, Hannover, Germany © 2019 Elsevier Inc. All rights reserved.

Introduction Structure–Mechanics Relationship Basics of Continuum Mechanics Hyperelastic Behavior Material Objectivity Stress–Strain Relationships Material Symmetry Invariant-Based Formulations Material Stability and Mathematical Requirements Incompressibility Condition The penalty method The augmented lagrangian method Application and final remarks on incompressibility State of the Art of Hyperelastic Constitutive Modeling Multiscale structural approach Comparison with experimental data Inelastic Behavior Viscoelasticity Active Response Plasticity Damage Growth and Remodeling Open Problems and Future Challenges Further Reading

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Introduction Constitutive modeling in mechanics is the art of describing the mechanical properties of materials through mathematical models, that is by means of mathematical problems formulated in connection with physical concepts and experimental evidence. The effectiveness of constitutive models plays a key role in the predictive capabilities of computational models of structures. In this framework, computational models of biological structures have proved themselves to speed up the rate of scientific discovery in medicine and to improve the effectiveness of clinical approaches. In silico analyses reduce indeed the need for expensive laboratory work and clinical trials because they are able to reproduce different pathophysiological scenarios in a rapid and low-cost way. Accordingly, the development of computational environments for biomedical research contributes to: (i) clarify the complex mechanobiological equilibrium that maintain the physiological behavior; (ii) identify relationships between histological and biochemical alterations with pathological responses; (iii) gain a better understanding of the etiology of diseases; (iv) support the tailoring of clinical treatments to patient-specific features. Soft tissues are biological tissues made up by cells, collagen, elastin, and ground matrix, not being mineralized. They provide the essential link and support for organs and biological structures throughout the whole human body, typical examples being tendons, ligaments, skin, muscles, blood vessels, heart, cornea, and intestine. Therefore, the mechanical response of soft tissues highly affects the functionalities of many body systems in health and disease (e.g., musculoskeletal, cardiovascular, digestive). Hence, the constitutive modeling of soft tissues is a frontier research challenge at the cutting edge of biomechanics. As a general rule of thumb, constitutive models shall be able to reproduce the in vivo or in vitro mechanical behavior of tissues, and in particular, the biomechanical phenomenon under investigation in the final numerical simulation. Moreover, the constitutive description shall be based on model parameters whose value can be determined from in vitro or in vivo observations. Finally, constitutive models shall respect mathematical requirements in order to be physically consistent and effective for conducting numerical simulations. The mechanical response of soft tissues is characterized by nonlinear stress–strain relationships associated with an incompressible or quasi-incompressible behavior due to the high water content (i.e., characterized by negligible or small volume variations). Moreover, tissue mechanics can be also affected by inelastic mechanisms, such as damage and viscous effects, as well as growth and remodeling. This article opens with a brief overview of the structure–mechanics relationship in soft tissues and with the description of basic ingredients in continuum mechanics needed for the development of constitutive models. The main part of this work

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addresses the constitutive modeling of soft tissues in terms of nonlinear hyperelasticity, anisotropic behavior, and incompressibility. Although the description of techniques for facing computational issues on numerical simulations of biological structures is beyond the scope of the present work, some hints on the treatment of the incompressibility constraint from a computational point of view will be provided, especially considering the coupling with an anisotropic behavior. Afterward, the state of the art in terms of some available constitutive approaches is presented with an insight on multiscale approaches. Finally, a general framework for the modeling of possible inelastic effects occurring in soft tissues will be introduced, addressing viscosity, active response, plasticity, damage, growth, and remodeling. Some remarks on open problems and future challenges conclude the article.

Structure–Mechanics Relationship Collagen and elastin are the constituents mainly responsible for tissue elastic, stiffness, and strength properties. The high water content in the ground substance also plays an important role from the mechanical point of view. Tissue constituents are organized following a precise and hierarchical multiscale arrangement. Elastin is assembled as a continuous network of fibers which are believed to have a key role in providing distensibility properties and elastic recoil to tissues. Nevertheless, the stiffness of the elastin network is significantly lower than the one of collagen. Accordingly, in elucidating the structure–mechanics relationship, the greatest attention is dedicated to collagen behavior. Collagen can be found in form of fibers arranged in agreement with a regular (e.g., tendons) or an irregular (e.g., skin) pattern. The arrangement of collagen fibers in regular tissues follows a predefined pattern and these can be conveniently classified in uni(e.g., tendons and ligaments) or multi- (e.g., arterial walls) directional. In unidirectional tissues, the main orientation of collagen fibers is unique and fibers can be retained as parallel one to each other. A multidirectional tissue is intended to be made up of a number of stacked thin layers, each of them with a regular unidirectional fiber arrangement. The basic building blocks of collagen fibers are tropocollagen molecules. The latter are proteins that can be found in more than 27 forms, depending on the structure. Type I collagen is the most abundant in the human body, being the most important for maintaining the structural integrity and a functional mechanics of soft tissues. Type I collagen molecules are made up by three polypeptide strands, each one being a left-handed helix (see Fig. 1). The three helices are twisted together into a triple helix quaternary structure about 300 nm long and 1–2 nm in diameter, stabilized by interstrand hydrogen weak bonds. Collagen molecules exhibit hydroxyproline-deficient sequences characterized by 60 residues, referred to as labile domains and indicated as molecular kinks (see Fig. 1). The overall length of molecular kinks is about 20 nm and it is comparable with the value of

Fig. 1 Hierarchical multiscale arrangement of collagen constituents in soft tissues (top). Mechanical response of soft tissues (bottom): J-shaped stress–strain relationship of a unidirectional tissue subjected to a uniaxial traction along the fiber direction, where the dominant mechanisms affecting the toe, heel, and linear regions are highlighted.

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the persistence length for collagen (about 14 nm). The persistence length represents the minimum contour length over which molecular segments fluctuate due to thermal energy. Therefore, molecular kinks are activated by thermal undulations, that is their configuration follows a statistical distribution due to thermal excitation. Molecular kinks can be straightened by forces applied at molecular ends that counteract thermal undulations, hence experiencing a transition from less ordered states (thermally activated kinks) to more ordered ones (nearly straight molecule). In this regime, usually referred to as entropic elasticity, the mechanical response of collagen molecules is mainly dominated by the flexural behavior of the polypeptide helices, rather than by the extensibility of intramolecular covalent bonds. This contributes to collagen extensibility up to molecular contour length. Nevertheless, collagen shows a significant level of molecular extensibility beyond its contour length. Accordingly, when molecular end-to-end length approaches the contour length, covalent bonds within the polypeptide strands are stretched out, inducing the transition from entropic elasticty to a different mechanism, which can be referred to as energetic elasticity. In soft tissues, tropocollagen molecules self-assemble to form long and continuous cylinder-like structures, named fibrils, characterized by a diameter between 50 and 500 nm (see Fig. 1). Collagen molecules are organized within fibrils by following a complex three-dimensional crystallographic pattern. Nevertheless, simple arrangement models are effective in capturing the key mechanical aspects of fibrils. For instance, according to the Hodge–Petruska scheme, fibrils can be successfully modeled as staggered arrays of parallel macromolecules with an axial offset of about 67 nm and an equilibrium center-to-center distance of about 1.5 nm between two transversally adjacent molecules. Within fibrils, molecules interact with each other by means of intermolecular covalent crosslinks (see Fig. 1). Accordingly, the elongation of collagen fibrils is affected by two mechanisms which can be retained as in series: the stretching of collagen triple helices (depending both on entropic and energetic mechanisms) and molecular rearrangement mechanisms (mainly, intermolecular sliding). Since cross-links prevent intrafibrillar sliding, the load transmission within fibrils is highly affected by the amount of cross-links. Indeed, the sliding-to-stretching ratio of the total fibril elongation highly depends on the biochemistry of cross-links production and renewal. Collagen fibrils are densely packed in bundles called fibers. Adjacent fibrils within fibers are stabilized by lateral fibril-to-fibril proteoglycan filaments. A controversial matter is if proteoglycans play or not a significant role in loading transfer among adjacent fibrils. Collagen fibers in soft tissues are characterized by a crimped microstructure (see Fig. 1). Fibers appear indeed as periodic-like curvilinear structures with characteristic length period in the order of mm. The crimp amplitude of collagen fibers is highly variable with tissue location and functional role, although generally in the order of one tenth of fiber period. The mechanical response of soft tissues is highly nonlinear and characterized by J-shaped stress–strain curves. In the case of a unidirectional tissue subjected to a uniaxial traction along the fiber direction, the stress–strain curve has been described as subdivided into three main regions where mechanisms occurring at very different length scales are dominant in each region (see Fig. 1): 1. Toe region (strain range z0%–2%): This is a low stiffness region associated with the straightening of the microscopic crimp in collagen fibers. 2. Heel region (strain range z 2%–4%): This is a region associated with a significant stiffening response due to the straightening of molecular kinks. 3. Linear region (strain range z4%–10%): This is a high stiffness region mainly related to fibril elongation (both collagen stretching and sliding). Accordingly, the structured and hierarchical arrangement of tissue constituents is the responsible for the peculiar mechanical behavior of soft tissues. In many applications, soft tissue mechanics under relatively large deformations (up to about 10% nominal strain) can be retained purely elastic. Approaches for the modeling of an elastic behavior will be introduced in what follows in the framework of nonlinear hyperelasticity. Nevertheless, the in vivo mechanical response of biological structures can be associated with the observation of mechanisms that cannot be described by employing a hyperelastic framework. The mechanical response of soft tissues depends indeed on the loading history. Addressing the cyclic loading–unloading testing of soft tissues, a hysteretic behavior (i.e., a stress-strain phase lag in the cyclic process) is generally experienced, accompanied by relevant strain-rate effects (see Fig. 2). By repeated cycling, eventually a steady state is reached at which no further change will occur. In this state, the tissue is said to be preconditioned.

Fig. 2

Inelastic behavior of soft tissues. Left: viscous-like response. Right: damage-related and plastic-like mechanisms.

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Furthermore, as schematically depicted in Fig. 2, damage-related and plastic-like behaviors can be obtained when soft tissues experience supraphysiological but subfailure loadings (e.g., during some surgical procedures such as balloon angioplasty in arteries). In particular, significant residual strains arise upon loading removal when plastic-like mechanisms occur, while damage is mainly associated with a degradation of tissue mechanical properties. In this framework, a behavior analogous to the Mullins effect in rubber materials can be revealed. The Mullins effect consists in a degradation of material stiffness whenever the load increases beyond its prior all-time maximum value. On the contrary, the final failure is associated with a drop in tissue stress. Moreover, tissue mechanics can be altered by a change in tissue structure driven by the biological activity of cells in response to both mechanical and biochemical stimuli. Finally, the presence of active elements in cells endows tissues with the ability to contract and relax in response to biochemical signals, affecting the mechanical response of biological structures. Some underlying nano- and microscale mechanisms responsible for the afore-introduced inelastic responses will be described when the modeling of an inelastic behavior will be introduced in the framework of generalized standard materials.

Basics of Continuum Mechanics Let E be a Euclidean space and V be the vector space associated with E. Moreover, denote by Lin the set of all linear transformations of V into itself, namely the space of all second-order tensors. As notation rules, first- and second-order tensors are indicated, respectively, in lowercase and uppercase boldface, while scalars are indicated as regular typeface. For a given (invertible) tensor A ˛ Lin, let A 1, AT, Tr(A), and Det (A) denote the inverse, the transpose, the trace, and the determinant of A, with Adj (A) ¼ Det (A) A 1 being the adjoint of A and Cof (A) ¼ (Adj (A))T the cofactor. Furthermore, given A, B ˛ Lin, the Frobenius inner product is denoted by A: B ¼ Tr(ATB). Finally, Linþ collects transformations in Lin with positive determinant, R is the set of real numbers and Rþ collects strictly positive real numbers. Introducing I ˛ Lin as the identity tensor, Q denotes the group of orthogonal tensors, i.e., 
(1) Q ¼ Q ˛ Lin s:t: QT Q ¼ QQT ¼ I ; Qþ, the group of special orthogonal (rotation) tensors, that is Qþ ¼ fQ ˛ Q s:t: Detð QÞ ¼ 1g; and U, the set of symmetric and positive definite tensors, i.e., 
U ¼ U ˛ Lin s:t: U ¼ UT and vT Uv > 0; cv ˛ V;vs0 :

(2)

(3)

A biological medium is regarded as a continuum body occupying region Uo ˛ E in the reference configuration, parametrized in xo ˛ V, and region U ˛ E in the current configuration, parametrized in x ˛ V (see Fig. 3). The deformation map 4 ˛ V maps point xo (material coordinates) onto points x (spatial coordinates), resulting x ¼ 4(xo). In what follows, the subscript o denotes quantities in

Fig. 3 Theoretical framework for the mechanics of biological structures. Basic ingredients of continuum mechanics, material symmetry principles, and representation of fiber reinforcements in soft tissues.

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the reference configuration, while quantities in the current configuration have no subscript. At this standpoint, Grad (•) is introduced as the gradient operator with respect to material coordinates and, introducing the time variable t, let x_ be the material time derivative of quantity x. The deformation map is locally described in terms of the deformation gradient tensor F, namely F ¼ Grad (4) ˛ Lin, which is a two-point tensor transforming vectors from the reference to the current configuration. Local invertibility enforces the nonsingularity of F, namely the Jacobian J ¼ Det (F) shall satisfy J > 0 or analogously F ˛ Linþ. Introducing unit vectors vo, no, v, and n, the infinitesimal line vodxo, area vector nodAo, and volume dUo in the reference configuration transforms in the corresponding quantities vdx, ndA, and dU in the current configuration through: vdx ¼ Fv o dxo ;

ndA ¼ Cof ðFÞno dAo ;

dU ¼ DetðFÞdUo :

(4)

Moreover, the (unique) polar decompositions R ˛ Qþ ;

F ¼ RU ¼ VR;

U; V ˛ U;

(5)

highlight that each deformation gradient can be regarded as the superimposition of a rotation (associated with a rigid-body motion) and a stretch, where U and V are named the right and left stretch tensors, respectively. Further common strain measures employed in constitutive models are the right Cauchy–Green deformation tensor C ¼ FTF and the Green strain E ¼ (C  I)/2. Equilibrium is enforced by local balance equations, namely balance of mass, balance of linear and angular momentum, and laws of thermodynamics. The balance of linear momentum and the second law of thermodynamics are here recalled, due to their relevance for defining constitutive laws. For a given kinematic quantity, chosen as reference measure for the development of the model, there exists a dual internal static quantity which produces power for a motion associated with the time derivative of the chosen kinematic quantity. For instance, the static quantity dual to the Green strain E is historically defined as the second Piola–Kirchhoff stress, denoted by tensor S. In the case of conservative applied loads, the balance of linear momentum can be expressed in terms of a stationary condition for the energy functional P ¼ P(4) ¼ !Uopint(4) dU þ Pext(4), namely Z  4 ¼ argmin pint ðhÞdU þ Pext ðhÞ ; (6) h ˛ C

Uo

where C collects kinematically admissible transformations, pint(4) is the potential internal energy accumulated in the deformed body per unit volume in the reference configuration, and Pext(4) is the potential energy of the external actions. The potential energy in a thermodynamic system can be expressed by means of the mass specific free-energy functional jfe. For reasons of material objectivity (detailed in the following), it is convenient to represent jfe as a function of the right Cauchy–Green deformation tensor C, whose dependence on the deformation map 4, that is C ¼ C(4), is implicitly accounted for in what follows. Moreover, in the framework of generalized standard materials, a set of internal variables V is introduced for describing possible internal mechanisms. The potential internal energy pint of a deformed continuum body at constant temperature can be represented in terms of the free-energy function jfe as: pint ð4Þ ¼ Jfe ðC; VÞ ¼ ro jfe ðC; VÞ ;

(7)

where Jfe is the free energy per unit volume and ro is the mass density in the reference configuration (here assumed to be constant in time, if not differently specified). Addressing an isothermal and reversible thermodynamic process, !UoJfe dU represents the largest quantity of work that can be gained from the deformed body. In case of irreversible processes, a certain amount of work is lost and internal dissipation occurs. The second law of thermodynamics enforces prescriptions on the internal dissipation Dint (per unit volume) which shall result nonnegative for any admissible thermodynamic process. Neglecting thermal effects, the second law of thermodynamics is referred to as the Clausius–Duhem inequality. In this framework, Dint represents the difference between the power produced by the internal static quantity and the rate j_ fe of the free energy released by the deformed body. Accordingly, by choosing the Green strain E as reference kinematic quantity (dual to the second Piola–Kirchhoff stress S), the Clausius–Duhem inequality prescribes: C_ Dint ¼ S : E_  ro j_ fe ¼ S :  ro j_ fe  0; (8) 2 for any admissible C_ and j_ fe . An admissible C_ corresponds to a motion which respects the constraints enforced by C, while conditions on the admissibility of j_ fe derive from the choice of internal variables V and of their evolution equations.

Finally, other useful stress measures derived from the second Piola–Kirchhoff stress S are the first Piola–Kirchhoff stress P and the Cauchy stress s: P ¼ FS;

1 1 s ¼ PFT ¼ FSFT : J J

(9)

The first Piola–Kirchhoff stress P is the static quantity dual to the deformation gradient F, while the Cauchy stress s produces _ 1 . Balance of angular momentum prescribes that both S and s are symmetric tensors. power for the spatial velocity gradient l ¼ FF Finally, it is worth pointing out that the Cauchy stress s is a true measure of the force per unit area dA in the current (deformed) configuration, the first Piola–Kirchhoff tensor P relates the force acting in the current configuration to the surface area element dAo in

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Biomechanics j Constitutive Modeling of Soft Tissues

the reference configuration (i.e., it is a two-point tensor), and the second Piola–Kirchhoff tensor S is the pull-back of P on the reference configuration (see Fig. 3).

Hyperelastic Behavior A material is said to be hyperelastic or Green elastic when the energy stored upon deformation is fully released upon the removing of the deforming cause. In other words, a hyperelastic material is conservative and a potential function exists, referred to as the strain energy, that physically represents the potential mechanical energy stored in the body in the current configuration at constant temperature. At this standpoint, function Jse is introduced as the strain energy density per unit volume in the reference configuration. It is worth highlighting that Jse shall depend on a measure of deformation and, due to its physical meaning, it shall result in a scalar-valued function taking positive values. Constitutive models differentiate themselves in terms of different choices of the strain-energy function. The properties of the strain-energy function directly follow physical and thermodynamical requirements, recalled in what follows. First of all, it is worth pointing out that the strain-energy density Jse results to be function of material point xo for two main reasons: (i) soft tissues are generally characterized by inhomogeneous mechanical properties within Uo; (ii) the deformation is generally inhomogeneous within the body, leading to inhomogeneous values of Jse. Nevertheless, in order to arrive at a more compact notation and if there is no danger of confusion, the dependence on xo of the individual functions is generally omitted.

Material Objectivity Material objectivity prescribes that the value of the strain-energy function is independent of superimposed rigid motions. AccordF of the deformation ingly, in order to rule out pure translations, the strain-energy density can be represented as a function Jse F F gradient F, namely Jse ¼ Jse(F). Moreover, the value of Jse shall result independent from rotations superimposed to F. Therefore, it results:

Jse ¼ JFse ðFÞ ¼ JFse ðQFÞ;

cQ ˛ Q:

(10)

A strain-energy function satisfying Eq. (10) is said to be objective. Use of the polar decomposition in Eq. (5) and the choice Q ¼ RT in Eq. (10) prescribe that the strain-energy function shall depend on the right stretch tensor U instead of F. In order to avoid the polar decomposition and since C ¼ U2, the strain-energy function can be conveniently regarded as function of the right Cauchy– Green deformation tensor C. Although choices based on different strain measures are also possible, material constitutive behavior will be defined in what follows in terms of the strain-energy represented as:

Jse : Linþ 1Rþ ;

Jse ¼ Jse ðCÞ:

(11)

Stress–Strain Relationships The relationship between stress and strain comes from energy definitions and thermodynamical requirements. For standard hyperelastic materials, the free energy Jfe is defined without considering any additional internal variable V, and equal to the strain energy Jse(C), namely

Jfe ¼ Jfe ðCÞ ¼ Jse ðCÞ:

(12)

By definition, the internal dissipation of a hyperelastic material shall be zero for any admissible deformation process. Accordingly, considering Eqs. (7), (8), and (12), it results:  _  C_ _ se ¼ S  2 vJse : C ¼ 0 c admissible C: _ (13) Dint ¼ S :  J 2 2 vC The choice: S¼2

vJse ; vC

(14)

allows to a priori fulfill the requirement prescribed by Eq. (13) that holds for hyperelastic materials. The relationship S ¼ S(C) obtained from Eq. (14) defines the constitutive law between tissue deformation and the resulting stress. In order to account for the heterogeneous nature of soft tissues, where fibers are immersed in a soft ground matrix, a common and general choice for the strain-energy density function in the homogenized body is to employ an additive decomposition for Jse of the form:

Jse ðCÞ ¼ VM JM ðCÞ þ VF JF ðCÞ;

SðCÞ ¼ 2VM

vJM vJF þ 2VF ; vC vC

(15)

Biomechanics j Constitutive Modeling of Soft Tissues

87

where JM and JF, respectively, represent the matrix and the fibers contributions to the strain-energy density function, being, respectively, averaged by means of the matrix VM and fiber VF volume fractions. If the reference configuration Uo is stress-free, then it is referred to as a natural configuration and S(I) ¼ 0 holds.

Material Symmetry Symmetries in the microstructural arrangement of constituents translate into symmetry properties of material behavior at the continuum level. For instance, mechanical tests could be not able to distinguish the properties of a material in two distinct reference configurations. Requiring the constitutive model to be consistent with material symmetries (if any), place further restrictions on the strain-energy density function Jse(C), or analogously on stress function S(C). Consider a change from Uo to a new reference configuration Uo* with material points identified by xo* and such that the transT formation gradient Grad (x*) o ¼ G belongs to the orthogonal group, namely G ˛ Q. Introducing F* as the nonsingular deformation gradient relative to Uo*, note that F ¼ F*GT, and F* ¼ FG, since G 1 ¼ GT (see Eq. 1). Moreover, the corresponding right Cauchy– Green deformation tensor C* associated to F* results C ¼ ðF ÞT ðF Þ ¼ GT CG;

cG ˛ Q:

(16)

Accounting for Eq. (16), the material symmetry group G with respect to the reference configuration Uo collects all the transformation gradients G such that the material response is independent from a change of the reference configuration, namely  
 G ¼ G ˛ Q s:t: Jse ðCÞ ¼ Jse GT CG ; cC ˛ Lin :

(17)

Equivalently, in terms of stresses, material symmetry prescribes (see Fig. 3):   GT SðCÞG ¼ S GT CG ;

cC ˛ Lin; cG ˛ G:

(18)

The symmetry group G reflects the symmetry of the physical properties of the material. Therefore, it has to be specified case by case, depending on the choice of the reference configuration and on microstructural properties. If constituents are arranged with no apparent preferred orientation in the reference configuration, the material response obtained from mechanical tests is identical for every rotation and reflection applied to Uo. In this case, the material is said to be isotropic in Uo and it holds G h Q. The strain-energy density function can be represented in terms of isotropic scalar-valued tensor function Jiso for which, by definition,   Jse ðCÞ ¼ Jiso ðCÞ ¼ Jiso GT CG ; cC ˛ Lin; cG ˛ Q : (19) If constituents are arranged by following preferred orientations in the reference configuration, the material response obtained from mechanical tests depends on the testing direction. In this case, the material is said to be anisotropic in Uo and it results G 3Q. Soft tissues generally present anisotropic properties. For the sake of modeling, their anisotropic behavior is commonly described by introducing a discrete collection of nF fiber families, each of them with direction described by the unit vector eo(a) ˛ V in the reference configuration with a ¼ 1, ., nF (see Fig. 3). In what follows, as a superscript rule, let a imply values in {1, ., nF}. Whenever nF ¼ 1 fiber family is employed, superscript (a) is omitted. In order to lose the dependence on the orientation of eo(a), that is a different response for e(a) o , the structural tensor (a) M ¼ eo(a) 5 eo(a) is employed in constitutive models. Geometrical symmetries in fiber arrangement are described by the invariance set GM of structural tensors with respect to the reference configuration Uo, defined as
 GM ¼ G ˛ Q s:t: GT MG ¼ M ; where M ¼ {M(1), ., M(nF)} and GTMG denotes{GTM(1)G, ., GTM(nF)G}. From the Rychlewski’s theorem, the condition   Jse ðCÞ ¼ Jse GT CG ; cC ˛ Lin; cG ˛ G;

(20)

(21)

is satisfied if and only if the strain-energy Jse can be represented by a function Jani whose list of arguments additionally includes the structural tensors, namely

Jse ðCÞ ¼ Jani ðC; MÞ; which results to be an isotropic scalar-valued tensor function, namely   Jani ðC; MÞ ¼ Jani GT CG; GT MG ;

cC ˛ Lin; cG ˛ Q:

Indeed, it is immediate to verify that the isotropic tensor function Jani (i.e., for which Eq. 23 holds) respects:   Jani ðC; MÞ ¼ Jani GT CG; M ; cC ˛ Lin; cG ˛ GM 4Q; and thereby, the invariance set GM in Eq. (20) is a material symmetry group for anisotropic materials.

(22)

(23)

(24)

88

Biomechanics j Constitutive Modeling of Soft Tissues

Invariant-Based Formulations The construction of constitutive equations takes advantage from the classical invariant theory which allows to build isotropic functions from a basis of invariants. The latter is made up by a minimal set of invariants from which all other invariants can be generated. By following the additive decomposition in Eq. (15), the matrix contribution JM is generally associated with an isotropic constitutive response, while the fiber term JF collects the strain-energy contribution of the constituents which endow the tissue with an anisotropic response. In turn, the latter is additively decomposed in order to account for the different fiber families. Accordingly, on the basis of functions Jiso and Jani introduced in Eqs. (19) and (23), respectively, it results:

JM ðCÞ ¼

1 Jiso ðCÞ; VM

JF ðCÞ ¼

nF 1 1 X ðaÞ ðaÞ Jani ðC; MÞ ¼ v Jani ðC; MðaÞ Þ VF VF a¼1 F

(25)

(a) where Jani is the strain-energy contribution of the single fiber and vF(a) is the probability of finding a fiber with orientation e(a) o such P that anF¼ 1vF(a) ¼ 1. It is worth pointing out that both Jiso(C) and Jani(C, M) are isotropic scalar-valued tensor functions. For the anisotropic term, in particular, this is true thanks to the introduction of the notion of structural tensors. Polynomial invariants are focused in the following, although other choices are possible. A polynomial basis for the isotropic term Jiso(C) is made up by the set I iso¼ {I1, I2, I3}, where:

I1 ¼ Tr ðCÞ ¼ kFk2

(26a)

I2 ¼ TrðCof ðCÞÞ ¼ kCof ðFÞk2 ;

(26b)

I3 ¼ DetðCÞ ¼ kDetðFÞk2 :

(26c)

which geometrically represent, respectively, a measure of line, area, and volume change, as shown in Eq. (4). Introducing function

^ iso with scalar-valued arguments, it can be chosen: J

^ iso ðI iso Þ; Jiso ðCÞ ¼ J

(27a)

and the corresponding second Piola–Kirchhoff stress tensor Siso results:  ^  ^ ^ iso vI2 vJ ^ iso vI3 ^ iso  ^ iso ^ iso vJiso vJiso vI1 vJ vJiso vJ vJ vJ Siso ðCÞ ¼ 2 ¼2 þ þ ¼2 I þ vI1 Cþ Cof ðCÞ ; vC vI1 vC vI2 vC vI3 vC vI1 vI2 vI2 vI3

(27b)

where the following identities have been considered: vI1 ¼ I; vC

vI2 ¼ I1 I  C; vC

vI3 ¼ Cof ðCÞ: vC

(27c)

A polynomial basis for the anisotropic term Jani(C, M) is made up by the set I ani¼ {I iso, I F(1), ., I F(nF), I M }, with I iso from (1) ,., IM (nF)}with: Eq. (26) and where I F(a) ¼ {I4(a), I5(a)} and I M ¼ {IM   ðaÞ ðaÞ ðaÞ I4 ¼ TrðCMðaÞ Þ; I5 ¼ Tr C2 MðaÞ ; IM ¼ TrðMðaÞ Þ (28) In particular, I M is generally ruled out since identically equal to the unitary set. Invariant I4(a) geometrically represents the square stretch along the direction identified by eo(a), resulting I4(a) ¼ kFeo(a)k2. ^ ðaÞ with scalar-valued arguments, it can be chosen: Introducing I (a) ¼{I , I (a)} and function J ani

iso

F

ani

ðaÞ 

ðaÞ



ðaÞ ^ ðC; MðaÞ Þ ¼ J ^ J ani ani I ani ;

(29a)

(a)

and the corresponding second Piola–Kirchhoff stress tensor Sani results: ðaÞ

Sani ðCÞ ¼ 2

 ^ ðaÞ ðaÞ  ^ ðaÞ ðaÞ ^ ðaÞ vIðaÞ ^ ðaÞ vJani e vJani vI4 vJ vJani ðaÞ vJ ðaÞ ðaÞ ani ani 5 e ¼ Siso ðCÞ þ 2 þ ¼ S ð C Þ þ 2 M þ ðCM þ M CÞ ; iso ðaÞ vC ðaÞ vC ðaÞ ðaÞ vC vI4 vI5 vI4 vI5

(29b)

where e Siso is defined as:  ^ ðaÞ ^ ðaÞ  ^ ðaÞ ^ ðaÞ vJani vJ vJ vJ ani ani e Siso ðCÞ ¼ 2 þ I1 ani I  Cþ Cof ðCÞ ; vI1 vI2 vI2 vI3

(29c)

and the following identities have been considered: ðaÞ

vI4 ¼ MðaÞ ; vC

ðaÞ

vI5 ¼ CMðaÞ þ MðaÞ C: vC

It is worth highlighting that e Sani ¼ 0 whenever I (ania) h I (Fa) .

(29d)

Biomechanics j Constitutive Modeling of Soft Tissues

89

In conclusion, the second Piola–Kirchhoff stress of soft tissues derived from the choice in Eq. (15) is SðCÞ ¼ VM SM ðCÞ þ VF SF ðCÞ;

(30a)

where, accounting for Eqs. (25), (27a), and (29a), it results SM ðCÞ ¼ 2

vJM 1 ¼ Siso ðCÞ; VM vC

SF ðCÞ ¼ 2

nF vJF 1 X ðaÞ ðaÞ ¼ v S ðCÞ; VF a¼1 F ani vC

(30b)

(a) given in Eqs. (27b) and (29b). with Siso and Sani Tendons and ligaments are made up by regular unidirectional tissues that can be modeled by introducing nF ¼ 1 fiber family. In this case, tissue behavior results transversely isotropic since only one preferred orientation is present. As a matter of fact, the invariance set in Eq. (20) is made up by all rotations about the eo-axis through an angle q, represented by tensor Rq (eo, q)˛ Q, as well as by a reflection with respect to the plane with unit normal equal to eo, represented by tensor Qr (eo) ˛ Q. Accordingly, the symmetry group Gti for a transversely isotropic tissue results

Gti ¼ f  I; Qr ðeo Þ;Rq ðeo ; qÞ

with 0  q < 2pg

(31)

On the other hand, a large number of biological structures (e.g., arterial walls, heart valves) can be regarded as planar and made up by regular multidirectional tissues that can be modeled by introducing nF ¼ 2 fiber families. The latter lie on tissue plane and form an angle 2bo between each other, namely eo(1) , eo(2) ¼ cos (2bo). For the analysis of material symmetries, the superposition of fibers may require a more sophisticated treatise than the analysis of the invariance set in Eq. (20), which is representative only for symmetries in the geometric arrangement of fibers. Starting from the orientation eo(1) and eo(2) of fibers, the local orthonormal coordinate basis (g1, g2, g3) can be introduced as: ð1Þ

ð2Þ

eo þ eo g1 ¼ ð1Þ

;

eo þ eoð2Þ

ð2Þ

ð1Þ

eo  eo g2 ¼ ð2Þ

;

eo  eð1Þ

o

g3 ¼

g1  g2 ; kg1  g2 k

(32a)

such that: ð1Þ

eo

ð2Þ

¼ cosðbo Þg1 þ sinðbo Þg2 ;

eo ¼ cosðbo Þg1  sinðbo Þg2 :

(32b)

Recalling that Qr (a) represents a reflection with respect to the plane with a as unit normal and introducing Qr,• ¼ Qr (g•), it results: ðiÞ

ð jÞ

ðiÞ

Qr;1 eo ¼ eo ;

ð jÞ

ðiÞ

Qr;2 eo ¼ eo ;

ðiÞ

Qr;3 eo ¼ eo ;

(33)

with i, j ˛ {1, 2} and i s j. Hence, the following relationships hold:    ðiÞ ðiÞ T ð jÞ ¼ TrðCMð jÞ Þ ¼ I4 ; Tr QTr;q CQr;q MðiÞ ¼ Tr CQr;q eo Qr;q eo

(34a)

  ðiÞ  ðiÞ T  ðiÞ Tr QTr;3 CQr;3 MðiÞ ¼ Tr CQr;3 eo Qr;3 eo ¼ TrðCMðiÞ Þ ¼ I4 ;

(34b)

where i, j, q ˛ {1, 2} and i s j. Without loss of generality, the anisotropic term is assumed to depend on the fourth invariant of ^ ðaÞ IðaÞ . Based on the representation in Eq. (25) and accounting for relationships in Eq. (34), it ^ ðaÞ ¼ J deformation only, namely J holds (for q ˛ {1, 2}):

ani

ani

4





ð1Þ 



ð2Þ 



ð1Þ ð2Þ ^ ð2Þ ^ Jani C; Mð1Þ ; Mð2Þ ¼ vFð1Þ J þ vF J ani I4 ani I4 ;





(35a)

ð1Þ 



ð2Þ 



ð1Þ 



ð2Þ 



ð2Þ ð2Þ ^ ð1Þ ^ Jani QTr;q CQr;q ; Mð1Þ ; Mð2Þ ¼ vFð1Þ J þ vF J ani I4 ani I4 ;





ð1Þ ð2Þ ^ ð2Þ ^ Jani QTr;3 CQr;3 ; Mð1Þ ; Mð2Þ ¼ vFð1Þ J þ vF J ani I4 ani I4

(35b) (35c)

ð1Þ ^ ð1Þ ð2Þ ^ ð2Þ Accordingly, if it holds vF J ani ¼ vF Jani (i.e., roughly speaking, fiber families have the same mechanical response), it is immediate to show that:     Jani C; Mð1Þ ; Mð2Þ ¼ Jani GT CG; Mð1Þ ; Mð2Þ ; (36)

for G ¼  I,  Qr,1,  Qr,2,  Qr,3. Therefore, in this case, material behavior is orthotropic according to axes g1, g2, and g3 with material symmetry group Gorth in the reference configuration Uo equal to:
 (37) Gorth ¼  I; Qr ðg1 Þ;Qr ðg2 Þ;Qr ðg3 Þ :

90

Biomechanics j Constitutive Modeling of Soft Tissues

Material Stability and Mathematical Requirements The notion of material stability is related to the physical reasonability of the material response. A material is said to be stable if any deformation increment from a given state needs a positive work to be produced. In the framework of an isotropic behavior, material stability can be expressed by the Baker–Ericksen inequality which requires that the maximum principal Cauchy stresses occur in direction of the maximum principal stretches. More generally, addressing both isotropic and anisotropic responses, material is said to be stable if the Legendre–Hadamard inequality is respected at each material point. The latter condition ensures the existence of traveling waves with real wave speeds in each direction. This requirement can be verified by the analysis of the ellipticity of the acoustic tensor, obtained from the second derivative of the strain-energy function JseF (F) with respect to the deformation gradient F. Nevertheless, the existence of minimizers of variational principles in finite elasticity is not guaranteed for stable materials. Therefore, further restrictions shall be enforced on the strain-energy function for ensuring the solution of boundary value problems associated with finite elasticity. Sufficient conditions for the existence of minimizers are coercivity and sequential weakly lower semicontinuity (s.w.l.s.) of the strain-energy function. F (F) / þ N for extremely high strains. A possible The coercivity condition corresponds to enforce the growth condition that Jse coerciveness formulation is to require that there exists a1 > 0, a2, p  2, q  p/(p  1) and r > 1 such that:

 JFse ðFÞ  a1 kFkp þ kcof ðFÞkq þ ðDetðFÞÞr þ a2 ; cF ˛ Linþ : (38) Moreover, a second growth condition is generally introduced in order to enforce that an infinite energy shall be required to annihilate volume:

JFse ðFÞ/ þ N for DetðFÞ/0þ ;

cF ˛ Linþ :

(39)

By employing the additive decomposition in Eq. (15) with positive additive terms, the strain-energy density function is coercive when, at least, one additive term respects the coercivity condition in Eq. (38). Accordingly, from a practical point of view, coercivity is enforced on the matrix contribution JM of the strain-energy function and several expressions classically employed in the state of the art for soft tissues respect this condition. The sequential weakly lower semicontinuity of a strain-energy function is implied by its polyconvexity in the sense of Ball. Polyconvexity is a stronger condition than s.w.l.s. but it results more feasible to handle from the mathematical point of view. The strainF (F) is said to be polyconvex if there exists a convex function fj (in general, not unique), energy function Jse fj ¼ fj ðA; B; cÞ; with A; B ˛ Lin; c ˛ R; (40) such that

JFse ðFÞ ¼ fj ðF; AdjðFÞ;DetðFÞÞ:

(41)

It is worth highlighting that all terms in Eq. (15) should satisfy the polyconvexity requirement. A construction principle for a polyconvex strain-energy function is to introduce sets P iso (resp., P (ania)), collecting elements of I iso (resp., I (ania)) that result to be convex functions in the list of arguments F, Adj (F) and Det (F). Then, accounting for Eqs. (27a) and (29a), an additive decompo^ ðaÞ is employed: ^ iso and J sition of strain-energy terms J ani X ^ iso;j Pj jPj ˛ P iso ðI iso Þ; ^ iso ðI iso Þ ¼ J J (42a) j ðaÞ 



ðaÞ ^ J ani I ani ¼

X

ðaÞ





ðaÞ ðaÞ ^ J ani;j Pj jPj ˛ P ani I ani



;

(42b)

j

^ iso;j and J ^ ðaÞ shall result convex and monotonically increasing functions of arguments in sets P and where the single terms J iso ani;j ( a) P ani . Addressing isotropic contributions, some convex terms in P iso are: pffiffiffiffi I1 I2 1 ; I3 ; ln I3 ; ; (43) I1 ; I1 ¼ 1=3 ; I2 ; 1=3 I3 I3 I3 where terms with I3 at the denominator are useful for respecting the growth condition in Eq. (39). On the other hand, possible convex anisotropic terms in P (a) ani are: ðaÞ

I4 ;

ðaÞ

I4 ¼ ðaÞ

ðaÞ

I4

1=3 I3

;

ðaÞ

K2 ¼ I1  I4 ;

ðaÞ

ðaÞ

ðaÞ

K1 ¼ I5  I1 I4 þ I2 ; ðaÞ

ðaÞ

ðaÞ

K3 ¼ I1 I4  I5 :

(44a)

(44b)

Terms K1(a), K2(a), and K3(a) are introduced to account for invariant I5(a) which is not elliptic and hence nonpolyconvex. Their physical meaning can be elucidated by showing that:

ðaÞ ðaÞ 2 (45a) K1 ¼ Cof ðFÞeo ;

Biomechanics j Constitutive Modeling of Soft Tissues ðaÞ

K2 ¼ TrðCðI  MðaÞ ÞÞ;

91

(45b)

ðaÞ ðaÞ 2 K3 ¼ kCof ðFÞk2  Cof ðFÞeo : (a)

(a)

(45c) (a)

Hence, K1 controls the deformation of the area element with unit normal eo , K2 the square stretch in the plane with unit normal eo(a), and K3(a) the deformation of an area element with unit normal perpendicular to eo(a).

Incompressibility Condition Soft tissues demonstrate a very slight volumetric compressibility (< 1%), and hence the constraint Det (F) z 1 shall be enforced. This brings inherited computational issues that require the implementation of special numerical techniques. Among others, two methods are here presented: the Penalty method and the Augmented Lagrangian method. As preliminary material for both methods, it is useful to consider the multiplicative split of the deformation gradient F into the volumetric part Fvol and the isochoric part Fdev such that: with Fdev ¼ J 1=3 F; Fvol ¼ J1=3 I;

F ¼ Fvol Fdev ¼ Fdev Fvol

(46)

with Det (Fvol) ¼ J and Det (Fdev) ¼ 1. The volumetric part Fvol is associated with a motion that changes the volume but preserves the shape. On the contrary, the isochoric part Fdev preserves the volume but changes the shape. For material objectivity requirements, let T Fdev ¼ I3 1/3C be introduced as the isochoric part of the right Cauchy-Green deformation tensor. Cdev ¼ Fdev For the sake of conciseness and without loss of generality, only nF ¼ 1 fiber family is employed in what follows when dealing with the anisotropic contribution. Moreover, let the deviatoric operator Dev ((•)) be defined as Dev ((•)) ¼ (•)  [(•): I]I/3.

The penalty method The Penalty method consists to enforce the incompressibility constraint by an additive splitting of the isotropic strain-energy contribution as ^ ðI3 Þ þ J ^ ðC Þ; Jiso ðCÞ ¼ kJ dev iso iso vol

^ vol J iso

where k > 0 is a penalty parameter, endowed by the following properties:

dev

is a penalty function, and

^ : Rþ 1Rþ Wf0g; J iso

^ dev J iso

is the isochoric strain-energy part. The term

^ is convex; J iso

vol

vol

^ ðI3 Þ ¼ 05I3 ¼ 1; J iso vol

(47) ^ vol J iso

shall be

(48a)

such that it immediately follows that: ^ vJ iso ¼ 05I3 ¼ 1: vI3 vol

(48b)

Addressing the isotropic matrix contribution of the Cauchy stress siso ¼ FSisoFT/J with Siso in Eq. (27b), the physical rationale vol dev and siso behind the constitutive assumptions in Eq. (47) is elucidated by considering the additive decomposition of siso in siso such that dev siso ¼ svol iso I þ siso ;

svol iso ¼

Trðsiso Þ ; 3

sdev iso ¼ Devðsiso Þ:

(49a)

Employing the constitutive choice in Eq. (47), the resulting stress can be split into a physically meaningful decomposition: ! ^ vol ^ vol ^ dev vJ vJ 2 vJ dev iso iso iso T svol Dev F ; s ¼ 2kJ ¼ k ¼ F : (49b) dev iso iso J vI3 vJ vCdev dev Accordingly, addressing exclusively the isotropic part, the split in Eq. (47) ensures that change of volume and change in shape are disjoint, being respectively related to a spherical or a deviatoric state of stress. The former corresponds to a stress state siso of the form siso ¼ Tr(siso)I/3, while the latter is such that Tr(siso) ¼ 0. Indeed, any spherical state of stress will produce only a change of volume vol and not a change of shape, since siso s 0 (see Eq. 49a) and hence I3 s 1 from Eq. (48b). On the other hand, any deviatoric state of vol ¼ 0 (see Eq. 49a) and hence I3 ¼ 1 from Eq. stress will produce a change of shape but no change of volume, since it shall result siso (48b). 1=3 is A convenient definition of the isochoric term is based on an invariant formulation where only the term I1 ¼ TrðCdev Þ ¼ I1 I3 employed: ^ Jdev iso ðCdev Þ ¼ Jiso ðI1 Þ dev

(50)

^ Since I1 is polyconvex (see Eq. 43), Jdev (Cdev) is ensured to be polyconvex if function J dev is convex and monotonically 2=3 increasing. On the other hand, the purely isochoric term I2 ¼ TrðCof ðCdev ÞÞ ¼ I2 I3 is not polyconvex.

92

Biomechanics j Constitutive Modeling of Soft Tissues Attention shall be paid when adopting the Penalty method in the presence of an anisotropic contribution due to the fiber term

JF. In agreement with the additive decomposition in Eq. (15) and positions in Eq. (25), the Cauchy stress contribution

sani ¼ FSaniFT/J, with Sani inEq. (29b), is added to the total Cauchy stress, which indeed results:  vol dev dev s ¼ siso þ sani ¼ svol iso þ sani I þ siso þ sani :

(51)

vol dev vol dev ¼ Tr(sani)/3 and sani ¼ Dev(sani) is introduced, while siso and siso are given in Eq. Here, the volumetric-isochoric split sani (49b). The analysis is here restricted to the case where the anisotropic contribution is associated only with the fourth invariant of deformation I4, namely when anisotropic stress contribution is related only to fiber elongation. As a consequence, stress sani is expected pffiffiffiffi to be one-dimensional and aligned with the unit vector u ¼ Feo = I4 . Following the rationale employed in Eq. (47) for the isotropic part, a natural choice would be to introduce the dependence of the strain-energy function on the deviatoric part Cdev of the right Cauchy–Green deformation tensor, namely

^ Jani ¼ Jdev ani ðCdev ; MÞ ¼ Jani ðI4 Þ; dev

(52)

^ dev of a scalar argument is introduced. Therefore, where I4 ¼ Tr(CdevM) is polyconvex (see Eq. 44) and the scalar-valued function J ani fiber stress results sani ¼

 ^ dev  2I4 vJ I ani ; u5u  3 J vI4

(53a)

not contributing to the spherical part of the stress tensor, since svol ani ¼

 ^ dev  2I4 vJ I ani ¼ 0; Tr u5u  3 3J vI4

(53b)

^ dev 2I4 vJ ani Devðu5uÞ: J vI4

(53c)

sdev ani ¼

vol ¼0 Therefore, a purely deviatoric state of stress does not change the volume since, from Eqs. (51) and (53b), it must result siso and this holds for I3 ¼ 1 (see Eqs. 48b and 49b). On the other hand, a purely spherical state of stress induces a spherical state of vol s 0), since the fiber contribution to the spherical stress state is a priori null (see Eq. 53b). Moreover, in stress in the matrix (i.e., siso case of slight compressibility, the purely volumetric deformation of the matrix can be attained, resulting in a null deviatoric stress dev dev dev ¼ 0. This leads to a null stress in the fibers because sani ¼  siso ¼ 0 (see Eqs. 49b, 51, and 53c). Therefore, in the formustate siso lation of Eq. (52), a purely spherical state of stress can be accompanied by a null fiber stress and a null change of shape, in contrast with the anisotropic properties of the material. Furthermore, as a further drawback, the resulting Cauchy stress sani in Eq. (53a) is by no means one-dimensional, contradicting the physical sense. To avoid these unrealistic physical responses, the dependence of the strain-energy function on the complete right Cauchy–Green deformation tensor C could be adopted, namely

^ ani ðI4 Þ; Jani ðC; MÞ ¼ J

sani ¼

^ ani 2I4 vJ u5u: J vI4

(54)

recovering the one-dimensionality of the resulting Cauchy stress. It is immediate to show that, in this case, fibers contribute to both the spherical and the deviatoric parts of the stress tensor, namely svol ani ¼

^ ani 2I4 vJ ; 3J vI4

sdev ani ¼

^ ani 2I4 vJ Devðu5uÞ: J vI4

(55)

Accordingly, accounting for Eq. (51) and in contrast with the choice in Eq. (52), a deviatoric state of stress might be accompanied vol vol by a volume change, because it results siso ¼  sani , and hence I3 s 1 from Eqs. (48b) and (49b). On the other hand, a purely vol vol s 0) and fibers (sani s 0). Moreover, spherical state of stress implies a stress state (and thereby, a deformation) in both matrix (siso dev dev dev dev a null deviatoric stress state implies siso ¼  sani , but not necessarily siso ¼ 0 and sani ¼ 0. Therefore, a spherical state of stress does not induce a spherical state of stress in the matrix and in the fibers, being accompanied by a change of shape which is in full agreement with the anisotropic properties of the material. In other words, in the formulation of Eq. (54), change of volume and change in shape are fully coupled and not related to a spherical or a deviatoric state of stress. Based on these considerations, in the presence of an anisotropic behavior, a formulation based on the complete right Cauchy–Green deformation tensor C for the anisotropic strain-energy term shall be adopted. As a concluding remark, the effectiveness of the Penalty method for enforcing the incompressibility constraint in numerical applications depends on the value of the penalty parameter k since, only if k / þN, the volumetric isotropic stress contribution results vol vol / þ N for I3 s 1 (under the condition siso ¼ 0 for I3 ¼ 1, see Eqs. 48b and 49b). As a matter of fact, the fulfillment of equilibsiso rium conditions in standard applications requires a finite norm of the total Cauchy stress, i.e. ksk  þ N, and then a finite norm of

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vol the single terms in Eq. (51). In particular, this has to occur for the volumetric isotropic term, i.e. siso  þ N, which can be satisfied only for I3 / 1 if k / þN. Accordingly, in numerical applications, high values of k shall be employed. It is worth highlighting that the simultaneous fulfillment of equilibrium and incompressibility cannot be guaranteed in an exact way by the Penalty method.

The augmented lagrangian method In the framework of Lagrangian duality and in agreement with Eq. (7), the Augmented Lagrangian method consists in introducing e int in Eq. (6) based on a free-energy Jfe generalized with respect to the one in Eq. (12). Considering an enriched potential energy p the strain-energy density function Jse, free-energy Jfe is enriched with a Lagrange multiplier ^p through the internal constraint func^ , resulting tion J h ^ ð J Þ: e int ð4; ^pÞ ¼ Jfe ðC; ^pÞ ¼ Jse ðCÞ þ ^p J pint ð4Þ ¼ p h

(56)

^ is such that: Here, function J h ^ : Rþ 1R; J h

^ ð JÞ ¼ 05J ¼ 1: J h

(57)

^ ð J Þ ¼ 0 corresponds to enforce incompressibility. From Eq. (8), the internal dissipation density Accordingly, the condition J h Dint for incompressible materials with the potential internal energy density in Eq. (56) results:  _  ^ vJ C vJ vJse (58) 2 Dint ¼ S  2^p h : ; 2 vJ vC vC and, in order to a priori enforce the nondissipative properties of hyperelastic materials (see Eq. 13), the constitutive choice: S ¼ ^pJ

^ vJ vJse h 1 C þ2 ; vJ vC

(59)

replaces Eq. (14) for incompressible materials. e int in Eq. (56), the Lagrangian Sufficiency Theorem shows that there Employing the enriched internal energy functional density p ^ ð JÞ ¼ 0, and hence J ¼ 1, fulfilling the exists a value p of the Lagrange multiplier ^p such that the elastic equilibrium is satisfied for J h incompressibility constraint (see Eq. 57). Value p is referred to as the optimal Lagrange multiplier. Since the optimal value p of the Lagrange multiplier is an unknown, this can be obtained either from an equilibrium condition (generally determined from stationarity conditions of the total potential energy functional), or by means of an iterative procedure. In the latter case, introducing constant kAL > 0, the iterative algorithm presented in Table 1 can be implemented. The following section presents an applicative case where the Penalty method and the Augmented Lagrangian method are compared, highlighting advantages and drawbacks of the two approaches.

Application and final remarks on incompressibility Addressing an exemplary application, a tissue with nF ¼ 1 fiber family is addressed, subjected to the uniaxial traction along the fiber direction. Due to material symmetry, the deformation gradient tensor F takes the form: F ¼ lF ðu5eo Þ þ lt ðut1 5et1 þ ut2 5et2 Þ;

(60a)

where lF and lt represent along-the-fiber and perpendicular-to-fiber stretches, respectively, (eo, et 1, et 2) forms an orthonormal basis in the reference configuration, and (u, ut 1, ut 2) in the current configuration. Due to the simplicity of the case under consideration, it results (u, ut 1, ut 2) h (eo, et 1, et 2). As constitutive choices, the additive split in Eq. (15) of the strain-energy density function Jse and the positions in Eq. (25) are ^ dev and the term kJ ^ vol , where k is the employed. The isotropic contribution Jiso is made up by the sum of an isochoric term J iso iso vol ^ penalty parameter and Jiso is a penalty function. The choices: pffiffiffiffi 2 ^ dev ðI1 Þ ¼ c1 ðI1  3Þ; kJ ^ vol ðI3 Þ ¼ k I3  1 J (60b) iso iso 2 Pseudocode of an iterative algorithm for the solution of elastic equilibrium problems with incompressible materials modeled via the Augmented Lagrangian method

Table 1

^ h ðJ Þr > TOL 1. Initialize p^1 ¼ 0 and J1 such that rJ 1 ^ h ðJ Þr > TOL 2. LOOP n > 1 WHILE rJ n1   (a) Solve: R e int ðh; pn1 Þd U þ Pext ðhÞ 4n ¼ arg min p h˛C

Uo

(b) Compute Fn ¼ Grad (4n), Jn ¼ Det (Fn), Cn ¼ Fn TFn ^ h ðJn Þ (c) Compute p^n ¼ p^n1 þ kAL J END LOOP 3. Assign p ¼ p^n , 4 ¼ 4n and C ¼ Cn

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Biomechanics j Constitutive Modeling of Soft Tissues

are employed, with c1 > 0 being a stress-like parameter and where conditions in Eq. (48a) are satisfied. The anisotropic contribution related to fibers is modeled either following a formulation based on Eq. (52), ^ ðI4 Þ ¼ Jani ¼ J ani

  k1  exp k2 ðI4  1Þ2  1 ; 2k2

(60c)

^ ani ðI4 Þ ¼ Jani ¼ J

  k1  exp k2 ðI4  1Þ2  1 ; 2k2

(60d)

dev

or on Eq. (54)

where k1 > 0 is a stress-like parameter and k2 > 0 is a nondimensional constant. The afore-introduced choices fully define the formulation employed for enforcing the incompressibility constraint via the Penalty method. If imposing the incompressibility constraint via the Augmented Lagrangian method (see Eq. 56), the same constitutive choices are employed for the strain energy density function Jse and, in agreement with conditions in Eq. (57), the constraint ^ is chosen as function J h ^ ð JÞ ¼ J  1: J h

(60e)

Employing the choices in Eq. (60), the Cauchy stress tensor s is characterized by the following relationships: s : ðut1 5ut1 Þ ¼ s : ðut2 5ut2 Þ;

(61a)

s : ðu5ut1 Þ ¼ s : ðu5ut2 Þ ¼ s : ðut1 5ut2 Þ ¼ 0

(61b)

Accordingly, two stress quantities are relevant for the present application: the along-the-fiber sF ¼ s: (u 5 u) and the perpendicular-to-fiber st ¼ s: (ut 1 5 ut 1) Cauchy stresses. Hence, condition st ¼ 0 (i.e., null stress in the direction perpendicular to fibers) ensures equilibrium together with Eq. (61b), the latter being a priori satisfied. In fact, the only nontrivial stress component shall result sF because the tissue is subjected to a uniaxial traction along the fiber direction. The functional dependencies  Penalty method sF ðlF ; lt Þ sF ¼ s : ðu5uÞ ¼ (62) sF ðlF ; lt ; ^pÞ Augmented Lagrangian method and

 st ¼ s : ðut1 5ut1 Þ ¼

st ðlF ; lt Þ st ðlF ; lt ; ^pÞ

Penalty method Augmented Lagrangian method

(63)

on stretches lF and lt, and eventually on Lagrange multiplier ^p, are highlighted. A displacement-based virtual test is here considered, by choosing stretch lF as control variable. Therefore, by adopting the Penalty method, tissue mechanical response is obtained by determining stretch lt, which can be characterized by adopting: strategy P1: Find lt such that st ðlF ; lt Þ ¼ 0 Clearly, by following strategy P1, the incompressibility constraint is not automatically fulfilled and should be a posteriori verified. Alternatively, since Det(F) ¼ lF(lt)2, stretch lt can be defined from pffiffiffiffiffi strategy P2: lt ¼ lF such that Det ðFÞ ¼ 1; in order to a priori fullfill the incompressibility constraint. Nevertheless, it is not guaranteed that the equilibrium condition st ¼ 0 is satisfied and this should be a posteriori verified. The obtained constitutive relationships between lF and sF and between lF and st are shown in Fig. 4, as well as the resulting relationship between lF and lt, for both strategies P1 and P2 and for both constitutive choices in Eqs. (60c) and (60d). ^ dev in Eq. (60c), the incompressibility constraint is violated since It results that, by imposing st ¼ 0 (strategy P1) and adopting J ani  1/2 , leading to a nonphysical behavior where the inverse proportionality between lt and lF is lost. On the other hand, if lt s lF ^ dev , then it results sts 0, violating equilibrium. the incompressibility constraint is a priori enforced (strategy P2), together with J ani These drawbacks are related to the non-one-dimensionality of the stress associated with fibers (see Eq. 53a) and are due to the constitutive choice based on the formulation in Eq. (52). ^ ani in Eq. (60d) as constitutive choice for the anisotropic term and thanks to the oneOn the contrary, by employing J dimensionality of the resulting stress, results obtained by employing strategies P1 and P2 are practically coincident such that the incompressibility constraint and equilibrium can be retained as simultaneously respected, although not exactly. By following the Augmented Lagrangian method, the values of the stretch lt and of the optimal Lagrange multiplier p can be determined by applying: strategy A1: Algorithm in Table 1 where step 2a reduces to :   Find lt;n such that st lF ; lt;n ; pn1 ¼ 0:

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95

Fig. 4 The Penalty method: tissue with nF ¼ 1 fiber family subjected to uniaxial traction along the fiber direction (Eq. 60). Top: along-the-fiber stretch lF versus along-the-fiber sF and perpendicular-to-fiber st Cauchy stresses. Bottom: lF versus perpendicular-to-fiber stretch lt. Results are ^ dev in Eq. (60c) or J ^ obtained by a priori enforcing equilibrium (i.e., strategy P1) or incompressibility (i.e., strategy P2), and by adopting either J ani ani in Eq. (60d) for the anisotropic term. Parameters: k ¼ 2 GPa, c1 ¼ 10 kPa, k1 ¼ 50 kPa, k2 ¼ 50.

The exact fulfillment of incompressibility and equilibrium is thereby paid in terms of a solution strategy that requires the implementation of an iterative procedure, on the contrary with respect to the Penalty method. Alternatively, the optimal Lagrange multiplier can be determined from equilibrium conditions. Due to the simplicity of the addressed application, an analytical solution strategy (here used as benchmark) can be pursued for a priori fulfilling incompressi2 , this is bility and equilibrium with the Augmented Lagrangian method. Introducing Ct ¼ l t pffiffiffi strategy A2: lt ¼ lF such that Det ðFÞ ¼ 1 and  ^ 1 vJse vJ h p ¼ 2Ct such that st ¼ 0: vCt vJ The obtained constitutive relationships between lF and sF and between lF and st are shown in Fig. 5, as well as the resulting relationship between lF and lt, for both strategies A1 and A2, compared with results obtained by means of the Penalty method (solved through strategy P1), adopting the same value for the penalty parameter k. Fig. 5 shows also the evolution of the optimal Lagrange multiplier p with lF, confirming the effectiveness of the iterative procedure in strategy A1 compared with the exact analytical solution from strategy A2. Contrary to the Penalty method, the Augmented Lagrangian method allows to simultaneously and exactly satisfy both equilibrium and the incompressibility constraint employing low values of the penalty parameter k. This outcome allows to point out a significant advantage of the Augmented Lagrangian method with respect to the Penalty method in terms of the (fourth-order) tangent stiffness tensor C, that is C¼p

^ v2 J v2 Jse h þ vCvC vCvC

(65)

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Biomechanics j Constitutive Modeling of Soft Tissues

Fig. 5 The Augmented Lagrangian method: tissue with nF ¼ 1 fiber family subjected to uniaxial traction along the fiber direction (Eqs. 59, 60a, 60b, 60d, and 60e). Top: along-the-fiber stretch lF versus along-the-fiber sF and perpendicular-to-fiber st Cauchy stresses. Bottom: lF versus perpendicular-to-fiber stretch lt (continuous lines) and optimal Lagrange multiplier p (discontinuous lines). Results are obtained by computing lt and p via an iterative procedure (i.e., strategy A1) or via an analytical equilibrium relationship (i.e., strategy A2) and are compared with the Penalty method by a priori enforcing equilibrium (i.e., strategy P1). Parameters: k ¼ kAL ¼ 10 kPa, c1 ¼ 10 kPa, k1 ¼ 50 kPa, k2 ¼ 50.

where p ¼ 0 for the Penalty method. Indeed, adopting the Penalty method, the tangent stiffness tensor C may result illconditioned in the initial strain-range (i.e., close to the reference configuration) due to the high values of the penalty parameter k requested to fulfill the incompressibility condition. This is not the case when employing the Augmented Lagrangian method. In order to show this outcome, the condition number of the tangent stiffness tensor obtained from the along-the-fiber uniaxial traction of tissues with nF ¼ 1 fiber family is reported in Fig. 6 as function of the applied stretch. Addressing the initial strain range, the condition number results significantly higher for the Penalty method (strategy P1) than for the Augmented Lagrangian one (strategy A1). In order to allow for the most fair comparison, the penalty parameter k employed in the Penalty method has been reduced as far as possible, under the constraint that, in the addressed simple uniaxial traction case, the incompressibility relationship is satisfied with the same level of accuracy than the one obtained from the iterative strategy A1. It is worth highlighting that, due to the anisotropic nature of soft tissues, the advantage on the condition number obtained with the Augmented Lagrangian method is limited to the initial strain-range because, with increasing strain, fibers generate large entries in the stiffness matrix, thus leading to a dramatic enhancement of the condition number. The negative implications for numeric connected to this outcome are open problems under investigation that have to be solved on the computational side. A final remark is dedicated to numerical issues related to computational problems which involve spatial discretization techniques (e.g., the finite element method) coupled with incompressibility conditions. As a matter of fact, employing the Penalty method based on pure displacement formulations, the constraint condition J ¼ Det (F) ¼ 1 can only be fulfilled with a considerable stiffening of the bending modes of individual elements, known as volume locking. In order to avoid this drawback, mixed finite element formulations can be employed by introducing additional unknowns with respect to displacements in the energy functional density. For instance, the expression in Eq. (56) for the Augmented Lagrangian method represents a possible choice for a two-field mixed formulation, where the Lagrange multiplier ^p is introduced as independent variable.

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Fig. 6 Tissue with nF ¼ 1 fiber family subjected to uniaxial traction along the fiber direction (Eqs. 59, 60a, 60b, 60d, and 60e). Relationship between lF and the condition number cond(C) of the tangent stiffness tensor (in log scale) obtained via the Penalty method (i.e., strategy P1 with k ¼ 10 MPa) and the Augmented Lagrangian method (i.e., strategy A1 with k ¼ 10 kPa). Parameters: c1 ¼ 10 kPa, k1 ¼ 50 kPa, k2 ¼ 50.

Following a Hu–Washizu formulation, a more general and robust choice is based on a four-field functional which additionally includes a volume change variable Q and a Lagrange multiplier pQ as additional unknowns to the enriched free energy in Eq. (56), namely: ^ ðQÞ þ ^p J ^ ðQÞ þ p ðJ  QÞ; Jfe ðC; ^p; Q; pQ Þ ¼ Jse ðCÞ þ kJ h Q iso vol

(66)

where Jse(C) ¼ Jiso (Cdev) þ Jani(C, M) is the strain-energy density function (split in an isochoric isotropic term Jiso and an ^ vol is a penalty function for the volumetric response (see Eq. 48a), and J ^ enforces incompressibility in anisotropic term J ), J dev

ani

dev

h

iso

a Lagrangian framework (see Eq. 57). In this case, accounting for Eq. (13), the constitutive relationship results: S ¼ pQ JC1 þ 2

vJse ; vC

with pQ ¼ k

^ vol ^ vJ vJ iso þ ^p h ; vQ vQ

(67)

where, in numerical applications, multiplier ^p can be updated following the rule in Table 1 (step 2c) or it is determined from equilibrium relationships determined from stationarity conditions of the total potential energy functional.

State of the Art of Hyperelastic Constitutive Modeling A large number of constitutive models for soft tissues available in the literature follows the additive decomposition in Eq. (15) of the strain-energy Jse into a matrix JM and fiber JF term, respectively associated with an isotropic and an anisotropic response (see Eq. 25). Collagen fibers surely contribute to material anisotropy, while the noncollagenous constituents (mainly, elastin) are generally associated with a purely isotropic response. Nevertheless, evidence supporting an anisotropic contribution of elastin is recently available, being an open issue under investigation. For the description of the isotropic behavior, R.W. Ogden has proposed an important class of polyconvex strain-energy density functions that satisfy also the coercivity condition. In agreement with the formulation in Eq. (42a), the strain-energy density function for Ogden-type materials, in the form presented by P.G. Ciarlet with invariants from Eq. (43), is:

JM ðCÞ ¼

I X i¼1

a =2

ai I1 i

þ

V X

b =2

bv I2v

þ Gð J Þ;

(68)

v¼1

where ai > 0, bv  0, ai, bv  1, and G:]0, þN[1 R is a convex function such that G(J) / þN as J / 0þ. A first example of Ogdentype material is the compressible Mooney–Rivlin model, characterized by Eq. (68) with I ¼ V ¼ 1, a1 ¼ b1 ¼ 2 and G(J) ¼ g1J 2  g2 ln(J) where g1, g2 > 0. A second example is represented by the compressible Neo-Hookean model, characterized by Eq. (68) with I ¼ V ¼ 1, a1 ¼ 2 and b1 ¼ 0. In order to enforce incompressibility and employing a volumetric-deviatoric split in agreement with Eq. (47), the first two terms dev of the strain-energy function in Eq. (68) represent the deviatoric term Jiso (replacing I1 and I2 with I1 and I2 , see Eq. 43), while the vol ^ penalty term J is enriched by function G(J). iso

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Biomechanics j Constitutive Modeling of Soft Tissues

Addressing the anisotropic behavior, important classes of strain-energy density functions in agreement with the formulation in Eq. (42b) are based on the sets of invariants I 4 ¼ {I4(1), ., I4(nF)} and are based on an exponential JFe or a polynomial JFp expression:

JF ðC; MÞ ¼ JeF ðI 4 Þ ¼

nF X a¼1

ðaÞ

ðaÞ

vF

k1

ðaÞ 2k2

JF ðC; MÞ ¼ JpF ðI 4 Þ ¼




ðaÞ  ðaÞ  ðaÞ 2  exp k2 I4  1  k3 1 ;

ðaÞ ðaÞ k1  ðaÞ I ðaÞ 4 k2 a¼1

nF X

vF

ðaÞ k2

(69a)

ðaÞ

 1  k3

:

(69b)

Here, k1(a) > 0 is a stress-like parameter associated with fiber stiffness; k2(a) > 0 dimensionless parameter governing fiber nonlinearities; k3(a)  0 is a measure of fiber slackness. The Macaulay brackets hxi ¼ (x þ| x | )/2 (such that hxi ¼ 0 for x  0 and hxi ¼ x for x > 0) are employed for both physical and mathematical requirements. Firstly, since the anisotropic response is generally associated with the presence of crimped collagen fibers, the fibrous constituents are assumed to bear load only in traction mode (i.e., for I4(a) > 1 þ k3(a)). Moreover, the polyconvexity of the ansatz in Eqs. (69a) and (69b) can be obtained thanks to the use of the Macaulay brackets, because the latter are instrumental for the increasing monotonicity of the employed functions. With reference to the existing literature, the well-established rationale, known as structural approach and proposed by G.A. Holzapfel, T.C. Gasser and R.W. Ogden in 2000, is recovered by Eq. (69a) with k3(a) ¼ 0. In particular, the Holzapfel–Gasser–Ogden ðaÞ model has been originally formulated by using the invariant I4 based on Cdev, instead of I4(a), on the lines of the ansatz in Eq. (52). As previously shown in the discussion of incompressibility, the latter choice leads to major drawbacks. Since I4(a) is a polyconvex invariant, the polyconvexity of the strain-energy term JFe immediately follows. Moreover, addressing the polynomial law JFp, the restriction k2(a)  1 ensures its polyconvexity, and k2(a)  2, a continuous fiber tangent stiffness. The stress obtained from Eqs. (69a) and (69b) satisfies a priori the stress free condition in the reference configuration, while relationships between model parameters in Eq. (68) have to be introduced for the stress computed from the matrix term JM. For instance, referring to a compressible Mooney–Rivlin formulation, the condition 2a1 þ 4b1 þ 2g1  g2 ¼ 0 is obtained. Furthermore, in order to fulfill the nonessential normalization condition J(I) ¼ 0, constants are generally introduced in the existing literature. Soft tissues with a multidirectional fiber arrangement (such as arterial tissues, heart valves, and myocardial laminae) might be characterized by a high dispersion in fiber orientation which determines a complex multiaxial response with a high degree of anisotropy. Accordingly, a high number of fibers families (i.e., nF [ 1) should be employed in order to reproduce a continuous angular distribution. The latter shall be integrated during the analysis for obtaining the stress applied by fibers at each material point and by employing the specific deformation at hand. This method, known as angular integration (AI) model, leads to issues in terms of computational costs. Alternatively, the generalized structural tensor (GST) approach has been introduced by T.C. Gasser, R.W. Ogden, and G.A. Hol^ ðaÞ , is defined as: zapfel in 2006. For each fiber family, the GST, denoted by M   ^ ðaÞ ¼ kðaÞ I þ 1  3kðaÞ MðaÞ ; M (69c) F F where M(a) ¼ eo(a) 5 eo(a) is the classical structural tensor associated with the mean fiber direction eo(a), and k(Fa)˛ [0, 1/3] is a dispersion parameter, specific for each fiber family. In the limiting cases, k(Fa)¼ 0 corresponds to no dispersion (transverse isotropy) and k(Fa)¼ 1/3 to a three-dimensional isotropic distribution of fiber orientation. In principle, the upper limit of k(Fa) is 1/2 (corresponding to a planar isotropic distribution of fibers with a preferred direction normal to that plane), but it has been shown, for instance by G.A. Holzapfel and R.W. Ogden in 2010, that the range k(Fa)˛ (1/3, 1/2] can lead to unphysical results with nonmonotonic stress responses for increasing deformation. ðaÞ ^ ðaÞ Þ is used in constitutive relationships (for instance, Within the GST approach, the generalized strain invariant ^I4 ¼ TrðCM (a) with the form of Eqs. 69a and 69b) instead of the classical one I4 , allowing to obtain a non-one-dimensional response for each fiber family. Therefore, the multiaxial response associated with fiber dispersion can be captured maintaining a low value for the fiber family number nF and avoiding AI techniques. Nevertheless, attention shall be paid on how to account for the switch between fibers in tension and the ones in compression when using the GST. Indeed, the Macaulay Brackets are not effective anymore, because all fibers are treated as a whole. Suitable computational or theoretical approaches should be employed, as recently proposed, for example, by G.A. Holzapfel and R.W. Ogden in 2015 or by M. Latorre and F.J. Montáns in 2016. Moreover, it is worth highlighting that, for a given strain-energy expression of an individual fiber used for AI, a different strain energy shall be employed within the GST approach for obtaining the same results in terms of stress–strain relationships. For instance, even considering the same functional form, a different set of parameters shall be employed for obtaining a correspondence between the AI and the GST approaches, especially for high fiber dispersions. In other words, for a given (known) mechanical behavior of an individual fiber, an ad hoc calibration of model parameters shall be performed for obtaining the corresponding behavior of the dispersed fiber pattern through the GST approach, like in a phenomenological framework.

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Multiaxial effects can be also captured by employing strain invariants which control the deformation of area elements in the strain-energy formulation (see Eq. 44). For instance, introducing the set K3 ¼ {K3(1), ., K3 (nF)} of invariants, a multiaxial stress response (see Eq. 29b) can be obtained by employing the polynomial expression JFp2 defined as: p JF ðC; MÞ ¼ Jp2 F ðI 4 ; K3 Þ ¼ JF ðI 4 Þ þ

ðaÞ ðaÞ k4  ðaÞ K3 ðaÞ k5 a¼1

nF X

vF

ðaÞ k5

ðaÞ

 1  k6

;

(69d)

on the lines of the model by D. Balzani, P. Neff. J. Schröder and G.A. Holzapfel proposed in 2006. A number of different soft tissues have been modeled by following expressions of the form in Eqs. (68) and (69). Without the claim of being exhaustive, some examples are the models: of arteries by Gasser, Ogden, and Holzapfel proposed in 2006; of veins by Alastrué, Peña, Martínez, and Doblaré in 2008; of the intestine by Ciarletta, Dario, Tendick in 2009; of the esophagus by Yang, Bach, Zheng, et al. in 2006; of the cornea by Pandolfi and Manganiello in 2006; of the intervertebral disc by Eberlein, Holzapfel and Schulze-Bauer in 2001; of ligaments by Holzapfel and Stadler in 2006; of cartilage by Pierce, Trobin, Trattnig et al. in 2009; and of the myocardium by Holzpafel and Ogden in 2009. The biomechanical properties of soft tissues are well reproduced by these models and, in general, by strain-energy functions in Eqs. (68) and (69) or based on a similar ansatz. Nevertheless, in order to fit experimental data, the models shall be suitably calibrated. Model calibration consists in the identification of the values of parameters that govern these laws (i.e., ai, ai, ., for the isotropic part and k1(a), k2(a), ., for the anisotropic term). Collecting all model parameters in set § and following a displacement-based approach, a suitable approach for model calibration consists in solving the optimization problem:   ^ ; (70a) § ¼ argmin fobj ð§Þ ^ §˛A

where A is the set of admissible values of parameters and the objective function fobj can be defined as: vffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi u X n ^ 2 ksexp ðCi Þ  sðCi ; §Þk u1 ^ ¼t fobj ð§Þ
 : n i¼1 max ksexp ðCi Þk2

(70b)

i¼1;.;n

Here, Ci are the n values of the Cauchy–Green deformation tensor at which experimental data sexp are known, while sðCi ; ; ^ of parameters. To fully characterize tissue anisotropic ^ are model outcomes at C ¼ Ci, obtained by employing value § ; §Þ properties, reference has to be made to triaxial mechanical tests, where planar information only (i.e., uni/bi-axial and, possibly, shear tests) are suitable only in presence of special material symmetries. Clearly, other choices on the objective function are possible. It is worth highlighting that, although the fitting capabilities of constitutive laws in Eqs. (68) and (69) are generally good, the values of the parameters do not have a straightforward relationship with structural features, such as collagen fiber radius and crimp amplitude, intermolecular cross-link density, or molecular defects in the polypeptide sequence. Nevertheless, pathogenesis is associated with an alteration in the structural properties of soft tissues. Computational models developed with the aim of analyzing the response of pathological biological structures shall incorporate the structural differences of constitutive properties, as otherwise simulations might not be good predictors for the actual in vivo stress and strain state. To reach this aim, multiscale approaches for the constitutive modeling of soft tissues have been proposed. These approaches are based on the development of models for mechanisms occurring down at the nanoscale (i.e., collagen triple helix elongation mechanisms), through the mesoscale (i.e., cross-linked molecular assemblies), up to the microscale (i.e., crimped fibers). Therefore, multiscale constitutive models allocate macroscopic stress to different micro- and nanostructural mechanisms with a special insight on the structure–mechanics relationship. Accordingly, multiscale constitutive models are powerful alternatives to the phenomenological expressions introduced in Eqs. (68) and (69). Early ideas on microstructural relations for developing constitutive models have been proposed by Y. Lanir in the 1980s. Recent advancements in this framework have been obtained by explicitly considering nanoscale mechanisms in the model, such as in the formulations developed by F. Maceri, M. Marino and G. Vairo in 2010 or by M.S. Sacks, W. Zhang and S. Wognum in 2016. In contrast to the structural approach introduced by G.A. Holzapfel and co-workers, these nano-micro-macro formulations can be referred to as developed within a multiscale structural approach. The multiscale formulation based on the results proposed by M. Marino, P. Wriggers and co-workers in 2017 is described in what follows.

Multiscale structural approach The multiscale approach herein presented focuses on the fiber contribution JF to the strain-energy function, assumed to be representative of the behavior of collagen fibers. The strain-energy term JF is defined as:

JF ðC; MÞ ¼ Jm F ðI 4 Þ ¼

nF X a¼1

ðaÞ

vF

Z

1þhl4 1i ðaÞ

1

Z

1þðx1Þ

EF ðhÞdh dx;

(71)

1

where l4(a) ¼ (I4(a))1/2 is fiber stretch and EF ¼ EF(l4(a)) is the equivalent tangent modulus of crimped collagen fibers. The latter is function of the change of fiber along-the-chord length, measured via stretch l4(a), and it is obtained by means of a multiscale

100

Fig. 7

Biomechanics j Constitutive Modeling of Soft Tissues

Multiscale constitutive approach: representation of stretch measures and structurally motivated parameters.

description of internal deformation mechanisms. For the sake of compactness, superscript (a), denoting different fiber families, is omitted in what follows, although all the following introduced quantities generally vary family by family. Collagen fibers are assumed to have a circular cross-section of radius rF and area measure AF ¼ prF2. As schematically depicted in Fig. 7, the crimped structure of collagen fibers is taken into account by considering locally periodic fibers of along-the-chord period length LF and amplitude HF in the current configuration (resp., LF,o and HF,o in the reference configuration). Geometric sources of mechanical nonlinearities are accounted for by the functional dependence of fiber period and amplitude on fiber stretch l4, namely LF ¼ LF (l4) and HF ¼ HF (l4). In particular, it is convenient to introduce fiber quarter-of-period ‘F (l4) ¼ LF (l4)/4, where ‘F,o ¼ ‘F (1) represents the value in the reference configuration. Material sources of mechanical nonlinearities are accounted for by means of fibril tangent modulus Ef that depends on fibril stretch lf, in turn related to l4 by the interscale compatibility relationship Ff between microscale and mesoscale (here meso means between micro and nano):

Ff ðl4 ; HF Þ ¼

F l4 ‘2F;o þ HF dH dlf dl4 ¼ rffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi   ffi; dl4 l2F ‘2F;o þ H2F ‘2F;o þ H2F;o

(72a)

formulated by considering lf as coinciding with the centerline stretch of a fiber with piecewise linear shape. Derivative dHF/dl4 will be given in the following Eq. (75c). Fibril stretch lf is associated with molecular stretch lm which, in turn, is function of entropy-related lms and energy-related lmh molecular stretches (see Fig. 7). These functional dependences are taken into account via the interscale compatibility relationships Ffm (from meso-to nano-scale), Fms, and Fmh (from nanoscale to atomistic scale). On the basis of simple equilibrium conditions formulated assuming mechanisms as in series, the interscale compatibility relationships are obtained from the tangent modulus of collagen fibrils Ef, of collagen molecules Em, and of entropy-related Ems and energy-related Emh mechanisms. Accordingly, it results    dl E ls ; lh (72b) Ffm lsm ; lhm ¼ m ¼ f  ms mh ; dlf Em lm ; lm 

  dlsm Em lsm ; lhm  s ; ¼ dlm Esm lm

(72c)

  dlhm Em lsm ; lhm ¼  h : dlm Ehm lm

(72d)

  Em lsm ; lhm Lc kc ‘m;o  ; Lc kc ‘m;o þ Am Em lsm ; lhm

(73a)



Fms lsm ; lhm ¼ 



Fmh lsm ; lhm ¼ where  Ef lsm ; lhm ¼

     Es lsm Ehm lhm Em lsm ; lhm ¼ m  s  ; Esm lm þ Ehm lhm

(73b)

   kB T‘m;o ‘3c Esm lsm ¼ þ 1 ;   ‘p ‘c Am 2 ‘c  ‘m;o ls 3 m

(73c)

and

Biomechanics j Constitutive Modeling of Soft Tissues

Ehm



lhm



‘m;o ¼ ‘c

) ^ E ^


   þ Eo ; 1 þ exp  h ‘m;o lhm  1 ‘c  εho

101

(

(73d)

with kB being the Boltzmann constant and T being the absolute temperature. Moreover, referring to the entropic behavior of collagen molecules (Eq. 73c), ‘p is the persistence length, ‘c is the contour length, ‘m,o is the end-to-end length in the reference configuration (resulting ‘m,o ¼ ‘c  ‘ks, with ‘ks being the length of molecular kinks), and Am is the cross-sectional area. Furthermore, ^ are, respectively, the low-strain and high-strain collagen tangent moduli, εoh is ^o and E addressing the energetic regime (Eq. 73d), E the uncoiling strain, and h is the uncoiling resistance. Finally, with reference to intermolecular sliding affecting fibril mechanics (Eq. 73a), Lc denotes the (mole fraction) density of intermolecular covalent cross-links, which are modeled with a linear elastic behavior with stiffness kc. Eqs. (72b) and (73a) (resp., Eq. 72c, 72d, and 73b) derive from the in series elasticity of molecular elongation and intermolecular sliding (resp., entropic and energetic molecular elongation mechanisms). Eqs. (73c) and (73d) are, respectively, based on theoretical results which recover the worm-like chain model for the description of entropic elasticity, and on atomistic computations that elucidate energetic mechanisms. The equivalent tangent modulus EF (l4) of crimped collagen fibers to be employed for the macroscale description of collagen mechanics (i.e., in Eq. 71) is defined as EF (l4) ¼ CF (l4)/l4, with CF being the along-the-chord tangent modulus of an Euler–Bernoulli curvilinear beam whose geometry corresponds to the one of fibers (i.e., quarter-of-period ‘F and amplitude HF) and whose material tangent modulus corresponds to the one of fibrils (i.e., Ef). Addressing a piecewise-linear beam shape, the incremental application of the principle of Virtual Works gives (dependences omitted):   1 ‘2F þ HF2 4H2  ffi ‘F þ 2 F ‘2F þ HF2 CF ¼ Ef qffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi ; (74) 3rF ‘F ‘2 þ H 2 F;o

F;o

where it is worth recalling that Ef ¼ Ef(lh , lm ) (see Eq. 73), ‘F ¼ ‘F (l4) and HF ¼ HF(l4). The functional dependences lms ¼ lms(l4) and lmh ¼ lmh(l4) are obtained from Eq. (72) and the application of the chain rule, via the interscale compatibility relationships:   dlsm ¼ Fms lsm ; lhm Ffm lsm ; lhm Ff ðl4 ; HF Þ; (75a) dl4 s

h

  dlhm ¼ Fmh lsm ; lhm Ffm lsm ; lhm Ff ðl4 ; HF Þ; dl4

(75b)

with lms(1) ¼ lmh(1) ¼ 1. Moreover, function ‘F (l4) is chosen as ‘F (l4) ¼ l4‘F,o, while HF (l4) is obtained from the solution of the geometry evolution equation:

   ‘F HF 4 ‘2F þ HF2  3rF2 dHF  ; ¼  2 2 (75c) dl4 l4 4HF ‘F þ HF2 þ 3‘2F rF2 with HF (1) ¼ HF,o, which gives the evolution of crimp amplitude upon fiber deformation and it is derived from a second application of the principle of Virtual Works on a curvilinear beam model. It is worth noting that the solution of Eq. (75) on the basis of the family-specific fiber stretch l4(a) gives the family-dependent internal mechanisms, that is:   ðaÞ   ðaÞ  sðaÞ hðaÞ ðaÞ ðaÞ (76) lm ¼ lsm l4 ; lm ¼ lhm l4 ; HF ¼ HF l4 :

Comparison with experimental data The effectiveness of available constitutive modeling approaches in fitting experimental data is clearly highly depending case by case (e.g., on the specific tissue under considerations, on the addressed mechanical tests, on the strain range of interest). Therefore, general conclusive statements on fitting capabilities cannot be traced. On the other hand, an exhaustive analysis on specific case studies is beyond the scope of present work which aims to be as general as possible. Following these considerations, an exemplary comparison between modeling approaches and experimental data is presented in one of the simplest available case study, that is one of the tendons. The latter are unidirectional tissues, with collagen fibers mainly aligned along the loading direction. Therefore, in this case, a single collagen fiber family can be considered (i.e., nF ¼ 1) and data on the uniaxial traction applied along-the-fiber direction are representative for the mechanics of these tissues. Tendons are here chosen because their simple organization allows to obtain data on the evolution of structural features with strain, such as collagen fiber crimp straightening upon tissue stretching. This allows for a full validation of the multiscale structural approach, although this modeling rationale has been applied in the literature also to tissues with a multidirectional fiber arrangement (e.g., arterial tissues). In the case under consideration, tissue deformation gradient tensor F results as in Eq. (60a). The strain-energy density function Jse is additively decomposed, as in Eq. (15), into a matrix JM and fiber JF term, respectively, associated with an isotropic response and an anisotropic response (see Eq. 25). Moreover, for dealing with the incompressibility constraint, an Augmented-Lagrangian ^ as in Eq. (60e) together with an analytical solution strategy to a priori fulfill approach as in Eq. (56) is employed, considering J h 2 both the kinematic condition lF ¼ l t and the equilibrium condition of null stress in the perpendicular-to-fiber direction (i.e.,

102

Biomechanics j Constitutive Modeling of Soft Tissues Table 2

Values of model parameters for fitting the data on the uniaxial traction of rat tail tendons in Fig. 8 for e p the exponential-based JF (see Eq. 69a) and polynomial-based JF (see Eq. 69b) terms

Model p F e F

J J

Table 3

k1 (MPa)

k2

k3

215 140

2.16 4.20

0.045 0.048

Values of structural features in rat tail tendons introduced in the multiscale constitutive m description for the term JF in Eq. (71) and employed to compare model results with experimental data in Fig. 8

Parameter

Value

Definition

rF HF,o LF,o Lc kc ‘ks ‘c ‘p Am E^o E^

1.63 mm 13.04 mm 240 mm 1400 mmol mol 1 5 nN mm 1 14 nm 287 nm 14.5 nm 1.41 nm2 1 GPa 80 GPa 22.5 0.1 310 K

Collagen fiber radius Collagen fiber reference amplitude Collagen fiber reference period Inter-molecular cross-links density Inter-molecular cross-links stiffness Collagen molecular kinks length Collagen contour length Collagen persistence length Collagen cross-sectional area Low-strain collagen tangent modulus High-strain collagen tangent modulus Collagen uncoiling resistance Collagen uncoiling strain Body temperature

h εo h T

Employed values belong to physiological ranges, as discussed in Marino M and Wriggers P (2017) Finite strain response of crimped fibers under uniaxial traction: an analytical approach applied to collagen. J. Mech. Phys. Solids 98, pp. 429–453.

strategy A2). Addressing the matrix term JM, a compressible Mooney–Rivlin formulation, that is Eq. (68) with I ¼ V ¼ 1 and a1 ¼ b1 ¼ 2, is chosen. Three possibilities are employed for the fiber strain-energy term JF: (i) the exponential-based law JFe in Eq. (69a); (ii) the polynomial-based law JFp in Eq. (69b); (iii) the multiscale structural approach JFm in Eq. (71). Results are obtained employing values of parameters in Tables 2 and 3 with nF ¼ 1, VM ¼ 0.43, VF ¼ 0.57, a1 ¼ b1 ¼ g1 ¼ g2/8 ¼ 10 kPa. The obtained relationships between along-the-fiber nominal strain and stress are shown in Fig. 8, and compared with experimental data obtained by K.A. Hansen, J.A. Weiss, and J.K. Barton in 2002. In addition, the multiscale approach gives also information on the dependence of fiber crimp with stretch, that is function HF (l4) which is a measure of the average amplitude of collagen fibers within the tissue. Function HF (l4) is compared with data on the amount of extinguished crimps at different strain levels, obtained by means of nondestructive imaging analyses conducted in conjunction with mechanical tests. The effectiveness of the employed constitutive approaches clearly arises, showing also that the multiscale approach is able to give a special insight on the relationship between structure and mechanics.

Inelastic Behavior The main inelastic effects that can be observed in soft tissues and that are specifically addressed in this work are:

• • • • •

viscoelasticity, that is a time dependence of the stress–strain relationship; active behavior, that is the occurrence of a stress field originated by self-contracting elements, also in the absence of external loads; plasticity, that is the occurrence of permanent strains, not associated with a degradation of the mechanical properties of tissue constituents, at the removal of the deforming cause; damage, that is a degradation of tissue mechanical properties (e.g., stiffness, strength); growth and remodeling, that is an alteration of tissue mass and microstructure driven by mechanobiological mechanisms.

A general framework for the modeling of the afore-introduced inelastic effects is based on the choice of a multiplicative decomposition of the deformation gradient F in an elastic Fe and an inelastic Fi term, holding (see Fig. 9): F ¼ Fe Fi :

(77)

Biomechanics j Constitutive Modeling of Soft Tissues

103

Fig. 8 Comparison between constitutive models and experimental data on the uniaxial traction of rat tail tendons. Modeling rationale in agreement with Eqs. (15) and (25): matrix term JM in Eq. (68) with I ¼ V ¼ 1 and a1 ¼ b1 ¼ 2; fiber term JF via the exponential-based law JF e in Eq. (69a), the polynomial-based law JF p in Eq. (69b), and the multiscale structural approach JF m in Eq. (71). Incompressibility constraint treated via an ^ as in Eq. (60e)) and an analytical solution strategy. Results: along-the-fiber first Piola–Kirchhoff Augmented-Lagrangian approach (Eq. (56) with J h stress P: (eo 5 eo) (top) and percentage amount of straight fibers, i.e., extinguished crimp function EC(l4) ¼ 1  HF (l4)/HF,o (only multiscale pffiffiffiffi approach, bottom), versus along-the-fiber nominal strain l4  1, with l4 ¼ I4 . Experimental data by Hansen KA, Weiss JA, Barton JK (2002) Recruitment of tendon crimp with applied tensile strain. J. Biomech. Eng. 124, pp. 72–77. Parameters: nF ¼ 1, VM ¼ 0.43, VF ¼ 0.57, a1 ¼ b1 ¼ g1 ¼ g2/8 ¼ 10 kPa (Tables 2 and 3).

The inelastic term Fi is associated with the onset and evolution of inelastic mechanisms and with an intermediate configuration between the reference one Uo and the current one U. In principle, the intermediate configuration is incompatible from the kinematic point of view and the elastic deformation Fe serves also to re-establish compatibility, together with describing the deformation due to applied loads (see Fig. 9). In order to obey material objectivity requirements, the elastic part of the right Cauchy–Green deformation tensor is introduced as Ce ¼ Fe TFe where the functional dependence Ce ¼ Ce(C, Fi) is conveniently highlighted since:  1 (78) Ce C; Fi Þ ¼ FT i CFi : In agreement with Eq. (7), inelastic mechanisms are described by means of a set V of internal variables, whose time dependence determines the onset and evolution of inelastic mechanisms (see Fig. 9). The inelastic contribution Fi to the deformation tensor clearly belongs to the set of internal variables. Moreover, for the sake of generality, tensor A and scalar a are here introduced to enrich the set of internal variables for describing additional internal mechanisms associated with an inelastic response, resulting V¼ {Fi, A, a}. Accordingly, free-energy Jfe can be defined as:

Jfe ¼ Jfe ðC; Fi ; A; aÞ: and the internal dissipation Dint in Eq. (8) results:  _  vjfe vjfe vjfe vjfe C _ Dint ¼ S  2ro :  ro : A_  ro a: : F_ i  ro 2 vC vFi vA va

(79)

(80a)

104

Biomechanics j Constitutive Modeling of Soft Tissues

Fig. 9 Modeling of the inelastic response of biological structures. Top: general modeling rationale represented by the multiplicative decomposition of the deformation gradient F in elastic Fe and inelastic Fi terms, and the introduction of internal variables A and a. Bottom: schematic representation of physical mechanisms originating inelastic effects in soft tissues.

On the basis of Eq. (80a), stress quantities Qi, Hi, and qi are introduced as dual static quantities to Fi, A, and a, respectively. Introducing the constitutive relationships: S ¼ 2ro

vjfe ; vC

Qi ¼ ro

vjfe ; vFi

Hi ¼ ro

vjfe ; vA

qi ¼ ro

vjfe ; va

(80b)

the Clausius–Duhem inequality reads as: Dint ¼ Qi : F_ i þ Hi : A_ þ qi a_  0;

_ a: _ c admissible F_ i ; A;

(80c)

_ and a_ are given by evolution equations for the internal variables. The definition of these evolution equations Admissible F_ i , A, closes the constitutive formulation. If a potential function F of the generalized static quantities S ¼ {Qi, Hi, qi} exists, then evolution equations can be derived from the associate normality rule: vF ; F_ i ¼ z_ vQi

vF A_ ¼ z_ ; vHi

vF a_ ¼ z_ ; vqi

(81a)

where z_ is an indeterminate multiplier and F (S) is a convex function that may also depend on internal variables. The formulation is completed by the classical Kuhn–Tucker conditions: z_  0;

F  0;

_  0; zF

(81b)

ensuring that internal mechanisms evolve on the potential surface F ¼ 0 in the space of generalized static quantities S. It is immediate to show that the Clausius–Duhem inequality in Eq. (80c) is a priori satisfied by the formulation in Eq. (81). Ad hoc definitions, not motivated by thermodynamic considerations, of the evolution equations are also possible and have been employed in the existing literature. For instance, in a strain-driven framework, the explicit definition of: F_ i ¼ F_ i ðC; Fi ; A; aÞ;

A_ ¼ A_ ðC; Fi ; A; aÞ;

a_ ¼ a_ ðC; Fi ; A; aÞ;

(82)

can be introduced, accompanied by the a posteriori verification of the dissipation inequality in Eq. (80c) in order to be thermodynamically consistent. Evidence and modeling approaches related to the afore-introduced inelastic mechanisms are presented in what follows. Whenever necessary and referring to the ansatz in Eq. (25), only nF ¼ 1 fiber family is introduced for the sake of compactness, resulting M¼ {M}, without loss of generality.

Biomechanics j Constitutive Modeling of Soft Tissues

105

Viscoelasticity The mechanical response of a viscoelastic material is characterized by a hysteretic and strain-rate behavior (see Fig. 2). Moreover, viscoelastic properties are associated with stress relaxation (i.e., a time step-wise applied strain results in a decreasing stress) and creep (i.e., a time step-wise applied stress results in an increasing strain). A main source of viscoelasticity in soft tissues is related to the high water content and it is due to the flow of free water in the ground matrix, driven by strain. This mechanism leads to a significant time-dependent behavior. A physically motivated approach for modeling viscoelasticity due to water movement is given by the rationale of poromechanics. This approach couples equations for the elastic response of the solid matrix with Navier–Stokes equations for the viscous fluid, as well as with Darcy’s law for the flow of fluid through the porous matrix. A detailed description of poromechanics modeling is beyond the scope of present chapter, and reference is made to the seminal papers and books by S. Cowin. A second source of viscoelasticity in soft tissues can be associated with an intrinsic time dependence of the behavior of tissue constituents. For instance, the deformation rate affects the unfolding pathways of the triple-helix structure of tropocollagen molecules. The presence of water directly bound to collagen molecules contributes to the latter evidence (see Fig. 9). A general framework for the description of tissue viscoelasticity is to employ a generalized Maxwell model where the presence of water is taken into account in a phenomenological way. The generalized Maxwell model is characterized by an additive decomposition of the free energy in two contributions: Jv eq responsible for the thermodynamic equilibrium state at t / þN, and Jv eq responsible for the thermodynamic nonequilibrium state of the material which may be seen as a “dissipative” potential (i.e., the behavior of relaxation and/or creep). The rheological description of a generalized Maxwell model consists in the parallel connection between a spring (associated with term Jv eq and the entire kinematics, described by C) and a Maxwell element made up by the series of a dashpot (associated with the inelastic part of the deformation tensor Fi) and a spring (associated with term Jv neq and the elastic part of the deformation tensor Fe). In agreement with these observations, it results Jv eq ¼ Jv eq(C) and Jv neq ¼ Jv neq(Ce). Therefore, the generalized Maxwell model motivates the ansatz:  neq Jfe ¼ Jfe C; Fi Þ ¼ Jeq (83a) v ðCÞ þ Jv ðCe ðC; Fi ÞÞ resulting in (see Eqs. 7 and 80b): neq S ¼ Seq v þ Sv ;

vJneq Qi ¼  v ; vFi

Hi ¼ 0;

qi ¼ 0;

(83b)

where (see Eq. 78) Seq v ¼2

vJeq v ; vC

1 Sneq v ¼ 2Fi

vJneq v FT : vCe i

(83c)

The constitutive framework is completed by the definition of Jv eq and Jv neq and of evolution equations for Fi, under the constraint prescribed by the dissipation inequality in Eq. (80c). It is worth pointing out that, since the stress tensor S in Eq. (83b) shall reach the thermodynamic equilibrium for t / þN, the constitutive model shall be defined in order to obtain Sv neq |t / þ N ¼ 0.

Active Response A specific class of cells (e.g., skeletal muscle cells, myocardial cells, smooth muscle cells) are able to contract and relax in absence of applied loads (see Fig. 9). Cell contractile features are due to internal structures where calcium ions activate the interaction between actin and myosin filaments. These mechanisms induce an inelastic strain field in the tissue, driven by biochemical and biophysical actions, which significantly affect the distribution of stresses and strains in biological structures. Tissues characterized by such mechanisms are referred to as endowed with an active response. Active response may be also associated with electroactive properties where mechanical forces and electric fields are coupled. These properties endow tissues with the ability to spontaneously deform upon the application of an electric field. The latter mechanism is of main relevance for the mechanical response, for instance, of the cardiac muscle and of the intestine. The mechanics of soft tissues with an active response can be described by considering the strain-energy function Jse, related to the purely elastic deformation (i.e., Jse ¼ Jse(Ce)), and function Ja describing the contractile nature of the active part and depending on the overall deformation (i.e., Ja ¼ Ja(C)). Function Jse can be defined as in a hyperelastic framework and function Ja accounts for biochemical and/or electrochemical fields. Assuming an additive decomposition of the free energy, the ansatz of Eq. (83) can be employed, replacing Jv eq with Ja and Jv neq with Jse. The formulation is completed by the definition of functions Jse and Ja, of evolution equations for Fi, under the constraint prescribed by the dissipation inequality in Eq. (80c). If other fields are introduced, such as the electrical field for cardiac or gastrointestinal tissues, further balance laws shall be accounted for.

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Biomechanics j Constitutive Modeling of Soft Tissues

Plasticity The phenomenological response of a material undergoing plastic mechanisms is well elucidated by the analysis of the stress–strain relationship obtained from a uniaxial traction test. Material response can be described as hyperelastic below a certain stress value, denoted as elastic limit. Beyond the elastic limit, a plastic (equivalently, yielding) regime is attained. With reference to Fig. 2 and addressing a positive stretching beyond the elastic limit, the slope of the stress–strain relationship significantly reduces, being equal to zero in the ideal plastic case. Hence, a change in the shape of the specimen at (almost) constant stress is obtained. Moreover, residual strains occur upon loading removal. In mechanics, plasticity is classically associated with the slip of dislocations in the crystalline structure of metals. Nevertheless, plastic-like mechanisms, leading to a nonreversible change of shape of biological structures, occur also in soft tissues. For instance, the alteration of the structure of the noncollagenous matrix may lead to the permanent slip of collagen fibers relative to each other, leading to a change of the reference state. Moreover, collagen fibers may experience nonelastic straightening upon overstretching, not returning to the original unstrained wave pattern when residual strain appears in the tissue. This might be the sign of a permanent alteration in the multiscale arrangement of collagenous structures within fibers. As schematically depicted in Fig. 9, atomistic computations based on molecular dynamics simulations of collagenous assemblies show that two inelastic plasticlike mechanisms may intervene: slip pulse and interstrand delamination. Slip pulse is a permanent intermolecular sliding where weak bonds between adjacent molecules break and continuously reform (among different residues) during sliding. Interstrand delamination is the loss of the structural integrity of collagen molecules due to the breaking of weak bonds between the polypeptide strands within a single collagen triple helix. In this case, the permanent sliding of one polypeptide strand, with respect to the other two, occurs. Furthermore, the rearrangement and alterations of tissue constituents due to the evolution of internal plastic-like mechanisms may lead to a stronger interaction between collagen molecules in fibers and between collagen fibers with the elastin network. This phenomenon might be the source of the strengthening of tissues by plastic deformation, that is of hardening effects in soft tissues. A plastic behavior can be described by introducing the additive decomposition of the free energy in a strain-energy function Jse, related to the purely elastic deformation (i.e., Jse ¼ Jse(Ce)), and the term Jp, depending on the hardening variable a and being a measure of tissue strengthening with plasticity. Accordingly, a possible ansatz is:  Jfe ¼ Jfe ðC; Fi ; aÞ ¼ Jse ðCe C; Fi ÞÞ þ Jp ðaÞ; (84a) resulting in (see Eqs. 7 and 80b): S ¼ 2F1 i

vJse T F ; vCe i

vJse Qi ¼  ; vFi

Hi ¼ 0;

vJp : qi ¼  va

(84b)

The constitutive framework is completed by the definition of material response in the elastic regime (i.e., function Jse), of hardening mechanisms (i.e., function Jp) and of evolution equations for Fi and a, under the constraint prescribed by the dissipation inequality in Eq. (80c). In this context, the evolution equation for Fi is generally referred to as the flow rule. If the formulation of Eq. (81) is employed, the surface identified by F ¼ 0 in the space of generalized stresses is denoted as the yield surface. The Kuhn–Tucker conditions ensure that the stress state never lies outside the yield surface. As a matter of fact, when plasticity evolves (i.e., z_ > 0), the condition F ¼ 0 holds from Eq. (81b) or, in other words, the stress state remains on the yield surface. Nevertheless, the shape and size of the surface may change as the plastic deformation evolves because F may depend on the internal variable a, leading to the description of hardening effects.

Damage Damage in materials initiates and evolves at strain/stress values beyond a certain limit which corresponds to the elastic limit. Below the elastic limit, the material is within the elastic regime and it is generally described as hyperelastic. The onset and evolution of damage is associated with a progressive reduction of material stiffness. In a strain-driven test, stress may also decrease, and this phenomenon is known as softening. If damage evolution in the material is fast and the stress drops, failure is described as brittle and tissue final failure is reached (see Fig. 2). Damage can be associated with the process of initiation and growth of cracks occurring at a length scale comparable to the one of the structure under consideration. This mechanism leads to the need of describing damage as the evolution of a discontinuity in the integrity of the structure. This modeling approach is known as fracture mechanics. Different theoretical and computational approaches exist for dealing with the discontinuous character of fracture mechanics (e.g., phase field approaches, the extended finite element method, cohesive zone model) but the description of these is beyond the scope of present work. On the other hand, the phenomenology of damage can be associated with an alteration of material features occurring at a length scale significantly smaller than the one of the structure under consideration (e.g., like nucleation and growth of voids, nanoscale and/or microscale defects of constituents). In this framework, stress measures obtained from constitutive models might represent a homogenized quantity which implicitly accounts for the deterioration of the material. This modeling approach is known as continuum damage mechanics.

Biomechanics j Constitutive Modeling of Soft Tissues

107

In this context, possible sources of damage in soft tissues are the rupture of covalent bonds within polypeptide strands of collagen molecules or the degradation of covalent cross-links between collagen molecules (see Fig. 9). For instance, the amount of intermolecular cross-links may decrease due to biochemical reactions mediated by enzymes. Collagen-related damage mechanisms determine a degradation of the mechanical properties of collagen fibrils, in turn associated with degraded mechanical properties at tissue level. In the framework of continuum damage mechanics, the mechanical response of soft tissues undergoing damage can be modeled by means of the strain-energy Jse, which describes the hyperelastic material response in the elastic regime, modulated by the continuous internal variable a which describes damage. Variable a is valued in [0, 1], being equal to 0 if the soft tissue is sound (i.e., nondamaged and behaving as purely elastic) and equal to 1 if it is fully damaged (i.e., not able to carry load). Values of a ˛ (0, 1) describe a partially damaged situation. It is worth highlighting that the inelastic part of deformation Fi, and hence the multiplicative decomposition in Eq. (77), is not needed for modeling damage effects associated with softening. Damage is described by introducing the damage function g, valued in [0, 1], and such that: gð0Þ ¼ 1;

gð1Þ ¼ 0;

g0ðaÞ ¼

vg  0: va

(85)

An ansatz for the free energy for describing damage in soft tissues is:

Jfe ¼ Jfe ðC; aÞ ¼ gðaÞJse ðCÞ;

(86a)

vJse ; vC

(86b)

resulting in (see Eqs. 7 and 80b): S ¼ 2g ðaÞ

Qi ¼ 0;

Hi ¼ 0;

qi ¼ Jse ðCÞ:

From Eqs. (85) and (86b), it is immediate to show that stress S corresponds to the one obtained in a hyperelastic framework for a ¼ 0 (i.e., when damage is not activated). Since function g(a) is monotonically decreasing (see Eq. 85), the onset and evolution of damage (i.e., when a increases) is associated to a stress decrease, up to the limit case S|a / 1 ¼ 0 (see Eq. 85), recovering a softening response. In order to move toward a structurally motivated framework, damage can be associated to a single constituent and introduced in the definition of the strain invariants. For instance, addressing collagen damage, let eI4 ¼ eI4 ðC; aÞ be a generalized fourth invariant of deformation, defined as: eI4 ðC; aÞ ¼ g ðaÞ½I4 ðCÞ  1 þ 1 ¼ g ðaÞ½TrðCMÞ  1 þ 1;

(87a)

where g(a) respects the conditions in Eq. (85). In this framework, a structurally motivated formulation of the free energy of soft tissues undergoing damage is:

Jfe ¼ Jfe ðC; aÞ ¼ VM JM ðCÞ þ VF JF ðeI4 ðC; aÞÞ;

(87b)

resulting in (see Eqs. 7 and 80b): SF ðeI4 ; aÞM; S ¼ VM SM þ VF e

qi ¼ VF g’ðaÞ½I4 ðCÞ  1 SF ðeI4 Þ;

(87c)

with Qi ¼ 0, Hi ¼ 0 and SM ¼ 2

vJM ; vC

e SF ðeI4 ; aÞ ¼ g ðaÞSF ðeI4 Þ;

SF ðeI4 Þ ¼

vJF : veI4

(87d)

Model response when damage is not activated (i.e., a ¼ 0) is purely elastic. In fact, it results eI4 ðC; 0Þ ¼ I4 ðCÞ (see Eqs. 85 and 87a) and the fiber stress tensor corresponds to the one that would be obtained in a hyperelastic framework, that is equal to VFSF (I4)M. Moreover, accounting for Eq. (85), fiber stress e SF in Eq. (87c) results inversely proportional to damage, up to the limit case e SF ja/1 ¼ 0 (see Eq. 85 and 87d) where only the matrix contribution SM contributes to the overall second Piola–Kirchhoff stress S. The definitions of the strain-energy function Jse (as in a hyperelastic framework), of damage function g(a), and of evolution equations for a complete the damage formulation, under the constraint enforced by the dissipation inequality in Eq. (80c).

Growth and Remodeling Living tissues continually undergo a turnover of constituents via a delicate balance between the production of new material and the removal of the old one (see Fig. 9). A change in the rates of turnover of constituents can lead to atrophy (i.e., decrease in the overall mass of material, also known as resorption) or hypertrophy (i.e., increase in the overall mass of material, also known as growth). In the literature, growth and resorption are collected under the term growth. A distinction is made between growth on the surface (i.e., appositional growth) and the one within the volume (i.e., interstitial growth) of tissues. Present work addresses volumetric growth, since the most relevant for soft tissues. Clearly, mass growth and volume variations are related through the mass density. On the other hand, remodeling describes the changes in the histology and biochemistry of tissues associated with an alteration of material mechanical properties (e.g., stiffness, strength, and anisotropy). It may be related to coarse changes in tissue

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composition (e.g., excess of fibrous material known as fibrosis), as well as to fine alterations of tissue microstructure (e.g., realignment, thickening, and crimp alterations of collagen fibers) and of biochemical properties (e.g., in density of intermolecular crosslinks). Remodeling is often used jointly with growth, although being two independent mechanisms. However, remodeling can be a manifestation of growth at spatial scales lower than the tissue one (e.g., in fibrosis). Moreover, deposed/resorbed mass associated with interstitial growth may be initially incompatible from a kinematic point of view with the preexisting material, and remodeling could smooth this incompatibility. Accordingly, an intimate connection between growth and remodeling arises. Since cells respond to mechanical stimuli, applied loads have a strong influence on growth and remodeling. For instance, cells align as a consequence of the applied stresses and/or strains, hence controlling the direction of the secreted collagen fibers which realign in agreement with the current stress and/or strain fields. Growth and remodeling in soft tissues can be modeled by following two different rationales: a constrained mixture model and a kinematic approach. The theory of constrained mixtures is based on the assumption that tissue constituents exhibit distinct reference configurations and rates of turnover. At this standpoint, full mixture equations for mass balance are used in combination with classical equations for linear momentum balance written in terms of rule-of-mixture relations for the stress response. The evolution of the reference configurations and rates of turnover of tissue constituents are defined in order to maintain a preferred (homeostatic) biomechanical state, incorporating biologically driven mass density productions and survival functions within constitutive relations. A detailed description of the constrained mixture approach is beyond the scope of present work. To this aim, reference is made to the seminal works by J.D. Humphrey and C. Cyron. On the contrary, the kinematic approach identifies a unique reference configuration. The overall deformation tensor is decomposed into elastic and growth components along the lines of Eq. (77). In the framework of open system thermodynamics, the kinematic approach focuses on volumetric growth, where the intermediate configuration reached by tensor Fi physically represents the (possibly incompatible) grown configuration without any elastic deformation. In order to describe growth, density ro in the reference configuration at the grown state (i.e., at current time) is introduced as an internal variable. Clearly, the balance of mass for open systems shall be taken into account, considering the rate of change of density r_ o in the reference configuration, a possible influx/outflux of matter R (e.g., cell migration) and a volumetric mass source R (e.g., cell proliferation, apoptosis). Therefore, denoting by Div (•) the divergence operator, it holds: Z t ro ¼ ro ðt Þ ¼ ro þ (88) r_ o dt; with r_ o ¼ DivðRÞ þ R: o

where ro is the density in the reference configuration at the initial state (i.e., at initial time t ¼ 0). It is worth pointing out that growth in soft tissues can be described by assuming that the density of the newly grown tissue has the same density as the initial tissue. Accordingly, a mass change is accompanied by a volume change. Hence, it results ri ¼ ro ¼ const, where ri ¼ ro/Det (Fi) is the density in the intermediate In this case, the source term R can be directly related to the deformation tensor Fi,  configuration.  resulting R ¼ ro DetðFi Þ Tr F_ i F1 . Clearly, a number of different choices are also possible, based on ad hoc hypotheses on tissue i growth. In order to describe remodeling, the rearrangement of fibers is herein considered as an illustrative example. To reach this goal, the structure tensor M is introduced as an internal variable, obtaining an alteration of anisotropic material properties with remodeling. In analogy with the ansatz of Eq. (79) with A ¼ M and a ¼ ro, the free-energy function jfe per unit mass is introduced as: jfe ¼ jfe ðC; Fi ; M; ro Þ

(89a)

Moreover, the Clausius–Duhem inequality shall be formulated considering open thermodynamical systems, allowing for the exchange of mass with the exterior, in order to account for the living nature of soft tissues. Accordingly, the dissipation inequality considers a mass specific entropy term s0 that takes into account external entropy flux and/or source. The Clausius–Duhem inequality in Eq. (8) for open systems reads as: Dint ¼ S :

C_  ro j_ fe ðC; Fi ; M; ro Þ  ro Ts0  0; 2

(89b)

for any admissible thermodynamical state. Accounting for the constitutive relationships in Eq. (80b), together with mass balance in Eq. (88), the Clausius–Duhem inequality in Eq. (89b) for growth and remodeling reads: _ þ Dm ðr_ o ; s0 Þ  0; Dint ¼ Qi : F_ i þ Hi : M

_ r_ o ; c admissible F_ i ; M;

(89c)

with the term Dm, associated with the exchange of mass, being: Dm ðr_ o ; s0 Þ ¼ qi r_ o  ro Ts0 ¼ qi ½Div ðRÞ þ R  ro Ts0 :

(89d)

A possible formulation for the free energy in growth and remodeling accounts for the elastic deformation from the intermediate to the current configuration by means of the (volume specific) strain-energy function Jse, depending on the elastic part of deformation Ce and weighted by the relative density (ro/ro )n, with n being a model parameter. Due to the remodeling of anisotropic

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directions, the explicit dependence of the strain energy on the structure tensor M is highlighted, that is Jse ¼ Jse(C, M). Therefore, the (mass specific) free-energy function jfe reads:   1 ro n jfe ðC; Fi ; M; ro Þ ¼ Jse ðCe ðC; Fi Þ; MÞ ; (90a) ro ro resulting in (see Eq. 80b): S¼2

 n ro vJse T F1 F ; i ro vCe i

 n r vJse Hi ¼  o ; ro vM

 n r vJse Qi ¼  o ; ro vFi

 n r qi ¼ ð1  nÞ o Jse ðCe ; MÞ: ro

(90b)

(90c)

Within the present ansatz, a possible definition of the extra entropy term s0 follows from requiring a null dissipation associated with the mass term Dm in Eq. (89d), obtaining:   1  n ro n s0 ¼ Jse ðCe ; MÞ½DivðRÞ þ R : (90d) T ro r0 The constitutive framework is completed by the definition of material response in the elastic regime (i.e., function Jse as in a hyperelastic framework), of evolution equations for Fi and M, as well as of mass flux R and source R, under the constraint prescribed by the dissipation inequality in Eq. (89c). Remarkably, a homogenized constrained mixture model, which gathers the mechanistic description of constituents-dependent growth and remodeling with the simplicity deriving from a gross description of the main kinematic effects, has been recently developed by C. Cyron and co-workers.

Open Problems and Future Challenges In silico analyses support nowadays the biomedical research in improving the effectiveness of the clinical outcome and in clarifying unknown and unexplored pathways that affect the functional behavior of biological structures. The understanding of the physical mechanisms behind the pathophysiological behavior of soft tissues, from the cellular to the organ scale, would open the door to ground-breaking results both in research and in clinics. Indeed, the ultimate goal is to turn heuristic knowledge into predictive capabilities, through the quantitative modeling of the fundamental interactions between mechanics and biology. The development of reliable constitutive models of soft tissues stands out as one of the major challenges in biomechanics. In this framework, both computational and theoretical issues arise. From the computational standpoint, the coupling of anisotropic and incompressible behavior in soft tissues inherits numerical stability problems. Moreover, especially when inelastic mechanisms are addressed, the implementation of constitutive models in numerical frameworks can be challenging. Finally, fast (ideally, real-time) simulations of large and realistic biological structures are needed for applying in silico approaches in clinics. Accordingly, special care shall be paid to verify the convergence rate of numerical algorithms, with constitutive models playing a major role on this issue. Issues on the theoretical formulation of novel constitutive models are associated with the need of capturing more and more experimental evidence that arise in tissue pathophysiology. Model generalizations shall not be paid in terms of loss of material stability mathematical requirements. Moreover, the description of a wide range of mechanical responses is generally accompanied by the increase of the number of parameters and these, generally, have a phenomenological meaning. Therefore, issues on model calibration appear, especially in a patient-specific framework. A common practice is to employ population-based mean data for material properties, even though this can introduce some inaccuracies when referred to specific scenarios. Furthermore, even assuming an accurate calibration, the evolution of the value of phenomenological parameters with the onset and evolution of pathologies is an open issue. In addition, possible constitutive inhomogeneities associated with pathological localized tissue defects are generally neglected. Typical disorders (e.g., aneurisms, keratoconus, arthrofibrosis) are indeed associated with structural defects at different length scales: in content of tissue constituents, in shape of collagen fibers, in collagen genetic pattern, in density of intermolecular cross-links. Microstructural and biochemical alterations reflect in a nonphysiological mechanical response of soft tissues, such as hyperextensibility or weakness. Multiscale approaches allow to incorporate structural information in a straightforward way, since model parameters have a clear physical meaning. Therefore, these models can be fed by patient-specific histological and biochemical information, accounting also for possible pathological alterations. The analytical rationale of available multiscale approaches, well-established in a hyperelastic framework, allows to maintain a low computational cost, despite the gained insight on the structure–mechanics relationship. Nevertheless, pathologies are often related to an inelastic tissue behavior but the application of a multiscale rationale to inelasticity is still an open issue. In this field, some results are available at the scale of collagen fibrils, starting from the modeling of damage- and plastic-related mechanisms affecting collagen mechanics (e.g., slip pulse, interstrand delamination), proposed by M. Marino and G. Vairo in 2014 and afterwards generalized by M. Marino in 2016. Advances on this issue, up to the tissue scale, would help to clarify the dominant viscous/damage/plastic mechanisms in pathological conditions across the length scales, contributing in developing targeted therapeutic approaches.

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Moreover, addressing growth and remodeling, multiscale approaches would allow to explicitly account for the strict relationship between tissue macroscopic response and remodeling mechanisms. Indeed, a number of classical approaches in growth and remodeling employ phenomenological expressions of fiber strain-energy functions, with fiber orientation being the only structural information. Therefore, apart from fiber angle, the remodeling-induced evolution of parameters governing strain energies is generally neglected. At most, only a phenomenological rationale, tough to be validated and calibrated, can be employed. On the other hand, multiscale approaches are promising alternatives, opening the door to the development of a new class of models for growth and remodeling. A multiscale rationale allows indeed to define remodeling laws directly from biological evidence and clinical cases, since model parameters are measurable properties describing tissue histology and biochemistry (e.g., the radius and crimp amplitude of collagen fibers, the density of intermolecular cross-links). Furthermore, one of the major challenges is related to the living nature of soft tissues, shedding the light on the role of biochemistry in tissue active response, as well as in growth and remodeling. Tissue mechanical response is indeed the result of biochemical processes regulated by cell–cell signaling pathways. Cells collect, produce, and release a wide range of ions and molecules (e.g., nutrients, hormones, enzymes, growth factors, drugs). The transport and diffusion of ions and molecules govern a cascade of biochemical reactions that may activate intracellular contractile elements and/or alter the composition, the properties and the arrangement of tissue constituents. Cell–cell signaling pathways are continuously active in soft tissues, governing the natural turnover of constituents. Pathologies develop when an imbalance of mechanobiological processes occurs, inducing nonphysiological histology and biochemistry. In turn, biochemical pathways are strongly affected by mechanics. For instance, the osmotic pressure in the interstitial fluid of soft tissues surely plays a role in the diffusion of ions and molecules. In addition, since biological cells respond to mechanical stimuli by altering the biochemical environment (possibly involving the nervous system), mechanical quantities also affect transport phenomena in terms of molecular sources and sinks. Accordingly, mechanics and biochemistry shall be taken into account within a unique framework, represented by a closed-loop feedback system. Furthermore, electrochemical mechanisms can also affect molecular pathways in specific tissues (e.g., myocardium, gastrointestinal tissue), adding the need of accounting for the electrical field in the formulation of the constitutive model. Therefore, the development of constitutive models of soft tissues within multiphysics frameworks is highly needed. In this context, chemo-mechano-biological modeling strategies for the coupling of arterial growth and remodeling with biochemical reactions have been traced, for instance, by P. Aparício, M.S. Thompson, and P. Watton in 2016. Along the same lines, an insight on the effects of the remodeling of fine tissue structural properties (i.e., the coupling with a multiscale constitutive approach), as well as on molecular transport mechanisms affecting cell–cell signaling pathways, has been provided by M. Marino, G. Pontrelli, G. Vairo, and P. Wriggers in 2017. Advances on the afore-introduced challenges arise as urgent priorities of the biomechanical research community. These would allow indeed to gain a novel insight on the living and adaptive properties of biological tissues, properties which distinguish the constitutive modeling in biomechanics from the one in the framework of standard “dead” materials.

Further Reading Ambrosi, D., Athesian, G. A., Arruda, E. M., Cowin, S. C., Dumais, J., et al. (2011). Perspectives on biological growth and remodeling. Journal of the Mechanics and Physics of Solids, 59, 863–883. Balzani, D., Brinkhues, S., & Holzapfel, G. A. (2012). Constitutive framework for the modeling of damage in collagenous soft tissues with application to arterial walls. Computer Methods in Applied Mechanics and Engineering, 213–216, 139–151. Cowin, S., & Doty, S. B. (2007). Tissue mechanics. New-York: Springer ScienceþBusiness Media, LLC. Cyron, C., & Humphrey, J. (2017). Growth and remodeling of load-bearing biological soft tissues. Meccanica, 52, 645–664. Fratzl, P. (2008). Collagen: Structure and mechanics. New York: Springer. Gasser, T. C., & Holzapfel, G. A. (2002). A rate-independent elastoplastic constitutive model for biological fiber-reinforced composites at finite strains: Continuum basis, algorithmic formulation and finite element implementation. Computational Mechanics, 29, 340–360. Holzapfel, G. A., & Ogden, R. W. (2010). Constitutive modelling of arteries. Proceedings of the Royal Society A: Mathematical, Physical and Engineering Sciences, 466, 1551–1597. Maceri, F., Marino, M., & Vairo, G. (2010). A unified multiscale mechanical model for soft collagenous tissues with regular fiber arrangement. Journal of Biomechanics, 43, 355–363. Marino, M. (2016). Molecular and intermolecular effects in collagen fibril mechanics: A multiscale analytical model compared with atomistic and experimental studies. Biomechanics and Modeling in Mechanobiology, 15, 133–154. Marino, M., & Wriggers, P. (2017). Finite strain response of crimped fibers under uniaxial traction: An analytical approach applied to collagen. Journal of the Mechanics and Physics of Solids, 98, 429–453. Marino, M., Pontrelli, G., Vairo, G., & Wriggers, P. (2017). A chemo-mechano-biological formulation for the effects of biochemical alterations on arterial mechanics: The role of molecular transport and multiscale tissue remodelling. Journal of the Royal Society Interface, 14, 20170615. Menzel, A., & Kuhl, E. (2012). Frontiers in growth and remodeling. Mechanics Research Communications, 42, 1–14. Sacks, M. S., & Wei, S. (2003). Multiaxial mechanical behavior of biological materials. Annual Review of Biomedical Engineering, 5, 251–284. Sacks, M. S., Zhang, W., & Wognum, S. (2016). A novel fibre-ensemble level constitutive model for exogenous cross-linked collagenous tissues. Interface Focus, 6, 20150090. Schröder, J., & Neff, P. (2003). Invariant formulation of hyperelastic transverse isotropy based on polyconvex free energy functions. International Journal of Solids and Structures, 40, 401–445.

Continuum Mechanics and Its Practical Applications at the Level of Scaling Laws Ko Okumura, Department of Physics and Soft Matter Center, Ochanomizu University, Tokyo, Japan © 2019 Elsevier Inc. All rights reserved.

Introduction Examples From Classic Problems Stokes’ Law of Drag Friction Hertz’s Law of Elastic Deformation Bending of a Plate Capillary Rise Practical Applications Microfluidic Devices for Liquid Mixing Kirigami Approach to Control Elastic Modulus Conclusion Acknowledgments References

111 111 111 112 112 113 114 114 115 117 118 118

Introduction In this article, we try to explain the usefulness of arguments at the level of scaling laws by examples in the context of continuum mechanics and its application relevant to biomedical sciences. Scaling laws are relations between physical quantities in which all the physical quantities appear in terms of powers, whereby a power of x is expressed in the form xa where a is a real number. For example, the relation U ¼ mV2/2, which gives the kinetic energy U of a particle of mass m moving at a velocity V, is expressed as U x mV2 at the level of scaling laws. In general, a scaling relation is often expressed by using the symbol “x.” For example, a scaling law “A x B” is a substitute of a more formal expression “A ¼ kB” with a dimensionless numerical coefficient k, the order of magnitude of which is not significantly different from one. In other words, the expression “A x B” means that the quantities A and B are dimensionally equal. This article is prepared for students and researchers who once studied the basics of continuum mechanics (theory of linear elasticity and hydrodynamics). However, this article does not require rich experiences in mathematical handling of equations in the textbooks. The basic knowledge of surface tension and capillary phenomena will help the readers to deepen their understanding. But this is not compulsory, and a brief introduction will be given. For readers who are tempted to study (again) continuum mechanics in a conventional way by following equations, we recommend textbooks such as Landau and Lifshitz (1987) (for Sections “Stokes’ Law of Drag Friction” and “Capillary Rise”) and Landau and Lifshitz (1975) (for Sections “Hertz’s Law of Elastic Deformation” and “Bending of a Plate”). This article is organized as follows. First, we derive some scaling laws for classic problems to show examples of scaling arguments. Second, we explain examples from recent studies, in which scaling laws are theoretically derived and experimentally confirmed. Finally, we summarize advantages of the arguments at the level of scaling laws.

Examples From Classic Problems Stokes’ Law of Drag Friction Let us consider a sphere with radius R, which is moving with velocity V in a fluid with viscosity h, as is shown in Fig. 1. In the limit in which the viscous effect is larger than the inertial effect, the drag force acting on the sphere from the liquid can be estimated at the

R V

η Fig. 1

Stokes’ friction: a sphere with radius R is moving at a velocity V in a liquid with viscosity h.

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level of scaling laws in the following manner. From Newton’s law of viscous force, the stress (s) acting on a surface is given by viscosity times velocity gradient at the surface: sxhVV in which the expression VV symbolically stands for the velocity gradient. In the present case, the velocity is of the order of V near the surface of the sphere, but it decreases as we go away from the sphere, and the characteristic decay length is R: dimensionally, the velocity gradient scales as V/R. The stress at the surface is thus given by hV/R. The drag force is given by the total force given by the stress times the area of the surface. In the present case, since the area of the sphere scales as R2, the drag force F scales as h(V/R)R2: FxhVR

(1)

More precise calculation gives a numerical coefficient 6p at the cost of lengthy calculation: F ¼ 6phVR.

Hertz’s Law of Elastic Deformation Let us consider a sphere with radius R in contact with a plate, where the distance between the center of the sphere and the plate is smaller than R, by the length d as in Fig. 2. Within the limit where R is very much larger than d, the elastic energy stored in the sphere can be estimated at the level of scaling laws in the following manner. Let us remember that the elastic energy u stored per unit volume in an elastic body with elastic modulus E subject to the strain ε scales as Eε2: uxEε2

(2)

In the present problem, the strain decays as we go away from the contact area and the characteristic decay length is the length of AC in the figure, which is called “a”: elastic deformation is localized in a volume xa3. Note that, since R [ d, the length a is much smaller than R, and this is the only length scale that characterizes the deformation of the sphere. This point would be better understood if one realizes that the boundary condition for the elastic deformation should be described by a (not d). Thus, the deformation is localized in a region near the contact area whose volume scales as a3. From this scaling view, the strain scales as d/a, and thus, the elastic energy per unit volume E(d/a)2 is stored in a region with a volume xa3: the total elastic energy stored in the sphere U is given by E(d/a)2a3. In fact, the length a is determined once R and d are given: for the right triangle OAC, Pythagorean theorem gives the relation R2 ¼ a2 þ (R  d)2, from which we obtain

With this relation, U is recast into the following form:

dxa2 =R

(3)

pffiffiffi UxE Rd5=2

(4)

Since the force F is given by the derivative dU/dd we obtain a nonlinear expression for F: pffiffiffi FxE Rd3=2 Note that at the level of scaling laws, the derivative with respect to d is equivalent to “division by d,” that is, d

(5) 5/2

/d ¼ d

3/2

.

Bending of a Plate Let us consider a plate with thickness h, whose length and width are a and b, respectively. When this plate is bent as in Fig. 3, we regard the plate as a collection of curved thin sheets of thickness dx whose surface S(x) is a part of the surface of a cylinder with radius R þ x ( h/2 < x < h/2), that is, S(x) ¼ bl(x) where lðxÞ ¼ ðR þ xÞq with a ¼ Rq

(6)

Assume that the strain of the sheet in the middle whose section is represented by the dashed curve in Fig. 3 (i.e. the sheet whose area is S(0)) is zero (if we set the strain of the surface S(0) to be ε ¼ ε0, the integral in Eq. (8) below has the minimum at ε0 ¼ 0): εðxÞ ¼ ðlðxÞ  aÞ=a ¼ x=R

R O A

a C δ

Fig. 2

Hertz’s contact: a sphere with radius R is pushed against the plate by the length d.

(7)

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h δ

R

θ

O Fig. 3 Bending of a plate. The length of the curved dashed line (a part of the arc of a circle with radius R) is a ¼ Rq, and the width of the plate (i.e. the length of the plate in the direction perpendicular to the sheet) is b.

The strain ε(x) of the surface S(x) is positive (negative) for x > 0 (x < 0). Since the energy per unit volume scales as Eε2 with Young’s modulus E, the total energy stored in the plate is calculated as Z h=2 dxEεðxÞ2 SðxÞxabEh3 =R2 (8) Ux h=2

in the small ε limit. Note that, at the level of scaling laws, the integration by x is equivalent to “multiplication by x.” By virtue of Pythagorean theorem, the deflection at the center d (see Fig. 3) satisfies the relation R2 ¼ (R  d)2 þ (a/2)2 in the small ε limit, which gives the relation Rd x a2. Thus, the energy is written as a function of d as U ðdÞxEbh3 d2 =a3

(9)

The force in the direction of d in this case is again given by “division by d,” which is linear in d: FxEbh3 d=a3

(10)

Capillary Rise Some basic knowledge of capillary phenomena will be, though not compulsory, helpful to understand this section. The key words are as follows: surface (interfacial) tension, surface (interfacial) energy, contact angle, capillary length, and capillary rise. These key concepts are briefly explained below. A liquid surface has a positive energy proportional to the area, and thus, the area always wants to shrink. Therefore, any line on the surface is pulled from both sides by the surface in the direction perpendicular to the line as in Fig. 4A (the force vectors are on the surface; since the two forces are equal in magnitude and directed in the opposite directions with each other, the total force on the line is zero). The force per unit length is called surface tension and denoted as g. Likewise, any solid surface or solid–liquid interface (usually) has a positive energy; because of the energy, any line on the surface is pulled from both sides by the surface in the perpendicular directions. The corresponding forces per unit length for the solid surface and the solid–liquid interface are denoted as gS and gSL, respectively. When a drop of a liquid with surface tension g is placed on a solid substrate with surface tension gS, the drop makes an angle q at the edge on the substrate as in Fig. 4B, and this angle is called the contact angle. As is clear from the force balance in the horizontal direction, we obtain Young’s relation: gS ¼ g cos q þ gSL. In general, capillary force acts strongly at small scales (compared with gravitational force), and gravitational force acts strongly at pffiffiffiffiffiffiffiffiffiffi large scales; the two forces become comparable at a length scale called the capillary length k1 ¼ g=rg with the liquid density r and the gravitational acceleration g. This scale can be estimated by balancing the surface energy and the gravitational energy of a spherical drop, at the level of scaling laws: gR2 x rR3gR (the latter scales as the gravitational potential energy of a drop at height xR). Solving this relation in terms of R, we obtain the size of liquid drop for which capillarity and gravity compete with each other, and this size is identified with the capillary length k 1. When a vertical tube of inner radius R is in contact with a bath of a liquid whose viscosity is h, the liquid starts to rise into the tube and then reaches the final height zf, if R  k 1 and q < p/2 (see Fig. 4C). This is understood in terms of force by considering interfacial forces acting on the liquid cylinder of length z. At the top, the column is pulled upward by the force 2pR(g cos q þ gSL); at

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(A)

γ

(C)

γSL

γ

γ

G

γSL

θ

γ

S z

(B)

G

L

γ γS

θ

L

γSL

S

2R G L γSL

γSL

Fig. 4 Capillary phenomena. (A) Surface tension. The dashed line of unit length on the surface is pulled from both sides by the forces with magnitude g. The two arrows representing the two forces g are perpendicular to the dashed line, and the force vectors are on the surface. (B) Force balance at the edge of a liquid drop placed on a solid substrate. Three forces are acting on the cylinder of unit length whose section is given by the dashed circle. L, S, and G stand for the liquid, solid, and gas phases, respectively. (C) Capillary rise. Interfacial forces acting on the liquid column between the two dashed straight lines.

the bottom it is pulled downward by the force 2pRgSL: in total, the column is pulled by the net interfacial force 2pRg cos q, which balances with the gravity acting on the column, at the equilibrium. The final height z ¼ zf is determined by the force balance 2pRg cos q ¼ rpR2zfg. When z is much smaller than zf, there is a regime in which the interfacial force and viscous friction become the two important players (in fact, when z is very small, we have another regime in which the interfacial force competes with inertia, de Gennes et al., 2013). In the so-called viscous regime, the force balance can be expressed as Rg cos q x h(V/R)Rz at the level of scaling laws. This is because the viscous friction per unit area on the surface of the liquid column scales as hV/R (at the center of the column, the liquid velocity scales as V; at the surface, which is away from the center by the distance R, the velocity is zero: the velocity changes by V over the length R, which means the velocity gradient scales as V/R) and the surface area of the cylinder scales as Rz. The balance Rg cos q x hVz can be rewritten as g cos q x h(z/t)z because at the level of scaling laws V x dz/dt x z/t (if z is given by some powers of t, the derivative with respect to t is equivalent to “division by t”). From this, we obtain pffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi zx tgRcosq=h (11)

Practical Applications Microfluidic Devices for Liquid Mixing Recently, in many biomedical applications, mixing of small amounts of liquids becomes an important issue, and microfluidics is promising in this context (Beebe et al., 2002; Squires and Quake, 2005). However, conventional microfluidic devices require pumps to transport liquids inside channels, and such a pumping system makes the dimension of the whole device considerably large. Here, we present examples of microfluidic devices that do not require pumps; thus, these devices can be easily used outside of laboratories. The key point of the device is to use open capillaries whose sections are rectangular (width w and depth d). Such a rectangular groove at submillimeter scale (x 0.1 mm) is now easy to make outside the clean room, by using a micromilling system. To use such rectangular open capillary channels for microfluidic devices, it is important to know the dynamics of capillary rise for the rectangular open capillary. By virtue of argument at the level of scaling laws, quite similar to the one we obtain in Eq. (11), we can show the following formula for the height z of the column after the time t in the deep channel limit (d [ w) (see Tani et al., 2015 for the details):

with

zxðe gwt=3hÞ1=2 for t  s

(12)

  s ¼ 12e gh= r2 g 2 w3

(13)

e ¼ gðcosq þ ðcosq  1Þw=2dÞ with g the surface tension of the liquid and q the contact Here, h is the viscosity of the liquid and g angle the liquid makes on the surface of channel walls. The scaling law is shown to be robust through agreement between theory and experiment. Although the formula is theoretically valid in the limits d [ w and t  s, the formula is shown to be practically correct as long as d > w and t < s, as demonstrated later.

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(B)

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(A)

2

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d w ν 0.2 0.225 500 0.4 0.230 100 0.4 0.222 500 1.0 0.220 50 1.0 0.220 100 1.0 0.221 500

10 5 0

0

5

10 15 t [100s]

20

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zLC / zf

zLC [mm]

20

d w ν 0.2 0.225 500 0.4 0.230 100 0.4 0.222 500 1.0 0.220 50 1.0 0.220 100 1.0 0.221 500 Theory

100 10−1 25 10−2 −3 10

10−2

10−1

100

101

102

t/τ Fig. 5 Data collapse showing the agreement between theory and experiment. (A) z vs. t. (B) z/zf vs. t/s. The parameters d and w are given in mm, and the kinematic viscosity n ¼ h/r is given in cS. From Tani, M., Kawano, R., Kamiya, K. and Okumura, K. (2015). Towards combinatorial mixing devices without any pumps by open-capillary channels: Fundamentals and applications. Scientific Reports 5, 10263.

The agreement was shown clearly via “data collapse.” Eq. (12) can be recast into the following dimensionless expression: z=zf xðt=sÞ1=2

(14)

with zf ¼

2e g rgw

(15)

The original relation given in Eq. (12) shows that, when z is plotted as a function of t, plot curves obtained for different parameters should look different. However, according to Eq. (14), when y h z/zf is plotted as a function of x h t/s by calculating the renormalization factors zf and s by using corresponding parameters, all the plot should look the same and represented by y ¼ kx1/2. This is clearly demonstrated in Fig. 5. The data collapse shown in Fig. 5 gives the value of numerical coefficient k, the only quantity that we do not derive theoretically. In addition, as mentioned earlier, the collapse shows that Eq. (12) is valid in a robust way: the data collapse well, as long as d > w and t < s while, theoretically, Eq. (12) becomes exact in the limit d [ w and t  s. Practically, the scaling law can be used as a guiding principle for designing complex open capillary devices. The robustness is practically important because deep channels are difficult to make (because micromilling tips are easily broken): availability of a simple law for easy-to-fabricate nondeep channels should be a great advantage. Potential weakness of open capillary system for contamination and evaporation effects was examined in Tani et al. (2015). It was shown that open capillary devices are practically useful, by means of two examples. The first example is shown in Fig. 6. This device has four mixing spots, in which we can put four different liquids. When we put, in the circular spot in the left, liquid M to be reacted with the four liquids, by virtue of capillary force, liquid M is transported and arrived at the four spots simultaneously, which is quantified by the plot in Fig. 6B. The second example is shown in Fig. 7. This device mixes two liquids placed at two spots that are transported to the central spot (shown in Fig. 7A) in order to initiate the mixing. As indicated in Fig. 7B, this device can estimate the reaction time without the influence of evaporation and contamination.

Kirigami Approach to Control Elastic Modulus The principle of kirigami is making many cuts on a sheet and stretching it to have a three-dimensional structure. The Japanese word “kiri” stands for cut, and “gami” stands for paper. Recently, many engineering applications of kirigami have been demonstrated, such as foldable actuators (Hawkes et al., 2010), self-folding shape-memory composites (Felton et al., 2014), stretchable lithium-ion batteries (Song et al., 2015), stretchable electrodes (Shyu et al., 2015), stretchable graphenes (Blees et al., 2015; Qi et al., 2014), and solar tracking batteries (Lamoureux et al., 2015). One of the features of kirigami is the high stretchability. Here, we discuss the physics of the high stretchability of kirigami at the level of scaling laws (Isobe and Okumura, 2016). We consider a simple example of kirigami illustrated in Fig. 8A, in which the parameters N , d , w, and h are defined. The force is plotted as a function of extension in Fig. 8B, which demonstrates clearly a first linear elastic regime terminated by a sharp transition. As illustrated in the photographs shown below the plot (Fig. 8B), this sharp transition corresponds to the transition from the inplane deformation to the out-of-plane deformation. We can estimate the corresponding energies as a function of the extension D at

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(A)

(B)

1.1 1

Normalized Mean Brightness

10.00s

5 mm

0.9 0.8 0.7 0.6

S1 S2 S3 S4

0.5 0.4 0.3

0

2

4

6 t [s]

8

10

12

Fig. 6 Mixing of liquids with four different acidities with BTB solution. (A) Result of the reaction. (B) Brightness analysis. S1–S4 stands for the four rectangular spots of the device shown in (A). From Tani, M., Kawano, R., Kamiya, K. and Okumura, K. (2015). Towards combinatorial mixing devices without any pumps by open-capillary channels: Fundamentals and applications. Scientific Reports 5, 10263.

(A)

(B)

40min

Mean Brightness

30

20

10

0

1 2 3 0

10

20

30

40

50

60

t [min] Fig. 7 Expression of GFP by an open capillary device. (A) Result of the expression. (B) Brightness analysis. From Tani, M., Kawano, R., Kamiya, K. and Okumura, K. (2015). Towards combinatorial mixing devices without any pumps by open-capillary channels: Fundamentals and applications. Scientific Reports 5, 10263.

the level of scaling laws in the limit w [ d by arguments similar to the one we employed for the derivation of Eq. (9) (see Isobe and Okumura, 2016 for the details). As a result, we obtain the stiffness in the first in-plane regime K1: 2NK1 xEd3 h=w3

(16)

with Young’s modulus E and the critical extension at the transition point Dc:

Dc =ð2NÞxh2 =d

(17)

These scaling laws are shown to be correct via “data collapse.” For example, Eq. (16) can be recast into the following dimensionless expression K1 =ðEhÞxðd=wÞ3

(18)

for a fixed N. The original relation given in Eq. (16) shows that, when K1 is plotted as a function of h, plot curves obtained for different parameters should look different. However, according to Eq. (18), when y h K1/(Eh) is plotted as a function of x h d/w, all the plot curves should look the same and represented by y¼ kx3, with k given by the agreement between theory and experiment. This is clearly demonstrated in Fig. 9. Similar data collapse is obtained for Eq. (17). The scaling laws given in Eqs. (16) and (17) clarify the physical mechanism of the high stretchability of kirigami. In addition, the laws established with numerical coefficients are useful for tuning the elastic modulus of sheet materials in the in-plane regime by adjusting the parameters: we can change the modulus of various sheet materials in a broad range by the kirigami structure without losing flatness (planeness). One interesting application would be the application of the kirigami structure for cell sheets.

Biomechanics j Continuum Mechanics and Its Practical Applications at the Level of Scaling Laws

(A)

(B)

w + 2d

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3 1

w

0.8

2.5 F [N]

d 2

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Initial Regime

0.4 Transition

F [N]

0.2 0

1.5

0 0.5 1 1.5 2 2.5 3 Δ [mm]

1 0.5

2Nd

0

d

0

20

40

60 80 Δ [mm]

Second Regime

Initial Regime

100

120

140

Final Regime

2 mm Fig. 8 (A) Simple model kirigami structure. (B) Force vs. extension. From Isobe, M. and Okumura, K. (2016). Initial rigid response and softening transition of highly stretchable kirigami sheet materials. Scientific Reports 6.

(B)

2 (d,w) (1,15) (1.5,15) (2,15) (3,15) (3.5,15) (2,10) (2,20) (2,30)

K1 [N/mm]

1.5

1

0.1

K1/Eh

(A)

0.5

0

0.01

(d,w) (1,15) (1.5,15) (2,15) (3,15) (3.5,15) (2,10) (2,20) (2,30) slope 3

0.001

0

0.05

0.1

0.15 h [mm]

0.2

0.25

0.3

0.01

0.1 d/w

1

Fig. 9 Data collapse showing the agreement between theory and experiment for N ¼ 10. (A) K1 versus h. (B) K1/(Eh) versus d/w. The parameters d and w are given in mm. From Isobe, M. and Okumura, K. (2016). Initial rigid response and softening transition of highly stretchable kirigami sheet materials. Scientific Reports 6.

Conclusion As seen in the classic examples, we see that some well-known expressions can be derived by a simple dimensional analysis if we give up deriving numerical coefficients. Note that, when developing arguments at the level of scaling laws, differentiation and integration can be replaced by division and multiplication, respectively. This approach at the level of scaling laws is useful in practical problems. Although such an approach is theoretically valid only in certain limits, scaling laws can be clearly checked by data collapse, and in some cases, they are robust, that is, valid for a wide range of parameters (Yokota and Okumura, 2011; Okumura, 2015); from the data collapse, we can experimentally derive numerical coefficients, which we do not obtain from theoretical analysis. It is stressed that finding scaling laws is not necessarily easy in general because scaling laws are valid in certain limits and we do not know a priori within which limits the scaling law in question becomes correct and we need to do experiments changing experimental parameters in wide ranges to find the answer. However, once a scaling law is established, the physics is clarified, and the law is useful as a guiding principle for practical applications because, without trials and errors, we can tune parameters for desired purposes. This approach is very popular especially in the field of soft matter physics, and monographs based on this approach are available for polymer physics

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(de Gennes, 1979) and capillary phenomena (de Gennes et al., 2005) (see also minireview articles on strength and toughness of biological composites (Okumura, 2015) and some other phenomena (Okumura, 2016, 2017)).

Acknowledgments This work was partly supported by Grants-in-Aid for Scientific Research (A) (No. 2014-PM01-02-01) of JSPS, Japan, and by ImPACT Program of Council for Science, Technology and Innovation (Cabinet Office, Government of Japan).

References Beebe, D. J., Mensing, G. A., & Walker, G. M. (2002). Physics and applications of microfluidics in biology. Annual Review of Biomedical Engineering, 4, 261–286. Blees, M. K., et al. (2015). Graphene kirigami. Nature, 524, 204–207. de Gennes, P. G. (1979). Scaling concepts in polymer physics. Ithaca, NY: Cornell University Press. de Gennes, P.-G., Brochard-Wyart, F., & Quéré, D. (2013). Capillarity and wetting phenomena: Drops, bubbles, pearls, waves. New York: Springer Science & Business Media. de Gennes, P.-G., Brochard-Wyart, F., & Quéré, D. (2005). Gouttes, Bulles, Perles et Ondes (2nd edn.). Paris: Belin. Felton, S., Tolley, M., Demaine, E., Rus, D., & Wood, R. (2014). A method for building self-folding machines. Science, 345, 644–646. Hawkes, E., et al. (2010). Programmable matter by folding. Proceedings of the National Academy of Sciences of the United States of America, 107(12441–12445). Isobe, M., & Okumura, K. (2016). Initial rigid response and softening transition of highly stretchable kirigami sheet materials. Scientific Reports, 6, 24758. https://doi.org/10.1038/ srep24758. Lamoureux, A., Lee, K., Shlian, M., Forrest, S. R., & Shtein, M. (2015). Dynamic kirigami structures for integrated solar tracking. Nature Communications, 6, 8092. https://doi.org/ 10.1038/ncomms9092. Landau, L., & Lifshitz, E. (1975). Elasticity theory. New York: Pergamon Press. Landau, L., & Lifshitz, E. (1987). Fluid mechanics (2nd edn.). Oxford: Pergamon. Okumura, K. (2015). Strength and toughness of biocomposites consisting of soft and hard elements: A few fundamental models. MRS Bulletin, 40, 333–339. Okumura, K. (2016). Simple views on different problems in physics: From drag friction to tough biological materials. Philosophical Magazine, 96, 828–841. Okumura, K. (2017). Viscous dynamics of drops and bubbles in Hele-Shaw cells: Drainage, drag friction, coalescence, and bursting. Advances in Colloid and Interface Science, 255, 64–75. https://doi.org/10.1016/j.cis.2017.07.021. Qi, Z., Campbell, D. K., & Park, H. S. (2014). Atomistic simulations of tension-induced large deformation and stretchability in graphene kirigami. Physical Review B, 90, 245437. Shyu, T. C., et al. (2015). A kirigami approach to engineering elasticity in nanocomposites through patterned defects. Nature Materials, 14, 785–789. Song, Z., et al. (2015). Kirigami-based stretchable lithium-ion batteries. Scientific Reports, 5, 10988. Squires, T. M., & Quake, S. R. (2005). Microfluidics: Fluid physics at the nanoliter scale. Reviews of Modern Physics, 77, 977. Tani, M., Kawano, R., Kamiya, K., & Okumura, K. (2015). Towards combinatorial mixing devices without any pumps by open-capillary channels: Fundamentals and applications. Scientific Reports, 5, 10263. Yokota, M., & Okumura, K. (2011). Dimensional crossover in the coalescence dynamics of viscous drops confined in between two plates. Proceedings of the National Academy of Sciences of the United States of America, 108, 6395–6398.

CT-Based Bone and Muscle Assessment in Normal and Pathological Conditions Paolo Gargiulo, Magnus K Gislason, Kyle J Edmunds, and Jonathan Pitocchi, Reykjavik University, Reykjavik, Iceland Ugo Carraro, IRCCS Fondazione Ospedale San Camillo Venezia-Lido, Venezia, Italy Luca Esposito, Massimiliano Fraldi, Paolo Bifulco, and Mario Cesarelli, University of Naples Federico II, Naples, Italy Halldo´r Jo´nsson, University of Iceland, Reykjavík, Iceland; and Landspitali University Hospital, Reykjavík, Iceland © 2019 Elsevier Inc. All rights reserved.

Introduction Muscle Research and Medical Imaging: A Focus on Muscle Changes and Degeneration Modeling Material Properties of the Bone Quantitative Muscle Assessment Nonlinear Trimodal Regression Analysis Applications Healthy, elderly, and pathological HU distribution parameters Muscle assessment in patients undergoing total hip arthroplasty Three-Dimensional approach: Virtual histology Quantitative Bone Assessment in THA Bone Profiling and Gain and Loss for 2-D-Based Assessment Applications Bone gain and loss for 3-D-based assessment Limitations CT-Based Finite-Element Modeling Loads and Constrains Aseptic Mobilization: Stress Shielding Topology Optimization Application: Preoperative Assessment Conclusion References Further Reading

119 119 120 121 121 123 123 125 126 127 127 127 127 130 131 131 132 132 132 132 133 133

Introduction Muscle Research and Medical Imaging: A Focus on Muscle Changes and Degeneration Muscle degeneration, characterized by the progressive loss of function, strength, and mass of the muscle, coupled with substitution of healthy muscle fibers with increased content of fibrous collagen and fat, has been consistently implicated as an independent mortality risk factor in aging individuals and those who suffer from neuromuscular primary pathologies and injuries (Goodpaster et al., 2006). When associated with aging, muscle changes are defined as sarcopenia, and its prevalence has been observed to incur significant declines in overall physical activity and quality of life. However, despite its growing interest in clinical aging research, a precise, sensitive, and quantitative method for defining etiology, diagnosis, and managements, in particular those related to prevention and rehabilitation, remains debated. Despite the absence of a universal definition, extant literature commonly associates sarcopenia with the loss of both skeletal muscle function and structure, and many mechanisms that govern these changes have been implicated (Fielding et al., 2011). The most prevalent of these mechanisms identifies the loss of muscle mass with infiltration of noncontractile tissues, which in turn confers an increased risk for frailty, disability, reduced mobility, hospitalization, and finally mortality. In an increasingly aging world, identifying a normative clinical definition for sarcopenia is of considerable importance. Many of the mechanisms that elicit muscle degeneration in sarcopenia have been analogously identified within the context of neuromuscular injury or pathologydmost notably with regard to spinal cord injuries (SCI). The dramatic deleterious changes in muscle function and composition exhibited in SCI patients have been suggested as accelerated analogues to the changes evident in sarcopenia, as paralysis from lower motor neuron denervation drastically and immediately reduces skeletal muscle function and then mass, whose extent is worsened by the associated increases of muscle adiposity and fibrosis. Severe skeletal muscle degeneration due to other metabolic and oncological illnesses, known as cachexia, has been likewise associated with muscle fiber loss and an increase in relative muscle adiposity and fibrosis, which are likewise been correlated to increased rates of cachectic patient morbidity and mortality. In general, the increasing prevalence of sarcopenic and cachectic muscle degeneration necessitates a gold standard for a quantitative assessment of muscle quality (Carraro et al., 2016).

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Fig. 1 Coronal view from a lower limb section underlining. Fat tissues
200 and
10 HU (A). Water, muscle, and connective tissues
9 and 200 HU (B). Bone cancellous 201 and 500 HU (C). Cortical bone 501 and 2000 HU (D). Hounsfield distributions associated to soft tissues (E) and the bone (F).

This article will focus on quantified analysis of X-ray computed tomography (CT) imagesda modality whose image matrices consist of linear attenuation values that are calculated from the specific X-ray absorption characteristics of present tissue. These linear attenuation coefficients may then be linearly transformed into a scale known as the Hounsfield unit (HU) scale. Fig. 1 shows how different tissues displayed with CT images are associated to their corresponding HU distribution. In detail, we have Fig. 1A–D showing a lower limb section, coronal view, where, respectively, fat tissues, muscle and connective tissues, trabecular bone, and cortical bone have been underlined with different colors. Fig. 1E and F shows their correspondent HU distributions. The number of pixels within the volume of interest is higher for muscle tissues compared to fat (Fig. 1E) and for trabecular bone compared to cortical bone (Fig. 1F). Therefore, when considering CT images of soft tissue, HU distributions typically range from negative values around
200 HU for fat, up to 200 HU including muscle tissues and connective tissue. Further sections of this article will illustrate a method for quantifying these HU distributions.

Modeling Material Properties of the Bone In order to derive the mechanical properties of the bone from the CT scan data, CT numbers or HU is first converted into bone densities, and then, the bone material propertiesdin terms of Young’s modulidare estimated from these data. The relationship between bone density and CT numbers has been assessed to be linear by scanning a phantom, with known material densities, together with the patient bones. The relationship between bone density and bone elasticity has been deeply discussed. In a recent

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121

literature review, Helgason et al. (2008) selected a total of 18 studies and 22 different density–stiffness relationships that have been derived from these experimental tests. Results appear largely dispersed, and the considerable differences between the investigated studies are related to the complexity of comparison of the empirical data. Different normalization criteriadin terms of apparent density and strain ratedwere used to decrease interstudy differences, but the testing, the measurement procedures and instrumentation, the specimen geometry, the holders, the boundary conditions, and the anatomical site of the samples still play an important role. Morgan et al. (2003) investigated the bone site specificity. To this purpose, they tested 142 specimens from 61 cadavers measuring on-axis elastic moduli and apparent densities from different anatomical sites with an experimental protocol that minimized the end artifacts and specimen misalignments. The observed site specificity is attributed to intersite variations in trabecular architecture, and when the site-specific structure was taken into account, there were no differences in terms of tissue moduli. Also, the relationship between bone density and strength of trabecular bone has been examined. Volume fraction and bone architecturedwhich in turn hugely depend on anatomical site, age, sex, and various pathologiesdaffect the strength of the trabecular bone causing a remarkable discrepancy in the value of the failure stresses. In fact, probing the effects of aging, McCalden et al. (1997), testing 255 specimens of cancellous bone from 44 human femora, highlighted the decrement in bone volume fraction of 8.5% per decade and the consequences on the mechanical competence of the bone due to quantitative changes rather than qualitative changes; moreover, the ultimate stress is reduced by almost 7% per decade for the human proximal femur and by almost 11% per decade for vertebral bone. Zysset et al. (1991) in a pioneer study reviewed the several efforts oriented to relate measures in the bone with density and strength also using a power-law function, even if in a static regime. Moreover, it is not currently known if this relationship depends on anatomical site. The relationship between bone density and CT numbers can be described by the following equation: rapp ðPÞ ¼ 0:000412HU þ 1:018668 where P denotes the generic point inside the femur. Moreover, the value of ash density, rash(P), is related to apparent density, rapp(P), by means of the following equation:
0:522 rapp ðPÞ þ 0:007 if rapp < 1 rash ðPÞ ¼ 0:779 rapp ðPÞ
0:025 if rapp  1 As a consequence, nonhomogeneous and isotropic behavior of bone femur models can be assumed, and at each point P, the following correlation between Young’s Modulus, E, and the local density, rash ¼ rash(P), can be chosen: 8 33900rash ðPÞ2:2 if rash  0:27 < EðPÞ ¼ 5307rash ðPÞ þ 469 if 0:27 < rash < 0:6 : 10200rash ðPÞ2:01 if rash  0:6 where E represents the actual local stiffness (MPa) at the generic point P. The choice of a very fine mesh in the FE model ensures that structural gradients over the representative volume element result in very small, avoiding conflicts in terms of the relation between structural gradients and material elastic symmetries. Postelastic behavior or bone tissue can be taken into account. Since values of the ultimate tensile stress of trabecular bone equate about 0.79 of the compressive yield stress and values of the ultimate tensile stress of cortical bone are about 0.76 of compressive yield stress, symmetrical bilinear isotropic hardening material models can be adopted, neglecting the tiny difference. The yield stress defined as that corresponding to the mean value of yield strain can be set to 0.7, while the tangent modulus can be set to five hundredths of the corresponding elastic modulus, depending on local density. Poisson’s ratio of the bone is typically set to 0.4.

Quantitative Muscle Assessment As previously mentioned, this article highlights the use of entire HU distributions to assess muscle quantity and quality. In this section, we will analyze soft tissue (fat, water and loose, and connective and healthy muscle) considering the information within the HU interval from
200 to 200. We will begin here by highlighting a method that allows quantifying Hounsfield distribution based on nonlinear trimodal regression analysis. Follows three examples from individuals representing whole spectrum of different soft tissue distributions: healthy control, elderly subject, and subject with degenerative muscle pathology. Finally, leading from the profiling, a fulldimensional approach is introduced to underline the utility of virtual histology.

Nonlinear Trimodal Regression Analysis Based on what we can see in Fig. 1E, we may define any distribution of soft tissue as based on standard Gaussian distributions: that is to say that the distribution of HU values correlating to that tissue varies according to a normal probability density function described by the following standard equation:

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Biomechanics j CT-Based Bone and Muscle Assessment in Normal and Pathological Conditions ðx
mÞ2 N 4ðx; m; s; aÞ ¼ pffiffiffiffiffiffi e
2s2 s 2p

(1)

where 4 represents the probability density function of the tissue type, N is the Gaussian distribution’s relative amplitude, m is the location of the distribution’s mean, and s is the width of the distributiondall of which may be evaluated as a function of each pixel or voxel’s computed HU value (variable x). Next, we may observe the aforementioned notion that segmented soft tissue in CT images generates HU distributions with three distinct tissue domains: fat, water-equivalent and loose connective tissue, and muscle and fibrous connective tissue. These three tissue types have linear attenuation coefficients that occupy distinct HU domains, namely, from
200 to
10 HU, from
9 to 40 HU, and from 41 to 200 HU, respectively (Edmunds et al., 2016). Indeed, if we take a CT image cross section and isolate all pixel values from
200 to 200 HU (the following linear transformation from CT bin number), the following characteristic curve is generated: It is clear from the shape of the earlier distribution that optimum regression analysis will require the definition of a generalized soft tissue distribution function that contains each of the three aforementioned tissue types (Fig. 2). Writing this in simple summation notation, we now have the following: X3

4ðx; mi ; si ; ai Þ ¼ i¼1

2

ðx
mi Þ Ni
pffiffiffiffiffiffi e 2si 2 si 2p

(2)

Finally, one may additionally observe the inwardly sloping asymmetries within the fat and muscle peaks. To optimize our regression analysis, it is crucial to account for this asymmetrydto do this, we employ the use of the Gaussian skewness parameter:   ðx
m Þ2 X3 Ni ai ðx
mi Þ
2s i2 i pffiffiffiffiffiffi pffiffiffi 4ðx; m ; s ; a Þ ¼ erfc e (3) i i i i¼1 si 2p si 2 where erfc is the error function and a is the distribution skewness. Note that, for the purposes of our definition, the skewness for the central peak (water-equivalent and loose connective tissue) may be assumed to be zero, thereby effectively removing its error function component. Utilizing this definition, we can now generate a theoretical curve via precise regression analysis. To do this, we begin by employing an iterative generalized reduced gradient algorithm via the minimization of the sum of standard errors at each CT bin value, x. After running hundreds of iterations, the standard error at each point will be gradually and progressively reduced. Finally, the sum of all standard errors may be utilized to compute the overall R2 value for the theoretical curve’s fit to the image datada procedure that utilizes the following set of equations:

Fig. 2 Diagram depicting the three components of the trimodal radio-densitometry distribution utilized in this study. This figure illustrates the location and skewness of each probability density function, with tissue types as previously defined.

Biomechanics j CT-Based Bone and Muscle Assessment in Normal and Pathological Conditions

R2 ¼ 1


RSS 2 RSS R ¼1
SST SST

SST ¼ RSS ¼

Xn

i¼1

(4)

ðyi
yÞ2

(5)

ðyi
f ðxi ÞÞ2

(6)

i¼1

Xn

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In summary a CT-scanned muscle volume generates 11-parameters that identify the muscle in unique way N, m, s, and a for fat, water-equivalent connective tissue, and muscle tissue. These distribution parameters may then be exported for further analyses and comparison.

Applications Healthy, elderly, and pathological HU distribution parameters To ascertain potential differences in muscle degeneration pathways, as evidenced by subtle changes in HU distributions, the aforementioned regression analysis method has been utilized on the CT scans of the three previously mentioned subjects: healthy control, elderly, and pathological. Results from the HU distribution profiling for each of these subjects are depicted in Fig. 3. As is evident from the results displayed in Fig. 3, there are significant qualitative differences between the shapes of the HU distributions of the healthy, elderly, and pathological subjects. The curve of the healthy subject exhibits a definitively high-amplitude muscle peak and a comparatively blunted fat peak, whereas the fat and muscle components in the elderly subject’s curve are decidedly the opposite in appearance. Contrastingly, the pathological subject elicited a distribution with heavily skewed fat and muscle peaks that were likewise closer together and shifted toward negative HU values. When compared according to the typical metric of average HU value, it is evident that the healthy subject’s average HU value was significantly shifted toward the muscle peak in the distribution. However, the average HU values of the elderly and pathological subjects were nearly indistinguishable from one another. To better explore the clearly obtuse differences in their distributions, each of the 11 regression analysis parameters was compiled and compared. The qualitatively distinguishable differences between HU distributions are further exemplified by these results in Fig. 4.

Fig. 3 Radio-densitometry distributions showing their respective nonlinear regression curves and average HU values. (A) The control subject’s curve showed a large muscle peak at around 55 HU, which directly contrasted with (B) the elderly subject’s distribution. (C) However, the pathological subject’s distribution was much lower in total pixel count (due to lower overall mass within the leg volume) and elicits fat and muscle peaks that are similar in amplitude with a large connective tissue regime between them. Likewise, the control subject’s average HU value was remarkably higher than the other subjects, whose values were similarly negative. All regression analyses yielded R2 values above 0.99 (Edmunds et al. 2016).

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Fig. 4 Results from the nonlinear trimodal regression analyses of each HU distribution. (A) The amplitude parameter, N; (B) the location parameter, m; (C) the width parameter, s; and (D) the skewness parameter, a. Note that the skewness for the connective tissue peak was assumed to be zero in accordance with our hypothesis of its being a normal distribution (Edmunds et al., 2016).

As is evident in Fig. 4, each distribution parameter confers its own distinct differences and relationships between the three subjects. For example, the elderly subject’s fat amplitude is at least fourfold larger than those of the other subjects but for the control subject’s muscle amplitude is largest by at least twofold. However, the pathological subject has the highest connective tissue amplitudeda parameter that, in general, increases nearly linearly between subjects, from the lowest value in the control subject. A potential physiological explanation for this variation in tissue amplitudes could reside with their precise definitionsdmost importantly, for the connective tissue distribution. This central tissue regime accounts for any water-equivalent and loose fibrous tissues, which, in addition to being generally present within healthy leg volumes, are increasingly evidenced in degraded, unhealthy muscle with significant fatty infiltration. An analogous linearity is present when observing fat tissue skewness, which is likewise lowest in the control subject and highest in the pathological subject. These data suggest that the fat tissue variance in positive distribution asymmetry may likewise be due to the progressive fatty infiltration into the much higher HU value muscle tissue. However, since this relationship is not commensurate in the muscle peak’s negative skewness, it is clear that skewness alone may not adequately describe present physiology. Moving to the location parameter, it is evident that fat location values are almost identical between subjects but muscle peak locations are singularly high in the elderly subject; these data associated to low amplitude in the muscle suggest muscle degeneration that have as results that the only remaining tissue within the muscle volume is a fibrous connective tissue. Similarly, the control subject has a singularly high connective tissue location value, although the difference is less extreme. While less significant, perhaps, it remains evident that the central connective tissue distribution shifts toward lower, “fatter” HU values in these elderly and pathological subjects, which is again evidence of muscle degeneration. Finally, the width parameter exhibits noticeable differences between subjects. The control subject has the widest fat distribution, but their muscle width is at least half of the other subjects’ values. Furthermore, while the elderly and pathological subjects exhibit similar fat and muscle widths, the elderly subject exhibits a much wider connective tissue distribution. A precise physiological interpretation of width as a parameter is difficult to obtain, but it is clear that sharper muscle and fat peaks confer to a more homogenous distribution of tissue HU values, suggesting that a reduction in the width parameter would confer to a commensurate reduction in muscle degeneration.

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Muscle assessment in patients undergoing total hip arthroplasty The potential utility of the regression analysis method previously detailed may be further tested with a cohort of total hip arthroplasty (THA) patients to assess changes in their upper leg muscle following hip replacement surgery. To do this, one may observe HU distributions from these patients from both presurgical and 1-year postoperative CT scans, followed by the assembly of distribution parameters for both their healthy and operative legs. Results from these analyses are depicted in Fig. 5. As is evident in Fig. 5, the results from the THA cohort analyses further support many of the aforementioned relationships between distribution parameters and the degree of muscle degeneration. To clarify this interpretation, it is useful to operate under the physiological assumption that these patients’ operative legs would be naturally underutilized compared with their healthy legs. Firstly, in general, all patients’ fat amplitudes are shown to decrease, while muscle amplitudes increase 1 year following surgery. Additionally, operative leg muscle amplitudes elicit a significant increase after this timespan. Furthermore, regarding the location parameter, there are minimal shifts evident in fat and muscle peaks, but there are notable increases in connective tissue location values in both the healthy and operative legs 1 year after surgery. This shift toward healthy muscle suggests the overall efficacy of corrected ambulation and 1 year of normative use. Indeed, once again, this is most evident and singularly significant in the operative leg.

Fig. 5 Results from the nonlinear trimodal regression analyses of the n ¼ 15 THA cohort. (A) The amplitude parameter, N; (B) the location parameter, m; (C) the width parameter, s; and (D) the skewness parameter, a. Note that $ and * denote P < 0.05 and the results are presented for before (b) and 1-year after (a) surgery (Edmunds et al., 2016).

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The width parameter elicits no significant or meaningful changes in either leg, but in accordance with our previous observation, the connective tissue peak widths are significantly larger than muscle or fat, which are both nearly identical in both legs. While not immediately useful in a physiological sense, this is further evidence that each distribution parameter seems to have its own sensitivity with respect to the entire population. Finally, Fig. 5 shows that the skewness parameter decreases in magnitude in both the healthy and operative fat peaks but remains relatively diminished and constant in the muscle peak, with one important exception: the preoperative muscle peak has significantly higher skewness than each of the others. We previously saw that the pathological example subject had a much more negative fat skewness than the other two example subjects. This suggested the presence of fatty infiltration and/or buildup of intramuscular fat in their degenerating muscle. This notion is similarly evident here, as the operative legs of the THA cohort elicit significantly more extreme negative skewness’s compared with other fat peak values. Altogether, the results from utilizing the reported regression analysis method indicate significant improvement in muscle quality in both legs following THA surgery. These data further support the notion that each HU distribution parameter may have particular specificities regarding muscle assessment, which further supports the method’s utility as a robust and straightforward indicator for muscle degeneration.

Three-Dimensional approach: Virtual histology Carrying out such analysis on volumetric data can yield a three-dimensional representation of the overall state of an anatomically defined full muscle (Gargiulo et al., 2010; Carraro et al., 2015). Fig. 7 shows the three-dimensional modeling of the tibialis anterior muscle for three different subjects: a healthy control subject, a healthy elderly subject, and an SCI patient with different degrees of muscle changes in the legs. Accordingly, the figures show that the relative content of the three main components of a human muscle varies to a great extent. Indeed, the red color represents healthy muscle tissue, the blue color represents loose connective tissue (with tissue density very near to water), and the yellow color represents fat. Interpretation of the images has to take into account that the longitudinal three-dimensional reconstruction of the full tibialis anterior in the three subjects is based on the cortical layer of the muscle in which the healthy muscle fibers of the full muscle are separated by loose connective tissue from the adjacent muscles. Indeed, even in a normal healthy person, the muscle shows large amounts of nonmuscle tissues. The true discriminative analysis is the one in which the percentage content in the 3-D reconstruction of the muscle are presented. In the case of a normal healthy person (left panels), the healthy muscle is of about 80%, the loose connective is around 20%, and fat is around 1%. In the elderly subject (central panels), the three values are near 50%, 45%, and 5%. Interestingly, the values in the SCI person (pathological subject) identify the unilateral lesion present in that case. It is worth stressing that similar percentage changes were observed in the content of different tissues when analyzed by morphometry of muscle biopsies harvested from the SCI patient (Fig. 6) (Kern et al., 2010).

Fig. 6 Segmented soft tissues and compositions within the tibialis anterior are from each subject’s 3-D upper leg volumes. Three tissue types of distinct radiodensitometric domains were utilized for the purposes of this study as follows: from
200 to
10 HU, from
9 to 40 HU, and from 41 to 200 HU representing, respectively, fat (yellow), loose connective tissue and atrophic muscle (cyan), and normal muscle (red). (A) The control subject’s muscle composition is composed primarily of healthy muscle tissue, whereas (B) the elderly subject exhibited markedly more fat and connective tissue, to the detriment of healthy muscle. (C) However, the pathological subject, whose left leg was affected by the sciatic nerve denervation, exhibited an almost identical healthy leg composition compared with the elderly subject but almost entirely fat and connective tissue in the pathological leg. It should be noted that, for the purposes of comparing pathological muscle degeneration with sarcopenic degeneration, only the radiodensitometric distributions from subjects’ left legs were utilized in this study (Edmunds et al., 2016).

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Quantitative Bone Assessment in THA The bone quality and the bone’s mechanical strength play a pivotal role in the success of THA. Many studies have looked at how the bone is likely to respond to external forces before and after THA, based on mechanical testing or computational studies. Clinical assessment of bone material distribution is an important factor in determining the treatment for patients or monitoring the changes in bone quality over a period of time. For THA, it is important to have information about the mechanical strength of the bone, prior to surgery and after surgery to monitor the possible aseptic loosening of the implant. Once a metal implant has been placed into the femoral canal or the acetabular cup, the stress state of the surrounding tissue will change as mechanical forces are proportionally distributed to the stiffness of the material. Since the stiffness of the prosthesis is in general an order of magnitude higher than for the bone tissue, the prosthesis becomes the main structural element, therefore reducing the load onto the surrounding bone tissue. The phenomenon is called stress shielding. This results in that bone tissue around the prosthesis becoming less dense through the remodeling process governed by Wolff’s law. The loss of density of the bone can have serious implications with regard to the fixation of the implant that can become loose. The loosening of the prosthesis is a common reason for revision surgery. It is inevitable that the bone will undergo changes in morphology and strength following a THA procedure. It is however important to be able to monitor where the changes occur to determine where the bone density decreases (loss) and where it increases (gain). Using CT scans that are taken 24 h and 1 year postoperatively, to create a three-dimensional model of the femur, can determine where the bone loss and gain have been occurring, respectively.

Bone Profiling and Gain and Loss for 2-D-Based Assessment Using the same methodology as described earlier and in Eqs. (1) and (2), it is possible to quantify the bone material distribution either at given cross sections (2-D) or over the whole bone (3-D). Like muscles, the bone will show one or two peaks at the density to which most pixels belong. For slices around the epiphyses of the long bones such as at the femur and the tibia, a cancellous and a cortical peak can be seen, but for slices around, the diaphysis will just show a single cortical peak since no cancellous bone is present at that location. Fig. 7 shows the different cross sections analyzed and the resulting bone density profiles from each cross section. From Fig. 7, it can be seen how the different profiles correspond to different cross sections around the femur. Slices A and B show a peak toward the low-density spectrum representing the cancellous bone at those locations, whereas slices C–E show the peak at the higher-density regions representing the cortical bone. The information obtained from the analysis of the peak will give information about the material characteristics such as the volume of the cancellous or the cortical bone, the stiffness, and the mass. Analysis of individual slices using image registration and processing could be beneficial for the clinical environment, both in terms of reducing radiation doses to the patients and the possibility of automatizing the bone segmentation procedure. The two scans from the 24 h and 1 year are superimposed using bony landmarks of the femur. The two slices from the same cross section are then compared, and if the difference in pixel values exceeds a given threshold (111 HU), it is determined that there has become a change in the bone material. If the difference is positive, there is a gain, and if it is negative, there is a loss. Fig. 8 shows how the formation and erosions of bone can be seen in a single slice around the lesser trochanter. Such analysis can give clinicians information both about the mechanical strength based on the bone mineral density distribution (Pétursson et al., 2015) of the bone at selected regions and information about the spatial progression of the bone remodeling process over a given time period to be able to assess how the bone is adapting to the prosthesis.

Applications Analysis was carried out on three patients undergoing THA, looking both at the bone profiles and using the Gaussian fit model and the image registration method. The results from the parameter fitting can be seen in Fig. 9. Looking at the fitted variable coefficients, it is possible to look at the material characteristics of the bone in each slice. The percentage difference between the value of the parameters representing the amplitude (changes in the amount of material) and the location (changes in the density) over a period of 1 year can be seen in Table 1. From Table 1, it can be seen that there is a tendency in the bone to shift toward the lower-density regions for the cancellous bone around the greater (subjects 2 and 3) and the lesser trochanter (subject 1). This can also be seen around the cortical region at the distal aspect of the bone but to a much lesser content. The results from the parametric fitting can then be compared with the finding using the 2-D image registration to identify the amount of gain and loss in each slice. The full results of the two-dimensional image analyses can be seen in Table 2. (See Table 3.) From Table 2, it can be seen that the area around the lesser and the greater trochanter is more susceptible to bone loss over a period of 1 year, whereas the distal aspect is more likely to show higher levels of bone gain.

Bone gain and loss for 3-D-based assessment For the 3-D analysis, image registration needs to be carried out to align the two scans so that the femur is in identical position between the scans. The image registration and realigning were carried out using Mimics (Materialize). The image registration was carried out to identify the following anatomical landmarks:

• •

The most distal tip of the stem Protuberance under the greater trochanter

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Fig. 7 Selected slices for the two-dimensional analysis of the bone profiles.

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Fig. 8 Slices at the lesser trochanter at 24 h postoperatively (left), 1 year postoperatively (middle), and the difference between the gain (green) and loss (red) of bone in the slice over a 1-year period (right).

Fig. 9

• • • •

Pixel distribution at the cross section and the fitted Gaussian curves through the dataset.

Superior aspect of the greater trochanter Lesser trochanter Gluteal tuberosity in the axial view Protuberance of the pectineal line in the axial view

The two scans were aligned using a common line defined using the longitudinal axis of the stem of the prosthesis. The two scans were then resliced according to the new orientation ranging from the proximal end of the femur to 2 cm distally to the distal tip of the prosthesis. Segmentations of both femurs were carried out incorporating pixel values between 255 and 3070 HU. Using Boolean operators, the prosthesis is deducted from the model, and finally, adjustments are made in order to remove image artifacts that arise

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Table 1

Percentage difference in the values of the fitted parameters describing the amplitude and density values of the bone material within each slice

Region

Subject 1 (%)

Greater trochanters Lesser trochanter 5 cm proximal to the distal part of the stem Distal part of the stem 2 cm distal to the distal part of the stem

Amp 38.8 11.4 9.8
2.8
5.7

Subject 2 (%)

Loc 10.3
23.7
5.2 1.0
1.2

Amp
8.7 2.0 12.0 2.6
10.5

Loc
32.5 8.3
22.7
2.2
0.5

Subject 3 (%) Amp 15.0
17.1 5.1
43.2
10.1

Loc
26.3 14.4 20.9
5.4
2.9

A positive value indicates an increase over a period of 1 year, and a negative value indicates a decrease.

Table 2

Two-dimensional analysis of the bone gain and loss of slices at various regions of the femur over a period of 1 year Subject 1 (%)

Region Greater trochanters Lesser trochanter 5 cm proximal to the distal part of the stem Distal part of the stem 2 cm distal to the distal part of the stem

Table 3

Loss 23.1 12.9 7.0 0.3 0.8

Gain 2.3 6.6 0.6 6.7 0.4

Subject 2 (%) Loss 5.6 8.4 12.4 7.8 8.8

Gain 2.9 3.3 0.1 3.8 3.5

Subject 3 (%) Loss 9.8 1.4 0.6 15.0 3.3

Gain 11.2 7.7 14.7 16.8 15.2

Results from the bone gain and loss from three different subjects Subject 1 (%)

Region Greater trochanter Lesser trochanter 5 cm proximal to the distal part of stem Distal part of the stem 2 cm distal to the distal part of the stem

Loss 1.8 5.0 5.2 1.5 5.5

Gain 2.7 3.3 3.0 0.2 3.6

Subject 2 (%) Loss 1.1 5.6 8.0 6.1 7.2

Gain 1.0 3.3 0.7 0.1 0.7

Subject 3 (%) Loss 1.6 4.6 5.5 8.6 8.2

Gain 4.0 2.4 1.7 1.4 5.7

from the presence of the metal implant during the scanning protocol. With the scans now aligned, resampled according to pixel size, a direct comparison is carried out on a voxel basis. To estimate the gain, the 24 h scan is subtracted from the 1-year scan where positive pixel values indicate bone gain. The opposite was done to work out the bone loss where the 1-year scan was subtracted from the 24 h and positive pixel values indicate the extent of the loss. A cohort of patients has been analyzed, and the results show that the medial proximal area of the femur around the lesser trochanter is most susceptible for bone loss. Some bone gain can be seen at the various regions of the femur, mostly at the distal part. The results from the gain and loss can be summarized in Table 1. The table represents the percentage of voxels that have decreased in density (loss) and increased density (gain). Using three-dimensional representation of the bone gain and loss will give the clinician detailed information about the location and the extent of the bone loss. Fig. 10 shows the extent of the bone gain and loss in 3-D for three subjects. From the image, it can be seen that the three-dimensional changes in the bone density vary greatly between the subjects. All show loss around the lesser trochanter area and some gain toward the distal aspect of the prosthesis. Three-dimensional analysis will give a detailed representation of the bone gain and loss, but the process is time-consuming, and therefore, two-dimensional analysis is more appropriate in the clinical practice.

Limitations For the analysis, there are several limitations with regard to the quantification of the gain and loss. The greatest limitation is that the scans between 24 h and 1-year postoperative scan can vary in terms of resolution, slice thickness, distance between slices, body posture of the patient, etc. Additionally, it is possible that the prosthesis has shifted, thus compromising the integrity of the image registrations based on the position of the prosthesis. The comparison between the voxels is based on reslicing and resampling of the 1-year postoperative scan that can introduce errors to the analysis. Additionally, partial volume effect can play an important role during the bone segmentation. As can be seen from the results of both the two-dimensional analysis and the three-dimensional analysis, the results between individuals vary to a great extent. A larger cohort and more vigorous subdivision of groups (fixation, gender, age, and prosthesis type) is needed to be able to get a more detailed view of the material and morphological changes in the bone over a period of 1 year.

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Fig. 10 Location of the bone gain and loss over a period of 1 year postoperatively for three subjects. The yellow color indicates no changes, red color indicates bone loss, and green color indicates bone gain.

CT-Based Finite-Element Modeling CT is a methodology used to acquire volumetric densities of tissues. CT-based finite-element (FE) models have been exploited for analyzing the micro- and macroscopic mechanical behavior of bone sites and for studying growth, remodeling, and morphogenesis phenomena. Many research groups have focused their activities on the development of rapid and automated FE preprocessing (Breusch et al., 2001). In principle, there are two basic concepts of generating FE models from CT scans: geometry-based and voxel-based. The voxelbased meshing technique is achieved by matching each CT voxel to a single FE, and the main advantage of this strategy is that it is a simple and automated technique that can deal with any shape of arbitrary complexity. However, curved surfaces cannot be properly represented by brick elements. Hence, where voxel meshes can provide accurate internal stresses and strains, the jagged-edged surface causes peak stresses and strains, which is a disadvantage when accurate mechanical data are needed at surfaces. Moreover, unstable elements (i.e., elements insufficiently anchored to the whole model and thus potentially involved in partial rigid-body motion) can be generated, which is a crucial problem in obtaining consistent FE models, hindering mechanical analyses (Lotz et al., 1990). In the following figure, two FE models are shown: on the left side geometry-based model and on the left side voxel-based model of prosthesized femur (Fig. 11).

Loads and Constrains Preclinical endurance tests on hip implants require defining realistic in vivo loads from younger and more active patients. These loads require simplifications to be applicable for simulator tests and numerical analyses. Bergmann et al. (2001) measured the contact forces in the joint with instrumented hip implants in 10 subjects during nine of the most physically demanding and

Fig. 11

Geometry-based (left) and voxel-based model of prosthesized femur.

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frequent activities of daily living. Typical levels and directions of average and high joint loads were extracted from the intra- and interindividually widely varying individual data. Because the magnitudes and orientations of peak forces substantially vary among the activities, load scenarios that reflect a collection of time-dependent high forces should be applied rather than using unidirectional forces.

Aseptic Mobilization: Stress Shielding Bad implant position or imprecise indexing of the prosthesis can determine aseptic mobilization phenomena that could result in collateral effects in the long period. After hip replacement, a frequent complication may also occur, represented by a mechanical loosening of the implant. This is revealed by implant movement and remodeling of the bone around the prosthesis, bone remodeling being the physiological dynamic response of the bone to the environmental stress level. From the mechanical point of view, the factors influencing the primary stability of the stem depend on biomechanical interaction between femur and prosthesis. The difference in stress before and after THA can be calculated and divided by the stress occurring pre-THA to determine the stress shielding increase (SSI) for that region. The before and after ratios are then volume-averaged over a specific region to calculate SSI for that region. Since the von Mises stress is strictly nonnegative, positive stress difference values indicate a decrease of the stress level in post-THA situation, therefore stress shielding. The explicit expression for SSI is the following: vanishing SSI means vanishing stress shielding and indicates an optimal condition. On the contrary, negative values of the SSI indicate an increase of stress when the prosthesis is present, and thus, they can be interpreted as a measure of stress concentrations, especially, if the actual stress in the bone exceeds yield strength or physiologically based thresholds.

Topology Optimization With particular interest on THA, optimization of orthopedic prostheses can be employed to minimize the probability of implant failure or maximize prosthesis reliability. This goal is often identified with the reduction of stress concentrations at the interface between bone and these devices. However, aseptic loosening of the implant is mainly influenced by bone resorption phenomena revealed in some regions of the femur when prosthesis is introduced. As a consequence, bone resorption appears due to stress shielding, that is to say the decrease of the stress level in the implanted femur caused by the significant load carrying of the prosthesis due to its higher stiffness. A maximum stiffness topological optimization (TO)-based strategy can be utilized for nonlinear static FE analyses of the femur–implant assembly, with the goal of reducing stress shielding in the femur and furnishing guidelines for redesigning hip prostheses.

Application: Preoperative Assessment Using the FE method as a clinical tool for surgical planning can give the clinician important information about the possible effects that the procedure is likely to have on the bone. During the press fitting, the prosthesis is hammered into the femoral canal, thus causing an elevated strain level onto the cortical shell. Should excess force be applied, there is a risk of causing periprosthetic fracture of the femur. The interference fit between the bone and the prosthesis was modeled using the FE method, by applying radially directed displacement boundary conditions on the surface of the femoral canal. By analyzing the strain field at the exterior surface of the bone, it is possible to determine which patients would be eligible for the press-fitting procedure and which should undergo cemented fixation. Fig. 12 shows the strain distribution of three individuals that show various mechanical response of the femur. From the figure, it can be seen that two of the three subjects show levels of strain that exceeds the theoretical fracture limit of the cortical bone. One of those two patients did suffer a periprosthetic fracture during the press-fitting surgery. The FE method used in conjunction with CT data can give important information to the clinician about the overall state of the bone. Such surgical planning tool not only can increase safety during the operation but also can help to identify the optimum treatment for the patient resulting in lower revision rates.

Conclusion Evaluating the material composition of both the muscle and bone in healthy subjects, it is possible to further understand the mechanisms that lead to the degradation of the tissues in pathological conditions and carry out the appropriate treatment. However, it must be taken into account the variability of the HU values that depend on different factors such as CT scan device, scanning protocol, region of interest, and artifacts. The methods presented here use thresholds and intervals based on the experimental data and calibration measurements performed with the CT devices employed in these studies. Finally, this article demonstrates how various different techniques can be used to assess the bone and muscle, both qualitatively and quantitatively.

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Fig. 12

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Three-dimensional strain distribution showing the anterior surface of the femur for three patients.

References Bergmann, G., Deuretzbacher, G., Heller, M., Graichen, F., Rohlmann, A., Strauss, J., & Duda, G. N. (2001). Hip contact forces and gait patterns from routine activities. Journal of Biomechanics, 34(7), 859–871. Breusch, S. J., Lukoschek, M., Kreutzer, J., Brocai, D., & Gruen, T. A. (2001). Dependency of cement mantle thickness on femoral stem design and centralizer. The Journal of Arthroplasty, 16(5), 648–657. Carraro, U., Edmunds, K. J., & Gargiulo, P. (2015). 3D false color computed tomography for diagnosis and follow-up of permanent denervated human muscles submitted to homebased functional electrical stimulation. European Journal of Translational Myology, 25(2), 129–135. Carraro, U., Kern, H., Gava, P., et al. (2017). Recovery from muscle weakness by exercise and FES: Lessons from masters, active or sedentary seniors and SCI patients. Aging Clinical and Experimental Research, 29, 579–590. Edmunds, K. J., Árnadóttir, Í., Gíslason, M. K., Carraro, U., & Gargiulo, P. (2016). Nonlinear Trimodal regression analysis of Radiodensitometric distributions to quantify Sarcopenic and sequelae muscle degeneration. Computational and Mathematical Methods in Medicine, 10. pp. Article ID 8932950. Fielding, R. A., Vellas, B., Evans, W. J., et al. (2011). Sarcopenia: An undiagnosed condition in older adults. Current consensus definition: prevalence, etiology, and consequences. International working group on sarcopenia. Journal of the American Medical Directors Association, 12(4), 249–256. Gargiulo, P., Kern, H., Carraro, U., et al. (2010). Quantitative color three dimensional computer tomography imaging of human long term denervated muscle. Neurological Research, 32(1), 13–19. Goodpaster, B. H., Park, S. W., Harris, T. B., et al. (2006). The loss of skeletal muscle strength, mass, and quality in older adults: The health, aging and body composition study. The Journals of Gerontology. Series A, Biological Sciences and Medical Sciences, 61(10), 1059–1064. Helgason, B., Perilli, E., Schileo, E., Taddei, F., Brynjolfsson, S., & Viceconti, M. (2008). Mathematical relationships between bone density and mechanical properties: A literature review. Clinical Biomechanics, 23, 135–146. Kern, H., Carraro, U., Adami, N., et al. (2010). Home-based functional electrical stimulation rescues permanently denervated muscles in paraplegic patients with complete lower motor neuron lesion. Neurorehabilitation and Neural Repair, 24, 709–721. https://doi.org/10.1177/1545968310366129. Lotz, J. C., Gerhart, T. N., & Hayes, W. C. (1990). Mechanical properties of trabecular bone from the proximal femur: A quantitative CT study. Journal of Computer Assisted Tomography, 14, 107–114. McCalden, R. W., McGeough, J. A., & Court-Brown, C. M. (1997). Age-related changes in the compressive strength of cancellous bone. The relative importance of changes in density and trabecular architecture. The Journal of Bone and Joint Surgery. American Volume, 79, 421–427. Morgan, E. F., Bayraktar, H. H., & Keaveny, T. M. (2003). Trabecular bone modulus–density relationships depend on anatomic site. Journal of Biomechanics, 36, 897–904. Pétursson, T. H., Edmunds, K. J., Gíslason, M. K., et al. (2015). Bone mineral density and fracture risk assessment to optimize prosthesis selection in total hip replacement. Computational and Mathematical Methods in Medicine, 2015, 162481, 7 pages. Zysset, P. K., Hayes, W. C., & Piazza, S. J. (1991). Biomechanics of fracture risk prediction of the hip and spine by quantitative computed tomography. Radiologic Clinics of North America, 29, 1–18.

Further Reading Carter, D. R., & Hayes, W. C. (1977). The compressive behavior of bone as a two-phase porous structure. Journal of Bone and Joint Surgery, 59, 954–962. Fischer, D., Kley, R., Strach, K., Meyer, C., Sommer, T., Eger, K., et al. (2008). Distinct muscle imaging patterns in myofibrillar myopathies. Neurology, 71(10), 758–765. Hansson, T. H., Keller, T. S., & Panjabi, M. M. (1986). A study of the compressive properties of lumbar vertebral trabeculae: Effects of tissue characteristics. Spine, 11, 56–62.

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Knowledge Extraction From Medical Imaging for Advanced Patient-Specific Musculoskeletal Models Marie-Christine Ho Ba Tho and Tien Tuan Dao, Sorbonne Universités, Paris, France; and Université de Technologie de Compiègne, Compiègne, France © 2019 Elsevier Inc. All rights reserved.

Introduction Geometrical Properties Extraction From Computed Tomography and MRI Modalities Mechanical Properties Extraction From Computed Tomography Modality Biochemical Properties Extraction From Advanced MRI Loading and Boundary Conditions Extraction From Advanced MRI Muscle Force Extraction From Magnetic Resonance Elastography Reliability and Uncertainty Data Conclusions and Perspectives Acknowledgments References

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Glossary Advanced medical imaging Relates to new imaging sequences and protocols, which are developed commonly for research purposes. However, these sequences are sometimes used in routine practice for preclinical tests. Advanced patient-specific modeling Is an engineering process to develop mathematical representation of a physical system with properties extracted directly from patient data. Patient-specific data may include geometry, mechanical properties, biochemical properties, loading, and boundary conditions. Finite-element modeling Is a continuum modeling approach to solve a physical problem by approximating the solutions on a set of finite elements. This approach is commonly used for studying the stress–strain relationships of the biomaterials. Knowledge extraction Is the information processing aiming to extract and establish useful knowledge from data. Material-driven meshing Is the meshing process to discretize the model geometry into finite elements based on the knowledge of involved materials. Rigid body modeling Is a mechanistic modeling approach to simulate the kinematics, kinetics, and muscle force behaviors of the musculoskeletal system.

Introduction The article will address the methodology we have developed in order to model bone and joints with appropriate geometric and mechanical properties derived from medical imaging. Medical imaging system such as MRI (magnetic resonance imaging) CT (computed tomography) is commonly used to evaluate musculoskeletal disease. Numerical methods are used for solving physical and mechanical engineering problems. These numerical methods are appropriate for modeling such complex system as human bone and joints and musculoskeletal systems. Literature review demonstrated the extensive use of finite-element modeling in biomechanics. During the last decade, the virtual physiological human a framework supported by the European Commission has allowed research based on personalized, predictive, and integrative medicine with in silico modeling at different scales of the body. Patient-specific computer modeling has been developed since the last decade, but still the specificity is not fully described or is limited to patient geometry. Concerning bone and joints modeling it is possible to obtain geometry and mechanical properties derived from CT, but using medical images one should be paid attention on the reliability of the images quality and the control of the acquisition parameters (Ho Ba Tho, 2003; Hellmich and Kober, 2006). Besides these extensive numerical models, most of models are derived from CT data and few from MRI. Few consider appropriate material properties derived from tissue characterization obtained from medical images, as they mostly are issued from the literature (data or relationships). The methodology we have developed is based on a semiautomatic generation of a three-dimensional geometric model of bone and joints, muscles anatomy derived from medical imaging CT, or MRI data (Ho Ba Tho, 1993). Predictive relationships obtained from the previous work demonstrated significant correlation between the material properties and quantitative measurements derived from imaging techniques (Rho et al., 1995). Then, from the same source of medical imaging data, numerical models with individualized geometric and mechanical properties were developed (Couteau et al., 1998). Concerning soft tissue same methods are applied (correlation between in vivo tissue characterization and in vitro) or direct assessment could also be performed

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using MRE (magnetic resonance elastrography) (Bensamoun et al., 2008). Concerning the in vivo forces, dynamic MRI (Dao et al., 2013) and MRE (Bensamoun et al., 2013) are able to provide data which could be exploited via a musculoskeletal model in order to predict in vivo loads by dynamic inverse analysis. This article aimed to present state-of-the-art methodologies for extracting knowledge from medical imaging for advanced patient-specific modeling of the musculoskeletal system.

Geometrical Properties Extraction From Computed Tomography and MRI Modalities Scientific and technological progresses of the medical imaging have advanced the knowledge of structure–function relationships inside the human body in a more quantitative and precise manner (Abramson et al., 2015; Ren et al., 2016). Computed tomography (CT) and magnetic resonance imaging (MRI) are two most important imaging modalities, which allows detailed 3D structures of the biological tissues and organs to be commonly acquired and used for numerical modeling (Ho Ba Tho, 2003). CT modality has been commonly used for hard tissue (e.g., bone) characterization, while MRI modality has been applied for soft tissue (e.g., muscle, tendon) characterization (Frangi et al., 2016). These conventional imaging modalities allow rigid body and finiteelement models of the human musculoskeletal system to be generated in a subject- or patient-specific manner (Ho Ba Tho, 2003; Blemker and Delp, 2005; Dao et al., 2012) (Fig. 1). Based on the raw images, image processing may be applied to extract geometrical properties of the musculoskeletal system like muscle volume, physiological cross-sectional area, or cartilage thickness (Dao et al., 2011).

Mechanical Properties Extraction From Computed Tomography Modality Mechanical properties of bone have been studied for over three decades in order to understand the mechanical behavior of bone in the process of fracture risk, repair, and bone-related disease. Besides, few data were available or insufficient for human bone modeling. Human bone is highly heterogeneous and anisotropic material. It can be compared to composite materials; it is made of two different tissue spongious bone (high porosity) and cortical bone (compact bone) depending on the anatomical location (Ho Ba Tho et al., 1991). In order to associate geometric and mechanical properties, we assume that measurements derived from medical imaging could predict material properties. We have investigated the relationships between CT number derived from CT imaging technique and mechanical properties of bone. The CT number characterize a linear coefficient of attenuation of X-ray within the tissue. For the CT scan, the pixel values are represented by an empirical number called CT number expressed in Houndsfield Units (HU). Predictive relationships between elastic properties and density and CT numbers for different human bone have been provided (Rho et al., 1995). Finally from CT images, one could model geometric and mechanical properties of the subject (Couteau et al., 1998). Recently, we developed material-driven mesh techniques allowing to drive the mesh with the knowledge of the material properties (Nguyen et al., 2016) (Fig. 2).

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Patient specific: geomtrical and FEM model (A) and rigid body model (B) derived from medical image (Dao et al., 2012).

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Material distributions on segmented CT slices (A), material-driven meshes (B), and cross sections with Young’s modulus mapping (C).

Biochemical Properties Extraction From Advanced MRI To provide patient-specific data at the tissue scale, advanced MRI sequences such as T2 mapping, diffusion weighted, and diffusion tensor imaging have been a potential solution (Dao et al., 2013). These techniques were commonly used to provide microstructural properties of the tissues of interest (e.g., cartilage, intervertebral disk (IVD)). In fact, T1r time reflects the loss of macromolecules. The T2 relaxation time estimated reflects the change of material (water and proteoglycan) properties. Apparent diffusion coefficient estimated from diffusion-based MRI reflects the molecule mobility of these material properties of the IVD. Diffusion-weighted image deals with the amount of water diffusion occurring within a voxel. It is well known that the diffusion in disk tissue is anisotropic. Consequently, this shows the direction-dependent character of disk structures. On the other side, diffusion tensor imaging quantifies the direction and magnitude of water diffusion in three dimensions in each voxel. In fact, diffusion-based imaging can be used as a potential microstructural marker for revealing early degeneration states of the IVD. Furthermore, these properties can be used to develop the accurate numerical IVD model at the tissue level. Among the available studies in the literature, the investigation of T1r-weighted MRI sequence on the IVD is complete. This MRI sequence has been used for assessing the in vivo and in vitro IVD tissue. In an in vitro study, Johannessen et al. (2006) reported a fair correlation (R coefficient ranges from 0.58 to 0.70) between T1r time and the biochemical properties of the IVD. In another study of the same research group (Nguyen et al., 2008), they showed a fair correlation (R ¼ 0.59) between T1r time and the mechanical properties of the IVD. Moreover, T1r time has been reported for the degenerated IVDs (Auerbach et al., 2006). This property has been considered as a novel biomarker for the assessment of disk degeneration and low back pain (Borthakur et al., 2011). However, despite its great potential capacity, this MRI sequence could not be used in any MRI machine without additional technical implementation (Taheria and Sood, 2006). For this purpose, T2 mapping and diffusion-based MRI sequences could be used as alternative solutions thank to their available implementation in classical medical imaging machines (Dao et al., 2013, 2014) (Fig. 3). Moreover, based on the correlation outcomes between in vitro image-derived properties and biomechanical property measurements, finite-element models may be developed with patientspecific biochemical property information (Fig. 4) (Nguyen, 2016).

Fig. 3 The T2 relaxation time estimated from the T2-mapping MRI images: (A and D) raw MRI image; (B and E) T2 maps of in vivo and in vitro IVD; (C and F) ADC maps of in vivo and in vitro IVD; (G) a dissected view of a cadaveric IVD. The T2 color codes mean that the outer AF is in blue color, the inner AF is in green color, and the NP is in red color. The ADC color code means that the outer AF is in blue color, the inner AF is in green color, and the NP is in yellow color.

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T2 image (A), meshed IVD model (B), and water content distribution on the meshed model (C) (Nguyen, 2016).

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T2 image (A), meshed IVD model (B) and water content distribution on the meshed model (C) (Nguyen, 2016).

From data obtained for CT and MRI, the patient-specific numerical model can be derived from extracted material properties of CT (vertebrae) and MRI (disk). According to our knowledge it is a first numerical model of patient-specific lumbar spine including patient material properties of the vertebrae and the IVD (Fig. 5). The next step would be to integrate the patient-specific forces for specific movement.

Loading and Boundary Conditions Extraction From Advanced MRI Finite-element simulation of the musculoskeletal system during dynamic movements requires complex loading and boundary conditions. The coupling between rigid body and finite-element models has been commonly performed to achieve this challenge (Halloran et al., 2012). Muscle forces and joint loading, estimated from rigid body modeling, are usually prescribed as loading and boundary conditions in finite-element models. These quantities may be computed in a straightforward manner for the upper and lower limbs. 3D motion capture systems like VICON has been commonly used for acquiring simulation kinematics (Dao et al., 2012). However, the kinematics of the lumbar spine system is still very hard to be acquired due to the complexity of the lumbar spine geometries. Lumbar spine ranges of motion are commonly acquired using medical imaging (e.g., 2D radiography Frobin et al., 1996 or biplanar radiography Rillardon et al., 2005) or dual fluoroscopic imaging (Xia et al., 2010) or Upright MRI (Alyas et al., 2008) or motion capture techniques (e.g., electromagnetic tracking system Wong and Lee, 2004) or computerized dynamic motion analysis devices (Mannion and Troke, 1999) or 3D motion tracking system with implanted bone pins (Rozumalski et al., 2008). Medical imaging techniques provide internal accurate lumbar spine ranges of motion, while motion capture provides external ranges of motion. However, imaging techniques provide only quasi-static motions rather than real dynamic motion (Powers et al., 2003). Moreover, due to limited range of motion and spatial/temporal image resolution, medical imaging approach needs further developments and investigations to provide accurate dynamic motion data. Furthermore, invasive character of some techniques limits their use in vivo (Frobin et al., 1996; Rillardon et al., 2005; Rozumalski et al., 2008) even they provide accurate motion data. Recently, we developed dynamic MRI for acquiring lumbar spine kinematics (Dao et al., 2015) (Fig. 6). Then, these data have been used for estimation of the lumbar spine muscle forces as well as for reassessment of kinematic of model outcomes (Fig. 7). In fact, noninvasive conventional dynamic MRI technique opens new perspectives to provide in vivo spinal kinematic and muscle force data reflecting the real lumbar spine motions. Then, these data may be used as loading and boundary conditions for the finite-element model in a patient-specific manner (Toumanidou et al., 2016).

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Fig. 6

Dynamic MRI images during hyperlordosis motion.

Fig. 7

Patient-specific lumbar spine model: frontal view (A) and lateral view (B).

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Muscle Force Extraction From Magnetic Resonance Elastography At the present time, the optimization technique has been accepted as a unique solution to estimate healthy muscle forces (Erdemir et al., 2007) but abnormal behavior of the musculoskeletal system, due to muscle diseases (Dao et al., 2012), cannot be modeled and simulated. Generic-parameterized musculoskeletal models are mainly composed of the Hill-based model, which is the most used rheological model to assess the muscle tensile forces. Some authors investigated the sensitivity study of abnormal behavior of muscle due to aging-affect (Thelen, 2003) or paralyzed effect (Law and Shields, 2005) by adjusting the contractile properties of the Hill-based model. However, there is no way to validate such parameters adjustment strategies. In fact, the quantification of these forces requires more intrinsic properties of muscles which can be determined in vivo with the determination of the sarcomere contractile dynamics properties (Llewellyn et al., 2008) or the characterization of the muscles mechanical properties (Ringleb et al., 2007; Bensamoun et al., 2007). Recently, we proposed a new direction to estimate lower limb muscle forces by introducing in vivo muscle elastic properties, leading to future simulation of abnormal muscles (Bensamoun et al., 2013) (Table 1, Fig. 8). In fact, the MR-elastography technique could provide a full database of active and passive elastic properties of different types of muscle allowing for the measurement of in vivo muscle forces, corresponding to each muscle, in order to better understand abnormal muscle behavior or to characterize the effect of age and growing processes on the musculoskeletal system.

Table 1

Tensile forces (mean  SD in Newton) of vastus medialis muscle at different phases during gait Tensile VM force during the stance phase (0%–60%)

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Midswing (73%–87%)

Terminal swing (87%–100%)

Voigt Springpot

791  2 793  1

220  25 157  22

170  24 108  21

511  73 272  106

294  61 104  81

93  31 14  24

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The musculoskeletal model of the lower limbs.

Reliability and Uncertainty Data The accuracy, validation of predictive numerical models needed to be investigated, quantified as the clinical diagnostic will depend on. In fact, data are extracted from medical images so one should pay attention on the reliability of the images quality and the control of the acquisition parameters (Ho Ba Tho, 2003). The reliability of assessment of the knowledge (geometry, mechanical properties, loads, and boundary conditions) derived from medical imaging (or literature data) are to be addressed and see their impact on the numerical developed and its propagation until the predictive results. Data are extracted from medical images, such as relationships with biochemical, physics, mechanical properties, their certification should be required as illustrated in Fig. 9. As an example, the effect of geometrical uncertainties due to segmentation error was performed on the quantification of joint loading and muscle forces of the lower limb system (Dao et al., 2012). In fact, the use of medical images for the patient-specific model leads to new error, which should be characterized and quantified. Concerning mechanical properties of the biological material investigated, it is quite difficult or impossible to say that it is the best description of the mechanical behavior or the law. There is obviously a need to validate or to establish statistical, uncertainty modeling. Random and epistemic uncertainties are two types of uncertainties of biomechanical data. Random uncertainty relates to the variability of repeatable acquisitions artifacts, hardware/software errors, or human errors (e.g., intersubject, intrasubject, interoperator, intraoperator) on the measurements. Epistemic uncertainty exists due to the lack of information/knowledge (e.g., modeling hypothesis or limited experiments). To model these uncertainties, precise and imprecise probabilities should be used. Recently, probability-box structure has been used for modeling heterogeneous data uncertainty such as medical imaging, mechanical properties, physiological and clinical data leading to quantify their propagation effect on the simulation outcomes (Dao and Ho Ba Tho, 2015, 2017) (Fig. 10). Furthermore, expert judgment may be used to assess the reliability of the biomechanical data from the literature. Advanced data modeling theories like belief theory may be applied for this purpose (Hoang et al., 2016).

Fig. 9

Same section location with different acquisition “hardware” parameters (A) and “software” parameters (B).

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Fig. 10

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Example of P-Box on thigh mass properties (A) and its impact on estimation of muscles forces expressed in range of values (B).

Conclusions and Perspectives Geometrical, material, forces knowledge derived from advanced medical imaging are of interest to provide full patient specificity. Besides, one should note that the accuracy, reliability of such models have to be investigated in order to provide an objective tool for aided decision to clinicians. In that context we have developed a methodology to model uncertainty related to heterogeneous data obtained experimentally and/or from the literature such as medical imaging, mechanical properties, forces, physiological and clinical data, and its propagation in the model. Precise and imprecise probabilistic approach have been applied to achieve this challenging issue. This new approach allowed to interpret predictive results with a level of confidence. In fact, a future generation of musculoskeletal models will be developed to provide reliable biomarkers for clinical decision support system of the musculoskeletal disorders in the framework of in silico and personalized medicine.

Acknowledgments The authors acknowledge the financial support of the Collegium-UTC CNRS INSIS and the Labex MS2T through the program Investments for the future managed by the National Agency for Research (Reference ANR-11-IDEX-0004-02). We are grateful and acknowledge the MYSPINE European project (FP7/2007–2013, no. 269909) for providing data for this present work.

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Gait & Posture, 28(3), 378–384. Taheria, S., & Sood, R. (2006). Spin-lock MRI with amplitude- and phase-modulated adiabatic waveforms: An MR simulation study. Magnetic Resonance Imaging, 24, 51–59. Thelen, D. G. (2003). Adjustment of muscle mechanics model parameters to simulate dynamic contractions in older adults. Journal of Biomechanical Engineering, 125, 70–77. Toumanidou, T., Dao, T. T., Ho Ba Tho, M. C., & Noailly, J. (2016). In Exploring the interactions between muscle function and intervertebral disc multiphysics in the healthy and the degenerated lumbar spine12th World Congress on Computational Mechanics (WCCM), 24–26 July 2016, Seoul, Korea. Wong, T. K. T., & Lee, R. Y. W. (2004). Effects of low back pain on the relationship between the movements of the lumbar spine and hip. Human Movement Science, 23(1), 21–34. Xia, Q., Wang, S., Kozanek, M., Passias, P., Wood, K., & Li, G. (2010). In-vivo motion characteristics of lumbar vertebrae in sagittal and transverse planes. 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Mathematical Quantification of the Impact of Microstructure on the Various Effective Properties of Bones Miao-Jung Y Ou, University of Delaware, Newark, DE, United States Annalisa De Paolis and Luis Cardoso, The Graduate School of The City University of New York, New York, NY, United States © 2019 Elsevier Inc. All rights reserved.

Introduction Quantification of the Influence of Material Microstructure on Its Effective Properties Integral Representation Formulas (IRFs) of Effective Properties Permeability and Tortuosity of Anisotropic Materials Fabric tensor Permeability tensor as a function of fabric Generalized IRF for anisotropic permeability Computation of Fabric Tensors and Static Permeability of Cancellous Bone Samples Sample Preparation Micro-CT Image Acquisition Image Processing Global Measurements of Microarchitecture Directional Measurements of Microarchitecture Realignment of Images to Principal Axes of Symmetry Pore Diameter Measurements Computational Fluid Dynamic Simulations Numerical Results Discussion Acknowledgments References

143 145 145 147 147 147 147 148 148 148 148 149 149 149 149 149 150 152 153 153

Introduction Poroelastic composites are two-phase composite materials consisted of elastic solid frames with fluid saturated pore space. The study of poroelasticity plays an important role in biomechanics, seismology and geophysics due to the nature of objects of research in these fields, for example, fluid saturated rocks, sea ice and cancellous bone. As the starting point of a systematically study of the microstructure influence on all the effective properties, we assume that constituent materials, that is, the solid matrix and the pore fluid are isotropic and that the anisotropy of the effective properties of the poroelastic composites is determined by the microstructure. For cancellous bones, this assumption was experimentally validated in Odgaard et al. (1997). We would like to remark that in the trabeculae of cancellous bone, there exist pores much smaller than the thickness of trabeculae (see e.g., Cowin, 1999; Gailani and Cowin, 2011; Morin and Hellmich, 2014; Scheiner et al., 2016) and the references therein. The two-phase model considered in this paper does not take into account the effects caused by the existence of these pores or any other structure smaller than these pores. For discussion of possible anisotropy of the trabeculae and the cortical bones (see Fritsch et al., 2009; Hellmich and Ulm, 2005; Ascenzi et al., 2008). The physical properties of these composites depend not only on the constituent materials but also on the microstructure of pore space and how the viscous pore fluid interacts with the solid frame. The bulk properties are described by various effective parameters. They play the role of coefficients in the integral-differential equations governing wave propagation through poroelastic materials when the wavelength is much bigger than the scale of microstructure. To be more precise, let f be the volume fraction of the fluid, u and U the displacement of the solid part and the fluid part, respectively. Define v d vtu (solid velocity), w d f(U
u) (fluid displacement relative to the solid), q d vtw, z d
V $ w. Let (x1, x2, x3) be the coordinates that coincide with the principal directions of the static permeability tensor K0, which is symmetric and positive definite (Biot, 1962). Biot (1956a, b) considered separately the low frequency case and high frequency case, which are defined by a cut-off frequency above which the Poiseuille flow assumption for the pore fluid breaks down. It is known that the drag force exerted on the viscous pore fluid is dominated by the inertia term for low frequency and by the viscous term for high frequency (Johnson et al., 1987). In Johnson et al. (1987), a description of the inertia/viscous drag effects valid over the entire frequency range was derived by a causality argument and hence it unifies Biot’s theory for low frequency and high frequency poroelastic wave equation. We refer to this set of poroelastic wave equations as the Biot-JKD equations. The Biot-JKD equations in three-dimensional space consist of the stress–strain relations and the six equations of motion. Let
 1 vui þ vuj , then the stress–strain relation is given by Biot (1962). 3 ij d 2 vxj vxi

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s11

1

0

cu11

B C B B s22 C B cu12 B C B B C B B s33 C B cu13 B C B B s23 C ¼ B cu B C B 14 B C B B s13 C B cu15 B C B B s12 C B cu @ A @ 16 p Ma1

cu12 cu22 cu23 cu24 cu25 cu26 Ma2

cu13 cu23 cu33 cu34 cu35 cu36 Ma3

cu14 cu24 cu34 cu44 cu45 cu46 Ma4

cu15 cu25 cu35 cu45 cu55 cu56 Ma5

cu16 cu26 cu36 cu46 cu56 cu66 Ma6

10 1 311 Ma1 CB C B C Ma2 C CB 322 C CB C Ma3 CB 333 C CB C B C Ma4 C CB 2323 C CB C Ma5 CB 2313 C CB C B C Ma6 C A@ 2312 A
z M

(1)

where p is the pore pressure, ciju are the elastic constants of the undrained frame, which are related to the elastic constants cij of the drained frame by ciju ¼ cij þ Maiaj, i, j ¼ 1, /, 6. In terms of the material bulk moduli ks and kf of the solid and the fluid, respectively, the fluid–solid coupling constants ai and M are given by 8 1 X3 > > c for j ¼ 1; 2; 3 1
> > 3ks k¼1 jk < aj d > > 1 X3 > > :
c for j ¼ 4; 5; 6 3ks k¼1 kj Md

ks    1
k=ks
f 1
ks kf

c11 þ c22 þ c33 þ 2c12 þ 2c13 þ 2c23 9 RN f ðx; t Þ e
st dt and let g denote the inverse Laplace transform of Define the Laplace transform of a function f(x, t) to be bf ðx; sÞd kd

0

g(x, s). The equations of motion in the Biot-JKD equations are (see e.g., p. 265 of Carcione, 2001) 3 X vsjk vvj vqj þ rf ; j ¼ 1; 2; 3 ¼r vx vt vt k k¼1


 vvj vqj vp rf  j  ; j ¼ 1; 2; 3
¼ rf a þ vxj f vt vt

(2)

(3)

pffiffiffiffiffiffiffi  j ðt Þ is the inverse Laplace transform of the dynamic tortuosity function aj(u) with sd
1u, rf and rs the density of the pore where a fluid and of the solid, respectively, and r d rs (1
f) þ frf. The convolution term denoted by * in (3) signifies that the dynamical interaction between solid and fluid are frequency depen j ðt Þ ¼ Tj dðt Þ þ khf Hðt Þ, dent (Biot, 1956b; Johnson et al., 1987). The low frequency Biot’s equation (Biot, 1956a) corresponds to a j rf where h is the dynamic viscosity of the pore fluid, kj the static permeability in the xj direction, Tj the infinite-frequency tortuosity in the xj direction, d(t) the delta function and H(t) is the Heaviside function. The dynamic tortuosity functions aj are the quantification of the fluid–solid inertia coupling and the viscous energy dissipation in the poroelastic material. Osteoporosis is characterized by a decrease in bulk strength of the cancellous bone matrix mainly due to the deterioration of the microstructure, Fig. 1. Currently, bone mineral density (BMD) is the gold standard for in vivo assessment of the fracture risk of bones and is measured using X-ray absorptiometric techniques (Chaffai et al., 2000). However, only 70%–80% of the variance of bone strength is accounted for by bone density. As the brittleness of bone depends on more factors than bone density, biologists believe that quantitative ultrasound techniques (QUT) could provide an important new diagnostic tool (Langton and Njeh, 2003; Fry et al., 1978; Fellah et al., 2004). Moreover, in contrast to X-ray densitometry, ultrasound does not ionize the tissue, and its implementation is relatively inexpensive. Since the loss of bone density and the destruction of the bone microstructure are most evident in osteoporosis cancellous bone, it is natural to consider the possibility of developing accurate ultrasound models for the isonification of cancellous bone, which is a poroelastic composite of elastic solid trabecula with pore space filled with fluid, Fig. 1. Biot’s theory for poroelastic waves (Biot, 1956a, b) predicts a fast- and a slow compressional wave and a shear wave in a poroelastic material. The slow compressional wave does not exist for elastic materials, so the detection of two different compressional waves signifies the poroelastic property of a specimen. Hosokawa and Otani (1997) (see also McKelvie and Palmer, 1991) identified fast and slow compressional waves in cancellous bone. The experiment suggested therein leads to a multiparameter inverse problem for the effective parameters in the Biot equations. It would be of enormous clinical advantage if an accurate method could be developed using ultrasound interrogation to determine whether one had osteoporosis. However, even in the low-frequency range, solving the multiparameter inverse problem is a daunting task due to the number of effective parameters involved and the ill-posed nature of the problem. Also, some of the parameters such as the tortuosity T and the diffusion correlation length L are difficult to measure. On the other hand, since all the effective parameters of a poroelastic sample are tied to the same microstructure, it is natural to study how the microstructure affect each of the effective parameters and use it as an underlying common feature in the aforementioned inverse problems. In this paper, we will focus on the dynamic permeability tensor K(u) and the dynamic tortuosity tensor a(u). For poroelastic materials, K(u) and a are defined in the frequency domain as

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145

Fig. 1 (A) A 12  12  12 mm3 cubic VOI was obtained at the center of the calcaneum. The fabric tensor, denoted by F, was measured and images were rotated in DataViewer until the principal directions were coincident with the faces of the cubic sample. The VOI was digitally cropped to 7  7  7 mm3 at the center of each rotated cubic sample to perform micro-computational fluid dynamic (mCFD) analyses. (B–D) A 7  7  7 mm fluid volume was added before and after the fluid mask on each of the three directions analyzed for all trabecular bone samples.



  K ðuÞ b
u b ¼ b
Vb p þ rf u2 u
iuf U h

 b
u b ¼
Vb b ; p þ rf u2 u aðuÞrf ð
iuÞ2 U

(4)

(5)

These relations hold if the fluid in the pore space is Newtonian, however, they are in general considered valid for Reynolds numbers, Re < 10. Eqs. (4), (5) imply the following relation between K and a aðuÞ ¼

ihf
1 K ðuÞ urf

for us0;

(6)

pffiffiffiffiffiffiffi where id
1. The infinite-frequency tortuosity T d limu / N a(u) and the static permeability K0 d K(0). It was shown in Johnson et al. (1987) and Auriault et al. (1985) that K and a stay the same if the elastic solid in its frame is replaced with a rigid solid.

Quantification of the Influence of Material Microstructure on Its Effective Properties Integral Representation Formulas (IRFs) of Effective Properties The IRF for the dielectric property of a composite material with two isotropic constituents was first derived in Golden and Papanicolaou (1983). Naturally, the dielectric property of a composite material should depend on the dielectric properties of the constituent materials and on the microstructure. The thrust of this IRF is that it quantifies the influence of the former through their ratio appearing in the integrand of the IRF while encoding all the microstructure influence in the positive measure in the IRF. The nth moment of the measure was shown to be related to the (n þ 1)-point correlation function of the microstructure, n ¼ 0, 1, 2,/. The IRF was derived by applying Herglotz functions theory. This result has been applied to study bone structures (Kenneth et al., 2011). The generalization of IRFs from dielectric composites to elastic composites has been discussed in Milton (2002) (see the references therein) and in Ou (2012). In this section, we will focus on the IRF for dynamic permeability derived in Avellaneda and Torquato (1991) by summarizing the key ingredients in its derivation and discuss some of its implications.

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The IRF of the dynamic permeability of an isotropic porous media of rigid solid matrix saturated with fluid with kinematic viscosity n is Z v Q1 QdGðQÞ (7) K ðuÞ ¼ F 0 1
iuQ where G(Q) ¼ 0 for all Q  0 and G(Q) ¼ 1 for Q  Q1 for some positive constant Q1, whose meaning will be made precise later. Similar to the features of the IRFs mentioned above, (7) provides a representation of K that quantifies its dependence on frequency through the integrand while encoding all its microstructure dependence through the measure dG F . Here F is the formation factor. One key ingredient in the derivation of (7) is the following spectral properties of the Stokes equation. Let Jn and 3 n be the eigen functions and eigen values of the Stokes equation, V 1 the region occupied by the pore fluid and vV the fluid–solid interface

DJn þ VQn ¼
3 n Jn and V,Jn ¼ 0 in V 1 ; Jn ¼ 0 on vV; Qn dðv3 n Þ
1 ; 0 < 3 1  3 2  / and

(8)

/ N as n / N. The eigenfunctions are orthonormal in the sense Z 1 Jm ðxÞ,Jn ðxÞdx ¼ dmn ; ðKronecker deltaÞ (9) jV 1 j V 1 RN In terms of these eigenfunctions, the Laplace transform b v ðx; sÞd 0 vðx; t Þe
st dt of the solution to the linearized Navier-Stokes equation with a forcing term in an arbitrary direction e with magnitude v0, ! vv p ¼
V (10) þ nDv þ v0 edðt Þin V 1 ; V,v ¼ 0 in V 1 ; v ¼ 0 on vV; vt rf 3n

can be represented as b v ðx; sÞ ¼ v0

N X n¼1

bn Jn ðxÞ

1 ; 1=Qn þ s

(11)

Another key ingredient in the derivation is the functional representation of K(u), derived from homogenization theory, as the spatial average of b v K ðsÞ ¼

N X n b2n b : v ðx; sÞ,e ¼ nf v0 1= Q nþs n¼1

(12)

where the second equality results from substituting (11) in to the functional expression. This leads to the IRF in (7) and the explicit P P b2n expressions for the formation factor F d (f nN¼ 1 bn2)
1 and GðQÞd PQNn Q 2 . b n¼1 n

It was proved in Miao-Jung Yvonne (2014) that the well-known JKD dynamic permeability 1 ,0sffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi 4iT 2 K 20 rf u iTK0 rf u D A
K ðuÞdK0 @ 1
hf hL2 f2

(13)



 can indeed be represented as the IRF in (12) with Q1 ¼ xp and a probability measure dGðuÞ ¼ cI ðuÞ jðuuÞ du þ xrp dxp , where pffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi pffiffiffiffiffiffiffiffiffiffiffiffiffiffi pffiffiffiffiffiffiffiffiffiffiffiffiffi 2C2 xp ðxp
C1 Þ C uðC
uÞ C þ C2 þ4C22 0 jðuÞdp C2 2 þuðC1
uÞ ; C2 dFkn 0 ; C1 d4CL2 Fk ; xp d 1 2 1 ; rd ; cI the characteristic function of the interval [0, C1], du 2 2xp
C1 ½ 2  1 the Lebesque measure and d is the Dirac measure. This implies that the microstructure information influences L, the electrically weighted average of volume-to-surface ratio of the dynamically connected pore space, through the first moment and the second moment of measure dG as follows vffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi u 2K0 T u i L¼t h m2 ðdGÞ f 2
1 ðm1 ðdGÞÞ qffiffiffiffiffiffiffiffi 0 It is interesting to note that the empirical formula for L suggested by JKD (Pride, 1992) is Lz 2TK f=4 . Another implication of (12) is that the dynamic tortuosity function a(u) also assumes an IRF which contains a simple pole at u and a Herglotz function. Consequently, given infinite-tortuosity T and data of K(u) at M distinct frequencies, a(u) can be approximate to very high accuracy as (Miao-Jung Yvonne, 2014). M X rj ihf 1 þTþ ; rj > 0;
< pj < 0 aðuÞz rf K0 u
iu
pj Q1 j¼1

This result provides an efficient numerical scheme for handling the memory term in (3).

(14)

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147

Permeability and Tortuosity of Anisotropic Materials The engineering community has a long history of trying to incorporate the influence of the microstructure, for example, anisotropy, of composite materials into the effective properties. Among them, the permeability tensor and the effective elasticity tensor are of special interest in the poroelasticity community. A number of experimental, analytical and numerical studies have investigated the relationship between the permeability tensor and measures of the pore architecture. Such studies have shown that the permeability tensor depends on the size, shape and directionality of pores. A simple relationship between the permeability tensor and a tensorial descriptor of the spatial directionality of pores, the fabric tensor, A, was previously proposed by our group, and it is presented below. We will also present in this section new results based on generalizing the IRF in (7) to arbitrary anisotropic cases.

Fabric tensor The fabric tensor serves as a measure of the degree of structural anisotropy of the porous medium (Cowin, 1985, 2004; Cowin and Cardoso, 2011, 2014). Use of the fabric tensor is restricted to materials with orthotropic or higher symmetry. The eigenvectors of F are the principal axes of material symmetry of the porous solid medium and the eigenvalues of F provide a measure of the distribution of porous volume fraction in the direction of the principal axes of material symmetry. As with any symmetric positive definite second order tensor in 3D, the fabric tensor may be represented as an ellipsoid. The ellipsoid is one with three unequal axes for an orthotropic fabric, one unique axis for a transversely isotropic fabric (forming an ellipsoid of revolution about the unique axis) and the ellipsoid becomes a sphere in the case of an isotropic fabric. The fabric tensor is a good measure of the pore structure anisotropy in cancellous bone tissue, as well as the mechanical and fabric main directions that coincide in cancellous bone. The second rank fabric tensor A is dimensionless, symmetric, and its invariants I, II and III are related to the traces of A, A2 and A3 by I ¼ tr ðAÞ; II ¼

      1 1 tr ðAÞ2
tr A2 ; III ¼ tr ðAÞ
3tr A2 þ 2tr A3 2 6

The fabric tensor is normalized by setting tr(A) ¼ 1.

Permeability tensor as a function of fabric The relationship between the second-rank intrinsic permeability tensor K, and the fabric tensor A is obtained by assuming that K is an isotropic function of A. The relationship between two second-rank symmetric tensors in which one is an isotropic function of the other then produces the relationship (Cowin and Cardoso, 2011). 0 1 3 X (15) K3 Aiq Aqj A Kij ð0Þ ¼ k0 @K1 dij þ K2 Aij þ q¼1

with Ki, i ¼ 1, 2, 3 being functions of f, II and III and k0 dp2 ð2K1 þ K2 Tr ðAÞ þ K3 TrðA,At ÞÞ represents the value of the permeability tensor when it is averaged over all possible directions at a point. The expression in (15) takes into account dissipation phenomena due to viscous losses, however, it is adequate only for low frequencies of fluid motion and needs to be corrected to take into account the change in fluid flow regime occurring between low and high frequencies of wave propagation (Cardoso and Cowin, 2011; Cowin and Cardoso, 2011). This correction results in the following expression for the dynamic permeability 0 1

3 X J1 ð fdÞ @ K1 dij þ K2 Aij þ (16) K3 Aiq Aqj A; Kij ðuÞ ¼ k0 1
2 fd J0 ð fdÞ q¼1 qffiffiffiffiffiffiffiffiffiffiffiffiffiffi f d iurf =h is the viscous skin depth, d the squared average diameter or pores, and J0, J1 are the zerothorder and first-order Bessel functions of first kind, respectively. This correction was originally introduced by Johnson et al. (1987) describing a dynamic permeability in a porous medium system characterized by cylindrical tubes. The local fabric tensor A can be obtained from processing the digitized images of samples for parameters by fitting the Mean Interception Length (MIL) in all directions by ellipsoids (Harrigan and Mann, 1984; Tabor, 2009; Tabor and Rokita, 2007). An interesting feature of this empirical approach is that other effective properties such as the drained elasticity tensor can also be regarded as certain combinations of the fabric tensor with coefficients dependent on f, II and III (Cardoso and Cowin, 2011; Cowin, 1985, 1986). Despite its empirical nature, it is shown in Tabor (2009) that fabric tensor is able to explain at least 90% of the variation of the apparent elastic constants. 2

Generalized IRF for anisotropic permeability To generalize the IRF for isotropic (or sometimes referred to as scalar-valued) K in (Avellaneda and Torquato, 1991) to the permeability function for anisotropic materials, we note the direction-dependent version of the Laplace transformed (10) is

.  i p r þ nDb v i ¼ 0 in V 1 b v i þ v0 ei ; V,b v ¼ 0 on vV (17) sb v i ¼
V b v i has the where ei is the unit vector in the ith direction. Similar to the expression in (11), the direction-dependent solution b following expression

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Biomechanics j Impact of Microstructure on the Various Effective Properties of Bones

b v i ðx; sÞ ¼ v0

N X n¼1

with bni defined as


bin ¼ ei ,

1 jV 1 j

Z V1

bin Jn ðxÞ

1 ; 1=Qn þ s

(18)



Jn ðxÞdx eei ,hJn i

(19)

Finally, using the functional representation of permeability by L. Tartar in Sánchez-Palencia (1980), we have the expression for the (i, j) component of K, kij ðsÞ ¼

j N E X nD i bin bn b v ðx; sÞ,ej ¼ nf : v0 1=Qn þ s n¼1

(20)

In matrix form, it is expressed in terms of the outer products hJni 5 hJni K¼

N E X nD i hJn i5hJn i b v ðx; sÞ,ej ¼ nf : 1=Qn þ s v0 n¼1

(21)

P Define F d (f nN¼ 1hJni 5 hJni)
1, which is termed the formation tensor in the physics literature, and note that hJni 5 hJni is a real-valued, symmetric, nonnegative matrix for each n. This means K is a matrix-valued function holomorphic in Cy(
N, 3 1] on the complex s-plane and can be represented as an IRF with real, positive semidefinite matrix-valued distribution G(Q) d P fF( Qn < QhJni 5 hJni) Z Q1 QdGðQÞ (22) K ðsÞ ¼ nF
1 1 þ sQ 0 Note that G(Q) ¼ 0 for Q  0 and G(Q) ¼ I for Q > Q1. This IRF provides a precise characterization of microstructure’s influence on K through the matrix-valued positive semidefinite measure dl(Q): ¼ F
1QdG(Q), which is clearly independent of u. Recall that s d
iu. Most importantly, this IRF separates K(s)’s dependence on the frequency u from its dependence on the microstructure. Hence it provides the foundation of dehomogenization, which is to obtain information on dl(Q) from values of K at distinct frequencies. We would like to emphasize that the complexification, which is necessary in deriving the IRF, is applied to the frequency u, not the space variable x and hence it does not restrict the applicability of our approach to problems in R2. Even though this approach works in Rn for any n  2, we are mainly interested in problems in R3. As was in the isotropic case (Miao-Jung Yvonne, 2014), with a change of variable x d
1/s and R(x) d
s(nF)K(s), we see that R(x) is a matrix-valued Stieltjes function (Fritzsche et al., 2015). Z Q1 QdGðQÞ (23) RðxÞ ¼ x
Q 0

Computation of Fabric Tensors and Static Permeability of Cancellous Bone Samples Sample Preparation Three human calcanei were obtained from the National Disease Research Interchange (NDIR; Philadelphia, PA) resource center. Bone donors were female who died of cardio/pulmonary failure, drug/alcohol abuse, cancer or natural causes. Calcanei were harvested within 24 h postmortem, and immediately fixed in 10% formaldehyde. Upon arrival at our laboratory, calcanei were carefully cleaned free of soft tissues and immersed in formaldehyde at 20 C for one more day and stored at 4 C.

Micro-CT Image Acquisition Calcanei were thawed to room temperature prior to scanning with an 1172 SkyScan high-resolution mCT system (SkyScan, Belgium). X-ray projections were acquired at a nominal isotropic voxel size resolution of 13.47 mm using a 0.5 mm aluminum filter to eliminate beam-hardening artifacts. X-ray projections were generated every 0.2 of rotation, obtaining 900 consecutive projections. To produce high-contrast, low-noise images the projections were averaged four times, and a median filter was used to prevent speckle noise formation. Due to the large size of specimens, four consecutive vertical connected scans were needed to scan each bone; each of these scans consumed approximately 11 h. Hydroxyapatite rods (HA, 2 mm diameter, 0.25 and 0.75 gHA/cm3) were scanned using the same protocol to calibrate images for tissue mineral density, TMD (Palacio-Mancheno et al., 2014).

Image Processing Approximately 4500 images per sample were reconstructed from X-ray projections using the back-projection reconstruction algorithm in NRecon software (Skyscan, v.1.6.1.1, SkyScan). Hounsfield unit (HU) and tissue mineral density calibration procedures

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were performed in CTAn software (CT Analyzer, v.1.6.1, SkyScan, Belgium). Four volumes of interest (VOIs) were selected from regions that contained water, air, 0.25 or 0.75 gHA/cm3. The mean grayscale index value from water and air was used to calibrate images in HU, and the mean grayscale index values from the 0.25 and 0.75 gHA/cm3 mineral rods were used to generate a calibration curve between the grayscale color in each pixel and the corresponding mineral density in gHA/cm3. After image density calibration, the separation (image segmentation) between mineralized and soft tissues in each scan was performed using a mean global threshold value. The threshold value was determined by analyzing the images using an edge detection algorithm (ImageJ v 1.37, National Institutes of Health). The tissue mineral density threshold obtained through this edge detection procedure was 0.40 gHA/cm3 (Souzanchi et al., 2012). More precise calibration methods have been proposed in bone biomechanics; these more precise calibration methods also account for the attenuation behavior of organics and water, rather than neglecting the latter two (see Blanchard et al., 2016 and references therein).

Global Measurements of Microarchitecture For each VOI, global architectural parameters were measured, including the porosity (f), trabecular number (Tb.N), trabecular thickness (Tb.Th), and trabecular separation (Tb.Sp). Since images were calibrated for mineral density, volumetric bone mineral density (vBMD) and TMD were obtained using built-in algorithms in CTAn software employing the guidelines for assessment of bone microarchitecture using mCT.

Directional Measurements of Microarchitecture The directional variation of pore orientation was calculated via the measurement of the mean intercept length (MIL) tensor (Harrigan and Mann, 1984) M using CTAn software. Then, values of the fabric tensor A eigenvalues F1, F2 and F3, were obtained by taking the inverse square root of the eigenvalues of MIL tensor M. Fabric components were normalized by dividing each one by the sum F1 þ F2 þ F3. The three eigenvectors of A represent the principal axes of material symmetry, which also correspond to the principal orientations of trabeculae.

Realignment of Images to Principal Axes of Symmetry A 12  12  12 mm3 cubic VOI was obtained at the center of the coronal plane (medial–lateral–inferior–superior directions in the calcaneum). The fabric tensor was measured on each VOI to determine the relative alignment between the principal directions of the fabric and the three orthogonal planes formed by the faces of the cubic sample. Images were rotated in DataViewer and the fabric was re-measured until the principal directions were coincident with the faces of the cubic sample. The VOI was digitally cropped to 7  7  7 mm3 at the center of each rotated cubic sample to perform micro-computational fluid dynamic (mCFD) analyses aligned to the principal directions of bone microarchitecture, see Fig. 1A. A 7  7  7 mm fluid volume was added using IrfanView (v.4.44) before and after the VOI on each of the three directions analyzed for all trabecular bone samples, see Fig. 1B–D, (Palacio-Mancheno et al., 2014).

Pore Diameter Measurements Because of the anisotropy of the porous media, it is unlikely that the cross section area and the equivalent pore diameter will be similar along the three interrogated directions in each sample. Therefore, we measured the average pore diameter in all the images (xy-, xz- and yz-planes) normal to the direction of the permeability test. The average pore diameter for each direction was obtained from 2D measurements of trabecular separation (Tb.Sp.) in CTAn.

Computational Fluid Dynamic Simulations The stack of images representing the VOI and the added fluid domains was imported in Mimics Research (version 19, Materialize, Leuven, Belgium), and the regions corresponding to bone and fluid were segmented apart using an automated thresholding algorithm (Fig. 1B–D). The 3D mask of the fluid comprehensive of the fluid domains and the fluid inside the trabecular bone was transferred from Mimics into 3Matic Research (version 11, Materialize, Leuven, Belgium) for the generation of a 3D adaptive volumetric mesh. The finite element mesh was then exported from 3Matic into Abaqus/CFD (v. 6.14 Simulia, Providence, RI), where the CFD problem was solved. The transient and steady state fluid dynamic behavior of the fluid inside the trabecular bone pores was investigated using an explicit time-marching integration approach. The fluid was assumed to be Newtonian, incompressible and viscous. The fluid flow behavior is governed by the Navier-Stokes equations for an incompressible, viscous fluid where g is the acceleration of gravity and v and p are respectively the flow velocity and the pressure, 
vv þ v,Vv ¼
Vp þ nDv þ rf g; V,v ¼ 0 rf (24) vt The CFD simulations were conducted using a static pressure difference of 0.1 Pa across the trabecular bone sample. The outer boundary of the fluid pore space was considered to be perfectly rigid, with no slip, and no penetration boundary wall conditions.

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Since the fluid saturating the cancellous bone structure in our experiments is water, the fluid mass density rf ¼ 1000 Kg/m3, bulk modulus Kf ¼ 2.25 GPa and viscosity n ¼ 10
3 Pa s. The time step increment was selected as 0.01 s, and the time integration parameter in Abaqus was set by default to q ¼ 0.5, producing a second order accurate semiimplicit method suitable for time accurate transient analysis. The total simulation time was 5 s in order to capture both transient and steady state response in all simulations. The CFD was performed using tetrahedral fluid elements (FC3D4). An adaptive algorithm (3Matic) was used to create the fluid mesh (De Paolis et al., 2017). CFD simulations were performed on a high-end workstation with 24 core Xeon CPUs 2.7GHz, graphic accelerator with 2700 GPU cores, and 512 GB of RAM. The CPU time for each simulation varied between 1–3 h per simulation.

Numerical Results Global and directional parameters of microarchitecture were measured on the trabecular bone samples analyzed in this study. The porosity of the three samples analyzed in this study was 76%, 84% and 94%, spanning a range of values that are characteristic of normal and osteoporotic bone. Trabecular thickness and separation are global measurements of the porous medium, and were found as Tb.Th. ¼ 182, 133, 125 mm, and Tb.Sp. ¼ 615, 398, 806 mm, respectively for the three bone samples. Directional measurements of permeability were obtained from mCFD numerical simulations on each of the trabecular bone samples tested along the three principal directions of microarchitecture. Plots of the pressure, the fluid velocity, and the fluid streamlines in the principal directions of these samples are shown in Fig. 2. The three principal components of the static permeability tensor were determined from averaged fluid velocity, fluid viscosity and pressure gradient applied on each test. Directional measurements of the pore architecture (i.e., pore diameter and fabric), fluid velocity and static permeability are reported as a function of the x, y and z direction in Table 1. In Fig. 3, the data for the fabric components, permeability. and pore size is reported for each direction on the three bone samples. It can be observed that the preferential alignment of pores, along direction z, has the largest magnitude for the fabric, permeability and equivalent diameter pore size, see Fig. 3A–C. Also, there is a strong linear trend between permeability and fabric, see Fig. 3D; however, we can observe that such linear correlation is

Fig. 2

(A–C) pressure, (D–F) fluid velocity, and (G–I) stream lines in the principal directions of three bones with 76%, 84% and 94% porosity.

Biomechanics j Impact of Microstructure on the Various Effective Properties of Bones Table 1

151

Fabric, static permeability and pore size as functions of directions x, y and z

Bone 1 f ¼ 76% Bone 2 f ¼ 84% Bone 3 f ¼ 94%

Direction

Pore diameter (mm)

Fabric

Fluid velocity (mm/s)

Static permeability (m2)

x y z x y z x y z

387 478 468 441 640 740 814 785 931

0.288 0.274 0.438 0.291 0.333 0.377 0.257 0.62 0.382

39.38 23.75 132.02 91.34 126.28 183.45 274.52 353.31 395.08

2.10E
09 1.26E
09 7.02E
09 5.37E
09 7.43E
09 10.79E
09 18.07E
09 23.25E
09 26.00E
09

Fig. 3 (A) Fabric, (B) permeability and (C) pore size reported as a function of directions x, y and z, and porosity. (D) Trend between permeability and fabric eigenvalues, and (E) trend between permeability and pore size for each bone sample.

dependent on the porosity of the sample, since the linear relationship for each sample seems shifted to the left for low porosities and towards the right for high porosities. Similarly, there is a clear inverse correlation between permeability and pore size, which is also modulated by the porosity in each sample, see Fig. 3E. The computed static permeability, the fabric tensor, the pore size, the porosity and k0 for each bone sample are then used to compute the corresponding values of k0K1, k0K2 and k0K3 in (15). With the computed pore size, the graphs of the dynamic permeability defined in (16) and the corresponding dynamic tortuosity are evaluated and shown in Fig. 4. In Fig. 5, the relationship between the average micro-velocities measured at the pore level in the mCFD simulations is compared with the apparent level fluid flux, corresponding to the product of the porosity times the apparent level macro-velocity (Darcys velocity).

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Fig. 4 Dynamic permeability defined in (16) as a function of frequency for the 76%, 84% and 94% porosity bone from left to right in the top row panels, and the corresponding tortuosity is also shown as a function of frequency in the three bottom panels for the same bone samples. From Johnson, D. and Myklebust, H. (1967). Learning disabilities: Educational principles and practices, New York, 1967, Grune and Stratton, Inc, p. 37.

Fig. 5

Relationship between pore micro-velocities and the product of the porosity and the apparent level macro-velocity (Darcys velocity).

Discussion The numerical results presented in this work suggest a strong trend between the fabric tensor and the static permeability tensor, which is different for each bone, possibly because their different porosity level. This numerical result supports the notion that anisotropy of the permeability tensor in trabecular bone is a consequence of the porosity and mainly the fabric tensor describing the directionality of the pores. Also, an inversely proportional trend was observed between pore size and permeability, suggesting that the pore shape is also affecting importantly the permeability of the porous media. The static permeability presented in Table 1 exhibits a variability (10
9–10
8 m2) that falls within the data in studies reporting experimental measurements of the permeability (10
12–10
8 m2) on cancellous bone (Abdalrahman et al., 2015; Accadbled et al., 2008; Baroud et al., 2004; Benalla et al., 2014; Birmingham et al., 2013; Bleiler et al., 2015; Cardoso et al., 2013; Coelho et al., 2011; Coughlin and Niebur, 2012; Cowin and Cardoso, 2014; Grimm and Williams, 1997; Kameo et al., 2016; Kohles et al., 2001; Kohles and Roberts, 2002; Kreipke and Niebur, 2017; Li et al., 1987; Metzger et al., 2015; Nauman et al., 1999; Pakula et al., 2008; Sandino et al., 2014; Souzanchi et al., 2013; Syahrom et al., 2015; Tae-Hong and Hong, 2000; Teo and Teoh, 2012). The variability of the intrinsic permeability in porous media is due to the dependence of the permeability on the porosity (Nauman et al., 1999; Grimm and Williams, 1997; Benalla et al., 2013) and the microstructure of the sample (Nauman et al., 1999; Kohles et al., 2001; Baroud et al., 2004; Kohles and Roberts, 2002). Further studies comprising a larger number of samples are needed to extend and generalize the results outlined by the examples treated here. Based on our experience with IRFs, the n-point correlation functions form a natural hierarchy for quantifying the difference between microstructures. The volume fraction (i.e., porosity) is related to only the 1-point correlation functions, which are low in the hierarchy in the sense that among the microstructures that share the same porosity, we can use their two-point correlation functions to further divide them into subgroups; the microstructures within each subgroup are more similar to each other than those with those outside the subgroup. This process can be continued by looking into higher order correlation functions. With the IRF

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approach, we expect the information of n-point correlation functions to be coded in the moments of the positive measure in the IRF. Based on the definition of permeability tensor from homogenization, we see that these moments depend not only on the microstructure but also on the spectral properties of the Stokes’ operator acting in the pore space. Similarly, the measure in the IRF for effective elasticity tensor is influenced by the microstructure and the linear elasticity operator. In other words, an effective property’s dependence on the microstructure is a consequence of the interaction between the underlying operator (e.g., Stokes operator for permeability) and the microstructure. To be more specific, for example, in the IRF of dielectric properties of two-phase composite materials, the nth moments of the measure is the n þ 1-fold convolution of the fundamental solution of the Laplacian operator and the characteristic function that defines the microstructure. We are currently working on generalizing the dielectric result to static permeability. Another future work from here is to extend the dehomgenization scheme in Miao-Jung Yvonne (2014) to anisotropic case described in (22) and compare the measure F
1qdG with the fabric tensor. Another future direction is to apply mCFD to compute the dynamic permeability from its definition (i.e., Stokes equation with time-harmonic pressure gradient), instead of the from (22). Using these data, we can then compute using the algorithm presented in Miao-Jung Yvonne (2014) to compute the rj and pj in (14) for each principal direction and use them to compute the moments of the measures in the IRFs; these moments will serve as characterization of the viscodynamic properties of the bone samples in addition to the infinite-frequency tortuosities.

Acknowledgments The work of Miao-jung Yvonne Ou was partially sponsored by the US National Science Foundation grant NSF-DMS-1413039. This work of Luis Cardoso and Annalisa De Paolis was supported by NSF (CMMI-1333560, MRI-0723027, and MRI-1229449), and National Institute of Health grant NIHDK103362.

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Multiphase Porous Media Models for Mechanics in Medicine: Applications to Transport Oncophysics and Diabetic Foot Pietro Mascheroni and Raffaella Santagiuliana, University of Padova, Padova, Italy Bernhard Schrefler, Technical University of Munich, Garching bei München, Germany; and Houston Methodist Research Institute, Houston, TX, United States © 2019 Elsevier Inc. All rights reserved.

Introduction Multiphase Porous Media Models for Tumor Growth Mathematical Model Results From the Model Tumor growing in vitro and ex vivo Mechanical compression of tumor spheroids Effects of interfacial tensions Melanoma growth and angiogenesis Diabetic Foot Mathematical Model Numerical Example for Gait Cycle Conclusions Further Reading

155 156 157 158 158 158 159 160 162 163 163 165 166

Glossary Conservation equations In general, this expression denotes a set of laws describing the balance of mass, momentum, and energy in a physical system. These equations are valid for a wide variety of systems and need to be closed by suitable constitutive relations to address a specific problem. Finite element method In the context of boundary value problems, the finite element method is a numerical technique for computing approximate solutions to partial differential equations. Mechanotransduction This term encompasses several mechanisms by which cells sense mechanical stimuli and convert them into biochemical signals, producing specific cellular responses. Partial differential equations These are mathematical equations that involve functions of multiple independent variables and their partial derivatives. Porous media A porous medium is a material containing a solid phase, denoted as the solid skeleton, and several interconnected pores, that is, voids. These pores are typically filled with one or more fluid phases, which could be in a liquid or gaseous state.

Introduction Biological systems are very complex and span many orders of magnitude, ranging from molecules, cells, organs, and up to the whole organism. These levels are tightly entangled, including multiple feedbacks and cross talks between the different parts. One way of deconvolution of this complexity is the use of modeling based on physical and chemical principles. Further, the knowledge of many quantities of interest is often not accessible through direct measurements, and some model is necessary to infer these from indirect measurements. Hence, models are compulsory for aiding our understanding of biological processes. This is now well understood, and a plethora of models have been developed. Here, we refer particularly to mathematical models, and, to restrict our field, we do not discuss purely data-driven approaches (black box models). Our focus is on the modeling of physical and chemical processes for which experimental data can be obtained. These models are usually cast in terms of partial differential equations. The field of application of this approach is extraordinarily large. As a few key examples, we mention cardiovascular diseases, lifetime assessment of stents, brain injuries, prosthetics, locomotion, injury biomechanics, and modeling of the intervertebral disk. Notably, in the last few years, several models for drug delivery and tumor growth have been introduced in the research field. Starting from the analyses for the design of injectable micro- and nano-devices for cancer therapy, additional methods have followed for the improvement of immunotherapies, imaging of tumor lesions, and design of biomedical devices for tumor removal. The mechanics of cellular processes, including mechanosensing and mechanotransduction, have been extensively described, and studies of cellular constitutive behavior through the use of microfluidic systems and other physical approaches are still an active field of research. The list of possible applications is far from being exhaustive, and new observations and discoveries appear almost every day. This endeavor

Encyclopedia of Biomedical Engineering, Volume 2

https://doi.org/10.1016/B978-0-12-801238-3.99925-2

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Biomechanics j Multiphase Porous Media Models for Mechanics in Medicine

brings together researchers in biology, medicine, engineering, physics, chemistry, material science, and applied mathematics, in which numerical modeling plays an important role. The successful development of such models is based on three major aspects: (i) the development of the nonlinear field theory of mechanics during the 1950s and 1960s; (ii) the introduction of the finite element (FE) method in the 1950s; and, subsequently, (iii) the rapid advances in computer technologies. Clearly, FE analysis is not the only one used in models dealing with mechanics in medicine, but is by far the most popular and acknowledged one. It is in fact by now the most used discretization method in applied sciences and mechanics. Proof of this is the sheer number of books that have been written or are still appearing, and the large number of general purpose codes based on this method. They address all fields of solid and fluid mechanics, heat transfer, electromagnetics, computational chemistry, and physics, and, now in increasing number, also interaction problems between these fields. This has made simulations possible that in the recent past were simply unthinkable. To such interaction problems belong also the porous media theories discussed in this article. Such theories consider a generally deformable solid skeleton and freely moving pore fluids, where the phases are fully connected. The nature of the fluid can be manifold, ranging from water and air to biological fluids and even cell populations which also can be treated as particular fluid phases. The kinematic quantities are a solid displacement vector, which tracks the movement of the porous solid with respect to a reference configuration, and specific discharge vectors describing the motion of the fluid phases relative to the solid. The specific discharge is defined as the rate of fluid volume crossing a unit area of porous solid. The interaction between the solid phase and the fluids is not the only one considered. Interactions with thermal, chemical, and electromagnetic fields are also currently taken into account. It is clear that the whole process is usually time dependent, and this aspect has to be taken into account in the development of the computational implementations. The governing equations for such media are now generally obtained through averaging theories such as the thermodynamically constrained averaging theory (TCAT) and are then discretized by means of the FE method. Indeed, the TCAT framework provides a rigorous methodology for developing multiphase, continuum models at any scale of interest. Larger scale variables are explicitly defined in terms of smaller variables at smaller scales. At the microscale, classical local conservation equations and thermodynamic expressions can be written. However, as the domains of many problems are too large, or the phase distributions are too complex, it is usually impossible to model the phenomena of interest directly on such a small scale. In fact, simulations would be possible only for very small domains. To overcome such problems, many porous media models are formulated at a larger scale, called the macroscale. Standard continuum mechanics techniques for the formulation of these models rely on a direct approach, in which the conservation equations are written at the larger scale. Usually, rational thermodynamic assumptions are enforced to obtain closure relations. Unfortunately, the use of such methods may fail to retain a connection between larger scale variables and their microscale precursors. Through a set of averaging theorems applied to conservation and thermodynamic equations at the small scale, TCAT avoids these shortcomings and leads to equations that are both thermodynamically and physically consistent. The theory consistently transforms microscale conservation and thermodynamic equations to the macroscale and converts averages of microscale derivatives into derivatives of macroscale average quantities. Note that a realistic description of a multiphase system must include dynamic conservation and thermodynamic equations for all the phases and for all the points in which they interact. Although TCAT has been primarily employed in hydrology, it impacts also biomechanical modeling, since the underlying physics and mathematics are related. Once the governing equations are obtained at the macroscale, their weak form is derived by means of the standard Galerkin procedure and is then discretized in space by means of the FE method. For the particular type of problems that we discuss, integration in the time domain is carried out by the finite difference method adopting the q-Wilson procedure. Within each time step, the equations are linearized by the Newton–Raphson algorithm. The system of equations has been implemented in the threedimensional FE code CAST3M (http://www-cast3m.cea.fr) of the French Atomic Energy Commission. A staggered scheme is adopted, with iterations within each time step to preserve the coupled nature of the system. In the following, we discuss two applications of these procedures. We report on a model for tumor growth and its interactions with the surrounding environment. Then, we show results for the case of ulceration of the diabetic foot.

Multiphase Porous Media Models for Tumor Growth Cancer is a collection of related diseases in which some cells of the body start to divide without stopping and eventually spread into their surroundings. In normal tissues, healthy cells (HCs) grow and divide according to the needs of the organism. When cells grow old or become damaged, they are eliminated, and new cells take their place. However, when cancer develops, this carefully controlled process breaks down. As multiple alterations accumulate, old or damaged cells survive when they should die, and new cells form even if they are not needed. These extra cells divide uncontrolled and may result in abnormal masses called tumors. Malignant cancerous tumors can spread into surrounding tissues, displacing the neighboring HCs. In addition, as the tumor develops, some cancer cells are able to detach from the original tumor mass and travel to distant organs in the body through the circulation. Eventually, these cancerous cells may form metastases, that is, new tumors far from the original formation. Nowadays, cancer figures among the leading causes of mortality worldwide, with approximately 14 million new cases and 8.2 million cancer-related deaths in 2012. Despite new technological advances and significant efforts (projected national expenditures for cancer care are expected to total nearly $157 billion in 2020 just in the United States), the initial hopes put in the war on cancer have been largely disillusioned. Since the 1950s, indeed, age-adjusted cancer mortality rates have decreased by only 11%. Prevention, screening, and treatment success with some cancers have saved millions of lives, but the prognosis for many with metastatic cancer remains still as gloomy as it was several years ago.

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Looking at these premises, researchers from quantitative disciplines such as physicists, mathematicians, and engineers have contributed to cancer research over the last few years. One contribution results from discoveries and technological developments, which have led to advances in medical imaging and radiation therapy for the diagnosis and treatment of tumors. A second important contribution is brought by bioinformatics, providing the tools to handle large datasets of genome sequences, gene expression patterns, and cell-signaling networks. Finally, a third contribution has recently gained interest. This direction involves a more quantitative investigation of the physical processes underlying the evolution of a tumor. Indeed, it is now recognized that the physicochemical properties of several biological barriers are responsible for the transport of cells, particles, and molecules across the tissues. Remarkably, this transport and its deregulation play a predominant role in cancer physics. Tissue invasion can be considered as mass transport deregulation at the interface between the cell and the microenvironment; metastasis is a deregulation of local and distant cellular transport at the scale of the organism; tumor angiogenesis completely alters mass and fluid exchange across the microcirculation; abnormalities in the signaling pathways that accompany the evasion of apoptosis, growth signal dependence, and growth inhibitory messages from the microenvironment are also disruptions in molecular transport, since molecular signaling directly depends on the transport of signaling molecules. In addition, several aspects related to transport control the delivery of therapeutic agents, such as chemotherapeutics or molecularly targeted therapies. To be effective, these substances must pass through different and heterogeneous tumor and healthy compartments (e.g., vascular, stroma) with distinct physical properties. In fact, drug delivery is an extremely complex process involving different spatial and temporal scales. It involves a series of events that take place over several levels, ranging from the organism to the intercellular environment. Throughout these levels, different phenomena may act as transport barriers, possibly contributing to poor survival rates in cancer therapy. As a first example, the mononuclear phagocyte system belongs to these barriers, removing foreign substances from the body such as nanoparticles (NPs) in the blood plasma. A second obstacle is given by the tumor neovasculature, resulting from tumor angiogenesis. The new vascular network is tortuous and distorted and shows large fenestrations, leading to poor perfusion of the tumor tissue, chaotic blood flow, and uneven supply of nutrients. However, these fenestrations also allow for the enhanced permeability and retention effect (EPR), discovered in the 1980s by Maeda and colleagues. Indeed, the increased accumulation of long-circulating macromolecules and NP by extravasation through the tumor blood vessels has been carefully characterized in the last few years. Nevertheless, the EPR effect may still be hindered because of the altered gradients of the interstitial fluid (IF) pressure in the local tumor environment. Furthermore, limited drainage of IF (due to the lack of functional lymphatics and extensive fibrosis in the tumor) brings about pressure rises, reducing the extravasation of therapeutic molecules through advection. Hence, poor perfusion through the local environment hinders the diffusion of therapeutic agents in the tumor interstitium and affects ultimately the cellular uptake of NPs. In general, NP uptake by the cells follows endocytosis; only in a second step the payload of the NP might be released into the cytoplasm. The concept of biological barriers has brought a new understanding of how transport can modulate cancer biology and efficacy of therapies. Indeed, several studies have confirmed that altered transport plays a crucial role in cancer and drug delivery, including the phenomenon of resistance. Transport oncophysics views hence cancer as a disease of multiscale mass transport deregulation, involving distinct biological barriers at different levels. Computational transport oncophysics provides a set of computational tools that, together with imaging, data analysis, and quantification, can contribute to rationalize the development of the tumor and optimize the delivery of therapies. This framework should complement classical tools used to study pharmacokinetic and efficacy relations, with the final aim of creating novel precision tools to rationally tailor individual treatments.

Mathematical Model We concentrate here on mathematical models for tumor growth. Such models describe the behavior of several tumor constituents, such as tumor cells (TCs), both viable and necrotic, HCs, extracellular matrix (ECM), IF, neovasculature and co-opted blood vessels, nutrients, and waste products. A great variety of models have been developed, often based on different starting assumptions and fields of application. Some of the components mentioned above may be neglected, resulting in different types of models accounting for diffusion, single-phase flow, and multiphase flow with or without a solid phase. Three major classes of models can be distinguished as evidenced in several review papers: discrete, continuum, and hybrid models. Discrete models follow the fate of a single cell, or a cohort of cells, over time. As such, this modeling framework is not able to capture aspects of tissue mechanics, nor are the modeled subdomains representative of the whole tumor. However, discrete models are suitable for explaining cell-to-cell cross signaling and cell response to therapeutic molecules. On the other hand, in continuum models, cancerous tissues are regarded as domains composed of multiple fluid and solid phases interacting one with the other. Partial differential equations based on conservation laws and thermodynamics describe the spatiotemporal evolution of the system, but no direct information is provided at the single-cell level. Finally, hybrid models incorporate different aspects of discrete and continuum models, according to the problem of interest. For instance, cells may be represented individually, and the IF may be described as a continuum. Within continuum models, we encounter biphasic solid-fluid models in which the ECM is lumped with the cells and the resulting porous scaffold is permeated by the IF. An alternative is multiphase flow models, sometimes in a deforming porous material. Among this last kind, we have developed a general multiphase flow model in an ECM, considered as a deforming porous solid which may undergo remodeling. This model comprises three fluid phases, namely TCs, divided into living and necrotic cells, HCs, and IF. The IF transports chemical species such as tumor angiogenic factors (TAFs), nutrients, and therapeutic agents. Transport of these substances within extravascular space takes place by convection and diffusion. Co-opted blood vessels are included as line elements, across which blood exchanges nutrients and therapeutic agents with the IF. Angiogenesis is represented by blood vessel density (density of newly created endothelial cells). The model accounts not only for growth and necrosis, but also for migration of cells

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through the ECM, for different stiffness of the cell population with respect to the ECM, buildup of cortical tension between healthy and tumor tissues, and possible invasion of the tumor tissue by the healthy tissue or vice versa (mediated by these cortical tensions). Further, it allows for modeling lysis and connected lymphatic outflow from the tumor, adhesion of the cells to their ECM, as well as adhesion among cells (through dynamic viscosities) and possible detachment. The mathematical model includes the mass balance equations of the four main constituents, namely the ECM, TCs, HCs, and IF. Also, mass balance of the different species within the constituents is needed, resulting in advection–diffusion–reaction equations for the transported substances in the IF (such as the nutrients). After that, the linear momentum balance equations of the phases are stated, including the one for the solid. The model is finally closed with the necessary constitutive equations. In particular, the ECM is considered to be Green-elastic or elasto-visco-plastic. The model has been extensively validated with respect to experiments either from literature or carried out at the Houston Methodist Research Institute.

Results From the Model Tumor growing in vitro and ex vivo We show now a few results from the model, starting from the case of a tumor spheroid growing in vitro and ex vivo. Tumor spheroids are spherical aggregates of TCs that can be grown in the laboratory, with a careful control of the culture conditions. For tumor spheroids grown in vitro, TCs deposit their own ECM during the tumor evolution. On the other hand, in the ex vivo case, the TCs are seeded in a de-cellularized ECM to mimic the in vivo environment. This has been carried out successfully by Mishra and colleagues, for an ex vivo 3D lung model in which it was possible to grow perfusable lung nodules. Fig. 1 shows the evolution of the solid volume fraction for these two cases. In particular, the case of ECM deposited by TCs is shown on the top of Fig. 1, whereas the case of a remodeling ECM scaffold is displayed at the bottom of the image. In the first case (spheroid grown in vitro), the initial solid volume fraction is set to zero, since there is no solid phase. At the final stage, the solid volume fraction differs from zero, where the TCs have grown and have deposited their ECM. In the second case (spheroid grown ex vivo), the initial and final average solid volume fractions are fixed at 0.2, because the ECM scaffold is already present as a de-cellularized matrix.

Mechanical compression of tumor spheroids In the second example, a reduced computational model is validated against data from tumor spheroid cultures. U87-MG cells, from a human glioblastoma cell line, are cultured with a standard protocol. Cells are seeded at different initial numbers (1000, 5000, 10,000) and rapidly form spheroids suspended in standard culture medium. The evolution of the spheroid radii is then recorded over time via optical microscopy, and the results are shown in Fig. 2. Here, points are experimental data, and error bars are the

Fig. 1 Solid volume fraction distribution at initial (A) and final (B) stages of a MTS growing in a ECM deposited by TCs and solid volume fraction distribution at initial (C) and final (D) stages of a MTS growing in a remodeling ECM scaffold. Reprinted with permission under the CC-BY Attribution License from Santagiuliana, R., Stigliano, C., Mascheroni, P. et al. (2015). The role of cell lysis and matrix deposition in tumor growth modeling. Advanced Modeling and Simulation in Engineering Sciences, 2(1), 19. (http://creativecommons.org/licenses/).

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Fig. 2 Growth curves recorded from the free growth experiments. Each curve represents a different initial condition in terms of seeded cells. In the experiments, N  4 spheroids are considered for each condition. Points are experimental data, and error bars are the standard deviations of the measurements; solid lines are the results of fits with the mathematical model. Reproduced from Mascheroni, P., Stigliano, C., Carfagna, M. et al. (2016). Predicting the growth of glioblastoma multiforme spheroids using a multiphase porous media model. Biomechanics and Modeling in Mechanobiology, 15(5), 1215–28. © Springer-Verlag Berlin Heidelberg 2016. With permission of Springer.

Fig. 3 Optical images of U-87 MG spheroids grown under the effect of the dextran solutions. The first row shows the control experiments and the second row a spheroid under the highest compression (10 kPa). The scale bar is 200 mm, and the initial seeding is 5000 tumor cells. Reproduced from Mascheroni, P., Stigliano, C., Carfagna, M. et al. (2016). Predicting the growth of glioblastoma multiforme spheroids using a multiphase porous media model. Biomechanics and Modeling in Mechanobiology, 15(5), 1215–28. © Springer-Verlag Berlin Heidelberg 2016. With permission of Springer.

standard deviations of the measurements. The solid lines in Fig. 2 are the results of fits with the mathematical model. The model is able to reproduce the results in the experiments, in which the spheroids reach a steady state after 20 days from cell seeding. As the spheroids grow freely in the culture medium, the only mechanism to stop cell proliferation is nutrient deprivation. Lack of nutrients in the inner regions of the spheroids provides a sufficient explanation for growth saturation and the following existence of an asymptotic spheroid radius. After these preliminary experiments, the application of a constant mechanical stress on the surface of the growing spheroids is then investigated. Following a technique developed by Montel and colleagues, dextran is added to the cell culture medium producing an osmotic pressure on the outermost layer of cells located on the spheroid surface. Three pressure conditions are explored, namely 1, 5, and 10 kPa, plus a control experiment with no external pressure. The growth of the spheroids is followed for 18 days after the addition of dextran. Fig. 3 shows optical images of sample spheroids referring to the control and to the most compressed condition at different time instants. The comparison between experimental data and numerical values is shown in Fig. 4. Points with error bars are experimental values while solid lines are the results of fits with the mathematical model. The model results are in good agreement with the experimental values for all the different external mechanical pressures.

Effects of interfacial tensions In a recent work, the computational model has been enriched to account for the features of fully developed three-phase flow, capturing the interactions between the different cell populations. This required the introduction of suitable constitutive relationships, describing the pressure differences between the cell phases and the IF. As a consequence, the interfacial tensions between the different fluids appear explicitly. Note that these relationships allow for a more realistic modeling of adhesion between the cells and invasion of different tissue compartments. Fig. 5 shows the last time step of a simulation in which a spherical tumor grows within an external host tissue. The three panels display the volume fractions of the cellular phases over the radius of the spherical domain, for three different TC-HC interfacial tensions. As shown in Fig. 5A, a rather high interfacial tension between TCs and HCs leads to a significant spreading of the original tumor. Indeed, even if the tumor region was initially occupied only by TCs, their original distribution is not sustained and no domain with only TCs is obtained. On the other hand, Fig. 5B refers to the case of a medium interfacial tension

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Fig. 4 Comparison between experimental data (dots) and numerical results (solid lines) for the compression experiments. Error bars represent the standard deviations of the measurements. The model is able to reproduce the experimental results for all the different compression regimes. For each condition, N ¼ 5 spheroids are considered. Reproduced from Mascheroni, P., Stigliano, C., Carfagna, M. et al. (2016). Predicting the growth of glioblastoma multiforme spheroids using a multiphase porous media model. Biomechanics and Modeling in Mechanobiology, 15(5), 1215–28. © Springer-Verlag Berlin Heidelberg 2016. With permission of Springer.

Fig. 5 Effect of different interfacial tensions between the cell populations. The three figures refer to the cases of high (A), medium (B), and zero (C) interfacial tension between the TCs and the HCs. The model provides different profiles for the host tissue invasion, depending on the surface interactions between the cells. Reproduced from Sciumè, G., Gray, W. G., Hussain, F. et al. (2016). Three phase flow dynamics in tumor growth. Computational Mechanics, 53(3), 465–484. © Springer-Verlag Berlin Heidelberg 2013. With permission of Springer.

between TCs and HCs. The interactions between the two cell populations support the lateral displacement of the healthy tissue, favoring a rapid growth of the malignant mass. In both the cases, a residual volume fraction of HCs remains present within the malignant mass. However, the model predicts a higher density of TCs for lower values of the TC-HC interfacial tension. Finally, in the limit case with zero interfacial tension between TCs and HCs (Fig. 5C), we observe the complete displacement of the HCs by the TCs. In this case, it is possible to describe partial displacement of the host tissue by the tumor only in two ways: (i) by considering different values for the dynamic viscosity of the two cell populations and (ii) by accounting explicitly for a heterogeneous adhesion to the ECM.

Melanoma growth and angiogenesis As last application of the model, we show results for the growth of a melanoma in the presence of angiogenesis. Here, the endothelial cell density is assumed proportional to the density of new vessels. The outer structure of skin is layered and three compartments can be evidenced: (i) the epidermis, an outer epithelium of stratified cells; (ii) the dermis, an intermediate cushion of vascularized connective tissue; and (iii) the hypodermis, the lowermost layer made of loose tissue and adipose cells. The dermis is separated from the epidermis by the basement membrane, a tough sheet of ECM. Two well-defined clinical stages characterize the progression of a melanoma. First, a radial expansion in the epidermis occurs, which may be followed in a second step by vertical growth. In this second stage, the tumor grows perpendicularly to the skin surface and may eventually penetrate the basement membrane. Angiogenesis occurs during this latter phase. Blood vessels are here assumed at the base of the dermis, as shown in Fig. 6 which depicts the geometry of the simulated domain. Fig. 7 shows the resulting volume fraction of TCs after 10 and 20 days. At the beginning, the growing TCs deform the ECM and create a small bulge on the skin surface. Once the base membrane is reached, the growth pattern changes into a more complex configuration. TCs consume oxygen due to cell proliferation and their own metabolism. Oxygen consumption leads to a decrease of its mass fraction in the tumor area, as depicted in Fig. 8. As the cancer develops, a region of necrotic cells forms in the tumor interior. Before undergoing necrosis, living TCs are subjected to hypoxic conditions, in which the poor mass fraction of oxygen hinders cell replication. Living TCs under hypoxia produce TAFs that diffuse into the domain. Fig. 9 shows the effects of TAF diffusion after 10 and 20 days. Then, diffusion of endothelial cells sensing the chemotactic gradient of TAFs is depicted in Fig. 10. After 10 days, the distribution of the mass fraction of endothelial cells shows a higher concentration near the tumor, where the TAF concentration is larger. This behavior is highlighted in the plot at 20 days.

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Fig. 6 Growth of a melanoma in the presence of tumor angiogenesis. Skin structure and geometry of the simulated case. Reprinted from Santagiuliana, R., Ferrari, M., Schrefler, B. A. (2016). Simulation of angiogenesis in a multiphase tumor growth model. Computer Methods in Applied Mechanics and Engineering, 304, 197–216. Copyright 2016, with permission from Elsevier.

Fig. 7 Volume fraction of TCs after 10 (left) and 20 days (right). Reprinted from Santagiuliana, R., Ferrari, M., Schrefler, B. A. (2016). Simulation of angiogenesis in a multiphase tumor growth model. Computer Methods in Applied Mechanics and Engineering, 304, 197–216. Copyright 2016, with permission from Elsevier.

Fig. 8 Mass fraction of oxygen after 10 (left) and 20 days (right). Reprinted from Santagiuliana, R., Ferrari, M., Schrefler, B. A. (2016). Simulation of angiogenesis in a multiphase tumor growth model. Computer Methods in Applied Mechanics and Engineering, 304, 197–216. Copyright 2016, with permission from Elsevier.

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Fig. 9 Mass fraction of TAF after 10 (left) and 20 days (right). Reprinted from Santagiuliana, R., Ferrari, M., Schrefler, B. A. (2016). Simulation of angiogenesis in a multiphase tumor growth model. Computer Methods in Applied Mechanics and Engineering, 304, 197–216. Copyright 2016, with permission from Elsevier.

Fig. 10 Mass fraction of endothelial cells after 10 (left) and 20 days (right). Reprinted from Santagiuliana, R., Ferrari, M., Schrefler, B. A. (2016). Simulation of angiogenesis in a multiphase tumor growth model. Computer Methods in Applied Mechanics and Engineering, 304, 197–216. Copyright 2016, with permission from Elsevier.

Diabetic Foot A severe complication of the diabetes mellitus is the pathology of diabetic foot. During the disease, the high levels of blood glucose in diabetic patients can damage the nerves and blood vessels. The diabetic neuropathy, the peripheral nerve dysfunction associated with diabetes, causes patients to have a reduced ability to feel pain. The loss of sensation leads to repetitive minor injuries from internal (calluses, nails, foot deformities) or external causes (shoes, burns, foreign bodies) that remain undiscovered for a long while and may eventually lead to ulcers and infections. Furthermore, associated damage to the blood vessels can also induce low levels of blood and oxygen in the feet. This may cause vasoconstriction and decreased sweating, resulting in a loss of skin integrity, and providing a site vulnerable to microbial infection. As a result, strong difficulties to feet healing arise, with serious cases of ulcer infections that may lead to amputations. Diabetic neuropathy is the common factor in almost 90% of diabetic foot ulcers. The prevention of diabetic foot is crucial, considering the negative impact on a patient’s quality of life and the associated economic burden on the healthcare system. The management of diabetic foot ulceration involves a multidisciplinary approach. To avoid ulcers and prevent amputation, some diabetic treatments are adopted as offloading the wound by using appropriate therapeutic footwear, debridement when necessary, antibiotic, optimal control of blood glucose, and evaluation and correction of peripheral arterial insufficiency. The neuropathy causes loss of protective sensation and loss of coordination of muscle groups in the foot and leg, muscle weakness, atrophy, and paresis. These effects increase mechanical stresses during ambulation. The excessive pressure in the foot tissue may be measured, but due to the differences between the plantar pressure measurement devices, a universal threshold is not set for ulceration. Furthermore, ulceration often starts at the interface between bone angularity and plantar tissue, probably due to internal stresses, which are not directly measurable with these experimental tools. In this context, an approach based on FE can be helpful, because it can predict the stress on the plantar tissue. FE models of the foot have become even more faithfully realistic, including the actual geometry of the ligaments, muscles, and other foot substructures. As clearly shown in several studies from the recent literature, enhancement of FE models can aid the understanding of this

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complex problem. However, most models have treated the plantar tissue, which is directly concerned with the foot ulceration, as a hyperelastic material, whereas the other constituents (ligaments, bones, cartilage, and plantar fascia) have been considered as linearly elastic. This simplification has led to a poor predictive capability of the models. Moreover, in recent models, the duration of the gait cycle has been taken into account by means of viscous constitutive laws for the plantar soft tissue. The model presented below can simulate one or more gait cycles and predict the associated plantar pressure. It can integrate experimental data of patient-specific foot kinematics into the modeling process, improving agreement between the computed and measured pressures. The model is time dependent and includes information on the real microstructure of the plantar tissue. The latter is modeled as a porous medium, in which cells and ECM constitute a solid skeleton that is permeated by IF.

Mathematical Model The choice of adopting porous media mechanics for modeling the plantar tissue did not come about by chance. In fact, as observed ultrasonographically, the plantar tissue consists of micro- and macro-chambers, filled by IF. Hence, theories developed for porous materials can provide a realistic description of the mechanical behavior of this tissue: stress transfer from fluid to solid phase and vice versa occurring inside the tissue can be taken into account, thus resulting in a global viscoelastic response of the material. The model is a biphasic system in the large strain regime. The tissue cells and the ECM form the elastic porous solid skeleton. IF is the fluid phase that completely fills the pore space. Transport of nutrients and possible drug delivery within the microvasculature can be introduced as diffusion of chemical species within the IF. The diffusion coefficient can be estimated from the real degree of vascularization of the zone of interest. In addition, the model can account for necrosis of the plantar tissue, depending on the stress level and vasculopathies. The presence of the IF in the pores allows mimicking the real global viscoelastic behavior of the plantar soft tissue. As in the tumor growth model presented before, the evolution of the solid and fluid phases is regulated by mass and momentum balance equations.

Numerical Example for Gait Cycle To mimic the evolution of the pressure in the foot by FE analysis, the gait cycle is modeled from the end of the contact phase (when the heel is the first area of the foot in contact with the ground) to the active propulsion phase (when the heel lifts from the support side and ends with opposite-side heel strike). Furthermore, the global forces measured experimentally by a force platform device, translated and applied with the opposite sign to the ankle, are the input load for the numerical model. Applying a patient-specific load history is essential because not only the foot morphology and tissue properties change because of diabetic diseases, but gait alterations may also occur. For example, it has been observed that patients with diabetes commonly walk slower, tend to take shorter steps with a wider base of support, and demonstrate a longer double support time. In the next example, a healthy foot is modeled and four gait cycles are simulated. This numerical example is useful to show the potentialities of the model and the interesting aspects that can be captured by modeling the plantar tissue as a fluid saturated porous medium. Fig. 11 shows the geometry of the problem. For a real case, the geometry of a patient’s foot (obtained from

Fig. 11 Geometry and load conditions for the case of a healthy foot. Reproduced with permission from Sciumè, G., Boso, D.P., Gray, W.G., Cobelli, C., Schrefler, B.A. (2014). A two-phase model of plantar tissue: a step toward prediction of diabetic foot ulceration. International Journal for Numerical Methods in Biomedical Engineering, 30:1153–1169. © 2014 John Wiley & Sons, Ltd.

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Fig. 12 (A) The black line is the input vertical force (P) and the gray line is the torque (C$d) (figure readapted from Natali, A. N., Forestiero, A., Carniel, E. L., Pavan, P. G., Dal Zovo, C. (2010). Investigation of foot plantar pressure: experimental and numerical analysis. Medical and Biological Engineering and Computing, 48:1167–1174). (B and C) Vertical stresses, SMYY, and interstitial fluid pressures, PF, computed in the points T1 and T2 (these points are represented in Fig. 11). In (B and C), the markers (disks) represent solutions obtained with ABAQUS (using a monolithic solver) while the continuous lines are those obtained with Cast3M using the presented staggered procedure. Reproduced with permission from Sciumè, G., Boso, D.P., Gray, W.G., Cobelli, C., Schrefler, B.A. (2014). A two-phase model of plantar tissue: a step toward prediction of diabetic foot ulceration. International Journal for Numerical Methods in Biomedical Engineering, 30:1153–1169. © 2014 John Wiley & Sons, Ltd.

an MR image) and patient-specific parameters must be used to obtain predictive results. Moreover, in vivo measurements of displacements and rotations of the bone segments can be useful to confirm that the kinematics of the foot obtained numerically is sound. Fig. 12A shows the evolution with time of the vertical force and of the torque. This can be computed from the experimental position of the plantar pressure resultant during the gait analysis and is applied to the foot as shown in Fig. 11. Fig. 12B and C displays the evolution of the vertical total stress and of the IF pressure, during the four gait cycles for the points T1 and T2 represented in Fig. 11. Note that the IF pressure reverses sign during a gait cycle, which would imply flow reversal. The intrinsic permeability has a dominant role because it governs the efficiency of the tissue in absorbing quite important loads; the loads initially supported by the IF are then progressively transferred to the solid phase. Fig. 13 shows the vertical total stress in the tissue and in the supporting surface at the same time. We can see how in the first part of the midstance phase (4.0 s), the load is mostly supported by the heel, while at the end of the midstance phase (4.4 s), it is mostly supported by the forefoot.

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Fig. 13 Vertical stress during the last gait cycle at 4.0, 4.2, and 4.4 s. Reproduced with permission from Sciumè, G., Boso, D.P., Gray, W.G., Cobelli, C., Schrefler, B.A. (2014). A two-phase model of plantar tissue: a step toward prediction of diabetic foot ulceration. International Journal for Numerical Methods in Biomedical Engineering, 30:1153–1169. © 2014 John Wiley & Sons, Ltd.

Conclusions For a long time, the mechanics community has focused on complex engineering problems in a wide set of fields, such as transportation, material behavior, and defense. A broad collection of theoretical, numerical, and experimental tools have been validated, following prior development and testing. In the last few years, this collection of techniques and methods has been exploited to answer open questions in medicine. With the tools from mechanics, intricate problems involving many degrees of freedom have been solved, and answers are on their way to the clinic. In this work, we reported on computational models developed in the framework of porous media mechanics and applied to biomedical topics of significant relevance. We started from the case of tumor growth, describing the evolution of the tumor mass under different conditions. We analyzed the evolution of tumor spheroids, first as freely growing in the culture medium and then subjected to an external compression. After that, we focused on the coupled dynamics of tumor and host cells, analyzing the effects of different surface interactions between the two cellular species. Finally, the case of a melanoma was investigated, with a few remarks about tumor angiogenesis and oxygen consumption. The second part of the work was devoted to the study of the diabetic foot. Contrary to the usual approach, which employs the hypothesis of hyperelasticity, we modeled the plantar tissue as a porous material. This enabled a more realistic description of the tissue mechanical properties, which were directly influenced by the time evolution of the loading conditions. The vertical stress and fluid pore

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pressure were analyzed at different times. Faced with the surge in cancer incidence and occurrence of diabetic pathologies, the approaches presented here should stimulate novel therapeutic strategies and treatment optimizations, with the aim of improving the prognosis, outcome of intervention, and quality of life for patients.

Further Reading Alexiadou, K., & Doupis, J. (2012). Management of diabetic foot ulcers. Diabetes Therapy, 3, 4. Araujo, R. P., & McElwain, D. L. S. (2004). A history of the study of solid tumour growth: the contribution of mathematical modelling. Bulletin of Mathematical Biology, 66(5), 1039–1091. Bellomo, N., De Angelis, E., & Preziosi, L. (2003). Multiscale modeling and mathematical problems related to tumor evolution and medical therapy. Journal of Theoretical Medicine, 5(2), 111–136. Blanco, E., Shen, H., & Ferrari, M. (2015). Principles of nanoparticle design for overcoming biological barriers to drug delivery. Nature Biotechnology, 33(9), 941–951. Byrne, H. M. (2010). Dissecting cancer through mathematics: from the cell to the animal model. Nature Reviews. Cancer, 10(3), 221–230. Carson, E., & Cobelli, C. (2014). Modelling methodology for physiology and medicine (2nd edn). London: Elsevier. Chen, W. M., Lee, T., Lee, P. V. S., Lee, J. W., & Lee, S. J. (2010). Effects of internal stress concentrations in plantar soft-tissue – a preliminary three-dimensional finite element analysis. Medical Engineering & Physics, 32, 324–331. Cheng, A. H.-D. (2016). Poroelasticity. Switzerland: Springer International Publishing. Ferrari, M. (2010). Frontiers in cancer nanomedicine: directing mass transport through biological barriers. Trends in Biotechnology, 28(4), 181–188. Gefen, A., Megido-Ravid, M., & Itzchak, Y. (2001). In vivo biomechanical behavior of the human heel pad during the stance phase of gait. Journal of Biomechanics, 34, 1661–1665. Gray, W. G., & Miller, C. T. (2014). Introduction to the thermodynamically constrained averaging theory for porous medium systems. Switzerland: Springer International Publishing. Hanahan, D., & Weinberg, R. A. (2011). Hallmarks of cancer: the next generation. Cell, 144(5), 646–674. Jain, R. K., & Stylianopoulos, T. (2010). Delivering nanomedicine to solid tumors. Nature Reviews. Clinical Oncology, 7(11), 653–664. Lewis, R. W., & Schrefler, B. A. (1998). The finite element method in the static and dynamic deformation and consolidation in porous media. Chichester: Wiley. Mascheroni, P., Stigliano, C., Carfagna, M., et al. (2016). Predicting the growth of glioblastoma multiforme spheroids using a multiphase porous media model. Biomechanics and Modeling in Mechanobiology, 15(5), 1215–1228. Michor, F., & Beal, K. (2015). Improving cancer treatment via mathematical modeling: surmounting the challenges is worth the effort. Cell, 163(5), 1059–1063. Michor, F., Liphardt, J., Ferrari, M., & Widom, J. (2011). What does physics have to do with cancer? Nature Reviews. Cancer, 11(9), 657–670. Moore, N. M., Kuhn, N. Z., Hanlon, S. E., Lee, J. S. H., & Nagahara, L. A. (2011). De-convoluting cancer’s complexity: using a “physical sciences lens” to provide a different (clearer) perspective of cancer. Physical Biology, 8(1), 10302. Pai, S., & Ledoux, W. R. (2011). The quasi-linear viscoelastic properties of diabetic and non-diabetic plantar soft tissue. Annals of Biomedical Engineering, 39(5), 1517–1527. Quaranta, V., Weaver, A. M., Cummings, P. T., & Anderson, A. R. A. (2005). Mathematical modeling of cancer: the future of prognosis and treatment. Clinica Chimica Acta, 357(2), 173–179. Risler, T. (2015). Focus on the physics of cancer. New Journal of Physics, 17(5), 55011. Santagiuliana, R., Stigliano, C., Mascheroni, P., et al. (2015). The role of cell lysis and matrix deposition in tumor growth modeling. Advanced Modeling and Simulation in Engineering Sciences, 2(1), 19. Santagiuliana, R., Ferrari, M., & Schrefler, B. A. (2016). Simulation of angiogenesis in a multiphase tumor growth model. Computer Methods in Applied Mechanics and Engineering, 304, 197–216. Sciumè, G., Gray, W. G., Ferrari, M., Decuzzi, P., & Schrefler, B. A. (2013). On computational modeling in tumor growth. Archives of Computational Methods in Engineering, 20(4), 327–352. Sciumè, G., Gray, W. G., Hussain, F., et al. (2013a). Three phase flow dynamics in tumor growth. Computational Mechanics, 53(3), 465–484. Sciumè, G., Shelton, S., Gray, W., et al. (2013b). A multiphase model for three-dimensional tumor growth. New Journal of Physics, 15(1), 15005. Sciumè, G., Ferrari, M., & Schrefler, B. A. (2014). Saturation–pressure relationships for two- and three-phase flow analogies for soft matter. Mechanics Research Communications, 62, 132–137. Sciumè, G., Boso, D. P., Gray, W. G., Cobelli, C., & Schrefler, B. A. (2014a). A two-phase model of plantar tissue: a step toward prediction of diabetic foot ulceration. International Journal for Numerical Methods in Biomedical Engineering, 30, 1153–1169. Sciumè, G., Santagiuliana, R., Ferrari, M., Decuzzi, P., & Schrefler, B. A. (2014b). A tumor growth model with deformable ECM. Physical Biology, 11(6), 65004.

Multiscale Bone Mechanobiology Stefan Scheiner, Vienna University of Technology, Vienna, Austria Maria-Ioana Pastrama, KU Leuven, Leuven, Belgium Peter Pivonka, Queensland University of Technology, Brisbane, Australia Christian Hellmich, Vienna University of Technology, Vienna, Austria © 2019 Elsevier Inc. All rights reserved.

Introduction to Bone Remodeling Basics and Involved Key Players Implications of Bone Remodeling Disorders Mechanobiological Regulation Major Regulatory Pathways in a Nutshell Mechanical Stimulation Regulatory Interplay Across Several Orders of Magnitude Mathematical Modeling of Bone Remodeling Expected Benefits Historical Overview New Development: Pore Space-Specific, Multiscale Modeling Summary and Outlook References Further Reading

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Glossary Bone cell population model A model taking into account the behavior of cell populations, quantified in terms of cell numbers of molar concentrations, within a defined volume of interest. Cortical bone The compact type of bone forming the outer shell of most bones, exhibiting superior mechanical properties (as compared to trabecular bone). Deterministic modeling A system, for example concerning the behavior of cells, is modeled as detailed as possible, explicitly taking into account all known subprocesses and contributing factors. Homogenization The utilization of mathematical methods for finding estimates of the overall behavior of composite materials, in terms of specific physical quantities. Mechanobiology The neologism made up for expressing that related phenomena are governed by both mechanical stimuli and biological processes, in coupled fashion. Model validation The process of testing the soundness of a model, for example through comparison of model predictions with independent experimental data (the more validation attempts a model survives without significant discrepancies to the respective reference data, the better the model). Phenomenological modeling A modeling strategy where only the overall behavior of a system is taken into account empirically, that is based on observations, without considering subprocesses and how additional factors influence the system’s behavior. Representative volume element A microheterogeneous domain that can be considered as macroscopically homogeneous if the characteristic length of the heterogeneities is significantly smaller than the characteristic length of the representative volume element. Tissue Engineering A field at the interface of engineering and medicine, where malfunctioning or pathological biological tissues are replaced by biocompatible, synthesized materials, engineered in terms of morphology, biological characteristics, and mechanical properties. Trabecular bone The spongious, highly porous type of tissue usually found in the interior of bones (particularly at the ends of long bones).

Nomenclature Abbreviations CT FE

Computed tomography Finite Element

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lac M-CSF OBA OBP OBU OCA OCP OCY OPG PTH RANK RANKL RVE TGFb vas VEGF Wnt

Lacunar pore space Macrophage colony stimulating factor Active osteoblasts Osteoblast precursor cells Uncommitted osteoblast progenitor cells Active osteoclasts Osteoclast precursor cells Osteocytes Osteoprotegerin Parathyroid hormone Receptor activator of nuclear factor kappa b Ligand of RANK Representative volume element Transforming growth factor b Vascular pore space Vascular endothelial growth factor Wingless gene

Latin symbols A vas OBA Cij vas CTGFb exvas COCY Dvas OBP fj fvas f OBP/OBA Krep,TGFb Ni peak plac peak pvas P vas OBP Vj t

Constant apoptosis rate of active osteoblasts Molar concentration of species i within domain j Molar concentration of TGFb in the vascular pore space Concentration of osteocytes across extravascular bone matrix Maximum differentiation rate of OBPs in the vascular pore space Volume fraction of domain j Vascular porosity Frequency at which loading peaks occur Repression coefficient of OBP differentiation due to TGFb Amount of species i Lacunar pore pressure at peak loading Vascular pore pressure at peak loading Maximum proliferation rate of OBPs in the vascular pore space Volume j Time variable

Greek symbols OBP/OBA prep,TGFb mech,vas Pact,OBP

mech;vas

b P act;OBP mlac ref mlac mvas ref mvas vas lOBP vas lWnt/scl

Repression function for differentiation of OBPs to OBAs, due to the presence of TGFb Activation function of OBP proliferation in the vascular pore space, due to mechanical excitation mech,vas Baseline value of Pact,OBP Mechanical stimulus in the lacunar pore space Mechanical stimulus in the lacunar pore space at reference loading Mechanical stimulus in the vascular pore space Mechanical stimulus in the vascular pore space at reference loading Anabolic strength parameter related to direct excitation of the OBPs Anabolic strength parameter related to excitation of the OCYs

Introduction to Bone Remodeling Basics and Involved Key Players Nature has equipped bone with intriguing features, which allow for fulfillment of its vital roles for the human body, such as provision of sufficient load-carrying capacity, protection of organs, locomotion, and calcium homeostasis. For these purposes, continuous maintenance of the microstructural integrity of the bone tissue is crucial, to avoid macroscopic bone fractures. The mechanism concerned with this important task is referred to as bone remodeling (Robling et al., 2006).

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In short, bone remodeling involves numerous biochemically and mechanically stimulated mechanisms (which are briefly discussed in the following sections of this article), together leading to removal of bone tissue by multinuclear cells deriving from hematopoietic stem cells, called osteoclasts, and to concurrent addition of bone tissue by mononuclear cells deriving from mesenchymal stem cells, called osteoblasts. More precisely, osteoclasts attach themselves to the bone surface, and create an acidic environment in order to dissolve the bone matrix, thus leaving behind a resorption cavity. Osteoblasts lay down osteoid to fill this cavity, a substance consisting mainly of collagen type I. Over time, osteoid gets mineralized, with the bone mineral mainly consisting of carbonated hydroxyapatite. Since the greater part of bone mineral is formed very fast (within a few days), osteoblasts are often, for simplicity, referred to as bone-forming cells, without explicitly mentioning the intermediate step of osteoid secretion. Moreover, osteocytes, a third cell type found in bone, have been identified as bone remodeling “conductor.” In particular, osteocytes are former osteoblasts that have been “buried” in newly secreted osteoid. They are the most abundant cell type in bone tissue, and they are considered capable of sensing changes in their environment, particularly concerning the mechanical loading. Subsequently, osteocytes transduce these changes into biochemical signals, which, in turn, lead to corresponding changes of cellular activities in the bone microenvironment.

Implications of Bone Remodeling Disorders In healthy bone, the processes involved in bone remodeling are finely tuned such that the amounts of removed and added bone tissue are balanced, and that the bone tissue composition, that is the proportions of collagen, mineral, and water across several length scale orders of magnitude, is constant, thereby entailing the required bone stiffness. However, a range of circumstances may induce disturbances of this balance. Bone disorders are usually caused by the dysfunction or dysregulation of one or several factors governing bone remodeling (Robling et al., 2006). This concerns both sudden perturbations of the hormonal balance (e.g., occurring at and after menopause) and hereditary disorders (being the cause for pathologies such as osteogenesis imperfecta). A prominent example of the former is postmenopausal osteoporosis, where a reduced estrogen concentration initiates a cascade of biochemical processes leading to increased osteoclast activity, and, eventually, to more porous bone tissue. On the other hand, also long-term changes of the physical activity are potentially effective in terms of triggering corresponding changes of osteoclast and/or osteoblast activities; for example, bed rest and a sedentary lifestyle have been associated with significant bone loss. Exhaustively unraveling the regulatory intricacies behind these effects would have huge impacts, both in terms of the well-being of affected people, and on a commercial leveldwhen considering, for example, the fields of tissue engineering or pharmacology.

Mechanobiological Regulation Major Regulatory Pathways in a Nutshell Ever since the discovery of RANKL, which is the protein acting as ligand for RANK, the receptor activator of nuclear factor kappa b, the RANK-RANKL-OPG-pathway is considered to be the regulatory pathway directing bone remodeling. In particular, binding of RANKL, which is produced by both osteoblasts and osteocytes, to RANK, is a transmembrane protein found on osteoclasts, is required for osteoclastogenesis. Osteoprotegerin, another protein produced by osteoblasts, is capable of preventing the binding of RANKL to RANK, by binding to RANKL itself; thus it is considered to be a so-called decoy receptor for RANKL. Furthermore, PTH, a hormone produced by the parathyroid glands, is able to influence the RANK-RANKL-OPG pathway in catabolic fashion as it promotes the production of RANKL and inhibits the production of OPG, respectively. Another important protein for the signaling pathways in bone remodeling is the cytokine usually referred to as transforming growth factor b, standardly abbreviated as TGFb. Besides the prominent roles TGFb plays in tumor growth and cardiovascular diseases, it is also known to influence the differentiation tendencies of osteoblasts and osteoclasts. In particular, TGFb has turned out to (i) promote the differentiation of uncommitted osteoblast progenitor cells (derived from mesenchymal stem cells) to osteoblast precursor cells, to (ii) inhibit the further differentiation of osteoblast precursor cells to active osteoblasts, which are the cells which actually secrete osteoid, and to (iii) promote the apoptosis (or, programmed death) of active osteoclasts. Furthermore, macrophage colony stimulation factor (standardly abbreviated to M-CSF), which is produced by osteoblasts, contributes to the regulatory pathways of bone remodeling as it has a positive effect on the differentiation behavior of early-stage osteoclasts. Apart from the abovementioned factors, bone remodeling is known to be influenced by a large number of further proteins and hormones (such as nitric oxide, prostaglandins, vitamin D, corticosteroids, interleukins, calcitonin, just to mention a few). More details on this matter can be found in respective review articles (Robling et al., 2006), some of which are listed in the further reading lists at the end of this article. Importantly, osteocytes are known to somehow sense changes of the mechanical loading, see the next subsection (on mechanical stimulation of bone remodeling) for more details in this regard. In case of increased mechanical loading, osteocytes accordingly downregulate the production of sclerostin, a glycoprotein, which itself inhibits the Wntb-catenin pathway, Wnt being the common abbreviation for wingless gene. The latter pathway upregulates osteoblast proliferation; thus osteocytes are able to indirectly promote bone formation. The production of RANKL by osteocytes is also modulated by the prevailing mechanical loading. In particular, the amount of RANKL produced by osteocytes is indirectly proportional to the mechanical loading; thus increased mechanical loading causes osteocytes to produce less RANKL, and vice versa.

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Mechanical Stimulation It has already been mentioned above that specific subprocesses (e.g., proliferation of osteoblasts and production of RANKL) occurring in the course of bone remodeling are modulated by the prevailing mechanical loading. Thus, scrutinizing how and to which extent mechanical loading is capable of influencing the aforementioned processes is key for understanding the related effects on a macroscopic level. The overall response of bone tissue to mechanical loading is well known and undisputed; usually, increased mechanical loading, with respect to a certain “normal” physiological loading relating to everyday activities, leads to gain of bone mass (or volume), while decreased mechanical loading leads to bone loss. It is also well accepted that changes of the prevailing mechanical loading induce corresponding changes of the activities of osteoblasts and osteoclasts. The underlying mechanisms of mechanosensing (i.e., how cells sense the mechanical loading and changes thereof) and mechanotransduction (i.e., how cells process changes of the mechanical loading) have been studied intensely. As a result, the eventual effects of changed mechanical loading are (at least qualitatively) known; for example, an increased mechanical loading causes, via the Wntsignaling pathway, increased proliferation of osteoblast precursor cells (and in further consequence elevated osteoblast activity), while a decreased mechanical loading causes increased production of RANKL (and in further consequence elevated osteoclastogenesis). However, the true routes along which mechanobiological regulation of bone remodeling occurs, including the question of how exactly the involved cells are able to perceive changes in their mechanical environment, still remain to be revealed. Nevertheless, it is instructive to briefly review the state of the art concerning the mechanical stimuli that have been hypothesized to effectively influence the activities of the bone remodeling-driving cells (Klein-Nulend et al., 2005; Robling et al., 2006; Sugiyama et al., 2012; Galea et al., 2013). For the latter task, a few specific mechanical stimuli have turned out as particularly promising candidates; these are discussed next. Undoubtedly, subjecting bone to mechanical loading causes, in some form, pressurization of the fluid filling the various pore spaces of bone tissue. In vitro studies have confirmed that bone cells (i.e., osteocytes, and also osteoblasts, osteoclasts, as well as their progenitors) exhibit altered activities when subjected to hydrostatic pressure at frequencies of up to 1 Hz, with amplitudes of several tens (to hundreds) of kilopascals. Remarkably, it could be shown, through poromicromechanical modeling, that physiological macroscopic loading implies lacunar pore pressures of adequate amplitudes, thus corroborating the potential relevance of pore pressures for mechanosensation of osteocytes. On the other hand, a heterogeneous distribution of stresses across the bone tissue (as, for instance, due to exposure to bending moments, or a heterogeneously distributed bone composition) leads to the occurrence of pore pressure gradients, which, in turn, induces the movement of the pore fluid. In the late 1980s, the hypothesis was brought forth that this flow of the fluid contained in the network made up by lacunar and canalicular pores leads to corresponding shear stresses acting on osteocytes and their cell processes, and that these shear stresses may be effective mechanical triggers of mechanotransduction. The bone research community has readily embraced this idea ever since. Direct experimental evidence is not available for the two pore pressure-related mechanisms of mechanosensation. It is thus instructive to analyze the characteristic loading times of bone when subjected to physiological loading regimes. According to a recent study (Scheiner et al., 2016), it turns out that for most physiological loading conditions, undrained conditions are to be expected in the lacunar–canalicular pore network, that is the pore fluid remains “trapped” in the respective pore spaces when exposed to the typically fast-occurring, dynamical loading. Thus, while the exact roles of both mechanisms for the mechanobiology of bone remodeling are not yet fully understood, it is likely that in the past the importance of pore pressures may have been underestimated, while the importance of fluid flow-induced shear stresses may have been overestimated. Apart from the aforementioned pore fluid pressure-related mechanical stimuli, several alternatives have been suggested. Firstly, once cells are attached to bone surfaces, they may (partly) undergo the same deformation as the surrounding bone tissue. This deformation was also hypothesized to trigger mechanotransductive responses of such cells to mechanical loading and changes thereof. However, as in vitro studies have shown, the strains required to induce effective responses in terms of cellular activities are quite high, reaching up to 10%. These high strains may be explained by amplification effects occurring in the immediate vicinity of holes, such as lacunar pores. Another potential factor contributing to the mechanical regulation of bone remodeling are microcracks occurring as consequence of (mild) overloading. It was suggested that, on the one hand, microcracks cause osteocyte apoptosis and disrupt the communication pathways of the lacunar–canalicular network. Interestingly, it has also been shown that nonapoptotic osteocytes surrounding the apoptotic ones upregulate the expression of RANKL and of VEGF (which is short for vascular endothelial growth factor) and decrease the expression of OPG. On the other hand, it has also been hypothesized that microcracks alter the fluid flow characteristics. However, to date, no conclusive experimental evidence could be found confirming the suggested effects of microcracks. Additionally, the piezoelectricity of collagen was suspected to be related to the mechanosensitivity of osteocytes. Importantly, not only the amplitude of the mechanical stimuli but also other aspects of load application (Turner, 1999) are key for mechanical regulation of bone remodeling. These include the load frequency, the number of load cycles, the rate of load application, the duration and amplitude of the dynamic part of the loads, as well as the particularity of the mechanical stimulus (with respect to “normal” loading).

Regulatory Interplay Across Several Orders of Magnitude An additional intricacy related to the regulation of bone remodeling manifests itself in the hierarchical organization of bone, spanning several orders of length scale magnitude. Taking into account the wealth of experimental data available in literature concerning the structural biology of bone, a multilevel organization emerges (Vass et al., 2018). At an observation scale of

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several tens of nanometers, a matrix of cross-linked collagen molecules hosts intermolecular pore space (the latter being filled with pore fluid), together forming wet collagen. One level up, wet collagen and crystals of impure hydroxyapatite form the so-called mineralized fibrils, with a characteristic length of several hundreds of nanometers. These fibrils, in turn, are embedded in a conglomerate of hydroxyapatite crystals and intercrystalline pore space, typically referred to as extrafibrillar space; transmission electron micrographs show the arrangement of collagen-free extrafibrillar space and collagen-rich fibrils, where the former is visible as dark regions and the latter as light regions. Together, extrafibrillar space and mineralized collagen fibrils form the bone ultrastructure, optionally known as extracellular bone matrix, which exhibits a characteristic length of a few micrometers. In extracellular bone, pore spaces are embedded, which are of great importance for the metabolism of bone. These are the osteocytes-hosting lacunar pores (with a characteristic length of approximately 10 mm) and the lacunae-connecting canalicular pores (being typically several hundreds of nanometers long, with a diameter of approximately 35 nm), together forming a dense pore network. On an observation scale of several hundreds of micrometers, extravascular bone is composed of extracellular bone and the lacunae–canaliculi network. Extravascular bone, in turn, is the matrix in which vascular pores can be found, with a characteristic length of 50–80 mm. In cortical bone, the vascular pores are synonymous to the network of Haversian and Volkmann canals, and in trabecular bone to intertrabecular pore space. In any case, the vascular space hosts many cells and biochemical factors initiating and driving bone remodeling. Importantly, exchange of factors between osteocytes and the vascular pore space is enabled through canaliculi extending to the surfaces of extravascular bone. Finally, bone microstructure, that is the tissue building up the substance of bone organs on a macroscopic observation scale, is the composite of extravascular bone and vascular pore space. Considering the multilevel organization of the bone tissue, it becomes apparent that for understanding the effects of specific influences (of both mechanical and biochemical nature) acting upon cells and other factors on the progress of bone remodeling, it is of utmost importance to take into account that the key players of bone remodeling are located in pore spaces of distinctively different characteristic lengths. In the following, this statement is underpinned by describing the cascade of events occurring in response to the mechanical loading to which a bone organ is subjected. The mechanical loading is applied macroscopically onto the respective bone. Then the loading is transferred, depending on the actual, site-specific bone composition, to the mechanosensitive osteocytes. The osteocytes then transduce the mechanical signal into biochemical signals that affect the behavior of osteoblasts and osteoclasts. The extent of bone formation by osteoblasts and of bone resorption by osteoclasts governs the bone composition evolution, that is whether the bone mass remains constant, increases, or decreases. Eventually, the bone composition determines the transfer of the mechanical loading from the macroscopic to the osteocytic length scale (Fritsch and Hellmich, 2007; Scheiner et al., 2013). Thus, the progress of bone remodeling is governed by several biochemical, mechanical, and bone composition-related factors, acting time-dependently, site-specifically, and length scale-specifically; many of these factors being actually linked with each other, in feedback-type fashion. Covering all these aspects in the framework of experimental studies is difficult, if not impossible. Typically, two kinds of experimental studies are performed: On the one hand, in vitro studies focus on exploring the specific aspects, for example related to the differentiation or expression behavior of cells, when subjected to specific factors. However, the “bigger picture” is lacking in such small-scale tests, mainly concerning an adequate physiological environment including the multiple interactions of the studied cell with factors other than the few which can be considered in the laboratory, and the embedding of the cells and factors in the respective pore spaces. On the other hand, in vivo tests, typically carried out in the form of animal studies (Sugiyama et al., 2012) or by making use of bioreactors, cannot provide continuous information on cell activities or morphological changes of the bone tissue, but only, if at all, at distinct time points, which is usually accompanied with termination of the respective test. Mathematical modeling of bone remodeling, making use of the wealth of experimental data collected in in vitro and in vivo studies, seems to be a reasonable remedy to the predicament sketched above. Describing the historical development and the current state of such modeling approaches will be the subject of the following section.

Mathematical Modeling of Bone Remodeling Expected Benefits Utilization of mathematical modeling for replication and prediction of the bone remodeling progress in response to specifically prescribed conditions holds a number of promises: First of all, simulation of bone remodeling based on a physically, biologically, and mechanically substantiated mathematical model is expected to significantly improve the understanding of the bone remodeling process itself. Ideally, computer simulations allow to study the effects of variations of specific factors on the bone metabolism. Considering that the underlying mathematical models take into account the results of experimental studies extensively, they provide some kind of sophisticated inter- and extrapolations of experimental data. Due to the complexity and nonlinearity of the bone remodeling process, this cannot be achieved by intuition or experience-guided guessing, at least not in a quantitatively reliable way. Once such a mathematical model has been established and validated, it can be applied as decision-support utility for a large variety of practically relevant problems. Examples include

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Pharmaceutical engineering, where the number of animal studies could be dramatically reduced by assessing the efficacy of drugs in silico (for example through predicting the effects of varying drug doses and administration regimes);

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Tissue engineering, where the effects of many different parameters of the respective tissue engineering materials can be studied by means of computer simulations, such as different morphologies of the material, or the effects of seeding cells and other factors on tissue engineering scaffolds made up by such materials; and Design of patient-specific treatment regimes, not only based on combining the aforementioned two fields, namely pharmaceutical and tissue engineering, but also based on recommendation of specifically tailored regimes of physical exercise.

Historical Overview Very often, until today, the seminal work of Julius Wolff is considered to be the basis for mathematical models of bone remodeling, at least when focusing on adaptation of bone to the prevailing mechanical loading. Julius Wolff summarized the main findings of his work as surgeon in the book “The Law of Transformation of Bone,” first published in 1892, where he postulated that “as a consequence of primary shape variations and continuous loading, or even due to loading alone, bone changes its inner architecture according to mathematical rules and, as a secondary effect and governed by the same mathematical rules, also changes its shape.”dnotably, a similar (yet more general) proposition was published already in 1638 by Galileo Galilei. Wolff’s hypothesis was warmly welcomed by the bone research community, and refined in the course of the 20th century. The 1960’s can be considered as formative years in the search for suitable (mathematical) formulations relating mechanical loading and bone metabolism-driven changes of the shape and morphology of bone (organs). Namely, Harold M. Frost published his famous “Utah Paradigm of Bone Physiology,” including the so-called mechanostat theory. The latter established a simple relation between the strain state bone is experiencing and the corresponding bone turnover. Due to its simplicity, the mechanostat theory has shaped the field of computational simulation of bone remodeling ever since. The development of mathematical models for simulating the (mechanical) behavior of complex structures is traditionally the specialty of civil and mechanical engineers. It is thus not surprising that mathematical modeling and computational simulation of bone remodeling phenomena was pioneered and expedited by researchers coming from the aforementioned engineering disciplines, and that related efforts first focused on the structural scale, that is the scale of whole bone organs. Facilitated by the everincreasing power and availability of computers systems, the respective state of the art between the 1970s and 1990s was dominated by large-scale Finite Element (FE) simulations (Webster and Müller, 2011; Schulte et al., 2013), complemented by simplified (and purely phenomenological) bone remodeling rules governed by mechanical stimuli, according to Frost’s mechanostat theory, while biological details were left aside. Clearly, the hierarchical organization of bone, spanning several orders of magnitude, could not (and still cannot) be considered in these FE simulations, due to the unmanageable computational effort implied by a sufficiently fine discretization of bone tissue. Hence, also the mechanical stimulus for bone remodeling could not (and still cannot) be chosen according to physiological insights on mechanical stimulation of bone remodeling. Instead, FE simulations of mechanically modulated bone remodeling events typically consider simplified strain-representative measures, such as the strain energy density occurring in the macroscopic bone tissue in response to specifically prescribed loading boundary conditions, or scalar damage variables. The flexibility of the FE method, in terms of studying structures of arbitrary, optionally changing shapes, and the over the years more and more increasing availability of commercial FE software (or, alternatively, of open-source FE codes) have led to the long-term establishment of the FE technique as the gold standard in the field. Recently, more and more detailed models, based on highresolution imaging techniques, could be successfully implemented, allowing even to take into account the exact mechanism by which osteocytes have been hypothesized to be mechanically stimulated, for example via the primary cilia, or through the response of the cell cytoskeleton to mechanical forces. However, due to the substantial computational effort related to FE simulations, organscale models mostly consider macroscopic features of the studied bone organ, entailing the danger of (quantitatively) inaccurate model predictions, see Fig. 1. The comparison of model-predicted and experimentally observed sites of bone formation and bone resorption due to mechanical stimulation of the 6th caudal vertebra of a mouse, as illustrated in Fig. 1, shows that the numerical model fails to predict the distribution of these sites correctly, while other model-predicted parameters, such as formation and resorption rates, agree reasonably well with the experimental data (Schulte et al., 2013). The reasons for this discrepancy could be that, on the one hand, the exact distribution of the (anisotropic) mechanical properties of the bone tissue was neglected, and, on the other hand, the exact

(A)

(B)

Resorbed bone

Formed bone Fig. 1 Exemplary results from bone remodeling studies on the 6th caudal vertebra of a mouse, performed (A) by means of the Finite Element method, and (B) experimentally (by means of micro-CT), highlighting that the spatial distribution of resorption and formation sites following from numerical simulations differs from the one obtained experimentally. Reprinted from (Schulte et al., 2013), with permission from Elsevier B.V.

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Terminally differentiated Sclerostin E2

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E2

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IGF-1R LRP

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AKT ERα ERβ ? Proliferation Differentiation

β-cat

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ERα ERβ

ERα β-cat

ERK

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E2

β-cat LRP Wnt

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Proliferation Differentiation Matrix production Antiapoptosis

Migration Antiapoptosis

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Osteoclast

Fig. 2 Illustration of the manifold ways mechanical stimulation of osteoblasts and osteocytes influences the role of estrogen receptors ERa and ERb, which, in turn, influence the behaviors of these cells; E2 stands for estrogen, LRP for low-density lipoprotein receptor, b-cat for beta-catenin, Wnt for wingless gene, ERK for extracellular signal-regulated kinase, IGF-1R for insulin-like growth factor 1 receptor. Reproduced from (Galea et al., 2013), by permission from Macmillan Publishers Ltd, BonekEy Reports, ©2009

biochemical and biomechanical regulation of bone remodeling was not taken into account. Particularly the latter is a challenging task, as the involved pathways are manifold and intricate, see for example Fig. 2 for the various ways mechanical stimulation acts on the estrogen receptors in osteoblasts and osteocytes, which, in turn, are important regulatory components for the behaviors of these cells. Mathematically modeling the multiply interrelated actions of different types of bone cells (osteoclasts, osteoblasts, and osteocytes) and the large number of biological factors, overall constituting the process of bone remodeling, was tackled in the early 2000s. In particular, the tool of bone cell population kinetics was discovered for quasi-deterministic simulation of bone remodeling in a temporarily evolving manner, within macroscopic representative volume elements of (cortical or trabecular) bone tissue. Taking into account only the major regulatory pathways of bone remodeling, that is explicitly considering only RANK, RANKL, OPG, PTH, and TGFb, systems of ordinary differential equations were compiled allowing for predicting the temporal evolutions of bone cell populations (quantified in terms of molar concentrations), in response to dynamically prescribed biochemical boundary conditions (relating e.g., to bone pathologies, such as osteoporosis). Thereby, the regulatory effects of the aforementioned biochemical factors were considered by means of respective (Hill-type) activation or repression functions. Additionally, efforts were undertaken to extend the merely temporal approaches to spatio-temporal models, by considering also spatial aspects (leading to sets of partial differential equations) see (Lemaire et al., 2004; Pivonka et al., 2008; Scheiner et al., 2013). So far, two very contrary, isolated approaches have been introduced, shaping the state of the art in computational simulation of bone remodeling over the past decades; on the one hand, the mostly mechanically driven large-scale numerical simulations, and, on the other hand, the biochemically driven bone cell population models. Thus, the obvious next step was to develop models considering both mechanical and biochemical regulation of bone remodeling, thus mathematically describing bone remodeling as a mechanobiologically driven process. This was achieved by coupling bone cell population models with continuum micromechanics models. The latter allow for downscaling the macroscopically applied mechanical loading bone is experiencing to the level of the mechanosensitive osteocytes. Such calculated local mechanical stimuli were then fed into a bone cell population model, according to experimental evidence. However, despite the conceptual advantages these models exhibit (compared to “conventional” modeling approaches), they do not take into account the fact that the players (cells and biochemical factors) are actually situated in pore spaces of separated characteristic length, namely in the vascular and lacunar pore spaces. This deficit provides the incentive for further developments, described next.

New Development: Pore Space-Specific, Multiscale Modeling In order to demonstrate how comprehensive multiscale modeling of bone remodeling can actually be achieved, this section is devoted to presenting a novel modeling approach for computational simulation of bone remodeling, developed recently by the authors of this article, and thoroughly obeying the multiscale paradigm, that is taking into account all relevant mechanisms on the appropriate length scales (Pastrama et al., 2018).

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The development of mathematical models is fundamentally based on reasonably simplifying the “real” material or structure that is to be modeled. Thereby, the guiding principle should be to keep the model representation as simple as possible, but at the same time as complex as necessary. In the present context aiming for a mathematical model allowing for simulation of bone remodeling, such model representation preferably spans three distinct observation scales, see Fig. 3A and B (Cardoso et al., 2013):

•

On the bone organ scale, macroscopically homogeneous regions of cortical or trabecular bone tissue build up whole organs. Within the organ-scale bone tissue, no further distinction between different material phases is made. This scale is particularly relevant for imposition of the mechanical loading conditions. On the scale of macroscopic bone tissue, two constituents can be distinguished, extravascular bone matrix and vascular pores. The latter pores, typically exhibiting a characteristic length of 50–80  10 6 m, can be found in cortical bone in the form of Haversian and Volkmann canals. Under normal physiological conditions, the cortical vascular porosity amounts to approximately 3%–5%, while in pathological bone, the cortical vascular porosity can be as high 35% (the porosity values relate to the respective pore volume fractions within macroscopic bone tissue). In trabecular bone, the intertrabecular pores exhibit porosities ranging from 50% to 90%. Vascular pores contain osteoblasts and osteoclasts (as well as their progenitor cells), and numerous biochemical factors affecting the behavior of these cells. On the scale of extravascular bone tissue, the extralacunar bone matrix hosts the lacunar pores, in which the mechanosensitive osteocytes reside. The lacunar pores exhibit a characteristic length of approximately 10 5 m. The exact value of the lacunar

•

•

Bone organization

(A)

Micromechanical representation

(C)

5 mm

OBU on

ati

ti en fer

Vascular pore space (pressurized)

Dif

OCP

tio nt re ffe Di

e ay ia th on v pathw G racti Inte KL-OP y PTH) b AN nced K-R RAN o influe (als

ion iat nt re

OBA

Proliferation

ia

Apoptosis

ffe Di

n

OBP

Wnt signalling pathway

Apoptosis OCA

Release of TGFβ

RANKL production

Cross section A-A

OCY A

A

Extravascular bone matrix (B)

10 μm

1 μm

dvas

dlac

Vascular pores

Lexvas

2 mm

Lmacro

10 μm

Lacunar pores (pressurized)

Extralacunar bone matrix

Micromechanical representation Fig. 3 (A) Hierarchical organization of cortical and trabecular bone, illustrated via an X-ray of a human femur showing a cross-section in the midshaft region acquired by means of microradiography, with the zoom into the cortical part showing a scanning electron microscopy (SEM) image highlighting the arrangement of vascular and lacunar pores, with the latter containing osteocytes, see the respective SEM image, while trabecular bone, illustrated by means of computed tomography (CT), has a differing microstructure, as shown by photomicrography, however also containing lacunae; (A) allows to deduce a respective micromechanical representation (B), spanning from macroscopic bone tissue to the lacunar pores; and (C) the cell population model taking into account cells and biochemical factors, as well as mechanical stimulation in the lacunar and vascular pore spaces (orange arrows indicate where mechanical loading eventually intervenes with specific, cell development-influencing processes; the X-ray image of a human femur and the photomicrograph of trabecular bone have been reproduced from Sinclair et al. (2013) with permission of Elsevier B.V.; the cross-section of the femur has been reproduced by courtesy of John G. Clement and David Thomas (taken from the Melbourne Femur Collection); the SEM image of cortical bone has been reprinted from Kessel and Kardon (1979), by courtesy of Randy H. Kardon; the CT image of trabecular bone has been reproduced from Padilla et al. (2008), with permission of Elsevier B.V.; the SEM of a single osteocyte has been reproduced from Pajevic (2009), by permission from Macmillan Publishers Ltd, IBMS BonekEy, ©2009.

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porosity is still a somewhat open issue; depending on the chosen imaging technique, porosities between 2% and 10% have been identifieddboth values relate to the pore volume fraction within the extravascular bone matrix. Morphological features discernible on lower observation scales, such as the canalicular pores, or the different arrangements of the extralacunar bone matrix components collagen, hydroxyapatite crystals, and water are not explicitly resolved in this model, but implicitly taken into account through an appropriate choice of the extralacunar bone matrix’s mechanical properties. It should be emphasized that the chosen model representation has been extensively and successfully validated based on comparing model predictions (in terms of stiffness, strength, poroelastic properties, and viscoelastic behavior) with corresponding experimental data, considering different types of bone, anatomical locations, species, and experimental protocols. Tying in with the aforementioned bone cell population models, the alluded-to modeling approach deals with temporal evolutions of cell numbers, or, when relating the cell numbers to the cell-hosting volumes, of molar cell concentrations, defined as j

Ci ¼

Ni ; Vj

(1)

with Cij denoting the molar concentration of species i, quantified with respect to the hosting domain j, Ni is the number or the amount of species i, and Vj is the volume of domain j. It is now instructive to differentiate Eq. (1) with respect to time, in order to obtain a mathematical expression describing how Cij changes with time, yielding     j dVj dCi d Ni 1 dNi ¼ Vj  Ni (2) ¼  2 dt Vj dt dt dt Vj Multiplying the right-hand side of Eq. (2) with the term (Vmacro/Vmacro), as well as considering the definition of pore volume fraction fj, fj ¼ Vj =Vmacro , leads to j

j

dCi 1 dNi Ci dfj ¼  : Vmacro fj dt dt fj dt

(3)

Eq. (3) expresses the fact that the concentration Cij changes, on the one hand, due to an increase of the amount of species i, and, on the other hand, due to the change of the species-hosting volume j. For the sake of demonstration, Eq. (3) is now applied to a specific subpopulation of osteoblasts, namely the active osteoblasts. The amount of these cells, NOBA, is known to increase due to differentiation from the committed osteoblast precursor cells, inhibited by the presence of TGFb, and to decrease due to cell apoptosis. Hence, the following mathematical expression emerges: dNOBA OBP=OBA vas ¼ Dvas OBP prep;TGFb NOBP  A OBA NOBA ; dt

(4)

OBP/OBA where Dvas OBP denotes the maximum differentiation rate of osteoblast precursors in the vascular pore space; prep,TGFb , the repression function of osteoblast precursor differentiation, governed by TGFb; NOBP, the amount of osteoblast precursor cells; and A vas OBA the constant apoptosis rate of active osteoblasts in the vascular pore space. Inserting Eq. (4) into Eq. (3) yields then

dCvas Cvas OBP=OBA vas vas vas OBA OBA dfvas ¼ Dvas ; OBP prep;TGFb COBP  A OBA COBA  dt fvas dt

(5)

OBP/OBA with fvas as the vascular porosity. Repression function prep,TGFb can be straightforwardly defined as so-called Hill-type function, reading as

OBP=OBA

OBP=OBA

prep;TGFb ¼

Krep;TGFb OBP=OBA

Krep;TGFb þ Cvas TGFb

;

(6)

OBP/OBA vas where Krep,TGFb is the repression coefficient of osteoblast precursor differentiation due to TGFb, and CTGFb is the concentration of TGF b in the vascular pore space. Analogously, all other relevant cell populations, biochemical factors, and mechanical stimuli can be considered. In line with the pioneering contributions in the field, see for example references (Lemaire et al., 2004; Pivonka et al., 2008), the mathematical framework of the multiscale bone cell population model briefly elaborated here includes the following processes and features, see also Fig. 3C:

•

The osteoblast lineage comprises uncommitted osteoblast progenitor cells, committed osteoblast precursors, and active osteoblasts. – Due to their abundance, the concentration of the progenitor cells can be assumed constant. – The amount of osteoblast precursors increases due to differentiation from the progenitor cells (promoted by the presence of TGFb), increases due to proliferation (promoted by mechanical stimulation), and decreases due to differentiation to active osteoblasts (inhibited by the presence of TGFb). – The concentration of active osteoblasts is governed as described above, see Eqs. (4)–(6).

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The osteoclast lineage comprises committed osteoblast precursors and active osteoblasts. – Due to their abundance, the concentration of osteoclast precursors can be assumed constant. – The amount of active osteoclasts increases due to differentiation from the precursor cells (promoted by the binding of RANK to RANKL, with the availability of RANKL being inhibited by mechanical stimulation) and decreases due to apoptosis (promoted by the presence of TGFb).

While the mathematical implementation of these aspects is described elsewhere, see references (Lemaire et al., 2004; Pivonka et al., 2008; Scheiner et al., 2013) , explaining how mechanical stimulation of bone remodeling can be efficiently considered is still pending. For this purpose, the proliferation of osteoblast precursors is considered. Similarly to the Hill-type repression function defined in Eq. (6), upregulation of osteoblast precursor proliferation can be considered by function Pmech,vas act,OBP ; taking into account only the change in the amount of the osteoblast precursors due to proliferation, one would obtain  vas  dNOBP mech;vas vas ¼ P vas (7) OBP Pact;OBP NOBP ; dt prolif mech,vas where P vas OBP is the maximum proliferation rate of osteoblast precursor cell in the vascular pore space. For definition of Pact,OBP , it is considered that the promotion of osteoblast precursor proliferation occurs in two ways: On the one hand, it has been shown experimentally that mechanical stimulation of osteoblasts leads to an increase of their proliferation, by up to 100%. On the other hand, it is also known that mechanical stimulation of osteocytes leads to downregulation of sclerostin expression, which, in turn, somehow causes upregulation of osteoblast precursor proliferation via the Wnt/b-catenin pathway. For capturing these processes mathematically, the following formulation has proven adequate: ! !# " mech;vas mvas mech;vas vas vas exvas 1  fvas mlac b Pact;OBP ¼ P act;OBP 1 þ lOBP ref  1 þ lWnt=scl COCY 1  1: (8) fvas mvas mref vas

b mech;vas is the baseline value of Pmech,vas (thus indicating systemic precursor proliferation, independent of the In Eq. (8), P act,OBP act;OBP vas mechanical loading), lOBP is the anabolic strength parameter related to direct excitation of the precursor cells (quantifying how vas is the anabolic strength parameter related to excitation of mechanical stimulation is translated into additional proliferation), lWnt/scl exvas is the concentration of osteocytes across the extravascular bone matrix, mvas and mlac are the mechanical stimuli the osteocytes, COCY ref ref and mlac are the respective reference values, relating to steady-state occurring in the vascular and in the lacunar pore spaces, while mvas conditions. For further definition of the mechanical stimuli, it is taken into account that the hydrostatic pressure has turned out to be a promising candidate as mechanoregulatory stimulus, that the prime properties of loading history are the peaks (and not the average loading), and that not only the magnitudes of the loading peaks are decisive but also the frequency at which they occur; thus peak peak peak peak and plac are the vascular and lacunar pore pressures related to the loading peaks, and f is mvas ¼ pvas f and mlac ¼ plac f, where pvas their frequency. Notably, the model representation illustrated in Fig. 3B satisfies the so-called separation of scales condition for definition of representative volume elements (RVEs), in a micromechanical sense, thus allowing to employ the theory of poromicromechanics. On this conceptual basis, it is possible to straightforwardly calculate the hydrostatic pressures in the vascular and lacunar pore spaces as they occur in response to macroscopically applied physiological loading. Next, the focus is on highlighting how such multiscale systems biology model of bone remodeling can be evaluated. Firstly, formation and resorption of bone was simulated considering the mouse tibia, for which the mechanical loading at which the bone composition remains constant was experimentally identified to be related to a macroscopic (compressive) peak strain of Epeak ¼  1056  106 , at a load frequency of 0.1 Hz, applied with 40 load cycles per day. Furthermore, both cortical bone, with an initial vascular porosity of 0.05, and trabecular bone, with an initial vascular porosity of 0.75, were considered, while for both types of bone a lacunar porosity of 0.1 was prescribed, quantified within the extravascular bone matrix. Then, the effects of different loading regimes were studied, considering peak loads corresponding to macroscopic peak strains of 0,  500  106 ,  1000  106 ,  1500  106 , and 2000  106 (all of which are initialized at t ¼ 0). Results are illustrated in Fig. 4, showing how the bone composition (quantified in terms of the vascular porosity) adapts to the changed mechanical loading. In particular, for both cortical bone, see Fig. 4A, and trabecular bone, see Fig. 4B, the vascular porosity increases over time if the prescribed loading is lower than the reference value (Epeak ¼  1056  106 ), while it decreases if the prescribed loading is higher than the reference value. The relative change is proportional to the difference of the prescribed loading to the reference value, and to the time span this adaptation process requires. Notably, in cortical bone, bone loss due to disuse loading turns out to be more pronounced (with respect to trabecular bone), whereas in trabecular bone, bone gain due to overloading turns out to be more pronounced (with respect to cortical bone). Moreover, the adaption to disuse loading occurs generally faster than to overloading. Secondly, as shown in histomorphometric studies, the vascular porosity increases with increasing age, while the lacunar porosity decreases with increasing age. Considering related experimental trends in a simulation of disuse loading reveals how the effect of loading may influence mechanical regulation of bone remodeling. In particular, human bone was studied, with “normal” loading resulting from walking (related to peak strains of Epeak ¼  500  106 , occurring at a frequency of 2 Hz, and a disuse loading represented by peak strains of Epeak;disuse ¼  450  106 , occurring again at a frequency of 2 Hz). The disuse loading was initiated at t ¼ 0, and after 5 years the loading was set back to normal loading conditions. Furthermore, two different starting ages were considered, namely 18 and 60 years, in order to investigate whether and to which extent the effect of aging varies between significantly different age groups. The simulation results show that the vascular porosity is much less affected by changes in the

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1.0 Epeak = −1000e–6 Epeak = −1500e–6

Epeak = 0 Epeak = −500e–6

Epeak = −2000e–6

Vascular porosity [-]

0.8 Trabecular bone 0.6

0.4

0.2

Cortical bone

0 0

20 40 60 80 Time after initialization of loading regime [days]

100

Fig. 4 Computed developments of the vascular porosity related to cortical bone (starting at an initial porosity of fvas ¼ 0.05) and trabecular bone (starting at an initial porosity of fvas ¼ 0.75), for various disuse and overloading scenarios.

Vascular porosity [-]

(A)

(B)

(C)

0.09

0.09

0.051

0.08

0.085

0.05075

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0.0505

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0.07 4.95

5

5.05

0.05 9.9

9.95

10

Time after initialization of the loading regime [years] No aging

Aging, 18-year old

Aging, 60-year old

Fig. 5 Effects of aging on the mechanoresponsiveness of human (cortical) bone, when subjected to disuse conditions at t ¼ 0, returning to normal physiological loading at t ¼ 5 years: comparison of nonaging bone and aging bone with disuse initiated at ages 18 and 60 years; (A) the development of the vascular porosity over 10 years, (B) a zoom-out to the time span around 5 years, and (C) a zoom-out to the time span immediately before 10 years.

mechanical loading conditions in an older subject compared to a younger subject (see Fig. 5). However, it becomes also apparent in these simulations that, in quantitative terms, the effect of porosity changes on the mechanoresponsiveness of bone is of minor importance, and consideration of additional mechanisms, such as osteocyte apoptosis, may be required in order to explain why in senescent bone the responses to changes in the mechanical loading differ significantly from the responses in young bone. Besides the presented parameter studies, the multiscale systems biology model could be successfully validated considering mechanical excitation of mouse tibiae (including both disuse and overuse), disuse in the tibiae of rats induced by tail suspension, and disuse in humans due to exposure to microgravity.

Summary and Outlook The intention of this article was to present an overview on the intricacies and various aspects that have to be considered when attempting to simulate bone remodeling computationally. On the one hand, the value of numerical methods (mostly of the Finite

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Element) stands undisputed, particularly in the context of studying the mechanobiological behavior of whole bone organs. The ever-increasing accuracy of high-resolution imaging of bone allows nowadays to create three-dimensional numerical models which are precise replicates of the bone organ under investigation. The challenges to be tackled in future research activities involve, to a large extent, the conditioning of numerical models such that the governing field equations are solvable within a reasonable amount of time. On the other hand, coupling the aforementioned large-scale numerical models with multiscale material models, taking into account the hierarchical organization of bone tissue, and also the biochemically as well as biomechanically driven dynamic evolution of the bone composition (in response to the immediate environment of the respective organ, see the section on “New Development: Pore Space-Specific, Multiscale Modeling” of this article), seems to be a promising strategy. In this way, the various aspects of bone remodeling (being of structural, compositional, biochemical, and biomechanical kinds), perceptible on distinctively separated characteristic lengths but together contributing to the mechanobiological behavior of bone, can be reconciled.

References Cardoso, L., Fritton, S. P., Gailani, G., Benalla, M., & Cowin, S. C. (2013). Advances in assessment of bone porosity, permeability, and interstitial fluid flow. Journal of Biomechanics, 46, 253–265. Fritsch, A., & Hellmich, C. (2007). ‘Universal’ microstructural patterns in cortical and trabecular, extracellular and extravascular bone materials: Micromechanics-based prediction of anisotropic elasticity. Journal of Theoretical Biology, 244, 597–620. Galea, G. L., Price, J. S., & Lanyon, L. E. (2013). Estrogen receptors’ roles in the control of mechanically adaptive bone (re)modeling. BoneKEy Reports, 2, 413. Klein-Nulend, J., Bacabac, R. G., & Mullender, M. G. (2005). Mechanobiology of bone tissue. Pathologie Biologie, 53, 576–580. Lemaire, V., Tobin, F. L., Greller, L. D., Cho, C. R., & Suva, L. J. (2004). Modeling the interactions of osteoblast and osteoclast activities in bone remodeling. Journal of Theoretical Biology, 229, 293–309. Pastrama, M. I., Scheiner, S., Pivonka, P., & Hellmich, C. (2018). A mathematical multiscale model of bone remodeling, accounting for pore space-specific mechanosensation. Bone, 107, 208–221. Pivonka, P., Zimak, J., Smith, D. W., Gardiner, B. S., Dunstan, C. R., Sims, N. A., Martin, T. J., & Mundy, G. R. (2008). Model structure and control of bone remodelling: A theoretical study. Bone, 43, 249–273. Robling, A. G., Castillo, A. B., & Turner, C. H. (2006). Biomechanical and molecular regulation of bone remodeling. Annual Reviews of Biomedical Engineering, 8, 455–498. Scheiner, S., Pivonka, P., & Hellmich, C. (2013). Coupling systems biology with multiscale mechanics, for computer simulations of bone remodeling. Computer Methods in Applied Mechanics and Engineering, 254, 181–196. Scheiner, S., Pivonka, P., & Hellmich, C. (2016). Poromicromechanics reveals that physiological strains induce osteocyte-stimulating lacunar pressure. Biomechanics and Modeling in Mechanobiology, 15, 9–28. Schulte, F. A., Zwahlen, A., Lambers, F. M., Kuhn, G., Ruffoni, D., Betts, D., Webster, D. J., & Müller, R. (2013). Strain-adaptive in silico modeling of bone adaptationdA computer simulation validated by in vivo micro-computed tomography data. Bone, 52, 485–492. Sugiyama, T., Meakin, L., Browne, W. J., Galae, G. L., Price, J. S., & Lanyon, L. E. (2012). Bones’ adaptive responses to mechanical loading is essentially linear between low strains associated with disuse and the high strains associated with the lamellar/woven bone transition. Journal of Bone and Mineral Research, 27, 1784–1793. Turner, C. H. (1999). Toward a mathematical description of bone biology: The principle of cellular accommodation. Calcified Tissue International, 65, 466–471. Vass, V., Morin, C., Scheiner, S., & Hellmich, C. (2018). Review of “universal” rules governing bone composition, organization, and elasticity across organizational hierarchies. In P. Pivonka (Ed.), CISM Courses and Lecture Series, Vol. 578: Multiscale Mechanobiology of Bone Remodeling and Adaptation (pp. 175–229). Berlin: Springer. Webster, D., & Müller, R. (2011). In silico models of bone remodeling from macro to nanodFrom organ to cell. Wiley Interdisciplinary Reviews: Systems Biology and Medicine, 3, 241–251.

Further Reading Aubin, J. E. (1998). Bone stem cells. Journal of Cellular Biochemistry Supplements, 30(31), 73–82. Buckwalter, J. A., Glimcher, M. J., Cooper, R. R., & Recker, R. (1995a). Bone biology. Part I. Structure, blood supply, cells, matrix, and mineralization. Journal of Bone and Joint SurgerydSeries A, 77, 1256–1275. Buckwalter, J. A., Glimcher, M. J., Cooper, R. R., & Recker, R. (1995b). Bone biology. Part II. Formation form, modeling, remodeling, and regulation of cell function. Journal of Bone and Joint SurgerydSeries A, 77, 1276–1289. Busse, B., Djonic, D., Milovanovic, P., Hahn, M., Püschel, K., Ritchie, R. O., Djuric, M., & Amling, M. (2010). Decrease in the osteocyte lacunar density accompanied by hypermineralized lacunar occlusion reveals failure and delay of remodeling aged human bone. Aging Cell, 9, 1065–1075. Colloca, M., Blanchard, R., Hellmich, C., Ito, K., & van Rietbergen, B. (2014). A multiscale approach for bone remodelling simulations: Linking scales from collagen to trabeculae. Bone, 64, 303–313. Cooper, D. M. L., Thomas, D. L., Clement, J. G., Turinsky, A. L., Sensen, C. W., & Hallgrímson, B. (2007). Age-dependent change in the 3D structure of cortical porosity at the human femoral midshaft. Bone, 40, 957–965. Coussy, O. (2004). Poromechanics. Chichester: John Wiley & Sons. Dormieux, L., Kondo, D., & Ulm, F.-J. (2006). Microporomechanics. Chichester: John Wiley & Sons. Jacobs, C. R., Temiyasathit, S., & Castillo, A. B. (2010). Osteocyte mechanobiology and pericellular mechanics. Annual Reviews of Biomedical Engineering, 12, 369–400. Kessel, R. G., & Kardon, R. H. (1979). Tissues and Organs: A Text-atlas of Scanning Electron Microscopy. San Francisco: W.H. Freeman and Co. Khayyeri, H., Isaksson, H., & Prendergast, P. J. (2015). Corroboration of computational models for mechanoregulated stem cell differentiation. Computer Methods in Biomechanics and Biomedical Engineering, 18, 15–23. Lauffenburger, D. A., & Linderman, J. J. (1993). ReceptorsdModels for Binding, Trafficking, and Signaling. New York: Oxford University Press. Lerebours, C., Buenzli, P. R., Scheiner, S., & Pivonka, P. (2016). A multiscale mechanobiological model of bone remodeling predicts site-specific bone loss in the femur during osteoporosis and mechanical disuse. Biomechanics and Modeling in Mechanobiology, 15, 43–67. Ozcivici, E., Luu, Y. K., Adler, B., et al. (2010). Mechanical signals as anabolic agents in bone. Nature Reviews. Rheumatology, 6, 50–59.

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Padilla, F., Jenson, F., Bousson, V., Peyrin, F., & Laugier, P. (2008). Relationships of trabecular bone structure with quantitative ultrasound parameters: In vitro study on human proximal femur using transmission and backscatter measurements. Bone, 42(6), 1193–1202. Pajevic, P. D. (2009). Regulation of bone resorption and mineral homeostasis by osteocytes. IBMS BoneKEy, 6, 63–70. Rubin, C. T., & Lanyon, L. E. (1985). Regulation of bone mass by mechanical strain magnitude. Calcified Tissue International, 37, 411–417. Sinclair C, Birch HL, Smith RKW, and Goodship AE (2013) Skeletal physiology: Responses to exercise and training. In: Equine Sports Medicine and Surgery, 2nd edn., pp. 145–165. Turner, C. (1992). Three rules for bone adaptation to mechanical stimuli. Bone, 23, 399–407.

Multiscale Mechanical Behavior of Large Arteries Claire Morin, Ecole Nationale Supérieure des Mines de Saint-Etienne, CIS-EMSE, SAINBIOSE, Saint Etienne, France; INSERM, Saint Etienne, France; and Université de Lyon, SAINBIOSE, Saint Etienne, France Witold Krasny, Ecole Nationale Supérieure des Mines de Saint-Etienne, CIS-EMSE, SAINBIOSE, Saint Etienne, France; INSERM, Saint Etienne, France; Université de Lyon, SAINBIOSE, Saint Etienne, France; and Université de Lyon, Ecole Centrale Lyon, France Ste´phane Avril, Ecole Nationale Supérieure des Mines de Saint-Etienne, CIS-EMSE, SAINBIOSE, Saint Etienne, France; INSERM, Saint Etienne, France; and Université de Lyon, SAINBIOSE, Saint Etienne, France © 2019 Elsevier Inc. All rights reserved.

Introduction General Information Structure–Property Relationships Characterization of the Arterial Tissue Structure Multiscale Observation Techniques of the Arterial Structure Macrostructure (100 mm–5 mm) Microstructure (5–100 mm) Ultrastructure Hierarchical Organization of the Arterial Tissue Macrostructure Microstructure Ultrastructure Quantitative Characterization Techniques of the Arterial Microstructure Collagen and elastin mass fractions at the macroscopic scale Volume fraction of fibers at the macroscopic scale Quantitative Assessment of the Arterial Microstructure Universal pattern for the arterial composition Quantitative parameters related to the fiber network arrangement Characterization of the Arterial Tissue’s Mechanical Function Multiscale Characterization Techniques of the Arterial Mechanical Response At the macrostructural scale At the microstructural scale At the ultrastructural scale Multiscale Mechanical Response of the Arterial Tissue At the macroscopic scale At the microstructural scale At the ultrastructural scale Conclusion References Further Reading

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Glossary Autoclave pressure chamber used to carry out processes requiring elevated temperature and pressure different from ambient air pressure. In the sequel, autoclaving is used to separate soluble proteins (such as collagen) from nonsoluble proteins (such as elastin). Constitutive relation in mechanics, material-specific relation between stress and strain that characterizes the response of the material to an applied mechanical loading. In vivo pre stretch and pre stress The in vivo state of the arterial tissue is not free of mechanical loading; in particular, it is subjected to an axial elongation and residual stresses. Staining auxiliary technique used in microscopy to highlight structures in materials. For biological tissues, it involves the use of structure-specific dyes.

Introduction General Information In the cardiovascular system, the main role of arteries is to ensure the circulation of blood from the heart to peripheral capillaries, in order to deliver oxygen and nutrients to all tissues of the body. This physiological function can be maintained as long as the arterial

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system can receive spurts of blood from the left ventricle and distribute them as steady flow through peripheral capillaries. This implies that flow and pressure pulsations remain confined to the larger arteries which act as cushions through elastic dilatation and recoiling at every cardiac cycle (O’Rourke and Hashimoto, 2007). In more details, during the systolic phase, the blood pressure rises and the large arteries distend and store blood; then, when the blood pressure falls during the diastolic phase, they recoil and discharge the previously stored blood to the smaller (muscular) arteries. This specific mechanical property of the large elastic arteries is usually referred to as the Windkessel effect (Wagner and Kapal, 1952). Fulfilling these different functions involves ensuring the mechanical integrity of the arterial structure. Conversely, a failure in the mechanical integrity of arteries can have lethal consequences, due to a lack in oxygen supply to different organs. As a matter of fact, the World Health Organization reports that cardiovascular disorders are the first cause of death worldwide, killing 17.5 million of people yearly (i.e., 31% of deaths) (Sidloff et al., 2014; Santulli, 2013). Many of these cardiovascular disorders are directly related to dysfunctions of the arterial mechanical role. For instance, an aneurysm is a permanent localized dilatation of an artery having at least a 50% increase in diameter compared to the expected normal diameter of the artery in question (Johnston et al., 1991). It is characterized by a modified and generally damaged arterial microstructure as well as by strongly modified mechanical load conditions applied on the arterial wall, jeopardizing the mechanical integrity of the artery. When detecting an aneurysm, clinicians need to decide whether or not a surgical intervention has to be scheduled. Surgical interventions consisting in the replacement of the aneurysmal aorta by a vascular prosthesis may have a mortality rate oscillating between 3% and 5%. Therefore, taking a clinical decision in such circumstances requires reliable and patient-specific criteria about the potential risks of both elective repair and conservative management of the aneurysm. Currently, surgery is clinically indicated by an invariable, purely geometrical criterion combining diameter and growth rate thresholds, representing a trade-off between surgery mortality and rupture (Davies et al., 2002; Fillinger, 2007). However, it is widely acknowledged that this criterion is not a reliable predictor of rupture risk (Gasser et al., 2010). As another example of a common cardiovascular pathology, atherosclerosis manifests as thickening and hardening of arteries, with the formation of a fatty, stiff plaque inside the arterial wall. Atherosclerotic plaques progressively obstruct the lumen and therefore hamper oxygen delivery to the organs. In some cases, debris of plaque may detach from the wall and may block the blood flow downstream, eventually causing a heart attack or stroke. If one wants to adapt and improve the management of these cardiovascular disorders, there is a pressing need to develop methods for assessing the mechanical integrity of blood vessels accounting for their loading conditions and composition (deBotton and Oren, 2013).

Structure–Property Relationships The mechanical integrity of biological tissues primarily depends on: (1) the mechanical loading to which it is subjected, (2) its specific geometry, (3) its mechanical properties. At the tissue scale, the mechanical properties of arteries appear as extremely scattered and were shown to vary with aging and pathologies. They vary also across the considered organs or species (Lally et al., 2004; Weizsäcker et al., 1983; Holzapfel et al., 2005). However, this variability may mainly result from the variability of microstructures and compositions in the tested tissues (Holzapfel et al., 2005). Due to its biological nature, the arterial composition and microstructure are not invariant over time but actually change due to biological processes called growth and remodeling (Gibbons and Dzau, 1994; Humphrey, 1995). The latter depend not only on biochemical conditions but also on the mechanical loading itself (Leung et al., 1976; Humphrey, 2006; Humphrey et al., 2014), such that the tissue actively adapts to its bio-chemo-mechanical environment. Tissue adaptation is thought to be governed by the different populations of cells in the arteries which try to reach a homeostatic bio-chemo-mechanical state (Humphrey, 2008). Cardiovascular pathologies or aging manifest with altered remodeling which may severely impact the mechanical integrity of arteries at long term (Jacob et al., 2001). For instance, abnormally high blood pressure (hypertension) results in wall thickening and stiffening (Berry and Greenwald, 1976; Jacob et al., 2001). These altered conditions may lead to accelerated fatigue and failure of the different arterial wall constituents, such as the elastic lamellae, and therefore to an increased stiffening of the arterial wall. This may end up with transmission of the pulsatile blood flow to smaller vessels and organs. Also, in the aorta, (mechanical) damage to or (chemical) degradation of the elastic fibers, in combination with the loss of smooth muscle function, is a common early contributor to the formation or expansion of all aneurysms. Then, both remodeling of collagen (i.e., turnover) and inflammation (i.e., invasion of mononuclear cells such as lymphocytes T and B) play a fundamental role in aneurysm enlargement and rupture (Jacob et al., 2001). All these examples show that the vascular pathologies may be translated at a given time instant into microstructural changes, whether composition, arrangement, or mechanical properties of the individual constituents. Still, up to now, mechanical and microstructural investigations were achieved separately and rarely bridged together. However, improving diagnosis and treatment of cardiovascular pathologies necessarily implies to account for the structure–property relationships (Humphrey, 2009). In this article, we review all relevant information about the structure–property relationships in healthy large elastic arteries. We first review in “Characterization of the Arterial Tissue’s Structure” section their microstructure organization, their characterization, and their composition. Then we review in “Characterization of the Arterial Tissue’s Mechanical Function” section their multiscale mechanical properties, showing the large variability across species and organs.

Characterization of the Arterial Tissue Structure Arteries are characterized by a great diversity and variability. Depending on the species or on the location in the organism, they have to sustain and regulate different levels of blood pressure. Still, the hierarchical structure of the arterial wall remains unchanged

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throughout the vertebrate kingdom. The arterial wall is made up by three main elementary constituents, namely elastin (see Fig. 1(J)), collagen (see Fig. 1(H)), and water, as well as by different other organic molecules (e.g., glycosaminoglycans, proteoglycans). These elementary constituents are arranged in morphological hierarchical structures occurring in most of the large arteries of vertebrates and described by means of the following three levels: - At a characteristic length of several millimeters, the macrostructure of the arterial wall is characterized by varying thicknesses and mechanical properties depending on the precise biological function of the considered arterial wall (see Fig. 1(A)). This arterial macrostructure is composed, at a few hundreds of micrometers scale, of three concentric tunicae: the intima, media, and adventitia, see Fig. 1(B). Each of these layers has a specific morphology as well as a specific function within the arterial wall. This scale will be referred to as the macrostructure of the arterial wall. - At a characteristic length of some micrometers to some tens of micrometers, each layer is made up by an arrangement of cells embedded in different imbricated fibrous elastic and collagenous networks (see Fig. 1(C–F)). This will be referred to as the arterial microstructure. - Zooming at these fibrous networks reveals the ultrastructure of the fibers, made of an arrangement of cross-linked fibrils, at a hundred of nanometers scale, see Fig. 1(G, I). Understanding the arrangement of the different constituents within the arterial wall has been the source of an impressive amount of scientific publications; the emergence in the last 15 years of multiphoton microscopy has led to a better understanding of the constituent arrangement and of the relationship to their mechanical function. After a brief overview of the diverse observation techniques, the multiscale description of the arterial wall microstructure will be reviewed in details. As a second step, the qualitative hierarchical description of arteries is completed by the extraction of quantitative parameters characterizing this microstructure, such as arterial composition, fiber orientations, and shape.

Multiscale Observation Techniques of the Arterial Structure The hierarchical character of arteries has been revealed by the use of different microscopy techniques, offering a wide range of resolutions, as well as different imaging characteristics (in-depth resolution, need for prior staining, etc.).

Macrostructure (100 mm–5 mm) The composite structure of the artery has been originally revealed by histological studies using optical microscopy. Optical microscopy is the gold standard to observe the composition of tissues. It uses visible light and a system of lenses to magnify images of biological samples. The specimens are previously sectioned (cut into a thin cross section with a microtome), stained, and mounted on a microscope slide.

Microstructure (5–100 mm) Different microscopy techniques may be used to visualize the arrangement of the vascular microstructure at the micrometer scale. One common optical sectioning method allowing in-depth imaging is called confocal microscopy or confocal laser scanning (Voytik-Harbin et al., 2003). It consists in the installation of a spatial pinhole at the confocal plane of the lens, which acts as a spatial filter and allows only the in-focus portion of the light to be imaged. Confocal microscopy is characterized by an increased spatial resolution but decreased signal intensity, inducing long exposure times of the biological samples to the imaging beam and the use of photomultipliers. More recently, multiphoton microscopy, also called nonlinear or two-photon microscopy, has been developed in order to increase axial resolution and penetration depth (van Zandvoort et al., 2004). Under multiphoton microscopy, the sample is illuminated at twice its normal excitation wavelength by a high-energy short-pulsed laser beam, allowing fluorescence to be precisely localized to the illumination region. When imaging biological tissues, the signal of collagen is generated from the second harmonic generation, while elastin signal on the other hand is recovered by autofluorescence, before being collected by two bandpass filters. Multiphoton microscopy can also reveal vascular cells such as fibroblasts or smooth muscle cells (O’Connell et al., 2008) upon prior staining for fluorescence. The both aforementioned techniques provide high optical resolution and contrast while eliminating out-of-focus light. Eventually, they enable the reconstruction of three-dimensional structures from the obtained images by collecting sets of images at different depths. Diffusion tensor imaging (Vilanova et al., 2006) relies on a different technology: it uses the diffusion of water molecules to generate contrast in magnetic resonance images. As the molecular diffusion of water in tissues is constrained, it can reflect interactions with many obstacles when tracked, such as macromolecules, fibers, and membranes. This tracking approach provides a mapping of diffusion patterns revealing microscopic details about the tissue architecture. Tomographic imaging of biological tissues provides section views of organic components through the use of various penetrating waves (e.g., X-ray). Tomographic slices can be reconstructed into 3D views by means of specific reconstruction techniques (Fujimoto et al., 1999). Finally, the vascular microstructure can also be imaged using polarized light microscopy (Gasser et al., 2012). Polarized light is obtained by means of a polarizer oriented at 90 to filter out the directly transmitted light. When imaging biological tissues under polarized light, previously stained collagen expresses different interference colors influenced not only by the fiber thickness but also by the way collagen fibers are packed.

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Fig. 1 Hierarchical structure of large elastic arteries: (A) macroscopic view of large arteries of a mouse, taken from Keyes et al. (2011)); (B) at the millimeter scale, the arterial wall is made of three concentric layers, the adventitia tunica (A), the tunica media (M) and the intima tunica (I); electron micrograph of a rabbit femoral artery from Ratz (2016); (C) at the micrometer scale, the intima is made of a continuous layer of endothelial cells (inverted phase microscope image of bovine aortic endothelial cells from Ives et al. (1986); the media is made of an arrangement of medial lamellar units, as seen in (D) by means of scanning electron microscopy on a rat abdominal aorta (taken from O’Connell et al., 2008), and (F1) from multiphoton microscopy stack of images; zooming further on the lamellar unit, multiphoton microscopy allows to distinguish (F2) the collagen fibers and (F3) the elastin network; the adventitia is made of (E1) an arrangement of collagen bundles (also E2) and elastin fibers (also E3); finally, the collagen fibers are made of a (G) staggered arrangement of collagen fibrils (scanning electron microscope image taken from Ushiki, 2002), themselves made of an (H) arrangement of cross-linked collagen molecules (scanning electron microscope image taken from Ushiki, 2002); while the elastin network is made of (I) elastic fibers, lamellae, and struts, themselves made of an (J) arrangement of elastin and cross-linking molecules (scanning electron microscope image taken from Ushiki, 2002). Images (E1)–(E3, (F1)-(F3) were obtained by a multiphoton microscope (IVTV Platform, ANR-10-EQPX06-01, FR) imaging a rabbit carotid artery.

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Ultrastructure At the scale of a few tens to hundreds of nanometers, the resolved structure of the different fiber networks can be revealed by means of scanning electron microscopy. This technique uses a beam of accelerated electrons as a source of illumination. The higher resolving power of scanning electron microscopes originates in the wavelength of electrons being up to 100,000 times shorter than that of visible light photons used in optical microscopy. Prior acid and elastase digestion of the tissue can be performed in order to highlight specific cells.

Hierarchical Organization of the Arterial Tissue Use of these different microscopy techniques allowed revealing the hierarchical structure of arteries.

Macrostructure At the macroscopic scale, the artery is a composite cylindrical structure made of three concentric layers: the adventitia, the media, and the intima, as seen in Fig. 1(B), an electron micrograph obtained by Ratz (2016). Each of these layers is characterized by specific microstructures, specific thicknesses (Wolinsky and Glagov, 1967b), different mechanical properties (Holzapfel et al., 2005), and different structural and biological functions (O’Connell et al., 2008). The relative proportions of layer thicknesses differ in particular between proximal and distal regions (Canham et al., 1989). First, the most inner layer of the arterial wall is the tunica intima, made of the endothelium and of an internal elastic lamina. The endothelium is a monolayer of endothelial cells, lining the luminal surface of blood vessels, as shown in Fig. 1(C), obtained from (Ives et al., 1986) by means of an inverted phase optical microscope. It plays the role of an interface between the vessel wall and the blood flow (Ives et al., 1986). The endothelium is subjected to both fluid shear stresses and pressure-induced strains (Ives et al., 1986). More specifically, shear stresses affect the morphology and function of endothelial cells morphology. Different studies highlighted the tendency of the endothelium to align parallel to the principal axis of strain. It was also shown that the morphology of endothelial cells is closely related to its cytoskeletal structure (Ohashi and Sato, 2005). Moreover, the mechanical forces applied by the blood flow are essential factors in cardiovascular diseases, such as atherosclerosis, as they can alter the morphology and function of endothelial cells (Ives et al., 1986). The internal elastic lamina (not shown in Fig. 1) follows the endothelium and is known to provide structural cohesion and support for axial pretension (Farand et al., 2007; Timmins et al., 2010). Then, the tunica media, the second concentric layer of the arterial wall, is a concentric set of superimposed medial lamellar units, as reconstructed in Fig. 1(F1) from a stack of multiphoton microscope images. A single medial lamellar unit, as shown in Fig. 1(D), obtained from O’Connell et al. (2008), by means of scanning electron microscopy, consists in a row of overlapping smooth muscle cells surrounded on the upper and lower sides by two concentric elastic lamellae. The overlapping smooth muscle cells lie parallel to tangential planes of the circumferential direction. The number of lamellar units in the media of adult mammalian aortas has been shown to be nearly proportional to the aortic radius regardless of species or of variations in measured wall thickness (Wolinsky and Glagov, 1967a, 1967b). The tunica media is therefore, from a morphological point of view, a composite material organized periodically: the radial transmural disposition of cells and matrix fibers on transverse sections of the media in well-developed aortas is proved to be cells, elastic lamellae, cells. Finally, the tunica adventitia, the third and most outer concentric layer of the arterial wall, is made of an arrangement of collagen bundles and few elastic fibers, as seen in Fig. 1(E1) obtained from reconstruction of a stack of multiphoton microscope images, together with embedded fibroblasts.

Microstructure At the micrometer scale, microscopy techniques reveal the precise morphology of the different fiber networks that exist in each arterial layer. Starting from the arterial lumen, in the endothelium, the actin filaments (F-actin) are one of the major cytoskeletal structures of the endothelial cells (Ookawa et al., 1992); this actin network also exists in the smooth muscle cells. They are organized in bundles and are usually grouped in the central part of the cells. Noticeably, the redistribution of F-actin filaments within the cells is one of the early cellular responses to the onset of shear stress: when experiencing low-shear forces coming from the blood flow, Factin filaments localize at the periphery of the endothelial cells (Ookawa et al., 1992); whereas when experiencing high shear forces, F-actin bundles are observed in the central part of the elongated cells. In parallel, the cells orient in the direction of the applied flow. A continuous sheet of elastin, called the internal elastic lamina (Farand et al., 2007; Timmins et al., 2010), surrounds the endothelium and marks a frontier between the endothelium and the tunica media. It takes the form of a dense elastin sheet equipped with a longitudinal network of elastic fibers coated on its continuous surface. Then, as mentioned earlier, the medial lamellar unit is made of an arrangement of elastic fibers, vascular smooth muscle cells, and collagen fibers, which are now successively described. From a morphological point of view at the micrometer scale, medial elastin takes three different forms: - First, lamellae, composed of a dense meshwork of elastic fibers oriented circumferentially (Farand et al., 2007; Timmins et al., 2010). They form a periodical concentric separation between medial lamellar units containing the vascular smooth muscles cells, see Fig. 1(F1) and Fig. 2(E1). These lamellae show large, round, reinforced fenestrations (see Fig. 2(E2), also visible on the upper lamella of Fig. 1(F1)), and house the superior and inferior anchorages of smooth muscle cells, allowing them to weave through the tunica media (Dingemans et al., 2000);

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Fig. 2 Schematic representation of the medial lamellar unit morphology. (E1) elastin lamella, (E2) fenestration of the elastin lamella, (E3) elastin interlamellar struts, (E4) elastin surface ridges, (SMC) smooth muscle cells, (C1) collagen fibers, (C2) cohesive collagen microfibril bundles, (C3) collagen envelope of SMCs. Inspired from illustrations by Dingemans, K. P., et al. (2000). Extracellular matrix of the human aortic media: An ultrastructural histochemical and immunohistochemical study of the adult aortic media. Anatomical Record 258 (1), 1–14; O’Connell, M. K., et al. (2008). The three-dimensional micro- and nanostructure of the aortic medial lamellar unit measured using 3D confocal and electron microscopy imaging. Matrix Biology 27 (3), 171–181.

- Second, thick radial elastin struts provide structural cohesion to the overall medial lamellar unit, while preserving an angular 20 tilt with respect to the smooth muscle cell orientation, see Fig. 2(E3) (Dingemans et al., 2000; Koch et al., 2014; Tsamis et al., 2013a); - Finally, thin radial elastic fibers take the form of ridges (Dingemans et al., 2000) or protruding ribs (Raspanti et al., 2006), connecting the smooth muscle cells to both lamellae, see Fig. 2(E4). Through their actively adaptive plasticity, the vascular smooth muscle cells are responsible for the regulation and maintenance of blood flow and for the regulation of stress across the arterial wall thickness. For this reason, vascular smooth muscle cells harbor actin–myosin filaments (myofilament) that permit rapid stress development and sustain stress maintenance and vessel constriction (Ratz, 2004). The relaxed smooth muscle cell is a very long and thin (fusiform) structure (Ratz, 2004) with a high surface area, and an ellipsoidal shape of its nucleus (O’Connell et al., 2008). Upon contraction, the vascular muscle cells can undergo dramatic shortening accompanied by shape change, surface rearrangements, and a loss of volume that is recovered upon relengthening. As for their spatial positioning among the constituents of the medial lamellar unit, it shows noticeable characteristics, namely, the cells weaving throughout the interlamellar elastin framework, resulting in approximately 20 radial tilt (O’Connell et al., 2008). Collagen is also organized in different morphologies within the tunica media. First, interlaced bundles of type IV collagen microfibrils (oxytalan fibers) of the immediate pericellular matrix contribute, along with elastin radial struts and interlamellar elastin protrusions, to smooth muscle cell cohesion and fixation on the lamellae, see Fig. 2(C2) (Clark and Glagov, 1985; Dingemans et al., 2000). It is understood that the smooth muscle cells preferentially adhere to these ill-defined streaks rather than directly to the solid lamellae, see Fig. 2(SMC) (Dingemans et al., 2000); second, wavy collagen fiber bundles are interposed between the facing elastin systems within the fibrous regions between cell layers, see Fig. 2(C1) and Fig. 1(F2) (Clark and Glagov, 1985). These collagen fibers are oriented circumferentially (O’Connell et al., 2008; Timmins et al., 2010; Roy et al., 2010; Hill et al., 2012) and are closely associated with the elastic lamellae (Dingemans et al., 2000) but not with the smooth muscle cells. Upon pressure, these medial collagen fibers decrimp and stretch to prevent over distension of the vessel. Collagen takes also the form of membranes enveloping the smooth muscle cells (see Fig. 2(C3)). Finally, the most outer arterial layer, the tunica adventitia, is made of an arrangement of networks of elastin and collagen with embedded fibroblasts. In the adventitia, elastin takes the form of a low-density meshwork made of variously oriented fibers showing bifurcations (transversely oriented segments), with a dominant longitudinal direction (Chen et al., 2011, 2013; see Fig. 1(E3)). Contrarily, collagen fibers pack into thick bundles of 10–30 fibers, folded (crimped) under the in vivo prestress and prestretch conditions. These wavy bundles are oriented helicoidally about the 45 angle and show a negligible transmurality (Roy et al., 2010; Rezakhaniha et al., 2012; Schrauwen et al., 2012; see Fig. 1(E1) and (E3)). For a flat piece of arterial tissue, the crimping appears considerably more important than for a cylindrical piece of arterial tissue, and the bundles are oriented closer to the axial direction (Tsamis et al., 2013a; Phillippi et al., 2014; Koch et al., 2014; D’Amore et al., 2010). The adventitial collagen network is capable of undergoing important morphology rearrangements under mechanical load, namely, decrimping, stretching, and reorientation in the load direction, with amplitudes that can exceed affinely predicted reorientation (Billiar and Sacks, 1997; Chandran and Barocas, 2006). These changes in morphology are known to be linked to the significant nonlinearity of the mechanical response and give rise to a very pronounced material stiffening under large strain (Chen et al., 2011; Schrauwen et al., 2012). These aspects are further detailed in section “Characterization of the Arterial Tissue’s Mechanical Function” of this article. The adventitial

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microstructure is also composed of fibroblasts (not shown on Fig. 1), which are mechanosensitive cells, participating to arterial remodeling and arranged circumferentially about collagen bundles (Esterly et al., 1968).

Ultrastructure At the hundreds of nanometers scale, the structure of elastic and collagen fibers is revealed, as well as the existence of cross-links between and within these fibers. The morphology of the elastic lamellae presents a fibrous texture suggestive of a “criss-crossed”, delicate filamentous scaffold (Ushiki, 2002; Raspanti et al., 2006). The elastic fibers and elastin meshwork however are made of 0.2 mm thick elastin fibrils and microfibrils that run in various directions (see Fig. 1(I) and (J)). Those microfibrils are coated together within an elastic fiber by proteoglycans and glycosaminoglycans (Ushiki, 2002). At the same scale, collagen bundles are made of closely packed, parallel, thin collagen fibrils, with a characteristic diameter of 30–100 nm, with altering number of coated fibrils depending on the region in the bundle (Ushiki, 2002; Raspanti et al., 2006). In the adventitia, these bundles are thicker than in the media due to a higher number of constituting fibrils, see Fig. 1(G). The fibrils present a regular, orthogonal lattice of surface-bound proteoglycans (Raspanti et al., 2006; Berillis, 2013) (see Fig. 1(H)). Finally, at the nanometer scale, the cohesion between fibrils constituting elastic or collagen fibers is realized by cross-linking proteoglycans, formed by covalent bonding between acid glycosaminoglycans and proteins (Eisenstein et al., 1975). At the microfibrillar level, collagen and elastin are cross-linked by a unique mechanism based on aldehyde formation from lysine or hydroxylysine side chains (Eyre et al., 1984; Sáez et al., 2016).

Quantitative Characterization Techniques of the Arterial Microstructure Many tests were performed over the last 70 years to determine the relative amount of elastin, collagen, and water, among different arteries, species, and at different ages, resulting in a broad variety of arterial compositions. The relative mass or volume fractions of collagen and elastin can be determined at different scales, depending on the chosen technique: chemical methods provide average weight fractions of collagen and elastin over a millimeter-sized dehydrated defatted arterial sample, while histochemical methods allow determining the volume fractions of collagen and elastin fibers in the arterial tissue.

Collagen and elastin mass fractions at the macroscopic scale Associated experimental endeavors involve primarily the following steps: - The arterial samples are generally excised from freshly sacrificed animals and analyzed directly after excision, or wrapped airtight in plastic films and stored at  20 C (Brüel and Oxlund, 1996; Looker and Berry, 1972). In some cases (Leung et al., 1977; Berry and Greenwald, 1976; Myers and Lang, 1946), the adventitia is removed by careful dissection, and only the compositions of the media and the intima are determined. - Defatting the arterial wall is performed by means of successive immersions in acetone and ether, for different durations depending on the precise protocol (Fischer and Llaurado, 1966; Harkness et al., 1957; Grant, 1967; Neuman and Logan, 1950; Feldman and Glagov, 1971); - The dehydration procedure consists in either drying to constant weight in vacuum (Harkness et al., 1957; Looker and Berry, 1972; Farrar et al., 1965; Hosoda et al., 1984; Berry and Greenwald, 1976; Feldman and Glagov, 1971; Spina et al., 1983) or drying to constant weight in an oven for few hours at a temperature higher than 50 C (Grant, 1967; Fischer and Llaurado, 1966; Lowry et al., 1941; Neumann and Logan, 1950; Leung et al., 1977). As a result, water represents about 70%–80% of the wet weight of the aortic tissue (Looker and Berry, 1972; Fischer and Llaurado, 1966; Dahl et al., 2007). - Although different methods exist for the detection and estimation of collagen mass fraction in a biological sample, the determination of hydroxyproline in the tissue seems to be the most widely used method. Hydroxyproline is an amino acid, whose content varies between 13.1% in collagen type I to 17.4% in collagen type III (Etherington and Sims, 1981). As elastin also contains hydroxyproline, collagen needs first to be removed from the tissue by autoclaving. This separation method between elastin and collagen accounts for the nonsoluble property of elastin. More precisely, a small piece of the dry defatted tissue is autoclaved twice for 3 h at 1 bar, [Pressure and duration of autoclaving were varied by different experimentalists: opted for 6 h at 1 bar (Grant, 1964; Grant, 1967); for 6 h at 2.75 bar (Feldman and Glagov, 1971); for 18 h at 1–1.4 bar (Fischer and Llaurado, 1966) followed by a second for 3 h; for 6 h at 2 bars (Harkness et al., 1957; Looker and Berry, 1972; Farrar et al., 1965; Berry and Greenwald, 1976); for 4 h at 3.5 bars (Lowry et al., 1941).] washed with water and centrifuged, and the resulting extracts are dried out (Neuman and Logan, 1950). Then, the following steps (with possible slight adaptations) are followed for hydroxyproline determination: (i) hydrolysis at 3.4 bar for 3 h with 6 mol L 1 hydrochloric acid to release the hydroxyproline from peptide linkage, (ii) oxidation with sodium peroxide, and (iii) formation of a reddish purple complex with p-dimethylaminobenzaldehyde (Harkness et al., 1957; Looker and Berry, 1972; Farrar et al., 1965; Berry and Greenwald, 1976; Neumann and Logan, 1950). Chloramine T together with Ehrlich’s reagent for colorimetric determination was preferred for use as an oxidant instead of sodium peroxide (Stegemann and Stalder, 1967; Hosoda et al., 1984; Brüel and Oxlund, 1996; Han et al., 2009). The collagen content follows from a multiplicative correction of the hydroxyproline content; this correction factor varies among the sources between 7.14, 7.46, or 7.8 depending on the exact content of hydroxyproline in the different collagen types (Neuman and Logan, 1950; Feldman and Glagov, 1971; Fischer and Llaurado, 1966; Grant, 1964, 1967; Harkness et al., 1957; Berry and Greenwald, 1976).

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- After autoclaving the arterial sample, the insoluble elastin remains in the residue, and its content is determined by a gravimetric method (Looker and Berry, 1972; Berry and Greenwald, 1976; Farrar et al., 1965; Leung et al., 1977; Hosoda et al., 1984; Lowry et al., 1941; Myers and Lang, 1946; Feldman and Glagov, 1971). In short, this method consists in purifying the residue remaining after collagen extraction, first by heating for 45 min at 100 C with a 0.1 mol L 1 sodium hydroxide solution (Lansing et al., 1952) and then by washing with water, dehydrating with acetone, and finally drying to constant weight in a hot air oven (Grant, 1967) or in vacuum over phosphorous pentoxide (Harkness et al., 1957; Looker and Berry, 1972) and weighing. Other methods for elastin determination include the same procedure of hydroxyproline determination as for collagen (Fischer and Llaurado, 1966; Harkness et al., 1957; Grant, 1964, 1967; Leung et al., 1977) and the determination of elastin-specific cross-linking amino acids (desmosine and isodesmosine) (Han et al., 2009). Finally, the volume fractions of cells and/or the cell content were also determined in several studies: dosage of DNA gives access to the total cell content of arterial tissue, being of 200 million of cells per milligram of dry defatted tissue in adult rats (Dahl et al., 2007), while histological observations with prior cell staining provide access to the volume fraction of some specific cell populations (O’Connell et al., 2008; Tonar et al., 2008).

Volume fraction of fibers at the macroscopic scale Histochemistry consists in studying the chemical constituents of a tissue by means of staining reagents. The arterial tissue is first fixed with paraffin and then stained with different dyes. In particular, collagen stains include (blue or green) Masson’s trichrome (Phillippi et al., 2014; Anidjar et al., 1990; Dahl et al., 2007), Picrosirius red (Phillippi et al., 2014), (yellow) Safranin-O staining (Azeloglu et al., 2008), (pink) eosin (Carmo et al., 2002), Van Gieson staining, which is a mixture of picric acid and acid fuchsin, staining collagen in red, nuclei in black, and cytoplasm in yellow (Carmo et al., 2002; Cattell et al., 1996). Elastin stains include (blue-black) Verhoeff’s staining (Tonar et al., 2003; Phillippi et al., 2014; Dahl et al., 2007), permanganate–bisulfite–toluidine blue reaction (Clark and Glagov, 1985; Fischer, 1979), (red) orcein staining (Hosoda and Minoshima, 1965; Anidjar et al., 1990; Scarselli and Repetto, 1959; Scarselli, 1959), Weigert’s (blue-black) resorcin-fuchsin staining (Hayashi et al., 1974). Histochemical methods also allow cells staining: toluidine blue stains nucleic acids in blue, hematoxylin (also called Weigert’s iron hematoxylin stain) stain cells nuclei in dark blue, eosin stains in pink the cell’s cytoplasm (Carmo et al., 2002); finally, Movat’s stain allows to identify glycosaminoglycans in blue (Dahl et al., 2007). The stained histological section is observed under a microscope and image processing techniques provide access to quantitative parameters. For example, several methods have been proposed to analyze the statistical fiber orientation based on microstructure imaging; they involved Hough transforms (Chaudhuri et al., 1993; Karlon et al., 1998), structure tensor-based texture analysis (Rezakhaniha et al., 2012), direct fiber tracking (Pourdeyhimi, 1999; Mori and van Zijl, 2002; Rezakhaniha et al., 2012; Hill et al., 2012; Ghazanfari et al., 2012), or 2D Fast Fourier transform (Ayres et al., 2006; Ayres et al., 2008; Timmins et al., 2010; Schriefl et al., 2012; Polzer et al., 2013).

Quantitative Assessment of the Arterial Microstructure Universal pattern for the arterial composition The mechanical properties of arterial tissues depend primarily on the tissue composition and on the arrangement of the constituents of its microstructure. Since access to the arterial composition is complex in vivo, we here propose to seek for such a general composition rule, valid over a wide variety of organs, species, and ages. To our best knowledge, no such relationship governing the arterial composition has ever been proposed. At most, correlations between the arterial composition and some other characteristics of the tissue have been established, but generally limited to very restricted sets of data. We here collected the mass fractions of elastin and collagen in various large elastic arteries from a great variety of species and ages: from rats to human, from the abdominal to the carotid arteries, etc. However, since important modifications in the arterial composition occur in the perinatal and early childhood periods (Bendeck and Langille, 1991), as well as in aging organisms (Tsamis et al., 2013b; Schlatmann and Becker, 1977), we restricted our selection to adult organs with no aging effects. In the perinatal period, a rapid accumulation of both elastin and collagen has been observed followed by a marked postnatal increase in arterial pressure, see (Bendeck and Langille, 1991). Upon aging, the absolute collagen content is increased, while the absolute elastin content remains constant, although the elastin becomes fragmented (Tsamis et al., 2013b; Schlatmann and Becker, 1977). We plot the collagen mass fraction as a function of the elastin mass fraction, for cases where those contents are related to the media and intima only (see Fig. 3) or to the whole arterial thickness (see Fig. 4). It is interesting to notice that collagen and elastin contents strongly correlate mutually, whether only the intima–media is considered or the entire thickness of the arteries. A closer look into the resulting graphs shows that, as a rule, the thoracic aortas are characterized by a larger content in elastin as compared to the abdominal aortas, which contain more collagen (compare the filled markers in Figs. 3 and 4 to the empty ones). This is in agreement with observations made on more restricted arterial tissue origins and numbers by Harkness et al. (1957), Sokolis (2007), Halloran et al. (1995), Cheuk and Cheng (2005): the elastin content decreases from proximal to distal regions. In parallel, it was also observed that the number of lamellar units also decreases in more distal aortas (Wolinsky and Glagov, 1969; Sokolis et al., 2002a,b). These changes in the arterial composition affect the mechanical behavior, as reviewed in section “Characterization of the Arterial Tissue’s Mechanical Function”.

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65 60 55

WF collagen

50 45 40 35 30 25 20 15 10 10

15

20

25

30

35

40

45

50

55

60

65

WF elastin Fig. 3 Composition of the media and intima layers of different arteries stemming from rabbit (upward pointing triangles), rats (diamonds), and human (right pointing triangles). Collagen and elastin weight fractions (WF) are related to the dry defatted weight of the arterial tissue. Filled markers correspond to thoracic arterial segments, while empty markers correspond to abdominal arterial segments as well as one pulmonary trunk tissue and one arch tissue. Data were taken from Leung, D. Y., Glagov, S., Mathews, M.B. (1977). Elastin and collagen accumulation in rabbit ascending aorta and pulmonary trunk during postnatal growth. Correlation of cellular synthetic response with medial tension. Circulation Research 41 (3), 316–323; Berry, C. L., Greenwald, S. E. (1976). Effects of hypertension on the static mechanical properties and chemical composition of the rat aorta. Cardiovascular Research 10 (4), 437–451; Brüel, A., Oxlund, H. (1991). Biosynthetic growth hormone changes the collagen and elastin contents and biomechanical properties of the rat aorta. Acta Endocrinologica 125 (1), 49–57; Han, W.-Q., et al. (2009). Changes in the composition of the thoracic aortic wall in spontaneously hypertensive rats treated with losartan or spironolactone. Clinical and Experimental Pharmacology and Physiology 36 (5–6), 583– 588; Andreotti, L., et al. (1985). Aortic connective tissue in ageing – A biochemical study. Angiology 36 (12), 872–879; Feldman, S. A., Glagov, S. (1971). Transmedial collagen and elastin gradients in human aortas: Reversal with age. Atherosclerosis 13 (3), 385–394; Apter, J. T., Rabinowitz, M., Cummings, D. H. (1966). Correlation of visco-elastic properties of large arteries with microscopic structure. Circulation Research 19 (1), 104–121; Tonar, Z., et al. (2003). Microscopic image analysis of elastin network in samples of normal, atherosclerotic and aneurysmatic abdominal aorta and its biomechanical implic ations. Journal of Applied Biomedicine 1 (3), 149–159; Spina, M., et al. (1983). Age-related changes in composition and mechanical properties of the tunica media of the upper thoracic human aorta. Arteriosclerosis, Thrombosis, and Vascular Biology 3 (1), 64–76.

Such a good correlation between elastin and collagen content could be used to determine the composition of healthy adult arteries, or to verify if a given arterial segment is healthy.

Quantitative parameters related to the fiber network arrangement Along with a detailed assessment of the vascular biological composition, many studies have focused on the extraction of particular geometrical characteristics of the microstructure (Humphrey and Holzapfel, 2012). Nonexhaustively, the implemented image analysis methods focused on the analysis of the fiber diameters (D’Amore et al., 2010; Phillippi et al., 2014), the fiber lengths (D’Amore et al., 2010; Rezakhaniha et al., 2011; Hill et al., 2012; Tsamis et al., 2013a; Cicchi et al., 2014), fiber volume fractions (Tonar et al., 2003; Hayashi et al., 1974; O’Connell et al., 2008; Verheyen et al., 1987), as well as on the evaluation of fiber waviness (Hill et al., 2012; Roy et al., 2010; Rezakhaniha et al., 2012; Schrauwen et al., 2012), and orientations (Holzapfel, 2006). More recently, for modeling purposes, more complex morphological features were quantified, such as fiber tortuosity (Koch et al., 2014), node connectivity and spatial intersections density (D’Amore et al., 2010; Koch et al., 2014), or density of transversely oriented segments (Koch et al., 2014). The number of elastic lamellae and of fenestrations inside the lamellae (Brüel and Oxlund, 1996) and appearance of collagen fibers (Wolinsky and Glagov, 1964) have also been investigated using image processing techniques.

Characterization of the Arterial Tissue’s Mechanical Function Arteries exhibit a highly nonlinear mechanical behavior, which has long been investigated from different point of views. As a first approach to arterial biomechanics, the macroscopic response of the arterial wall has been characterized through different uniaxial, biaxial, and tension–inflation tests. More recently, experimental setups coupling mechanical testing with live microscopy have been developed to decipher the microstructural mechanisms that are behind the nonlinear character of the response. Finally, at the 100 nm scale, the constitutive responses of collagen and elastic fibers have been characterized. We here restrict the review to in vitro mechanical testing techniques.

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60 55 50

WF collagen

45 40 35 30 25 20 15 10 10

15

20

25

30

35

40

45

50

55

60

WF elastin Fig. 4 Composition of the whole arterial segment of different arteries stemming from pigs (hexagrams), dogs (circles), sheep (downward pointing triangles), rats (diamonds), goats (upward pointing triangles), bovines (squares), puppies (pentagrams), and humans (right pointing triangles). Collagen and elastin weight fractions (WF) are related to the dry defatted weight of the arterial tissue. Filled markers correspond to thoracic and arch arterial segments, while empty markers correspond to abdominal and other arterial segments. Data were taken from Grant, R. A. (1967). Content and distribution of aortic collagen, elastin and carbohydrate in different species. Journal of Atherosclerosis Research 7 (4), 463–472; Grant, R. A. (1964). Estimation of hydroxyproline by the autoanalyser. Journal of Clinical Pathology 17, 685–686; Neuman, R. E., Logan, M. A. (1950). The determination of collagen and elastin in tissues. Journal of Biological Chemistry 186 (2), 549–556; Fischer, G. M., Llaurado, J. G. (1966). Collagen and elastin content in canine arteries selected from functionally different vascular beds. Circulation Research 19 (2), 394–399; Harkness, M. L. R., Harkness, R. D., McDonald, D. A. (1957). The collagen and elastin content of the arterial wall in the dog. Proceedings of the Royal Society B: Biological Sciences 146 (925), 541–551; Looker, T., Berry, C. L. (1972). The growth and development of the rat aorta. II. Changes in nucleic acid and scleroprotein content. Journal of Anatomy 113 (Pt 1), 17–34; Farrar, J. F., Blomfield, J., Reye, R. D. K. (1965). The structure and composition of the maturing pulmonary circulation. Journal of Pathology and Bacteriology 90 (1), 83–96; Hosoda, Y., et al. (1984). Age-dependent changes of collagen and elastin content in human aorta and pulmonary artery. Angiology 35 (10), 615–621.

Multiscale Characterization Techniques of the Arterial Mechanical Response At the macrostructural scale Due to the biological nature of arterial tissue, the storage conditions and the temperature prevailing during the mechanical tests may have an impact on the in vitro mechanical response of the arterial tissue. Concerning storage procedures, the arteries are usually harvested in freshly sacrificed animals or shortly after death. After excision, the arterial sample may be stored for 2–3 days in a saline solution at 4 C, or kept frozen at  20 C or at  80 C (Collins and Hu, 1972; Pham et al., 2013), to avoid sample drying and accelerated sample degradation. The impact of these different protocols on the mechanical properties was investigated by comparing the uniaxial tensile response of specimens stemming from the same tissue sample but subjected to different preparation protocols. Concerning the temperature of the mechanical test, two choices are classically made: either the ambient temperature of the room or the physiological temperature. Again, the uniaxial tensile responses relative to different temperatures of tests are compared. In order to characterize the arterial constitutive behavior, biomechanics imported the well-established uniaxial and biaxial tensile tests existing for metallic or inert materials to the arterial tissues: across the last 17 years, different arterial tissues stemming from different species and organs, at different ages, and different healthy or pathological states have been mechanically characterized. More precisely, for uniaxial tests, a dog bone-shaped sample is cut from the excised arterial segment, with the circumferential or longitudinal direction as main axis. Because uniaxial tensile tests do not reproduce the in vivo load conditions to which arteries are subjected, biaxial tests on flat squared samples have been developed. The load is applied simultaneously in both directions, maintaining a constant stress ratio between the two directions. Finally, tension inflation tests make use of a cylindrical arterial segment, subjected to a fixed axial stretch and a variable internal pressure.

At the microstructural scale Several studies have dealt with the possible relation between the observed variability in the macroscopic mechanical response and the variability in the arterial wall composition or organization. By combining mechanical tests to histology or to the chemical analysis of the arterial composition, such studies seek for relations between arterial composition/morphology and subsequent mechanical response. The influence of each fiber network on the overall mechanical response has also been investigated: enzymes such as

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collagenase or elastase are used to (partially) degrade one constituent, and the mechanical response of the degraded system is then compared to that of the original system, and to those of systems with different degrees of degradation. Besides, different studies proposed to physically separate the three arterial layers, so as to study the contribution of each arterial layer on the overall mechanical response (Weisbecker et al., 2012; Holzapfel et al., 2005; Sommer et al., 2010). Aiming at deciphering the microstructural mechanisms that lead to the nonlinear mechanical response, original testing devices have been developed, permitting live imaging of the arterial microstructure with a multiphoton microscope during the application of the load. Such tests, coupled to the previously described image processing techniques, permit the characterization of the rearrangements of the different fiber networks.

At the ultrastructural scale At the ultrastructural scale, attempts were made to characterize the mechanical behavior of the elementary constituents of the arterial wall. We here focus on collagen and elastin. Concerning collagen, this task is more complex because of the existence of different collagen types, with different mechanical properties. Our review is restricted to the characterization of collagen type I. This is the most abundant collagen type in bone, tendon, as well as in the tunica adventitia of arteries. Different characterization techniques have been employed for retrieving its mechanical characteristics at different length scales. Brillouin light scattering (Harley et al., 1977; Cusack and Miller, 1979) gives access to the velocity of elastic waves having a wavelength of several hundreds of nanometers; hence, it allows for characterizing the arrangement of several collagen molecules within fibrils. At a higher scale, nanoindentation tests performed by means of an Atomic Force Microscope provide access to the mechanical properties of collagen fibrils (Yadavalli et al., 2010; Strasser et al., 2007; Wenger et al., 2007; Graham et al., 2004; van der Rijt et al., 2006). Finally, the mechanical response of collagenous tissues at higher scales, from the fibrils to the tissue scales, is characterized by tensile tests, with different types and characteristic sizes of samples. The mechanical properties of collagen fibrils and fibers primarily depend on the hydration degree of the tissue under consideration. The picture is quite different regarding the mechanical characterization of elastin or elastic fibers. Elastic fibers are often characterized by the mechanical properties of collagen-free biological tissues (such as arterial tissue or nuchal ligaments), considering that the ground substance surrounding the elastic fibers is not mechanically active. Collagen is removed either by autoclaving (Aaron and Gosline, 1981) or by alkali treatment (Hass, 1942b). Then, uniaxial tensile tests are performed on the extracted elastin fibers.

Multiscale Mechanical Response of the Arterial Tissue The biological nature and the physiological functions of the arterial tissue render the characterization of the mechanical properties as a delicate task: harvesting the tissue implies the death of cells, with different consequences: in vitro characterization only investigates the passive response of the tissue and therefore cannot account for the active role of smooth muscle cells in distributing the load across the tissue thickness; the termination of the constituent turnover and the progressive degradation of the organic constituents making up the tissue have consequences on the mechanical properties of the tissue; harvesting the tissue also implies the relaxation of some of the prestress and prestretch existing in the arterial tissue in vivo, leading to fiber rearrangements in the microstructure with consequences on the macroscopic mechanical response.

At the macroscopic scale Before reviewing the macroscopic mechanical properties of arteries, efforts were made toward revealing the influence of sample freezing or refrigerating on the resulting mechanical behavior. Several studies (Zemánek et al., 2009; Adham et al., 1996; Armentano et al., 2006) reported no significant variations of the mechanical response after storage, while, in other studies (Venkatasubramanian et al., 2006; Chow and Zhang, 2011; Stemper et al., 2007), the variations in the mechanical behavior encompass variations in the initial and final stress–strain slopes, as well as changes in the knee point of stress strain curves, and in the ultimate stress. These variations may be explained by some damage occurring in the sample during freezing or refrigerating: formation of ice crystals, bulk water movement (Venkatasubramanian et al., 2006) can induce fiber cracking, loss of cross-links, networks disruption, and death of cells. These variations may also be explained by the decrease in the collagen content after 48 h cold storage (Chow and Zhang, 2011), as well as by the exact procedure followed to freeze the sample. Still, it is impossible to decide on the directions of variations, since the different relevant studies come to apparently contradictory results. It is however generally admitted that freezing better maintains the mechanical properties of arterial samples than refrigeration (Chow and Zhang, 2011; Stemper et al., 2007). Zemánek et al. (2009) studied the influence of the temperature choice on the mechanical response of the arterial wall and showed that samples are stiffer at ambient temperature than at in vivo temperature, since a temperature increase by 1 C results in a 5% stiffness decrease; this result is in good agreement with Fung (1993). Furthermore, arteries are subjected in vivo to residual stresses and prestretch, as evidenced by Bergel (1961), Fung (1993), and Vaishnav and Vossoughi (1987), the amount of which depends on the organ and species under consideration; this in vivo stress– strain state originates in the growth and remodeling processes undergone by arteries and allows achieving a homeostatic stress state, being nearly uniform and equibiaxial across the arterial wall thickness (Humphrey, 2009). As a consequence, excising and cutting open arterial segments partially release this existing in vivo stress–strain state, but there is no certainty that the load free configuration corresponds to a stress- and strain-free configuration.

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Another important feature of the arterial constitutive behavior is the existence of a transient mechanical response: during the first mechanical cycles, the mechanical response of the arterial wall exhibits an important hysteresis which is reduced after several load cycles; the stabilized mechanical response barely shows any hysteresis. Experimentalists usually get rid of this transient response by performing several preconditioning cycles. The number of preconditioning cycles varies with the precise protocol, and the loading path and maximum load generally coincide with the further applied loading (see column 6 of Table 1). The microstructural underlying mechanisms occurring during preconditioning have not been elucidated yet. The mechanisms are probably related to viscous effects, since the transient response is observed after a prolonged stop of the mechanical loading. Interestingly, Zemánek et al. (2009) noticed that no preconditioning was necessary for equibiaxial tensile tests. The previously described macroscopic testing procedures allow characterizing the mechanical response of a millimeter-sized arterial sample (see Table 1 for a literature review). The arterial wall exhibits a highly nonlinear mechanical response, which was already described in the 1880s (Roy, 1881): while at low applied stresses arteries are very easily deformed, the arterial response becomes much stiffer at higher applied stresses. This nonlinear response occurs in any load direction. As many other biological tissues, arteries exhibit an important variability in their mechanical response, both across species, organs, location, or interindividual (see Fig. 5). As a result, there is an important variability in the stretch and stress levels at which the change of stiffness reaches its maximum. Concerning tension–inflation tests, the arterial response is characterized by the variations of the arterial diameter as a function of the applied pressure as well as by the variations of the axial reaction force (Cox, 1975; Weizsäcker et al., 1983; Dobrin, 1986). The tension–inflation tests evidence the salient feature of the in vivo prestretch level; indeed, for axial prestretches being smaller (resp. larger) than the in vivo prestretch, the axial reaction force decreases (resp. increases) when the pressure increases. But, when the axial prestretch is equal to the in vivo prestretch, the axial reaction force does not depend on the applied inner pressure (Weizsäcker et al., 1983; Sommer et al., 2010) and remains constant during the pressure cycle (see Fig. 6, 4th curve from above). Similar results exist for isotonic tests, in which the axial force is kept constant but the axial prestretch varies with the applied pressure (Cox, 1975). Anisotropy of arterial walls has also been investigated through these experimental setups (Vosshougi and Weizsäcker, 1985; Lally et al., 2004). But no general conclusion can be drawn regarding the anisotropy of the arterial tissue. Still, carotid arteries are found stiffer in the circumferential direction (Cox, 1975; Patel and Janicki, 1970), while coronary arteries are stiffer in the longitudinal direction (Papageorgiou and Jones, 1988; Patel and Janicki, 1970). It is however remarkable that there is no statistical difference in the mechanical responses of arteries whether tested in the circumferential or in the longitudinal directions, when the tests are performed in conditions close to the physiological ones: this was observed, e.g., by Sato et al. (1979) on dog abdominal aortas, by Dobrin (1986) on dog carotid arteries, and by Sommer et al. (2010) on human carotid arteries. Another much debated feature is the strain-rate dependence of the arterial mechanical response. It is now admitted that at low strain rates, the mechanical response of arteries does not vary with the strain rate (Sato et al., 1979; Zemánek et al., 2009; Tanaka and Fung, 1974). The viscous character of arteries is however a much more complex question, since also creep and relaxation tests should be investigated. Finally, these macroscopic mechanical tests also allow for checking the widely accepted assumption of incompressibility of arterial tissues. Incompressibility implies that the changes in sample thickness can be deduced from the changes in the circumferential and axial dimensions. To our best knowledge, Carew et al. (1968) was the only study that experimentally checked this assumption, by evaluating the ratio between bulk and shear moduli based on a tension–inflation test. As a result, the arterial bulk modulus was about three times larger than the Young’s modulus determined by Bergel (1961), while the hydrostatic and deviatoric stresses were of the same order of magnitude. They could therefore conclude that arteries are only slightly compressible.

At the microstructural scale As a transition between purely mechanical characterization and mechanical testing coupled with live imaging, we here focus on correlations existing between mechanical characteristics of arteries and their composition. In this respect, we report two investigation results. First, several studies (Sato et al., 1979; Hayashi et al., 1974) emphasized the dependence of the mechanical response on the sample location along the aortic tree. The aortic stiffness is found to be higher in distal regions than in proximal regions (Zeinali-Davarani et al., 2015; Sokolis, 2007; Haskett et al., 2010; Tanaka and Fung, 1974), which correlates with a higher collagen content in the distal aortic regions. This property is directly related to the difference in mechanical function between proximal and distal regions: proximal regions of the aortic tree directly receives blood from the heart and therefore needs to exhibit larger damping properties, which is microstructurally translated through a larger elastin content and more undulated collagen bundles in the proximal aortas (Zeinali-Davarani et al., 2015). Second, several mechanical tests were performed after partial enzyme degradation of either elastin or collagen. By imparting compressive stresses to the collagen (Chow et al., 2014), the presence of elastin increases collagen folding, resulting in a more compliant response of the tissue (Ferruzzi et al., 2011), since straightened collagen fibers are stiffer than elastin fibers. As a consequence, degrading elastin leads to vessel enlargement, by relaxation of the internal compressive stress. This results in a mechanical response being softer in the low stress regime, and stiffer in the large stress regime (Fonck et al., 2007; Weisbecker et al., 2013). Elastin degradation also leads to an earlier recruitment of collagen fibers, the mechanical response being sooner stiffer (Rezakhaniha et al., 2011; Zeinali-Davarani et al., 2013). On the other hand, collagen degradation leads to the disappearance of the progressive stiffening of the mechanical response, while the initial slope of the stress–strain curves remains unchanged (Weisbecker et al., 2013).

Uniaxial, biaxial, and tension inflation tests on large elastic arteries. LONG (resp. CIRC) stands for the longitudinal (resp. circumferential) direction

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Table 1

Animal

Artery

Layer

Sollicitation

Preconditioning

Direction

Number of tests

(Balzani et al., 2006)

Human

Abdominal aorta

Media

1

Human

Abdominal aorta

Media

Circ

1

(Choudhury et al., 2009) (Chuong and Fung, 1984) (Cox, 1975)

Human

Ascending thoracic aorta

All

5 loading-unloading cycles at 1 mm min 1 5 loading-unloading cycles at 1 mm min 1

Long

(Balzani et al., 2006)

Rabbit

Thoracic aorta

All

4

Carotid arteries

All

Loading until 600 g and then unloading Various number of inflation/ deflation cycles until a stabilized response is found

Radial

Dog

Uniaxial tensile test on strips in saline solution at 37 C Uniaxial tensile test on strips in saline solution at 37 C Equibiaxial test on flat square samples Compression on square flat samples in air Inflation tests at constant tensile stretch or constant tensile load, controlled temperature

10 animals

(Dobrin, 1986)

Dog

Carotid arteries

All

(Hill et al., 2012)

Rabbit

Carotid arteries

All

8 samples

(Holzapfel, 2006)

Human (80 years old woman) Human

Abdominal aorta

Circ and Long

2 per layer

Left anterior coronary artery

5 cycles at 1 mm. min 1

Circ and Long

Porcine

Descending coronary arteries

Adventitia/ media/intima Adventitia/ media/intima All

5 inflation/deflation tests until stabilization of the response 5 cycles at 1 N under quasistatic loading 5 cycles preconditioning

Circ at fixed Long stretch And Circ at fixed Long force Circ at fixed Long stretch Circ

10 cycles tensile inflation tests

Long/Circ

78 strips (6  13) 5

In the longitudinal and circumferential direction 5 cycles at 1 N at a strain rate of 60%/min 5 cycles at 0.5 N at a strain rate of 60%/min 5 cycles

Circ at prescribed long. stretch Long

21

Long/Circ

8

Long and Circ

18 each

Long/Circ

14 animals

Circ

12

One cycle pressure up to 160 kPa

Circ followed by Long

18

Not reported

Long and Circ

20

Repeated pressure loading and cyclic stretching

Long at fixed Circ pressure

5 animals

(Holzapfel et al., 2005) (Keyes et al., 2013) (Kim and Baek, 2011) (Lally et al., 2004)

Porcine

Thoracic Aorta

All

Porcine

Coronary arteries

All

(Lally et al., 2004)

Porcine

Coronary arteries

All

(Mohan and Melvin, 1982) (Patel and Janicki, 1970) (Pandit et al., 2005)

Human

All

Dog

Mid-thoracic descending aorta Coronary and carotid artery

All

Porcine

Coronary arteries

All

(Papageorgiou and Jones, 1988)

Human

Iliac arteries

All

(Samila and Carter, 1981) (Sato et al., 1979)

Human

Carotid arteries

Dog

Abdominal aorta, carotid arteries, femoral arteries

Media and Intima All

Inflation test at fixed tensile stretch Tension of flat dog bone samples in a saline solution Tensile test in a saline solution on strips Tensile test in a saline solution on strips Planar biaxial testing at different stress ratios and tubular testing Inflation tests under fixed axial stretch Tensile tests on strips in a saline bath Equibiaxial tests on square samples in a saline bath Tensile tests at room temperature sprayed with saline solution Inflation test at in vivo stretch Inflation test in saline solution at imposed axial stretches Inflation test at in vivo length, static tests, kept wet; followed by axial tensile test on the whole sample Uniaxial tensile tests in saline solution Tensile test at fixed pressure, 37 C, sample kept wet

Response stabilizes after two load cycles Not reported

5 healthy

120 tests

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Biomechanics j Multiscale Mechanical Behavior of Large Arteries

Reference

Human

aorta

adventitia

Tensile test immerged in a saline bath Tensile test in a saline bath on strips Tensile test on strips at 37 C immerged in a saline solution Tensile test on strips at 37 C immerged in a saline solution Tensile test on strips at 37 C immerged in a saline solution Inflation under controlled axial force

(Silver et al., 2003)

Porcine

Aorta, carotid and vena cava

All

(Sokolis et al., 2002b) (Sokolis et al., 2002b) (Sokolis et al., 2006)

Rabbit

Abdominal aorta

media

Porcine

Abdominal aorta

media

Rabbit

Descending thoracic aorta

all

(Sommer et al., 2010)

Human

Carotid arteries

All and layer specific

(Storkholm et al., 1997)

Porcine

Abdominal aorta

All

Inflation test under controlled axial stretch

(Tanaka and Fung, 1974) (van de Geest et al., 2006) (von Maltzahn et al., 1984)

Dog

Aorta

All

Human

Abdominal aorta

All

Bovine

Carotid arteries

(Vorp et al., 2003) (Vosshougi and Weizsäcker, 1985) (Weisbecker et al., 2013) (Weizsäcker et al., 1983)

Human Rat and pig

Ascending aorta Aorta

Human

Thoracic aorta

Intact and media/intima only All Media and Intima Media

Uniaxial tensile test on strips in a saline solution Flat biaxial tests at different stress ratios Inflation tests under different stretch ratios

Rat

Carotid arteries

All

(Zeinali-Davarani et al., 2015) (Zemánek et al., 2009)

Porcine

Descending thoracic aorta

All

Porcine

Thoracic aorta

all

Uniaxial tensile in a saline solution Uniaxial tensile test in a saline solution at 37 and 24 C Uniaxial tensile test in a saline solution at 37 C Inflation test at constant axial stretch at 37 C, in a saline bath and axial stretch at different pressures Planar biaxial tests at different stress ratios Equibiaxial tensile test in a saline solution

3 quasistatic loading/unloading cycles Not reported 10 tensile cycles to the maximum tensile strain 10 tensile cycles to the maximum tensile strain 10 tensile cycles to the maximum tensile strain 5 cycles axial preconditioning followed by 5 inflation–deflation test Cyclic loading/unloading test between 0 and 25 kPa until stable response Cyclic loading unloading until stabilization 9 loading/unloading cycles for each tension ratio Inflation/deflation tests between 0 and 250 mmHg at 1 kPa s 1

Long and Circ

2 samples

Long and Circ

6 each

Long

15

Long

20

Long

35

Circ

20

Circ

5

Long and Circ Long/Circ Circ

Not reported Three loading/unloading cycles at constant extension rate Not reported

Long and Circ Long and Circ

10 inflation/deflation tests

Long and Circ

Eight loading cycles up to 30 N m

Long/Circ

Cyclic loading of the specimen at 0.5 N until stabilization of the response

Long/Circ

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Fig. 5 Mechanical response of different arterial tissues subjected to uniaxial tension in the (A) circumferential and (B) longitudinal direction: porcine coronary arteries (crosses, Lally et al., 2004), human ascending aorta (circles, Choudhury et al., 2009), human mid-thoracic descending aortas (diamonds, Mohan and Melvin, 1982), and rabbit and pig aortas (triangles, Sokolis et al., 2002b, 2006).

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Intraluminal pressure [kPa] Fig. 6 Variation of the axial reaction force as a function of the applied luminal pressure, for different prescribed axial prestretches. Experimental data from Weizsäcker, H. W., Lambert, H., Pascale, K. (1983). Analysis of the passive mechanical properties of rat carotid arteries. Journal of Biomechanics 16 (9), 703–715.

Besides, arteries exhibit a layer-specific mechanical response, which depends on the layer morphology. Mechanical tests on the tunica intima were performed by Holzapfel et al. (2005) and Weisbecker et al. (2012): intima exhibits a stiffer mechanical response when loaded in the longitudinal direction than in the circumferential one, which is in good agreement with the longitudinal orientation of the fibers of the internal elastic lamina (Farand et al., 2007). However, it is generally agreed on that the tunica intima barely contributes to the mechanical response of arteries. As regards the tunica media, the circumferential direction shows a stiffer uniaxial response than the longitudinal direction, which correlates with the preferred circumferential orientation of the fiber networks in the media. Contrarily, in the tunica adventitia, the uniaxial mechanical response is stiffer in the longitudinal direction than in the circumferential one (Holzapfel et al., 2005; Weisbecker et al., 2012) because of the longitudinal orientation of the fibers of the strips in the load-free configuration (Chen et al., 2013), see Fig. 7. Finally, the larger elastin content in the media as compared to the adventitia makes the media more compliant, while the adventitia resists to larger loads. At the scale of fiber networks, the microstructural origin of the nonlinear mechanical behavior could be evidenced using multiphoton microscopy: namely, in uniaxial tensile tests, when the load is increased, fiber networks tend to rearrange significantly. In situ mechanical testing showed the ability of the different fiber networks to resist the applied mechanical loading by progressively

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aligning with the load direction (see Fig. 8). At rest, the collagen bundles of the tunica adventitia form a dense and crimped network, which is progressively straightened and then stretched. The straightening mechanism of fibers is generally referred to as the fiber recruitment and coincides with the response at low stresses (Hill et al., 2012; Schrauwen et al., 2012; Sokolis et al., 2006): as long as collagen fibers are not straight, they cannot sustain load and the softer constituents making up the arterial tissue sustain the applied load. Furthermore, whatever the load direction in both uniaxial and biaxial tests, adventitial collagen bundles also have the faculty to reorient and to align with the load direction (Chen et al., 2013); the stretching of collagen bundles takes place after fiber realignment, as shown by tracking collagen bundle deformation, either by means of fluorescent microspheres (Chen et al., 2011) or by means of X-ray diffraction (Schmid et al., 2005). Nevertheless, the mechanisms driving the reorientation of collagen networks remain unclear. Two hypotheses are currently discussed: affine rotation, where the fibers follow the matrix deformation (Wan et al., 2012), or larger rotations of the collagen network, where the collagen bundles are able to generate shear stresses to rotate faster than the matrix (Jayyosi, 2015; Billiar and Sacks, 1997). It might well be that both assumptions are valid but in different strain ranges, the affine rotations being validated over the physiological deformation range (Wan et al., 2012). The collagen bundles recruitment is driven by the presence of the elastin network: in the adventitia, the elastin network is, at rest, aligned with the collagen network (Chen et al., 2011), and it tends to reorient and align by application of a mechanical loading. Still, the realignment of the adventitial elastin network is less pronounced than the collagen realignment (Chen et al., 2013). In the media, elastin lamellae, collagen fibers, and smooth muscle cells also undergo load-induced reorientation. Under uniaxial and biaxial load cases, the medial fiber networks and the collagen network tend to align with the load direction (Timmins et al., 2010). Also, the engagement of collagen fibers occurs first in the media and starts later in the adventitia (Zeinali-Davarani et al., 2015; Chow et al., 2014). However, fiber rotations are larger for the collagen bundles of the adventitia (see Fig. 8): rotations of the medial constituents as well as of the adventitial elastin remain limited (Chow et al., 2014); this may be related to the very different nature of the networks, the elastin network of the media being for instance a very dense and interconnected network, with elastic segments oriented in all directions. At a larger scale, the elastic lamellae progressively unfold with load application and then stretch, as observed by means of polarized light microscopy by (Sokolis et al., 2006). The cohesive pericellular interlaced bundles also straighten and reorient by application of a load (Sokolis et al., 2006), and this recruitment process was shown to be faster than the recruitment of the circumferentially oriented, parallel bundles covering the elastic lamellae (Sugita and Matsumoto, 2016); the latter authors propose the stiffness difference between elastin and smooth muscle cells as the possible explanation, the more compliant surrounding medium allowing faster recruitment of the initially crimped fibers.

At the ultrastructural scale In this section, the mechanical properties of the collagen and elastin structures are reviewed. Concerning collagen, the literature reports a very broad range of variations for the mechanical properties of collagenous tissues: from several hundreds of MPa for a hydrated arrangement of collagen fibers to a few tens of GPa for an arrangement of air-dried collagen molecules (see Table 2 for a literature review). At an intermediate scale, collagen fibrils have an elastic modulus of several GPa, whereby this exact value depends on the hydration degree of the sample. Concerning elastin, its mechanical behavior is purely elastic, with neither hysteresis nor transient effect up to more than 100% strain, eventually showing a brittle failure mode (Aaron and Gosline, 1981). Young’s moduli reported in the literature exhibit less variability: the elastic modulus of intact purified elastic fibers from aortas was reported to be 0.4 MPa (Hass, 1942a, 1942b, 1943;

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