Materials for Biomedical Engineering: Hydrogels and Polymer-Based Scaffolds 012816901X, 9780128169018

Table of contents :
Cover
Hydrogels and Polymer-based Scaffolds
Copyright
List of Contributors
Series Preface
Preface
1 Interactions between tissues, cells, and biomaterials: an advanced evaluation by synchrotron radiation-based high-resolut...
1.1 Conduction, Induction, and Cell Transplantation in Tissue Engineering: The Limitations of Cross-talk Studies by Convent...
1.2 X-Ray Computed Microtomography: A Challenging Diagnostic Tool
1.3 Innovative Approaches to High-Resolution Tomography by Synchrotron Radiation
1.4 Skeletal Tissue Engineering
1.4.1 Bone
1.4.2 Cartilage
1.4.3 Tendons
1.5 Muscle Tissue Engineering
1.5.1 Skeletal Muscles
1.5.2 Heart
1.6 New Frontiers
1.6.1 Central and Peripheral Nervous System
1.6.2 Vascularization
1.7 Conclusions
References
Further Reading
2 Bioprinted scaffolds
2.1 Introduction
2.1.1 Prebioprinting
2.1.2 Bioprinting
2.1.3 Postbioprinting
2.1.4 Geometry of Scaffolds
2.1.5 Surface Properties
2.1.6 Pore Size
2.1.7 Adherence and Biocompatibility
2.1.8 Degradation Rates
2.2 Mechanical Properties
2.2.1 Hydrogel-Derived Scaffolds
2.2.2 Agarose hydrogel
2.2.3 Alginate hydrogel
2.2.4 Chitosan hydrogel
2.2.5 Cellulose hydrogel
2.2.6 Fibrin hydrogel
2.2.7 Gelatin/collagen hydrogel
2.2.8 Hyaluronic acid hydrogel
2.2.9 Matrigel hydrogel
2.2.10 Synthetic Hydrogels
2.3 Fibrous Polymer-Derived Scaffolds
2.4 Porous Polymer-Derived Scaffolds
2.5 Conclusion and Perspectives
Acknowledgment
References
3 Fundamentals of chitosan-based hydrogels: elaboration and characterization techniques
3.1 Introduction
3.2 Chitosan Nature and Main Properties
3.3 Fundamentals of Chitosan Hydrogels
3.3.1 Physical Hydrogels
3.3.2 Chemical Hydrogels
3.4 Characterization Techniques
3.4.1 Structural Analysis
3.4.1.1 Microstructural and spectroscopic analysis
3.4.1.2 Ultraviolet–visible spectroscopy and Fourier-transform infrared spectroscopy
3.4.2 Property Measurements
3.4.2.1 Active compound release assessment
3.4.2.2 Mechanical resistance
3.4.2.3 Viscosity (sol–gel analysis)
3.4.2.4 Swelling index
3.4.2.5 Contact angle
3.4.2.6 Thermal analysis
3.4.3 Specific Properties for Biomedical Engineering Applications
3.4.3.1 Degradability
3.4.3.2 Cytotoxicity
3.5 Potential Applications and Future Trends of Chitosan Hydrogels
References
4 Bioreabsorbable polymers for tissue engineering: PLA, PGA, and their copolymers
4.1 Tissue Engineering
4.2 Scaffolds
4.3 Biomaterials
4.3.1 Polymeric Biomaterials
4.3.2 Bioreabsorbable Biopolymers
4.4 Poly(α-Hydroxy Acids)
4.5 Poly(α-Hydroxy Acids) Synthesis
4.6 Copolymerization of Poly(α-Hydroxy Acids)
4.7 Mechanisms of Degradation of Poly(α-Hydroxy Acids)
4.8 Biocompatibility
4.9 Toxicity of Poly(α-Hydroxy Acids)
4.9.1 In Vitro Cytotoxicity Tests
4.9.2 In Vitro Hemocompatibility Test
4.9.3 In Vivo Biocompatibility Tests
4.9.3.1 General tests for bone implants
4.9.3.2 General tests for stents
4.10 Applications of Poly(α-Hydroxy Acids)—PLA and PGA
4.10.1 Nonmedical Applications of Poly(α-Hydroxy Acids)—PLA and PGA
4.10.2 Medical Applications of Poly(α-Hydroxy Acids)—PLA and PGA
4.11 Future Trends in Biofabrication
4.11.1 Electrospinning
4.11.2 3D Bioprinting Rapid Prototyping
4.11.3 Bioresponsive Hydrogels
4.11.4 Biopolymer Composites in Tissue Engineering
4.12 Conclusions
References
Further Reading
5 Technological challenges and advances: from lactic acid to polylactate and copolymers
5.1 Lactic Acid
5.1.1 Factors That Influence Lactic Acid Production
5.1.2 Culture Medium for Lactic Fermentation: Alternative Sources of Carbon and Nitrogen
5.1.3 Production of Lactic Acid by Fermentation
5.1.4 Microorganisms Involved in the Production of Lactic Acid
5.1.5 Extraction and Purification of Lactic Acid
5.2 Poly(lactic Acid)
5.2.1 PLA Chemical and Physical Properties
5.2.2 PLA Synthesis
5.2.2.1 Chemical polymerization
5.2.2.2 Enzymatic polymerization: production of PLA directly by genetically modified microorganism
5.2.3 Kinds of Polymers, Copolymers, and Their Features
5.2.4 PLA Applications
5.2.5 PLA Market Development
5.2.6 PLA Biodegradation, Biocompatibility, and Toxicity
5.3 Conclusion
References
6 PLGA scaffolds: building blocks for new age therapeutics
6.1 Challenges in New Age Therapeutic Strategies
6.2 Poly(Lactide-co-Glycolide): General Introduction
6.3 Poly(Lactide-co-Glycolide) Synthesis
6.4 Poly(Lactide-co-Glycolide) Properties
6.5 Poly(Lactide-co-Glycolide) Scaffolds for Bone Tissue Engineering
6.5.1 Porous Scaffolds
6.5.2 Fibrous Scaffolds
6.5.3 Hydrogels
6.5.4 Injectable Microparticles
6.6 Poly(Lactide-co-Glycolide) Scaffolds in Anticancer Therapy
6.7 Poly(Lactide-co-Glycolide) Interventions in Central Nervous System Delivery
6.8 Poly(Lactide-co-Glycolide) Strategies for Gene Therapy and Vaccine Delivery
6.9 Miscellaneous Poly(Lactide-co-Glycolide) Therapeutics
6.10 Conclusions and Future Trends
Acknowledgments
List of Symbols and Abbreviations
References
7 Electrospun biomimetic scaffolds of biosynthesized poly(β-hydroxybutyrate) from Azotobacter vinelandii strains. cell viab...
7.1 Introduction
7.1.1 Polymers as Medical Devices
7.1.2 Shape Memory Polymers
7.1.3 Smart Polymeric Coatings
7.1.4 Electrospun Fibrous Scaffolds
7.1.5 Poly-β-Hydroxybutyrate
7.2 Methods of Characterization
7.2.1 Materials
7.2.2 Scaffold Fabrication
7.2.3 Fourier-Transformed Infrared Spectroscopy
7.2.4 Thermal Analysis
7.2.5 X-Ray Scattering
7.2.6 Small-Angle Light Scattering
7.2.7 Contact Angle
7.2.8 Polarized Optical Microscopy
7.2.9 Scanning Electron Microscopy
7.3 PHB Electrospun Fibrous Scaffolds
7.3.1 Scaffolds Morphology
7.3.2 Wetting Behavior
7.3.3 Aging
7.3.4 Sterilization Methods and Influence on Physical Properties
7.4 Cell Viability and Bone Tissue Regeneration
7.4.1 Cell Viability and HEK293 Cells
7.4.2 Bone Tissue Regeneration and Human Osteoblast Cells
7.5 Concluding Remarks
Glossary of Terms
References
Further Reading
8 Polyurethane-based structures obtained by additive manufacturing technologies
8.1 Introduction
8.2 Bioresorbable Polyurethanes in Biomedical Devices
8.3 Additive Manufacturing for Biomedical Polyurethane Processing
8.3.1 Inkjet Printing
8.3.2 Extrusion-Based Methods
8.3.3 Particle Binding
8.4 Additive Manufacturing of Composite Polyurethanes
8.4.1 Inkjet Printing
8.4.2 Extrusion-Based Methods
8.4.2.1 Direct ink writing
8.4.2.1.1 Liquid-frozen deposition manufacturing
8.4.2.1.2 Double-nozzle low-temperature deposition manufacturing
8.4.2.1.3 Integrated organ printing
8.4.2.2 Fused deposition modeling
8.4.3 Particle Binding
8.5 Remarks and Perspectives
Acknowledgment
References
9 Composites based on bioderived polymers: potential role in tissue engineering: Vol VI: resorbable polymer fibers
9.1 Introduction
9.2 Polyesters
9.2.1 Poly(Lactic Acid)
9.2.1.1 Poly(lactic acid) fabrication
9.2.1.2 Poly(lactic acid) processing
Drying and extrusion
Injection molding
Stretch blow molding
Cast film and sheet
Thermoforming
Foaming
9.2.1.3 Poly(lactic acid) properties
Physical proprties
Thermal properties
Mechanical properties
9.2.1.4 Poly(lactic acid) medical applications
Wound healing and stents
Scaffolds for tissue engineering
Orthopedic implants and fixation devices
Drug delivery
3D printing
9.2.2 Poly(lactic-co-glycolic acid) (PLGA) copolymers
9.2.2.1 Synthesis of PLGA
9.2.2.2 Properties of PLGA
9.2.2.3 Medical Applications of PLGA
9.3 Collagen
9.3.1 Collagen Bioactive Ceramic Composites
9.3.1.1 Collagen–HAP composites
9.3.1.2 Collagen TCP/BCP composites
9.3.1.3 Collagen-bioglass based composites
9.3.2 Medical Applications of Collagen
9.4 Silk Fibroin
9.4.1 Structure of Silk Fibroin
9.4.2 Processing of Silk Fibroin
9.4.2.1 Hydrogelation
9.4.2.2 Electrospinning
9.4.2.3 Porogen leaching
9.4.2.4 3D bioprinting
9.4.2.5 SF composites
9.4.3 Medical Applications of Silk Fibroin
9.4.3.1 SF scaffolds for tissue engineering
9.4.3.2 Delivery of bioactive molecules
9.4.3.3 Fixation devices
9.5 Biocellulose
9.5.1 Biocellulose Fibril Structure
9.5.2 Properties of Biocellulose
9.5.2.1 Mechanical properties
9.5.2.2 Biocompatibility
9.5.2.3 Hemocompatibility
9.5.2.4 Biodegradability
9.5.2.5 Nontoxicity
9.5.3 Biomedical Applications of Biocellulose
9.5.3.1 Substitute biomaterials for medical applications
9.5.3.2 Biocellulose-based scaffolds for bone tissue regeneration
9.5.3.3 Scaffolds for cell culture
9.5.3.4 Antimicrobial biomaterials
9.5.3.5 Drug delivery applications
9.5.3.6 Other biomedical applications
9.6 Conclusions
References
10 Composite scaffolds for bone and osteochondral defects
10.1 Introduction
10.2 Biodegradable Matrices
10.3 Bioresorbable Matrices
10.4 Applications in Tissue Engineering
10.4.1 Composite Scaffolds for Bone
10.4.1.1 Calcium phosphate particle loaded porous/nonporous composites
10.4.1.2 Fiber-loaded composites
10.4.1.3 Collagen-HA hybrid nanocomposite for bone
10.4.2 Composite Scaffolds for Osteochondral Defects
10.4.2.1 Multilayer porous scaffolds
10.4.2.2 Gradient porous/nonporous composites
10.4.2.3 Magnetic bioinspired hybrid nanocomposites for osteochondral tissue
10.5 Conclusions
References
Further Reading
11 Plasma treated and untreated thermoplastic biopolymers/biocomposites in tissue engineering and biodegradable implants
11.1 Introduction
11.2 Structure of PLA and PHAs
11.3 Synthesis of PLA and PHAs
11.4 Properties of PLA and PHAs
11.4.1 Mechanical Properties
11.4.2 Thermal Properties
11.4.3 Transparency
11.4.4 Biocompatibility
11.4.5 Processability
11.5 Application of PLA and PHAs in Tissue Engineering
11.6 Biodegradability of PLA and PHAs
11.7 Plasma Treatment of PLA and PHAs
11.7.1 Plasma and Plasma–Surface Interactions
11.7.2 Characterization Techniques for Plasma Treated Polymer Surfaces
11.7.3 Plasma Treatment of PLA
11.7.4 Plasma Treatment of PHAs
11.7.5 Disadvantages of Plasma Treatment
11.8 Conclusions
References
12 The design of two different structural scaffolds using β-tricalcium phosphate (β-TCP) and collagen for bone tissue engin...
12.1 Introduction
12.2 Collagen-Based Porous Scaffold
12.2.1 Fabrication and Characterization of Particle Distributed Scaffold
12.2.1.1 Fabrication of particle distributed scaffold
12.2.1.2 Characterization of particle distributed scaffold
12.2.2 In Vitro Cell Experiment
12.2.2.1 Cell culture
12.2.2.2 Compression test
12.2.2.3 Microstructural characterization
12.2.2.4 Evaluation of cell number and alkaline phosphatase activity
12.2.2.5 Gene expression analysis
12.2.2.6 Statistics
12.3 Experimental Results
12.3.1 Characterization of Particle Distributed Scaffold
12.3.2 Results of In Vitro Cell Experiment
12.4 Mechanism of Variational Mechanical Behavior Between Scaffold Structure and Cell Response
12.5 β-TCP-Based Porous Scaffold
12.5.1 Fabrication and Characterization of Two Phase Structural Scaffold
12.5.1.1 Fabrication of two phase structural scaffold
12.5.1.2 Characterization of two phase structural scaffold
12.6 In Vitro Cell Experiment
12.6.1 Cell Culture
12.6.2 Evaluation of Mechanical Characteristics
12.6.3 Microstructural Characterization
12.6.4 Evaluation of Cell Number and Alkaline Phosphatase Activity
12.6.5 Gene Expression Analysis
12.6.6 Alizarin Red S Staining
12.6.7 Statistics
12.7 Experimental Results
12.7.1 Characterization of Two Phase Structural Scaffold
12.7.2 Results of In Vitro Cell Experiment
12.8 Mechanism of Variational Mechanical Behavior Between Scaffold Structure and Cell Response
12.9 Summary
12.10 Present Study
12.11 Future Work
Acknowledgment
References
13 Composite materials based on hydroxyapatite embedded in biopolymer matrices: ways of synthesis and application
13.1 Types of Biopolymer Matrices (Collagen, Gelatin, Chitosan, Alginate, and Their Combinations)
13.2 Calcium Phosphates as an Essential Part of Composite Materials
13.3 Formation of Composite Materials
13.4 Biomedical Applications of Obtained Composite Materials
References
Further Reading
14 Study of microstructural, structural, mechanical, and vibrational properties of defatted trabecular bovine bones: natura...
14.1 Introduction
14.2 Bone Composition
14.2.1 Cortical Bone
14.2.2 Trabecular Bone
14.2.3 Bone Porosity
14.2.4 Hydroxyapatite
14.2.5 Biohydroxyapatite
14.2.5.1 Structural properties of BIO-HA
14.2.5.2 Mineral composition of BIO-HA
14.2.5.3 Thermal properties of BIO-HA
14.2.5.4 Methods to obtain HA and BIO-HA
14.2.6 Collagen
14.2.7 Osteocalcin
14.2.8 Water
14.2.9 Fat
14.3 Study of Spongy Bone
14.3.1 Collection and Preparation of Samples
14.3.2 Morphological Characterization
14.3.3 X-ray Tomography
14.3.3.1 Imaging
14.3.4 Structural Properties
14.3.4.1 Transmission electron microscopy
14.3.4.2 X-ray diffraction
14.3.5 Vibrational Characterization: Raman Spectroscopy
14.3.6 Mechanical Properties
14.4 Synthetic Scaffolds Versus Trabecular Bone
14.5 Conclusions and Perspective
Acknowledgments
References
Further Reading
Appendix A
15 Laser processing of biopolymers for development of medical and high-tech devices
15.1 Introduction
15.2 Structure and Raman Spectrum of Polydimethylsiloxane
15.3 Experimental and Analytical Techniques
15.4 Optical Properties of Polydimethylsiloxane during ns-laser treatment
15.5 Fs-Laser Nanostructuring
15.6 Ps-Laser Processing
15.7 Comparison Between Fs- and Ns-Laser Processing
15.8 XPS Study of Ns-Laser Processing of Polydimethylsiloxane
15.9 Electroless Metallization Directly After the Laser Treatment
15.10 Ns-Laser Processing in Different Environments
15.11 Conclusion and Perspectives for Future Investigations
Acknowledgments
References
Further Reading
Index
Back Cover

Citation preview

Hydrogels and Polymer-based Scaffolds

Materials for Biomedical Engineering

Hydrogels and Polymer-based Scaffolds Edited by

Alina-Maria Holban Faculty of Biology, University of Bucharest, Bucharest, Romania

Alexandru Mihai Grumezescu Faculty of Applied Chemistry and Materials Science, University Politehnica of Bucharest, Bucharest, Romania

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 © 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). Notices 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 must 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. British Library Cataloguing-in-Publication Data A catalogue record for this book is available from the British Library Library of Congress Cataloging-in-Publication Data A catalog record for this book is available from the Library of Congress ISBN: 978-0-12-816901-8 For Information on all Elsevier publications visit our website at https://www.elsevier.com/books-and-journals

Publisher: Matthew Deans Acquisition Editor: Gwen Jones Editorial Project Manager: Emma Hayes Production Project Manager: Debasish Ghosh Cover Designer: Greg Harris Typeset by MPS Limited, Chennai, India

List of Contributors Gustavo A. Abraham Research Institute for Materials Science and Technology, INTEMA (UNMdP-CONICET), Mar del Plata, Argentina Satish Agrawal Division of Pharmaceutics & Pharmacokinetics, CSIR-Central Drug Research Institute, Lucknow, India Hafsa Ahmad Pharmacognosy & Ethnopharmacology Division, CSIR-National Botanical Research Institute, Lucknow, India Luigi Ambrosio Institute of Polymers, Composites and Biomaterials, National Research Council of Italy, Naples, Italy Takaaki Arahira Section of Bioengineering, Department of Dental Engineering, Fukuoka Dental College, Fukuoka, Japan Stephan Armyanov Institute of Physical Chemistry Rostislaw Kaischew, Bulgarian Academy of Sciences, Sofia, Bulgaria Abhishek Arya Division of Pharmaceutics & Pharmacokinetics, CSIR-Central Drug Research Institute, Lucknow, India Petar A. Atanasov Institute of Electronics, Bulgarian Academy of Sciences, Sofia, Bulgaria Maria Ingrid Rocha Barbosa Institute of Biofabrication, School of Chemical Engineering, University of Campinas, Campinas, Sa˜o Paulo, Brazil Rejane Andrade Batista Instituto Tecnolo´gico e de Pesquisas do Estado de Sergipe, Rua Campo do Brito, Aracaju, Brazil Susan Michelz Beitel Department of Biochemistry and Microbiology, Institute Bioscience, Sa˜o Paulo State University (UNESP), Sa˜o Paulo, Sa˜o Paulo, Brazil Ana Carolina B. Benatti Institute of Biofabrication, School of Chemical Engineering, University of Campinas, Campinas, Sa˜o Paulo, Brazil; School of Medical Sciences, Surgery Department, University of Campinas, Campinas, Sa˜o Paulo, Brazil Binay Bhushan Department of Physics, Birla Institute of Technology, Mesra, Patna Campus, India

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List of Contributors

S. Bolshanina Sumy State University, Ministry of Education and Science of Ukraine, Sumy, Ukraine Pablo C. Caracciolo Research Institute for Materials Science and Technology, INTEMA (UNMdP-CONICET), Mar del Plata, Argentina Luciana Fontes Coelho Department of Biochemistry and Microbiology, Institute Bioscience, Sa˜o Paulo State University (UNESP), Sa˜o Paulo, Sa˜o Paulo, Brazil Jonas Contiero Department of Biochemistry and Microbiology, Institute Bioscience, Sa˜o Paulo State University (UNESP), Sa˜o Paulo, Sa˜o Paulo, Brazil; Associate Laboratory IPBEN-UNESP, Rio Claro, Sa˜o Paulo, Brazil Anil Kumar Dwivedi Division of Pharmaceutics & Pharmacokinetics, CSIR-Central Drug Research Institute, Lucknow, India Paula J.P. Espitia Nutrition and Dietetics School, Universidad del Atla´ntico, Atla´ntico, Colombia Rubens Maciel Filho Institute of Biofabrication, School of Chemical Engineering, University of Campinas, Campinas, Sa˜o Paulo, Brazil Alessandra Giuliani Department of Clinical Sciences, Universita` Politecnica delle Marche, Ancona, Italy Vincenzo Guarino Institute of Polymers, Composites and Biomaterials, National Research Council of Italy, Naples, Italy Florin Iordache Institute of Cellular Biology and Pathology “Nicolae Simionescu” of Romanian Academy, Bucharest, Romania; Faculty of Veterinary Medicine, University of Agronomic Sciences and veterinary Medicine, Bucharest, Romania Andre´ Luiz Jardini Institute of Biofabrication, School of Chemical Engineering, University of Campinas, Campinas, Sa˜o Paulo, Brazil Andreas Kaasi Institute of Biofabrication, School of Chemical Engineering, University of Campinas, Campinas, Sa˜o Paulo, Brazil; School of Medical Sciences, Surgery Department, University of Campinas, Campinas, Sa˜o Paulo, Brazil; Eva Scientific Ltda, Sa˜o Paulo, Sa˜o Paulo, Brazil

List of Contributors

Paulo Kharmandayan Institute of Biofabrication, School of Chemical Engineering, University of Campinas, Campinas, Sa˜o Paulo, Brazil; School of Medical Sciences, Surgery Department, University of Campinas, Campinas, Sa˜o Paulo, Brazil Konstantin Kolev Institute of Physical Chemistry Rostislaw Kaischew, Bulgarian Academy of Sciences, Sofia, Bulgaria Rakesh Kumar Department of Biotechnology, Central University of South Bihar, Gaya, India Sandra M. London˜o-Restrepo Posgrado en Ciencia e Ingenierı´a de Materiales, Centro de Fı´sica Aplicada y Tecnologı´a Avanzada, Universidad Nacional Auto´noma de Me´xico, Quere´taro, Me´xico Nayla J. Lores Research Institute for Materials Science and Technology, INTEMA (UNMdP-CONICET), Mar del Plata, Argentina Adrian Manescu “Victor Babes” University of Medicine and Pharmacy, Timisoara, Romania Serena Mazzoni Department of Clinical Sciences, Universita` Politecnica delle Marche, Ancona, Italy Nikolay N. Nedyalkov Institute of Electronics, Bulgarian Academy of Sciences, Sofia, Bulgaria Caio Gomide Otoni National Nanotechnology Laboratory for Agribusiness, Embrapa Instrumentac¸a˜o, Sa˜o Carlos, Brazil Nidhi Pareek Department of Microbiology, Central University of Rajasthan, Ajmer, India Kunwar Paritosh Centre for Energy and Environment, Malaviya National Institute of Technology, Jaipur, India Ana Fla´via Pattaro Institute of Biofabrication, School of Chemical Engineering, University of Campinas, Campinas, Sa˜o Paulo, Brazil Cristian F. Ramirez-Gutierrez Posgrado en Ciencia e Ingenierı´a de Materiales, Centro de Fı´sica Aplicada y Tecnologı´a Avanzada, Universidad Nacional Auto´noma de Me´xico, Quere´taro, Me´xico

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Ana Ame´lia Rodrigues Institute of Biofabrication, School of Chemical Engineering, University of Campinas, Campinas, Sa˜o Paulo, Brazil Mario E. Rodriguez-Garcı´a Departamento de Nanotecnologı´a, Centro de Fı´sica Aplicada y Tecnologı´a Avanzada, Universidad Nacional Auto´noma de Me´xico, Quere´taro, Me´xico Angel Romo-Uribe Johnson & Johnson Vision Care, Inc., Advanced Science & Technology, Jacksonville, FL, United States Monica Sandri CNR-National Research Council of Italy, Institute of Science and Technology for Ceramic Materials (ISTEC), Faenza, Italy Silvia Scaglione CNR-National Research Council of Italy, IEIIT Institute, Genoa, Italy Simone Sprio CNR-National Research Council of Italy, Institute of Science and Technology for Ceramic Materials (ISTEC), Faenza, Italy Nadya E. Stankova Institute of Electronics, Bulgarian Academy of Sciences, Sofia, Bulgaria Anna Tampieri CNR-National Research Council of Italy, Institute of Science and Technology for Ceramic Materials (ISTEC), Faenza, Italy Mitsugu Todo Research Institute for Applied Mechanics, Kyushu University, Fukuoka, Japan Giuliana Tromba Sincrotrone Trieste S.C.p.A, Trieste, Italy Eugenia Valova Institute of Physical Chemistry Rostislaw Kaischew, Bulgarian Academy of Sciences, Sofia, Bulgaria Herminso Villarraga-Go´mez Nikon Metrology, Inc., Brighton, MI, United States Vivekanand Vivekanand Centre for Energy and Environment, Malaviya National Institute of Technology, Jaipur, India Mariana Vitelo Xavier Institute of Biofabrication, School of Chemical Engineering, University of Campinas, Campinas, Sa˜o Paulo, Brazil; School of Medical Sciences, Surgery Department, University of Campinas, Campinas, Sa˜o Paulo, Brazil

List of Contributors

Monika Yadav Centre for Energy and Environment, Malaviya National Institute of Technology, Jaipur, India A. Yanovska Sumy State University, Ministry of Education and Science of Ukraine, Sumy, Ukraine

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Series Preface In the past few decades there has been growing interest in the design and implementation of advanced materials for new biomedical applications. The development of these materials has been facilitated by multiple factors, especially the introduction of new engineering tools and technologies, emerging biomedical needs, and socioeconomic considerations. Bioengineering is an interdisciplinary field encompassing contributions from biology, medicine, chemistry, and materials science. In this context, new materials have been developed or reinvented to fulfill the need for modern and improved engineered biodevices. A multivolume series, Materials for Biomedical Engineering highlights the most relevant findings and discusses key topics in this impressive research field. Volume 1. Bioactive Materials: Properties and Applications, offers an introduction to bioactive materials, discussing the main properties, applications, and perspectives of materials with medical applications. This volume reviews recently developed materials, highlighting their impact in tissue engineering and the detection, therapy, and prophylaxis of various diseases. Volume 2. Thermoset and Thermoplastic Polymers, analyzes the main applications of advanced functional polymers in the biomedical field. In recent years there has been a revolution in thermoplastic and thermosetting polymers with medical and biological uses, which are currently being developed for medical devices, drug delivery, tailored textiles, packaging, and tissue engineering. Volume 3. Absorbable Polymers, describes the main types of polymers of different compositions with bioabsorbable and biodegradable properties. The biomedical applications of such materials are reviewed and the most innovative findings are presented in this volume. Volume 4. Biopolymer Fibers, highlights the applications of polymeric fibers of natural biological origin in biomedical engineering. Such materials are of great utility in tissue engineering and biodegradable textiles. Volume 5. Inorganic Micro- and Nanostructures and Volume 6. Organic Micro- and Nanostructures, deal, respectively, with the preparation and properties of inorganic and organic nanostructured materials with biomedical applications. Volume 7. Hydrogels and Polymer-Based Scaffolds, discusses the recent progress made in the field of polymeric materials designed as scaffolds and tools for tissue engineering. The technological challenges and advances in their production, as well as current applications in the production of scaffolds and devices for regenerative medicine are presented. Volume 8. Bioactive Materials for Antimicrobial, Anticancer, and Gene Therapy, offers an updated perspective regarding new bioactive materials with potential in the therapy of severe diseases such as infections, cancer, and genetic disorders.

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Volume 9. Nanobiomaterials in Tissue Engineering, provides valuable examples of recently designed nanomaterials with powerful applications in tissue engineering and artificial organ approaches. Volume 10. Nanomaterials-Based Drug Delivery, discusses the most investigated types of nanoparticles and nanoengineered materials with an impact in drug delivery. Applications for drug-therapy, and examples of such nanoscale systems are included in this volume. This series was motivated by the need to offer a scientifically solid basis for the new findings and approaches relevant to the biomedical engineering field. This scientific resource collects new information on the preparation and analysis tools of diverse materials with biomedical applications, while also offering innovative examples of their medical uses for diagnoses and therapies of diseases. The series will be of particular interest for material scientists, engineers, researchers working in the biomedical field, clinicians, and also innovative and established pharmaceutical companies interested in the latest progress made in the field of biomaterials. Michael R. Hamblin1 and Ioannis L. Liakos2 1

Harvard Medical School, Boston, MA, United States 2 Istituto Italiano di Tecnologia, Genoa, Italy

Preface Scaffolds can be composed of various materials, depending on their intended use. Polymers of various origins are intensively investigated for the design of scaffolds and materials for tissue engineering. In fact, the whole field of regenerative medicine relies on the development of polymeric materials with particular properties, such as great biocompatibility, resistance, plasticity, and perspectives of use. Natural and synthetic polymeric materials were recently developed or reinvented to fulfill the requirements of the constantly expanding field of scaffolds with biomedical applications. In this context, hydrogels have arisen as valuable alternatives to traditional polymeric materials, offering unique properties for tissue engineering. This book aims to reveal the latest tools and applications of scaffolds developed with the help of hydrogels and polymeric materials. The volume contains 15 chapters prepared by outstanding authors from Italy, the United States, Mexico, Brazil, India, Romania, Argentina, Ukraine, Bulgaria, and Japan. Chapter 1, entitled Interactions between tissues, cells and biomaterials: An advanced evaluation by synchrotron radiation-based high-resolution tomography, was prepared by Alessandra Giuliani et al. The chapter discusses the most recent applications of synchrotron radiation based microtomography in the study of interactions between tissues, cells, and biomaterials. Synchrotron imaging has proven to be a fundamental characterization tool for understanding the mechanisms of biological behavior of biomaterials proposed as tissue substitutes. These innovative techniques allow not only the visualization and quantification of regenerated tissues, but also the eventual presence and distribution of the neovascularization. Chapter 2, Bioprinted scaffolds, by Florin Iordache et al., presents the recent progress in the field of three-dimensional (3D) bioprinting applied to fabricate tissue, organs, and biomedical parts that imitate natural tissue architecture. This technology combines cells, growth factors, and biomaterials to create a microenvironment in which cells can grow and differentiate in tissue structures. Chapter 3, Fundamental of chitosan-based hydrogels: Elaboration and characterization techniques, prepared by Rejane Andrade Batista et al., reviews the recent breakthroughs on hydrogels featuring biological properties, as well as their capabilities of carrying bioactive compounds. Fundamentals of hydrogel elaboration and the nature of biopolymers used are presented, as well as the future trends regarding the use of chitosan-based hydrogels. Chapter 4, Bioreabsorbable polymers for tissue engineering: PLA, PGA, and their copolymers, prepared by Ana Carolina Benatti et al., discusses biomaterials with characteristics compatible to the sites of medical application and do not present toxicity. This chapter aims to conduct a review of different synthesis routes, modification of polymeric structures by copolymerization, and to present the characteristics necessary for use in different medical areas and for applications in tissue engineering.

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Chapter 5, Technological challenges and advances: From lactic acid to polylactate and copolymers, prepared by Susan Michelz Beitel et al., reviews lactic acid, an organic acid which has been extensively used worldwide in a variety of industrial and biotechnological applications. Lactic acid can be obtained chemically or by microbial fermentation. Production by fermentation results in the formation of D( ) or L(1) lactic acid, or a racemic mixture, depending on the microorganisms used. PLA has been considered as one of the most promising biodegradable plastics by having physical characteristics similar to the polymers derived from nonrenewable sources, such as elasticity, stiffness, transparency, thermoresistance, biocompatibility, and good moldability. Chapter 6, PLGA scaffolds: Building blocks for the new age therapeutics, prepared by Hafsa Ahmad et al., presents recent progress of biodegradable interventions as 3D scaffolds for tissue engineering. Their biodegradation rate can be controlled and they are suitable for clinical applications. They have also been successfully used for delivery of several growth factors and in vaccine and gene delivery, besides being popularly used as suture and suture anchors for surgeries and as orthopedic-fixation devices. Chapter 7, Electrospun biomimetic scaffolds of biosynthesized poly (β-hydroxybutyrate) from Azotobacter vinelandii strains. Cell viability and bone tissue engineering, prepared by Angel Romo-Uribe et al., gives an up-to-date overview on the characteristics of fibrous polymeric scaffolds which comprise capabilities of biomimetics to the native tissue architecture and are promising for achieving functional tissue-engineered products with minimal surgical implantation. In vitro studies showed that PHB scaffolds are also suitable for bone tissue engineering by supporting adhesion and proliferation of normal human osteoblast cells. Chapter 8, Polyurethane-based structures obtained by additive manufacturing technologies, prepared by Pablo Caracciolo et al., reviews the use of novel additive manufacturing (AM) as microfabrication tools for bioresorbable segmented polyurethanes (SPU), elastomers, and SPU composites. Current advances in 3D printing of SPU are described and discussed. Advantages and shortcomings of the current approaches as well as future perspectives are outlined; a vision of possible future research on this topic is also presented here. Chapter 9, Composites based on bioderived polymers: Potential role in tissue engineering, prepared by Monika Yadav et al., reviews the role and research efforts for bioinspired designing of composites targeted toward biomedical applications. Studies of the structure function relationship of natural biological materials have inspired material scientists for the development of biomimetic designs of new materials. Chapter 10, Composite scaffolds for bone and osteochondral defects, prepared by Vincenzo Guarino et al., offers an overview of recently developed composite porous and nonporous platforms used for repair or regeneration of bone and osteochondral tissue. Chapter 11, Plasma treated and untreated thermoplastic biopolymers/biocomposites in tissue engineering and biodegradable implants, prepared by Binay

Preface

Bhushan et al., reviews the properties and applications of pristine and plasma-treated PLA and PHAs in tissue engineering and biodegradable implants. The surface modification of PLA and PHAs through plasma treatments are being undertaken to fit the requirements of various fields in biomedical engineering. Plasma-assisted surface modification offers a very suitable strategy to incorporate reactive functional groups on polyester surfaces. The devices made from these biopolymers can be implanted in human beings without necessitating a second surgery to remove the device. Biocompatible tissues made from PLA and PHAs could be employed to replace damaged or diseased tissues in reconstructive surgery. Chapter 12, The design of two different structural scaffolds using β- tricalcium phosphate (β-TCP) and collagen for bone tissue engineering, prepared by Takaaki Arahira et al., described two different structural scaffolds fabricated by using β-TCP and collagen. One is collagen-based scaffolds with β-TCP particles fabricated by freeze-drying methods. Collagen/β-TCP scaffold is named “particle distributed scaffold”. The other is β-TCP-based scaffold with porous collagen structures fabricated using the polyurethane template method and a subsequent freeze-drying method. β-TCP/collagen scaffold is named “two-phase structural scaffold.” Chapter 13, Composite materials based on hydroxyapatite embedded in biopolymer matrices: Ways of synthesis and application, prepared by A. Yanovska et al., discusses composite materials based on various biopolymers’ combination with HA and each other provide excellent mechanical properties, biocompatibility, osteoconductivity, biodegradability, and cell proliferation and could be successfully used as materials for bone substitution. Chapter 14, Study of microstructural, structural, mechanical, and vibrational properties of defatted trabecular bovine bones: Natural sponges, prepared by Sandra London˜o-Restrepo et al., dissected the properties of bovine bones, useful to be applied as sponges for tissue engineering. Chapter 15, Laser processing of biopolymers for development of medical and high-tech devices, prepared by Nadya Stankova et al., discusses the processing properties and applications of laser processing technologies applied for the development of advanced devices. Processing by nanosecond, picosecond, or femtosecond laser pulses of medical-grade polydimethylsiloxane (PDMS) are discussed. Alina-Maria Holban1 and Alexandru Mihai Grumezescu2 1

Faculty of Biology, University of Bucharest, Bucharest, Romania 2 Faculty of Applied Chemistry and Materials Science, University Politehnica of Bucharest, Bucharest, Romania

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Interactions between tissues, cells, and biomaterials: an advanced evaluation by synchrotron radiation-based highresolution tomography

1

Alessandra Giuliani1, Serena Mazzoni1, Adrian Manescu2 and Giuliana Tromba3 1

Department of Clinical Sciences, Universita` Politecnica delle Marche, Ancona, Italy 2 “Victor Babes” University of Medicine and Pharmacy, Timisoara, Romania 3 Sincrotrone Trieste S.C.p.A, Trieste, Italy

1.1 CONDUCTION, INDUCTION, AND CELL TRANSPLANTATION IN TISSUE ENGINEERING: THE LIMITATIONS OF CROSS-TALK STUDIES BY CONVENTIONAL TECHNIQUES Tissue loss or damage due to congenital defects, disease, and injury are major clinical problems. The majority of anatomical districts are comprised of several tissues for which the preferred method of replacement is through autologous grafting. For instance, autografts are standard for bone grafts because of their biocompatibility, immunogenic characteristics, and because they offer all the essential properties required: they achieve osteoinduction through bone morphogenetic proteins (BMPs) and other growth factors (GFs); osteogenesis, by means of osteoprogenitor cells; and osteoconduction, because autografts are implanted in the shape of 3D porous matrixes (Amini et al., 2012). However, there is often insufficient host tissue for a complete repair of a defect; or for specific diseases, sites, and tissues, the replacement is clinically forbidden. Furthermore, autologous transplants entail extremely expensive procedures, often inducing significant donor site injury and morbidity (Alsberg et al., 2001; Iezzi et al., 2016).

Materials for Biomedical Engineering: Hydrogels and Polymer-based Scaffolds. DOI: https://doi.org/10.1016/B978-0-12-816901-8.00001-8 © 2019 Elsevier Inc. All rights reserved.

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In this context, tissue engineering aims to restore function to or replace damaged or diseased tissues through the application of biological and engineering rationales (Alsberg et al., 2001). The approaches normally adopted, alone or in combination, include: conduction, induction, and cell (normally stem cell) transplantation (Langer and Vacanti, 1993; Putnam and Mooney, 1996). The choice of the correct approach depends on several factors, including the site and the size of the defect, the availability of cells in surrounding areas, cell migration kinetics, and the presence or absence of sufficient vascularization. In conductive approaches (Fig. 1.1, panels A and B), the biomaterial acts as a 3D matrix for endogenous cells migration, grafting, proliferation, and differentiation. These cells form the new tissue that, hopefully, will be integrated with the host tissue and with the grafted biomaterial that, depending on its composition, may or may not degrade over time. Sometimes, however, endogenous cell migration, differentiation of these cells, and the following tissue formation are processes that need to be controlled. In this case, an inductive approach to tissue engineering appears to be the best solution (Fig. 1.1, panels C and D). Bioactive scaffolds, GFs, drug, or plasmid DNA delivery are used to induce cell migration and control cellular behavior. Conductive and inductive methods are usually used in cases of limited defects, as in most maxillary or mandibular bone sites, when the presence of optimized biomaterials promote cell migration from the host tissue into the scaffolds. However, the repair of large defects often also requires the direct transplantation of cells. This approach (Fig. 1.1, panels E and F) is also required when there are not enough available cells in the tissue surrounding the defect, or when endogenous cell migration to the damaged site would require too long a time (Alsberg et al., 2001). In these cases, a biopsy is usually taken from a donor source, cells are isolated and expanded in vitro, and then seeded on a duly prepared scaffold to proliferate and form new tissue. In turn, this tissue is normally implanted into the damaged area. Autologous cells give insignificant immune response, but have the disadvantage of requiring a long time to expand to the needed quantities. Allogeneic cells, which are genetically different cells from the same species as the patient and even more so xenogeneic cells derived from a different species than the patient, present the opposite problems: they are ready available but strongly increase the possibility of immunological reactions. This is the case, for example, when trying to repair muscle damage in Duchenne muscular dystrophy (DMD) by transplanting myogenic progenitors directly into the muscles. In fact, this procedure has shown to suffer from the problems of limited cell survival and reduced migration of these cells in the muscles (Farini et al., 2012). In particular, cell therapies consist of the use of stem cell populations that have been previously manipulated and cultured in vitro, with the objective of repairing and regenerating damaged tissues. In this context, the presence in the heart of primitive cells, able to generate the different structures of the myocardium, has been recently documented (Giuliani, 2012).

1.1 Conduction, Induction, and Cell Transplantation

FIGURE 1.1 Description of the three approaches to tissue engineering. (A and B) Conduction—A scaffold controls and selects cells infiltrating the defect site from the outside in order to repair it. The scaffold may be resorbed in time or surgically removed. (C and D) Induction—Bioactive molecules bind with selected host cells (with receptors for the molecules) that migrate to the defect site forming a new extracellular matrix. (E and F) Cell transplantation—Cells are transplanted from a donor source to a scaffold. Afterwards, the cell/scaffold construct is grafted into the defect site to achieve tissue regeneration in conjunction with host cells that migrate to the defect site.

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The monitoring of the longitudinal outcomes of conduction, induction, and cell therapy requires the use of nondestructive methods that are capable of identifying the location, amount, and extent of cellular survival and fate, as well as of evaluating, qualitatively and quantitatively, the tissue growth under different conditions. This type of study has been frequently performed in the recent literature, for instance to engineer bone (Appel et al., 2013; Giuliani et al., 2011, 2014a; Olubamiji et al., 2014), cartilage (Olubamiji et al., 2014; Zehbe et al., 2010a), and tendon (Gigante et al., 2013). In this context, imaging techniques are assuming an increasingly important role, not only for a rigorous characterization of the properties and functions of biomaterials, but also to investigate the kinetics of their biological behavior in conduction, induction, and cell transplantation processes. Several advanced 2D imaging technologies are available to complement histological evaluation and to study complex biological events occurring at the interface between tissues and biomaterials (Appel et al., 2013; Nam et al., 2015). However, 2D imaging technologies present several limitations. In fact (Zehbe et al., 2010a), the current methods of optical microscopy require a sectioning of serial sections that takes a long time and they are severely limited in their ability to analyze opaque 3D biological structures. Furthermore, although electron microscopy (EM) methods offer good 3D topographic representation (by EM scanning), highresolution imaging is restricted to extremely thin samples (by transmission EM). Moreover, for the previously mentioned techniques, sample preparation sometimes causes significant damage to the specimen, practically inhibiting the visualization and quantification of the cells and tissues spatial distribution, present and formed within porous biomaterials in vitro and in vivo conditions. Furthermore, for regeneration of vascularized tissues, such as bone or muscle, there is a need for imaging techniques able to quantify 3D vascular ingrowth, particularly for recent innovative studies focused on exploring the potential to enhance regeneration via therapeutic angiogenesis strategies (Langer et al., 2009; Appel et al., 2013; Giuliani et al., 2017). The imaging modality most extensively applied for this purpose, especially for bone tissue engineering studies (Giuliani et al., 2013; Manescu et al., 2016a; Komlev et al., 2009; Belicchi et al., 2009), is high-resolution X-ray computed tomography (microCT). In particular, the use of synchrotron-produced X-rays has several advantages with respect to X-rays produced by laboratory or industrial sources. In this chapter, several major experiences applying synchrotron radiation (SR)-based microCT to explore regeneration in bone, tendons, cartilage, and skeletal and cardiac muscle sites were reviewed. Conductive, inductive, and cell transplantation approaches were overviewed, although different emphasis was given to each strategy depending on the specified engineered tissue. In addition, innovative protocols were explored to track donor cells after transplantation in order to clarify their role in tissue regeneration.

1.2 X-Ray Computed Microtomography: A Challenging Diagnostic Tool

1.2 X-RAY COMPUTED MICROTOMOGRAPHY: A CHALLENGING DIAGNOSTIC TOOL A new concept of diagnostic by imaging was developed at the beginning of the 1970s, when the first equipment for X-rays computed tomography (CT) was produced. The CT technique overcame several limitations of conventional X-ray radiology, many of which were mainly due to the 2D nature of the radiological images that had been used up to that moment in medical diagnostics. Indeed, conventional and digital X-ray radiology are imaging methods with constraints linked to their two-dimensionality: radiographs provide a 2D image of a 3D object, not accurately replicating the anatomy that is being assessed. Anatomical structures may superimpose causing misleading signals for radiograph interpretation. Indeed, 2D radiographs usually show minor structure damages compared to those actually present and do not reveal the interactions between soft and hard tissues (Manescu et al., 2016b). For these reasons, the diagnostic contribution provided by the CT was pioneering, allowing for the visualization of internal details of the sample with unprecedented precision, in a nondestructive way and achieving a contrast up to 1000 times better than conventional radiography (Claesson, 2001). Tomography data are a collection of cross-sectional images obtained from either transmission or reflection configurations and collected by illuminating the sample from many different directions (Kak and Slaney, 1988). The first applications were in diagnostic medicine, but today CT is a characterization technique also used in several nonmedical imaging applications. X-ray microtomography (microCT) exploits the same physical principles of conventional CT usually used in medical diagnostics, but, unlike this, reaches a spatial resolution of up to 0.1 μm (Weitkamp et al., 2010), that is, about three orders of magnitude higher. The number of projections and of data points per projection define the spatial resolution: thus, large datasets contain more information (i.e., more pixels in a smaller object is equivalent to better spatial resolutions). Indeed, the choice of spatial resolution versus overall sample size is a crucial issue in microCT. Both 3D conventional CT and microCT employ X-rays to virtually reconstruct the samples of interest based on their attenuation coefficient. In an ideal setup, the sample must absorb a value close to and at most equal to 90% of the incident photons: in this configuration, it is expected to achieve the best signal-to-noise ratio. Tomographic images are reconstructed starting from its projections, that is, the radiographies acquired over 180
sample rotation (or reciprocal rotation of the detection system around a fixed object). Hundreds of 2D projection radiographs are taken at several different angles: each of them is a projection of absorption density distribution, in the direction of the photons, onto the plane perpendicular

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to the direction of the beam. This means that, if the sample is imaged in different orientations, 3D (i.e., volumetric) information on the sample structure can be obtained in a second phase using computer algorithms. This second phase, referred to as “tomographic image reconstruction,” is based on solving an inverse problem, estimating an image from its line integrals on different directions, and is theoretically equivalent to the inversion of the Radon Transform of the image. Indeed, in 1917 Radon solved the problem to reconstruct a function from radiographic projections: this finding was exploited as Hounsfield’s invention of the X-ray computed tomographic scanner for which the same Hounsfield received the Nobel Prize in 1972. Nowadays, there are two types of algorithms with different approaches (Penczek, 2010) for the reconstruction phase: transform-based methods which exploit analytic inversion formulae; and series expansion methods based on linear algebra. The traditional reconstruction algorithm used in most practical applications of microCT is the filtered back projection (FBP), a Fourier-based technique. FBP is derived from the Fourier slice theorem, described in detail elsewhere (Kak and Slaney, 1988). The advantages of FBP are that its implementation is straightforward and execution is relatively fast. Another approach is based on iterative reconstruction algorithms: these give an initial guess of the attenuation coefficients (Webb, 2003) and compare such estimations with those actually acquired. The correction of the initial matrix is made in an iterative way for each projection in a first step and for the whole dataset in a second step until the residual error between the measured data and estimated matrix falls below a predesignated value. Iterative schemes are scarcely used in standard CT. 3D renderings of the data obtained after the reconstruction are easily obtained by stacking up the slices and may be sectioned in arbitrary ways to better locate and quantify the details. Indeed, if the 2D slices and the 3D reconstructions are fascinating tools to perform qualitative observations of the internal structure of biomaterials, cells and tissues, the real benefit is the quantitative information that can be extracted from 3D datasets (Ohgushi et al., 1989). Different methods may be applied to extract quantitative architectural parameters from tomographic images. For instance, in the field of bone research, different protocols have been proposed for bone microarchitecture quantification. The 3D mean intercept length method provides a good estimation of trabecular thickness and spacing based on structural geometry assumptions, for example, parallel plate model (Hildebrand and Ruegsegger, 1997a). However, using 3D images such assumptions can be avoided, allowing for the achievement of new model-independent quantitative parameters (Hildebrand and Ruegsegger, 1997b).

1.3 HI-RES Tomography

1.3 INNOVATIVE APPROACHES TO HIGH-RESOLUTION TOMOGRAPHY BY SYNCHROTRON RADIATION By properly selecting the photon energy, the interaction between X-rays and biological structures provides semitransparency of tissues, allowing penetration of even large specimens. Based on the algorithms described in the previous paragraph, angular projections can be used for tomographic imaging. The imaging of cells inside biological tissues is challenging with conventional microCT devices, because of several experimental conditions linked to the limited spatial and structural resolution. A major problem is the low difference in density and absorption contrast between cells and surrounding tissue. Therefore, a monochromatic X-ray beam, sufficiently high photon flux, and coherent beam properties are key requirements, currently only achieved with synchrotron light (Zehbe et al., 2010a). SR is an electromagnetic light created when charged particles (for instance electrons) are emitted by an electron gun and are then linearly accelerated by an electric field. Next, the particles are further accelerated to near the speed of light in two connected rings (the first, named the “booster ring,” and the second the “storage ring”). In the storage ring, as the electrons travel round the ring, they pass through different types of magnets and, in the process, they produce X-rays. Indeed, these magnets cause the electrons to change direction: this results in a change in their velocity vector and, consequently, in the emission of SR. In particular, when the electron is moving fast enough, the emitted energy is at X-ray wavelength (Giuliani et al., 2014b; Olubamiji et al., 2014; http://www.esrf.eu/ about/synchrotron-science/synchrotron). The properties of SR significantly improve contrast sensitivity in X-ray imaging systems (Cancedda et al., 2007; Giuliani et al., 2010), as specifically discussed in detail in the next paragraphs for the different anatomical districts. X-rays produced by synchrotron facilities have several advantages compared to those delivered by desktop laboratory sources. In particular, they offer several possibilities: 1. to get a high photon flux, achieving measurements with high signal-to-noise ratio as well as high spatial resolutions; 2. to tune the X-ray source, allowing measurements at different energies; 3. to set a specific monochromatic X-ray radiation, eliminating the beam hardening effects; 4. to perform parallel beam acquisitions, allowing the use of specific tomographic reconstruction algorithms. Due to the listed advantages offered by synchrotron light sources, several 3D imaging techniques have been developed in order to serve and support biomedical research.

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In a pure absorption experimental setup, the sample is as close as possible to the detector, with a minimization of image blurring. At synchrotron facilities, parallel photon beams are made available from a wide cross-sectional source; they pass through the sample and are collected by a 2D detector. Being the X-ray beam parallel, the projection of each slice of the object on the detector is not dependent of all the other slices (Stock, 2009). Absorption tomography (Fig. 1.2, panel A) is based on the Beer-Lambert law, which describes that the intensity of monochromatic X-rays that transmit the specimen decreases exponentially as a function of the line integral of the linear attenuation coefficients along the X-ray path:  ðN  I 5 I0 3 exp 2 μðxÞdx

(1.1)

2N

where μ is the linear absorption coefficient at position x along a particular beam direction. Tomography solves the problem of assigning the correct values of μ to each position along the photon path, knowing only the values of the line integral but for a large number of paths through the sample (i.e., rays sent through the sample from different angles). As they pass through the different phases of the sample, the photons interact in the electron shells of the atoms they pass. Absorption tomography exploits that the attenuation of X-ray beams of a given energy varies with the atomic electron density of the imaged material and its bulk density. The differences in the X-ray attenuation rate within the samples are represented by different peaks in the gray-scale histogram, corresponding to the different phases. However, by modulating the sample-to-detector distance, contrast is also generated by phase differences among the scattered X-ray waves (Fig. 1.2, panel B). In particular, this phase-contrast (PhC) effect puts into evidence the interface and edges between two materials, and it is particularly useful when media with similar absorption coefficients should be discriminated (Fiori et al., 2012). Indeed, differently from conventional X-ray tomography, in the PhC approach, the image contrast is not based solely on attenuation of the beam. In this case, the effect of an X-ray beam penetrating the sample is described by the refractive index, n 5 1 2 δ 1 iβ

(1.2)

where δ is the refractive index decrement and β is the attenuation index. δ is actually proportional to the mean electron density of the specific phase, which in turn is nearly proportional to its mass density. Moreover, the δ value is much larger than the imaginary part β in nonmineralized tissues, suggesting that the phase approach provides greater sensitivity than the absorption approach when studying these tissues. The methods used for the reconstruction of the refractive index n are typically based on a two-step approach: the phase projections are extracted in the first step while the refractive index decrement δ is reconstructed by applying a conventional FBP or similar algorithms in the second step (Giuliani et al., 2014a,b).

1.3 HI-RES Tomography

FIGURE 1.2 (A) Standard (absorption-based) tomography—The sample is mounted on a translationrotation stage (standard SR-microCT setup). The detector (made up of a scintillator, light microscope optics, and a CCD) is mounted on a translation stage. Projections are acquired with the detector close to the sample and, afterwards, are processed with the filtered back projection algorithm for the reconstruction of the 3D absorption index. (B) Propagation-based phase-contrast imaging—The imaging source consists, also in this case, of monochromatized SR X-rays. The long source to sample distance yields a high degree of spatial coherence. The detector, mounted on a translation stage at a long distance from the sample, allows free space propagation of the beam after the sample. In this case, the application of the FBP algorithm produces an edge-enhancement effect, which is proportional to the Laplacian of the refractive index. (C) In-line phase tomography (holotomography), from Ref. (Giuliani et al., 2013)—The tomographic acquisition is performed with the detector distances (D1, D2,. . ., Dn) from the sample. For each rotation angle, a phase retrieval algorithm is applied to projections acquired at each distance, providing the phase maps, in turn, are then processed with the FBP algorithm to recover the 3D refractive index decrement.

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For the first phase (Langer et al., 2010), two main classes of algorithms can be identified in literature: they are alternatively based on a linearization with respect to the propagation distance, which yields what is known as the transport of intensity equation (TIE) (Teague, 1982a,b), or on a linearization with respect to the object, which yields the contrast transfer function (Cloetens et al., 1999). Furthermore (Manescu et al., 2016a,b), phase retrieval usually implies the reconstruction of the two different real-valued 3D distributions, δ and β; such reconstruction generally requires the acquisition of 2D projections, at least at two different sample-detector distances at each view angle (Fig. 1.2, panel C). However, in some cases, it can be shown a priori that the real and imaginary parts of the refractive index are proportional to each other, that is: β5ε δ

(1.3)

where the proportionality constant ε does not depend on the spatial coordinates. This assumption is possible only for special classes of objects, such as pure phase (i.e., weakly absorbing) objects, or homogeneous objects, such as objects consisting predominantly of a single material (possibly with a spatially varying density) (Gureyev et al., 2004, 2006, 2009). Moreover (Giuliani et al., 2014a,b), Bronnikov suggested an algorithm providing a direct reconstruction of the refractive index and avoiding the first phase retrieval step. It established a fundamental relation between the 3D Radon transform of the object function and the 2D Radon transform of the phase-contrast projection (Bronnikov, 2000). Thus, a reconstruction algorithm is derived in the form of a FBP. In recent years, there has been increasing interest in the mentioned approaches to evaluate different biomaterials’ performance by means of SR-microCT. Tissue regeneration derived from hosting sites grafting with different types of biomaterials (with or without stem cells seeding), was recently explored using SRmicroCT (Cancedda et al., 2007; Giuliani et al., 2014a, 2016; Rominu et al., 2014; Gigante et al., 2013). Evaluation of newly formed tissue is usually based on histology, by observation of one or more sections; however, conventional histological evaluation and corresponding histomorphometric measurements provide only 2D information with the consequent risk that the selected sections do not properly represent the entire biopsy specimen (Giuliani, 2016). Furthermore, if the involvement of neighboring tissues with different morphology (bone, unmineralized extracellular matrix, regenerated vessels, etc.) on the regeneration process of defects or tissues is unknown or still not clearly verified, 3D analyzing methods, including high-resolution SR-microCT, are indicated to explore the dynamic and spatial distribution of regenerative phenomena in these anatomic structures. Traditionally, absorption imaging with SRmicroCT in medical applications was performed with almost no distance between sample and detector, obtaining significant information on morphometric distribution of the bioengineered structures (Cancedda et al., 2007; Giuliani et al., 2014a, 2016; Renghini et al., 2013; Pozdnyakova et al., 2010). However, homogeneous materials,

1.4 Skeletal Tissue Engineering

with a low attenuation coefficient (like collagen, polymers, thermoset and thermoplastic matrices, unmineralized extracellular matrix, vessels, nerves, etc.), or heterogeneous materials, with a narrow range of attenuation coefficients (like the case of heterologous bone scaffolds or graded mineralized bone), produce insufficient contrast for absorption imaging. For such materials, the imaging quality can be enhanced through the use of PhC-microCT, with an increased distance between sample and detector (Manescu et al., 2016a,b; Giuliani et al., 2014b; Albertini et al., 2009). In addition, whereas PhC-microCT can be based on a single distance between the detector and the sample (for the special classes of objects previously mentioned), holotomography (HT) involves PhC imaging at several distances (Fig. 1.2, panel C), then combining the phase shift information to generate 3D reconstructions. HT is helpful when the material of interest has extremely small variations in attenuation coefficients, which lead to unsatisfactory imaging results even with phase-contrast techniques on a single-distance and the phase retrieval algorithms previously described (Giuliani et al., 2013; Giuliani, 2016).

1.4 SKELETAL TISSUE ENGINEERING 1.4.1 BONE Bone structures have fundamental functions in the body. When congenital defects, trauma, or diseases are present, there is a significant need for bone replacement (Alsberg et al., 2001). The combination of living cells, biologically active molecules, and a structural scaffold to form a construct able to promote the repair and regeneration of bone is the fundamental concept underlying bone engineering. The scaffold plays a crucial role, being expected to support cell colonization, migration, growth, and differentiation. In parallel with bone formation, the scaffold may undergo degradation, releasing products that have to be biocompatible or that are easily excreted or subjected to metabolism (Hutmacher et al., 2007). In this context, imaging techniques, including SR-PhC-microCT investigations, were extensively applied to investigate the properties of several biomaterials proposed to act as scaffolds. Since each bone site performs multiple functional roles, it is unlikely that a single scaffold would serve as a universal support for the regeneration of bone tissue. The considerations for scaffold design are, hence, complex and include material composition, architecture, structural mechanics, surface properties, degradation properties and products, together with the composition of any added biological component and, of course, the changes in all of these factors with time (Hutmacher et al., 2007). Allowing an accurate 3D examination of samples, SR-based microCT was not only employed to reconstruct, at high resolution, the complex architecture of bone

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tissue at different scales (Langer et al., 2012; Peyrin et al., 2014; Giuliani et al., 2018a) and in different genetic and environmental conditions (Tavella et al., 2012; Costa et al., 2013; Canciani et al., 2015; Giuliani et al., 2018b), but it is also increasingly becoming a powerful tool for engineered bone characterization in different skeletal sites. In this scenario, interesting microCT studies have been performed on different biomaterials that have previously been indicated as bone-substitute (Olubamiji et al., 2014). SR-microCT was exploited by Yue (Yue et al., 2010) to characterize the scaffold morphology, the mineral distribution within scaffold pores, and the tissue ingrowth in 4-week-old explants of a bioactive glass foam scaffold implanted between the muscle and tibia of a mouse. Similarly, SR-microCT was used to successfully identify scaffold architecture and bone ingrowth into cell-loaded hydroxyapatite scaffolds implanted in immunodeficient mice for 8 weeks (Mastrogiacomo et al., 2004). The bone ingrowth was estimated in terms of total volume fraction, distribution, and thickness in the pores of the implant and the scaffold architecture was analyzed in terms of the porosity and spatial distribution of walls. The same group explored the ability of SR-microCT to examine the progressive resorption of and bone ingrowth into scaffolds implanted in immunodeficient mice for repair times of 8, 16, or 24 weeks (Cancedda et al., 2007; Papadimitropoulos et al., 2007; Komlev et al., 2006). When using a hydroxyapatite scaffold (Engipore), a single peak was observed in the X-ray absorption histogram before implantation, corresponding to the biomaterial used for the manufacturing of the scaffold itself (Fig. 1.3, panel A). After implantation, an additional peak was observed at lower X-ray absorption values, corresponding to the newly formed bone (Fig. 1.3, panel B). It is possible to observe that the newly formed bone peak shifted to higher values of linear attenuation coefficient with the increasing of the implantation time: this is explained by the progressive mineralization of the bone. In dental districts, successful bone regeneration using biphasic calcium phosphate materials was reported in some clinical applications for maxillary sinus elevation (Mangano et al., 2013; Ohayon, 2014). Special morphologies of 3D scaffolds, in the shape of granules or structured blocks, were shown to realize promising scaffolds to be used either in an acellular strategy (pure scaffold grafting and its colonization by endogenous cells) (Mangano et al., 2013) or combining the biomaterial with cells in vitro (Barboni et al., 2013). While previously reported studies were usually based on single time points, the long-term kinetics of bone regeneration being not fully investigated, a recent clinical study (Giuliani et al., 2016) reported a quantitative kinetics evaluation of blocks versus granules in biphasic calcium phosphate scaffolds carried out by SR microCT. Twenty-four bilateral sinus augmentations were performed and grafted with HA/β-TCP 30/70, 12 with granules and 12 with blocks. The samples were retrieved at 3, 5/6 and 9 months from grafting and were evaluated for bone regeneration, graft resorption, neovascularization, and morphometric parameters. Big

1.4 Skeletal Tissue Engineering (A)

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FIGURE 1.3 Hydroxyapatite scaffold (Engipore) before (A) and after 8 (empty dots), 16 (full dots), and 24 (triangles) weeks of implant (B). Profiles of gray levels for the whole sample. The peaks on the right (xD6.5 cm21) correspond to the hydroxyapatite scaffold while in (B), the central peak (xD2.5 cm21), corresponds to the new bone. This peak shifts to the right after 24 weeks of implantation, due to an increase in mineral content. From Cancedda et al. (2007).

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quantities of newly formed bone were detected in the retrieved biopsies, together with a good rate of biomaterial resorption and the formation of new vessels (Fig. 1.4). While the morphometric parameters were comparable up to 5/6 months from grafting, 9 months after grafting microCT revealed that the sites grafted with blocks (Fig. 1.4, panels E and F) mimicked the healthy native bone of the maxillary site slightly better than those grafted with particulates (Fig. 1.4, panels B and C). Other significant studies in literature (Cancedda et al., 2007; Komlev et al., 2009) have demonstrated, by SR-microCT, the bioactive role of TCP or of TCP in combination with HA in bone regeneration even if, in these cases, the regeneration was limited to ectopic sites of animal models. Naturally produced bioceramics are an interesting alternative to biphasic calcium phosphate materials. For instance, coral has been used for a long time as a scaffold for bone tissue engineering because of its porous and interconnected

FIGURE 1.4 MicroCT 3D reconstruction of (A) granule-based TCP/HA scaffold before grafting, (B and C) granule-based TCP/HA scaffold 9 months after grafting, (D) block-based TCP/HA scaffold before grafting and (E and F) block-based TCP/HA scaffold 9 months after grafting. (BE) Residual scaffold 9 months after grafting. Legend for panels C and F— dark gray: regenerated vessels; light phase: newly formed bone, and white phase: scaffold. From Giuliani, A., Manescu, A., Mohammadi, S., Mazzoni, S., Piattelli, A., Mangano, F., et al., 2016. Quantitative kinetics evaluation of blocks versus granules of biphasic calcium phosphate scaffolds (HA/β-TCP 30/70) by synchrotron radiation X-ray microtomography: a human study. Implant. Dent. 25 (1), 615.

1.4 Skeletal Tissue Engineering

architecture, its high compressive breaking stress, and high biocompatibility and resorption properties. All these characteristics make coral a promising candidate for use as delivery vehicles for cells. (Piattelli et al., 2016; Iezzi et al., 2016). The calcium-carbonate coral-derived material, named Biocoral, is a naturally produced ceramic biomaterial, such as animal and plant skeletons (hydroxyapatite or calcium carbonate). SR-microCT revealed in 3D that 6 months after grafting, a huge amount of newly formed bone was present in the retrieved Biocoral-based samples, coupled with a good rate of biomaterial resorption and the formation of a homogeneous and rich net of new vessels. The morphometric parameters were comparable to those obtained in the biphasic calcium phosphate-based samples after the same amount of time from grafting, with the exception of the connectivity index for which biphasic calcium phosphate-based samples exhibited the most well-connected structure (Giuliani et al., 2014a). The previously reported clinical cases demonstrate that mineralized-scaffold grafting is a major strategy for the repair and reconstruction of bone defects and that the SR-microCT technique plays a fundamental role in advanced characterization of such bone tissue-engineered constructs. However, the reconstruction of the mandible is still a challenge for oral and maxillofacial surgeons, at least until now. Indeed, the various methods that have been used, including insertion of bone grafts and allogeneic materials, are still not completely satisfactory: alloplastic materials carry the risk of bacterial infection and can perforate skin or oral mucosa (Engstrand, 2012), the harvesting of bone grafts is associated with morbidity and possible functional impairment at the donor site (Pieske et al., 2009), and the products used in some procedures can have costly, time-consuming manufacturing processes and controversial ethical issues (Sukumar and Drı´zhal, 2008). In this context, some researchers (Giuliani et al., 2013) sought to study the bone regeneration process in injured human mandibles repaired with autologous dental pulp stem cells (DPSCs)/collagen sponge biocomplex implants. They reported that stem cells unexpectedly regenerated a compact rather than a spongy bone type. This was assessed by SR-based HT, fully exploiting the sensitivity of phase-contrast imaging as discussed in Section 1.3. SR-HT reconstructed 3D images and analyzed tissues at resolutions comparable to histology. However, with respect to morphometric analyses conducted through histology, SR-HT generated additional and more reliable quantitative information because the entire 3D samples, rather than only selected 2D sections, were assessed.

1.4.2 CARTILAGE Degenerative, rheumatic, or traumatic processes are the predominant causes of articular cartilage damage. Biologic agents can block such deterioration, but the tissue has only limited regenerative potential and drugs are still unable to rebuild cartilage. Such limited ability of articular cartilage to regenerate often makes joint arthroplasty unavoidable (Makris et al., 2015). Consequently, cartilage tissue

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engineering was attempted by autologous cells conduction, sometimes filling the defect area with cell-scaffold constructs, or enhancing the regeneration through biochemical targeting mechanisms (Zehbe et al., 2010a). Zehbe et al. (2009) showed the efficiency of SR-microCT to image volumetric cartilage morphology, evaluating the spatial distribution of single cells inside the tissue and their quantification, and comparing their findings to conventional histological techniques. Recently, a series of acellular and cellular regenerative products and techniques were tested in order to promote the development of functional articular cartilage which could revolutionize joint care over the next decade. Acellular scaffolds are usually made of collagen or hyaluronic-acid-based materials, while cellular strategies use either primary cells or stem cells with or without the support of a scaffold. In their research on scaffolds for cartilage tissue engineering, Zehbe et al. (2009) used a directional freezing process to structure water-based solutions of gelatin. The electrolysis of water prior to freezing resulted in the introduction of gas bubbles inside the gelatin solution before its directional freezing. The addition of other components, like salts, acids, ceramics, or polymer particles, allowed for the synthesis of composite scaffolds and for achieving different pore morphologies. The scaffolds, after seeding with porcine chondrocytes, were stained with a combined Au/Ag stain to enhance the absorption contrast in SR-microCT. While only some cells showed enhanced absorption contrast, most cells did not show any difference in contrast to the surrounding scaffold and were consequently not detectable using conventional greyscale threshold methods. Therefore, an imagebased 3D segmentation tool was used on the tomographic data, revealing a multitude of nonstained cells. However, although a lot of studies referred to cartilage engineering concentrate on in vitro examinations, SR-PhC-microCT is currently used for the visualization of tissue-engineered repair in exercised tissues. In fact, providing the phase-contrast X-ray imaging a substantially enhanced contrast resolution for soft tissues compared to conventional absorption techniques, some works (Bravin, 2003; Coan et al., 2005, 2010) demonstrated that SRPhC-microCT is able to show structural properties of the cartilage matrix in excised tissues. The level of detail within the tomographic images is sufficiently consistent to allow for the differentiation between osteoarthritic and healthy cartilage (Coan et al., 2010). Indeed, the edges and interfaces between cartilage and calcified cartilage/subchondral bone are clearly discernible because of the beam refraction occurring at the edges between tissues with different refraction indexes. The osteoarthritic samples showed significantly lower chondrocyte distribution homogeneity, less chondrocyte alignment, lower height of tangential, transitional, and radial zones, and a higher prevalence of superficial cartilage damage (Coan et al., 2010). Practically, if in the future phase-contrast imaging will be translated into the clinical scenario, then it could become an important tool for the evaluation of

1.4 Skeletal Tissue Engineering

osteoarthritis (especially for early detection and monitoring of the disease) and, in general, a valuable noninvasive alternative to routine histological examination of tissue-engineered cartilages.

1.4.3 TENDONS Research on tissue engineering recently involved some attempts at tendon repair. This field is currently facing a significant challenge because of the urgency for developing strategies that will lead to a clinically effective and commercially successful product (Shearn et al., 2011). Indeed, this challenge is primarily motivated by the fact that injured tendons have limited repair ability after full-thickness lesions. Treatments normally involve surgical repair but usually the damaged tendon is unable over time to bear the mechanical stress of daily activities, with a consequent high risk of refailure. Autografts, allografts, xenografts, and synthetic prostheses have been tested for tendon augmentation. Although autografts have always been considered the best solution, the substantial morbidity of the donor site has pushed research toward alternative solutions (Longo et al., 2010). Silk, carbon fibers, Gore-Tex, and LARS1 have been tested as alternatives to autografts, but several cases have highlighted endurance and durability problems related to frictional problems (Gigante et al., 2013; Thomas et al., 2011) and side effects ranging from immunogenic reactions to necrotic phenomena (McKibbin, 1984). Better results were achieved in the early 1990s with the introduction of collagen membranes. In this context, some authors (Gigante et al., 2013) performed an ex vivo study, assessing the tendon regeneration ability of a new oriented collagen-I membrane in an experimental animal model, using conventional histological analysis and SR-microCT examination. Ten New Zealand White rabbits were sectioned in the central third of the patellar tendon that was grafted with a membrane obtained from purified equine Achilles tendon. The contralateral patellar tendons were cut longitudinally and evaluated as sham-operated control. Histological and SR-microCT findings showed satisfactory graft integration with native tendon without adverse side effects. SR-microCT experiments were performed in absorption configuration at the ID19 beamline of the European Synchrotron Radiation Facility (ESRF, Grenoble, France) and in phase-contrast setup at the SYRMEP beamline of the ELETTRA Synchrotron Radiation Facility (Trieste, Italy). Morphometric data and fiber thickness distribution analysis, obtained by the ESRF experiments, showed a more connected fiber stacking in the treated tendon than in the lesion site of the control tendon. A more uniform distribution of fiber thickness was found in treated compared to control tendons, suggesting possible better biomechanical performances, after healing, of the treated tendons. Furthermore, the microCT data allowed to evaluate also the mass density of treated tendons in terms of density. The implant site of treated tendons and the lesion area of control tendons were not significantly different, confirming good

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FIGURE 1.5 New Zealand White rabbit tendon grafted with collagen-I membrane. SR-PhC-microCT 3D reconstruction. Micrometric fiber distribution 6 months after surgery.

integration of the oriented collagen-I membrane in the treated site. A 3D reconstruction of a treated tendon is shown in Fig. 1.5. However, tendon tissue engineering also needs to employ more clinically relevant models of tendon injury, such as degenerative tendons. There is an urgent need to translate the previously described successes from small to larger animal models, with the objective of beginning to explore the clinical implications of these treatments (Shearn et al., 2011). In this direction, the use of SR-microCT was demonstrated to be, without a doubt, a fundamental tool because of its unique ability to quantitatively provide, at 3D level, the longitudinal outcome of a particular treatment based on the use of innovative biomaterials and the interaction of these with the host tissues.

1.5 MUSCLE TISSUE ENGINEERING 1.5.1 SKELETAL MUSCLES Skeletal muscles usually have good regeneration capacities, but, in cases of severe trauma or diseases, the loss of muscle functionality is inevitable (Qazi et al., 2015).

1.5 Muscle Tissue Engineering

In the past years several mechanisms of skeletal muscle repair have been explained thanks to research in the field of skeletal muscle tissue engineering. In this direction, various types of cells and bioactive factors, playing an important role during regeneration, were identified (Edmunds and Gargiulo, 2015). For instance, the delivery of myogenic stem cells to the sites of muscle lesions through systemic circulation is a promising approach to treating DMD. This is considered a valid alternative to the attempts to repair muscle damages by transplanting myogenic progenitors directly into muscles, this solution being inhibited by limited cell survival and limited migration of donor cells in the muscles (Belicchi et al., 2009). In the study reported in Torrente et al. (2006), the authors used SR-microCT for 3D visualization of stem cells, previously labeled with FeO (Endorem) nanoparticles, transplanted via intra-arterial infusion. SR-microCT showed the distribution of intra-arterially delivered stem cells within muscle biopsies, providing biological insights into the early processes of muscle stem cell homing. As shown in Torrente et al. (2006), labeled CD133 1 stem cells were distributed around the vessels of muscle tissues within 24 h of their intra-arterial transplantation. The same authors repeated the previous SR-microCT experiment on live animals (Farini et al., 2012), scanning and visualizing the leg injected with labeled CD133 1 stem cells at different times from cell administration. It was found that the majority of intra-arterially injected stem cells accessed the muscle tissues after several rounds of recirculation, but within the first 2 h of the injection. The originality of this study lies in the fact that it was the first investigation of the kinetics of the distribution of intra-arterially injected human stem cells into the capillary system of downstream dystrophic muscles. The efficient transplantation of stem cells into the muscle of dystrophin-deficient mice further supported the approach of intra-arterial delivery of cells for cell-based clinical therapies of neuromuscular diseases, such as DMD (Farini et al., 2012). However, delivering stem cells through systemic routes also presents some obstacles; indeed, the injected cells may become trapped intravenously in other organs, like the liver, spleen, or lungs, so that only a small portion enters the muscle microvasculature and migrates to the dystrophic muscles (Belicchi et al., 2009). Therefore, in order to improve the therapeutic effects of cells and GFs, several biomaterials have been tested on animal models in order to find a suitable biomaterial to serve as a template for guided tissue reorganization. This biomaterial should be a matrix that provides optimum microenvironmental conditions to cells, a delivering system capable of releasing bioactive factors in a controlled manner, and be able to act as local niches for in situ tissue regeneration (Qazi et al., 2015). Several materials have been tested as scaffolds for muscle tissue engineering (Albertini et al., 2009): natural materials (like collagen and alginate), cellular tissue matrices (like bladder submucosa and small intestinal submucosa), and

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synthetic polymers [like polyglycolic acid (PGA), polylactic acid (PLLA), and polylactic-co-glycolic acid (PLGA)]. In this context, SR-microCT is a promising imaging modality for the assessment of such engineered scaffolds and their morphometric parameters, such as porosity, pore size, and interconnectivity. The study of cellular integration and interaction with engineered scaffolds is achieved through microCT, as images can be acquired sequentially over time with minimal negative affects to cells from X-ray dose. Based on data from the literature, and considering that synthetic polymers have the advantages of degrading by nonenzymatic hydrolysis and producing nontoxic molecules, some researchers (Albertini et al., 2009) selected PGA/PLLA (a mixture of 50% PGA and 50% PLLA) as a bioscaffold to test cell viability after their loading. They used SR-PhC-microCT to visualize in 3D the ECM organization, after in vitro seeding on PGA/PLLA of bone marrowderived human and murine mesenchymal stem cells induced to myogenic differentiation and previously labeled with iron oxide nanoparticles. X-ray microCT allowed for the detection, with high spatial resolution, of the 3D structural organization of ECM within the bioscaffold, clarifying how the presence of cells modified the construct arrangement. Furthermore, species-specific differences between the matrix produced by human and murine cells were observed. However, in order to achieve sufficient X-ray contrast, muscle tissues and the previously listed biomaterials often need to be freeze-dried or kept under dry conditions instead of standard culture conditions, sometimes significantly affecting certain cell or tissue types (Edmunds and Gargiulo, 2015). For this reason, the study by microCT of skeletal muscle engineered biopsies is still challenging and requires an optimization of experimental phase-contrast setups.

1.5.2 HEART The engineering of heart myocardium has experienced exciting progress in the past 10 years. Advances in stem cell biology, tissue engineering, and knowledge of biomaterials suggest that cardiac tissue engineering techniques will be strongly used in preclinical research and drug development in the coming years. In order to reach these objectives, several steps need to be taken, including the standardization of myocyte production methods, the establishment of simple and efficient protocols for the vascularization of biomaterials (patches), systems for maturation of myocytes, and, finally, thorough selection of predictive diagnostic techniques and tools for preclinical investigations (Hirt et al., 2014). Traditionally, the engineering of 3D cardiac tissue was motivated by the need to produce in vitro tissue surrogates for cardiac repair, but also by the fascinating experience of observing a heart muscle beating in a dish. Unfortunately, despite the numerous papers that are published on the subject every year, cardiac tissue engineering has not yet entered clinical practice and has still not found wide application in preclinical drug development (Hirt et al., 2014).

1.5 Muscle Tissue Engineering

Published reports have contributed to identifying possible cellular therapy approaches to generate new myocardium, involving transcoronary and intramyocardial injection of progenitor cells (Giuliani et al., 2011). While light fluorescence scanning and transmission EM techniques attempt to visualize the tissue-rebuilding process but are limited to 2D local information, in vivo imaging methods, like MRI, PET, and conventional CT, could play a major role in achieving the quantification of the rebuilding process, including longitudinal cell tracking. However, these 3D techniques present intrinsic limitations to identifying the localization and fate of the injected cells in both clinical and experimental settings, as exhaustively described by Terrovitis et al. (2010). In this context, some researchers (Giuliani et al., 2011) explored the use of microCT in absorption and phase-contrast setups as experimental techniques with high spatial resolution for the detection of rat cardiac progenitor cells (CPCs) previously labeled with iron oxide nanoparticles inside infarcted rat hearts, 1 week after injection and in ex vivo conditions. Through microCT, they were able to observe, in 3D, the presence of these cells after migration to the damaged cardiac tissue, with important structural details that are difficult to visualize using conventional bidimensional imaging techniques. Indeed, 3D visualization of the spatial distribution of the grafted cells with respect to the myocardium and vascular system was obtained. In particular, the X-ray absorption of the labeled cells was higher than that of host tissues, allowing their visualization as bright spots in the 3D images (Fig. 1.6). One week after injection, labeled cells were distributed mostly in proximity and toward the damaged infarcted area (Fig. 1.6F), demonstrating migration of CPCs from the injection site (Fig. 1.6D), that is, around the coronary binding. It was also possible to identify finger-like cell structures in the inner part of the left ventricular wall (Fig. 1.6F). This is another example of useful information not detectable by conventional histological methods. Together with these particular structures, attributable to cell clusters, single smaller units were also observed in all areas of the heart, such as in the atria, the large vessels (image not shown), and in the right ventricle (Fig. 1.6D). These are important new data; in particular, they constitute a confirmation that these cells can migrate through the myocardium through a biological mechanism which is still unknown. This new 3D imaging approach, combining absorption and phase-contrast data with the fusion method (Stokking et al., 2003), appears to be an important way of investigating cellular events involved in cardiac regeneration and represents a promising tool for future clinical applications. However, cardiac tissue engineering techniques have yet to reveal their full potential prior to the introduction of stem cells for drug screening and modeling of a patient’s specific disease. Indeed, the limited success of cardiac cell therapy studies highlights the requirement for further research focused on the objective of finding improved cell delivery methods (Hirt et al., 2014). SR-based microCT techniques are able to give unevaluable support in this direction. Indeed, when imaging soft tissues with hard X-rays, like for an engineered heart, phase contrast is often preferred over conventional attenuation

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FIGURE 1.6 MicroCT reconstruction of an infarcted rat heart injected with 5 3 105 rat CPCs. (A) Upper part of the heart (base). (B) Same volume as in (A), with all the phases but the rare (Continued)

1.5 Muscle Tissue Engineering

L

contrast due to its superior sensitivity. However, it is unclear which of the numerous phase tomography methods yields the optimum results for given experimental conditions. Therefore, some authors (Lang et al., 2014) compared three phase tomography methods implemented at the beamline ID19 of the European Synchrotron Radiation Facility: X-ray grating interferometry (XGI), propagationbased phase tomography, that is, single-distance phase retrieval (SD-PhC), and HT, using the entire heart of a rat. They showed that the spatial resolution available to detect morphological features is about a factor of two better for HT and SD-PhC compared to XGI, whereas the XGI data generally exhibited much better contrast-to-noise ratios for anatomical features and in density measurements. In this direction, as a crucial goal of cardiac tissue engineering is the development of implantable constructs, such as cardiac patches, which provide physical and biochemical supports for myocardial regeneration, a key problem is the quantitative monitoring in situ, in a nondestructive way, of the success of these constructs. This is fundamental for longitudinal assessments necessary to translate studies from ex vivo to in vivo, in animal models and human patients. Izadifar (2016) produced experimental nanoparticles with the aim of temporally modulating the GF release in cardiac patches. In the same study, SR-PhC methods for visualization and quantitative assessment of 3D-printed cardiac patches implanted in a rat myocardial infarction model were explored. The results showed that polymer and external aqueous phase concentrations are the most significant process parameters affecting nanoparticle physical and release characteristics. The experimental nanoparticles, produced in order to overcome the limitations of PLGA nanoparticles, were made of a protein-encapsulating PLGA core and a poly(L-lactide) (PLLA)-rate regulating shell, thus allowing for low burst effect, protein structural integrity, and time-delayed release patterns. Izadifar (2016) also created patches from alginate strands using a 3D printing technique which were surgically implanted on rat hearts for their assessment by SR-PhC-microCT. Phase-retrieved images depicted visible and quantifiable structural details of the patch at a low radiation dose. The microstructural features of fibrin and alginate, which were low-density ( . 97% water) constituents of the patch, were clearly visualized and quantitatively characterized from the phaseretrieved PhC-microCT slices of the implanted patch. The MicroCT technique

labeled cells virtually made transparent. CPCs migrated here from the site of injection. (C) Middle part of the heart (equatorial portion). CPCs were injected here. (D) Same volume as in (C), with all the phases but the labeled cells virtually made transparent. (E) Bottom part of the heart (apex). This portion was largely involved by infarction, as shown by the thinning of the wall. (F) Same volume as in (E), with all the phases but the labeled cells virtually made transparent. From Giuliani, A., Frati, C., Rossini, A., Komlev, V.S., Lagrasta, C., Savi, M., et al., 2011. High-resolution Xray microtomography for three-dimensional imaging of cardiac progenitor cell homing in infarcted rat hearts. J. Tissue. Eng. Regen. Med. 5 (8), e168e178.

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provided anatomical details including the microvessels surrounding the implanted patch. Notably, these microstructural and anatomical features of the patch and heart were achieved without the usage of any contrast agent, suggesting that SRPhC-microCT is of great potential for noninvasive assessment of patch structural changes, myocardium regeneration, and vascularization in live animal studies for cardiac tissue engineering.

1.6 NEW FRONTIERS 1.6.1 CENTRAL AND PERIPHERAL NERVOUS SYSTEM Conventional therapy approaches to neurological diseases present several problems, mainly linked to the restricted intrinsic regeneration capacity of neurons. Therefore, currently, many researchers in various fields, including surgery, internal medicine, pharmacology, medical device technology, chemistry, and cell biology, are actively attempting to find solutions and to establish new therapies for curing neurological diseases. Recently, cell-based regenerative medicine has appeared as one of the most promising approaches for treating such diseases. Regenerative therapy by the direct injection of dissociated cells has been performed in clinical settings, but several published works (Ibarretxe et al., 2012; Lu et al., 2004; Bertani et al., 2005) confirmed only modest therapeutic benefits. Regarding the delivery of drugs into the central nervous system (CNS), the main obstacle is the presence of the bloodbrain barrier (BBB), since it forms a barrier, hindering the delivery of therapeutic agents from the bloodstream. To circumvent it, various drug carrier systems were developed. Carriers of drugs proven to be more efficient as systemic delivery systems include liposomes, polymeric nuclei, polymer micelles, NPs, and ceramic dendrimers (Oliveira et al., 2010). Of these, only liposomes and polymeric NPs have been widely exploited in the administration of brain drugs. A promising strategy to increase therapeutic efficacy while reducing systemic side effects is based on the local administration of therapeutic agents through a polymeric and biocompatible delivering system implanted at the target site (Jain et al., 2006). This approach offers many advantages: in addition to bypassing the BBB, it avoids systemic side effects and toxicity, inactivation of the peripheral drug, and the need to modify the surface of the vehicle. There are also disadvantages of local administration: the dosage cannot be adjusted after implantation, the rate of drug release usually decreases over time, and it may be necessary to repeat the implant for long-term release, requiring invasive surgery. Restoring peripheral nerve damage is another crucial issue in tissue engineering. One promising approach is based on nerve conduits, consisting of biodegradable polymers attempting to realize a preformed structure, serving as guidance for axonal growth in vivo.

1.6 New Frontiers

Some authors (Zehbe et al., 2010b) proposed a rather simple methodology to control cell growth by applying an electrical potential. They demonstrated that SR-microCT can be used to confirm the efficacy from a 3D morphological point of view. In fact, as cells usually grow in vitro in a nonordered way, they developed a technique to print microelectrodes on polymer sheets via a method named “inverse inkjet printing.” The method uses a sputter coater to establish a gold metallization, allowing a coded position of vital cells with the protein fibrin on the anode part of the electrodes to be obtained (Zehbe et al., 2010a). The electrodes presented parallel-aligned rows of thin gold made by the sputter coating process. All rows were contacted anodically allowing for the deposition of cells and fibrin. Afterwards, several microelectrodes, with the top cells, were stacked, fixated using glutaraldehyde, rinsed in distilled water, and freeze-dried for further SR-microCT investigations. The microCT results were rather interesting and unexpected. Being that the experiments were performed in absorption setup, the authors expected to observe strongly absorbing deposited gold structures and a weekly absorbing polymer substrate. Further, as the deposited cells were not stained with any metal stain, they did not expect to observe the cells themselves, as they have a low density and are thin compared to the polymer substrate. Nevertheless, interestingly and unexpectedly, cells were imaged well and in good accordance with the fluorescence microscopic images acquired previously. Another fundamental challenge in neuronal tissue engineering is the imaging of the human brain at cell level and in 3D. In vivo methods lack spatial resolution, and optical microscopy has a limited penetration depth. Research (Hieber et al., 2016) showed that PhC-microCT was able to visualize a volume of up to 43 mm3 of human post mortem or biopsy brain samples. The method was demonstrated on the cerebellum. They automatically identified 5000 Purkinje cells with an error of less than 5% at their layer and determined the local surface density to 165 cells per mm2 on average. Moreover, they highlighted that microCT allowed, by segmentation, the discrimination of subcellular structures, including the dendritic tree and Purkinje cell nucleoli, without the need of a staining process. Once again, SR-PhC-microCT was shown to be successful in the 3D analysis of soft tissues, achieving automatic cell feature quantification, with isotropic resolution and in a label-free manner. This result is expected to open, in the near future, new possibilities in the study of engineered tissues through different types of biomaterials in the CNS and peripheral nervous system districts.

1.6.2 VASCULARIZATION Satisfactory vascularization is a required condition for a clinical outcome of tissue-engineered sites. Thus, it is no wonder that several imaging modalities have been tested for the visualization of new vessels in tissue engineering. Currently, vascular imaging modalities are classified into three major groups (Upputuri et al., 2015): nonoptical methods (X-ray, magnetic resonance,

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ultrasound, and positron emission imaging), optical methods (optical coherence, fluorescence, multiphoton, and laser speckle imaging), and hybrid methods (photoacoustic imaging). However, referring to the first group, the observation of blood vessels using absorption-based X-ray imaging is challenging. To image vessel network structures, contrast agents, or corrosion casts, generating absorption contrast can be introduced and imaged using conventional X-ray techniques. In these cases, radiopaque contrast agents, such as barium sulfate or Microfil, are injected into the vasculature of the animal prior to sacrifice to allow for ex vivo imaging of the samples (Nam et al., 2015). Some researchers (Arkudas et al., 2010) created an automatic observer-independent quantitative method to analyze vascularization using microCT. An arteriovenous loop was created in the medial thigh of 30 rats and was placed in a particulated porous hydroxyapatite and β-TCP matrix, filled with fibrin with or without the application of fibrin-gelimmobilized angiogenetic GFs, vascular endothelial GF, and basic fibroblast GF. In both groups, microCT showed that the arteriovenous loop had led to the generation of dense vascularized connective tissue with differentiated and functional vessels inside the matrix. Quantitative analysis of microCT data also allowed for the assessment of several complex vascularization parameters within the 3D constructs, demonstrating an early improvement of vascularization through the application of exogenous GFs. However, in agreement with several papers, the authors commented that the proposed method presented two limitations, namely the limited resolution of microCT and the incomplete filling of the vessels by the contrast agent. Due to these inherent limitations, complete reliance on absorption contrast leads to significant challenges when attempting to simultaneously identify multiple tissue features in a single sample (e.g., microvascular and calcified tissue structure). Depending on vessel properties and conditions, phase-contrast techniques can provide details on vascular structure in the absence of exogenous contrast agents (Appel et al., 2011). Interferometric, analyzer-based and propagation-based tomography allowed for the detection of the vascular structure within the liver, the first method capable of resolving vessels as small as 50 μm in diameter. Instead, in-line holography micrographs were capable to visualize blood vessels 20 μm thick in diameter in the auricle region of live mice. An in-line technique, named pseudo-holotomography, allowed for the observation, in 3D, of the neovascularization of a porous ceramic scaffold, subcutaneously implanted in mice (Komlev et al., 2009). Silicon-stabilized TCP/MSC composites were implanted subcutaneously on the back of immunodeficient mice. The mice were sacrificed 24 weeks after implantation and the extracted constructs were investigated by absorption-based microCT and HT. MicroCT 3D images were reconstructed from a series of 2D projections using a 3D FBP algorithm, while the “holographic” acquisitions were treated, at each sample-to-detector distance, with a phase retrieval procedure based on TIE with a successive

1.6 New Frontiers

reconstruction of the 3D images using the FBP algorithm. To combine both δ and β maps, an optimized alignment-and-matching procedure for one tomographic and three HT volumes was used (pseudo-HT process). Using this method, the hard and soft tissues and the vascular network were simultaneously imaged and quantified. Another work (Giuliani et al., 2013) assessed the quality of the regenerated vessel network, with in-line HT, in mandible grafts (made of DPSCs seeded on collagen-I scaffolds) 3 years after the grafting intervention. In this study, an innovative method for phase retrieval, experimented by Langer (Langer et al., 2010) and described in Section 1.3, was applied. HT allowed not only to acquire qualitative and quantitative information on hard tissue, but also assessed the presence and the distribution of the blood vessels within the investigated specimens, as shown in Fig. 1.7. Besides the specific results previously described, phase-contrast methods are expected to be extremely useful when considering the vascularization of tissue engineered constructs (Giuliani et al., 2017). In particular, the progress associated to these X-ray imaging techniques could be extrapolated to angiogenesis and microvasculogenesis studies in pathologies characterized by inflammation and tissue damage, such as diabetes and osteoporosis (Giuliani et al., 2013).

FIGURE 1.7 Synchrotron X-ray HT. Portion of a 3D reconstruction for a human mandible revealing the vascularization net. From Giuliani, A., Manescu, A., Mohammadi, S., Mazzoni, S., Piattelli, A., Mangano, F., et al., 2016. Quantitative kinetics evaluation of blocks versus granules of biphasic calcium phosphate scaffolds (HA/β-TCP 30/70) by synchrotron radiation X-ray microtomography: a human study. Implant. Dent. 25 (1), 615.

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1.7 CONCLUSIONS The previously reported studies have applied SR-microCT with great success at imaging different engineered tissue and for the characterization of the chosen scaffolds. They show unequivocally that SR-microCT is a promising methodology to investigate tissues or biomaterial-cell engineered constructs, due to the superior photon quality generated in modern synchrotron facilities and to the multiple methods of preparation of the experimental setup (not always feasible with conventional X-ray tube microCT devices). However, the majority of these examinations focused on tissue samples, sometimes from humans (dental districts) and often from animal models or else they are in vitro studies (Olubamiji et al., 2014). This is mainly due to the high radiation dose combined to the large exposure times to achieve extremely high resolution. In this context, it is important to consider that phase contrast imaging was also effective using low radiation doses; thus, these procedures were fully compatible with preclinical and clinical standards (Hagen et al., 2014). By decreasing CT scan time and ROI, the effective dose could be further reduced in the near future, possibly without significant changes in the image quality. This means that PhCmicroCT has the potential to ultimately become a fundamental diagnostic tool for testing the effectiveness of regenerative medicine therapies, also achieving the monitoring of biomaterial behavior and functionality in vivo after transplantation into live organisms.

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FURTHER READING Garcia-Bennett, A.E., Kozhevnikova, M., Ko¨nig, N., Zhou, C., Leao, R., Kno¨pfel, T., et al., 2013. Delivery of differentiation factors by mesoporous silica particles assists advanced differentiation of transplanted murine embryonic stem cells. Stem Cells Transl. Med. 2 (11), 906915.

CHAPTER

Bioprinted scaffolds

2 Florin Iordache1,2

1

Institute of Cellular Biology and Pathology “Nicolae Simionescu” of Romanian Academy, Bucharest, Romania 2Faculty of Veterinary Medicine, University of Agronomic Sciences and veterinary Medicine, Bucharest, Romania

2.1 INTRODUCTION Three-dimensional bioprinting uses 3D printing techniques to fabricate tissue, organs, and biomedical parts that imitate natural tissue architecture. It combines cells, growth factors, and biomaterials to create a microenvironment in which cells can grow and differentiate in tissue structures. In 3D bioprinting biomaterials are printed layer by layer to produce structures similar to a desired organ or tissue. The first patent related to this technology was proposed in the United States in 2003 and granted in 2006 (Thomas, 2016). Bioprinting involves several steps which include selecting biomaterial, creating a bioprinting model using CAMs software, printing, and analysis of the printed constructs. Bioprinting is a method that uses 3D automated printers that can deposit cells embedded in a biomaterial using a computer aided layer-by-layer system for the fabrication of living tissues and organs. The process of fabrication can be described in three parts:

2.1.1 PREBIOPRINTING In this step the model to be printed is created. Models can be created using CAM software’s resulting 2D or even 3D structures or by importing computed tomography (CT) and magnetic resonance images that are converted into STL (stereolithography) files. Furthermore, the cells that will be printed are cultivated and encapsulated in hydrogels taking into account the number, type, and conditions of cultivation. In some cases the cells are cultivated until they form spheroids and are printed without a scaffold using an extrusion technique (Murphy and Atala, 2014; Chua and Yeong, 2015). Materials for Biomedical Engineering: Hydrogels and Polymer-based Scaffolds. DOI: https://doi.org/10.1016/B978-0-12-816901-8.00002-X © 2019 Elsevier Inc. All rights reserved.

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2.1.2 BIOPRINTING Bioprinting technologies, according to their working mechanism, can be classified into four major modalities, including: (1) extrusion-based bioprinting; (2) inkjet-based bioprinting; (3) laser-based bioprinting; and (4) stereolithography. During extrusion, bioink is deposited as filaments in a precise manner forming targeted 3D custom structures. The structure is stabilized by physical- or chemical-crosslinking which facilitate rapid solidification maintaining the geometrical fidelity of the bioprinted structure. Using this technology, alginate poly(lactic-co-glycolic acid) (PLGA) scaffolds are used for drug delivery applications. Collagen, hyaluronic acid, and gelatin hydrogels are used for cardiovascular tissue engineering for the manufacture of blood vessels, cardiac valves, and myocardial tissue. By mixing cells, nutrients, and biomaterials it forms a solution known as bioink that is placed in the printer cartridge and printed using different techniques. Technical parameters are extremely important for precise and defined scaffolds. These parameters include: speed rate, needle diameter, number of layers, viscosity of biomaterial, and temperature (Fig. 2.1).

2.1.3 POSTBIOPRINTING In this step the constructs should be maintained in good condition for cells to grow and proliferate through the scaffold and to create the tissue specific structure. To create these conditions the constructs are placed in bioreactors which

FIGURE 2.1 Schematic representation of bioprinting methods.

2.1 Introduction

FIGURE 2.2 General steps in 3D bioprinting. CAD, computer aided-design; CT, computer tomography image; MRI, magnetic resonance imaging; STL, stereolithography.

provide nutrients, oxygen, dynamic flow, and microgravity which try to mimic the organisms’ environments (Chua and Yeong, 2015; Hinton et al., 2015) (Fig. 2.2). In general, bioprinted scaffolds can be classified into: (1) hydrogel scaffolds; (2) fibrous scaffolds; or (3) porous scaffolds, but a more useful characterization is dependent on the application (e.g., for soft tissue, for hard tissue, for connective tissue) or composition of the biomaterial (e.g., organic, inorganic, polymeric, natural) (Table 2.2). In 3D bioprinting it is necessary to use an exhaustive approach, and knowledge and methodology from different fields, such as engineering, tissue engineering, stem cell biology, and biomaterials science, in order to create the ideal scaffolds. An ideal scaffold should be biocompatible, nonimmunogenic, nontoxic, antithrombotic, with vasoactive properties, compliant, and should permit remodeling of the host tissue postimplantation. To create such a scaffold its physicochemical parameters must be analyzed, such as: geometry, surface properties, pore size, adherence, degradation, and biocompatibility (Chua and Yeong, 2015; Datta et al., 2017).

2.1.4 GEOMETRY OF SCAFFOLDS The reconstruction of a part of a tissue should be done by respecting its anatomical properties in order to achieve its function. Thus, a scaffold should be constructed with a geometry similar to the target tissue. This can be achieved using CAD/CAM programs for simple structures or by converting CT or MRI images into STL files that are recognized by the bioprinter. Internal geometry is also important in the formation of a tissue because it creates a network with a defined surface-to-volume ratio, supporting cell attachment and proliferation (Chua and Yeong, 2015; Harnett et al., 2007).

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2.1.5 SURFACE PROPERTIES The main surface properties that are analyzed include surface energy, chemistry, charge, topology, and area. For various biomaterials, such as polystyrene, silicon, silicon dioxide, and indium tin oxide, surface energy was calculated after they were treated with adhesion molecules, such as collagen, fibronectin, poly-L-ornithine, and poly-D-lysine. A fibronectin coating was found to produce a monopolar acidic surface while poly-D-lysine, poly-L-ornithine, and collagen coatings were found to produce monopolar basic surfaces (Harnett et al., 2007). Surface free energy is the measure of “unsatisfied bond energy” which results from bonds exposed at a material’s surface. This energy affects protein adsorption and cell attachment. When proteins, water, or cells approach a surface, their surface domains align in order to minimize the overall surface free energy of the interface. Quantifying these interactions is not simple; contact angle measurements can predict surface free energy-related interactions, but when surface topography is modified these measurements are no longer accurate (Gentleman and Gentleman, 2014). Another parameter important for cell attachment, proliferation, and differentiation is wettability. The hydrophobicity/hydrophilicity of surfaces influence the interaction between proteins and the surface of scaffolds, changing the distribution, mass, and conformation of adsorbed proteins (Sarkar, 2013). The chemistry of the surface is important for cell attachment and development. Integrins, and cytoskeletal and extracellular matrix (ECM) proteins, like collagen, fibronectin, and RGD peptides can attached covalently, be electrostatically adsorbed, or self-assembled on biomaterial surfaces. Surface charge has effects on cellular behaviors, such as orientation, adhesions, and proliferation and inflammatory responses. Hunt et al., showed that when poly(ether)urethanes were charged with negatively sulfonate groups an early phase acute inflammatory response appear dependent on the number of charges. Furthermore, negatively charged scaffolds promoted significantly higher vessel ingrowth than positively charged scaffolds. The surface charge density has been described to alter colony formation through osteoblast and the orientation of neuroblastoma cells (Verma et al., 2011).

2.1.6 PORE SIZE Pore size and porosity are vital in the development of scaffolds for tissue engineering. Pores must be interconnected to permit cell growth, migration, and proliferation. If pores are too small cell migration is limited, this can also limit the diffusion of nutrients and the removal of waste resulting in necrotic regions within the construct. Equally if pores are too large there is a decrease in surface area which limits cell adhesion. The structure of the ECM offers curs for cellular communication sustained by the interaction of integrins with ECM proteins. Scaffolds with a pore size between 20 and 1500 μm have been used in bone tissue engineering; significant bone growth was observed when pores were greater than

2.1 Introduction

300 μm while pores smaller than 300 μm are involved in osteochondral ossification. Collagen glycosaminoglycan scaffolds used for skin regeneration, have the ability to promote cell growth and tissue development if the pore size ranges between 20 and 120 μm (Murphy et al., 2010). Furthermore, in addition to pore size other parameters are extremely important in scaffold designing. Such parameters include: distribution, pore shape, pore volume, pore wall roughness, pore connectivity, and pore throat size. The role of pore size on tissue regeneration has been highlighted by experiments that demonstrate an optimum pore size of 5 μm for neovascularization, 20 μm for the ingrowth of hepatocytes, 200 350 μm for osteoconduction, 5 15 μm for fibroblast ingrowth, and between 20 and 125 μm for regeneration of adult mammalian skin (Murphy and O’Brien, 2010).

2.1.7 ADHERENCE AND BIOCOMPATIBILITY In the 1950s the main goal of the first generation of biomaterials was to be bioinert, then in the 1980s it was for the second generation of biomaterials to be bioactive, and since the 2000s the third generation of biomaterials aim to regenerate functional tissue. A biocompatible scaffold should be nontoxic, noncarcinogenic, nonimmunogenic, antithrombotic, compliant, vasoactive, and permit the regeneration of target tissue. Preparation and characterization of scaffolds, including morphology, in in vitro tests on eukaryotic cells and in vivo in the host body represent the most important parameters that play a key role in the establishment of the biomedical application of scaffolds. The first step in the evaluation of biocompatibility is to assess cellular morphology and viability. There are numerous methods that can reveal the state of cells in the presence of scaffolds. Simple methods for visualizing the cells in scaffolds include optic or fluorescent microscopy. Microscopy provides information regarding cellular shape, attachment to substrate, integrity of cells, and the number of vacuoles. Cellular viability can be assessed by fluorescent microscopy. Cells are stained with fluorescent dyes that enter the cells and are transformed by the action of cellular enzymes in the fluorescent dyes, suggesting that cells are metabolically active and are, therefore, viable. For the detection of apoptosis cells can be stained with trypan blue, DAPI, Hoechst, propidium iodide, or annexin V, and then analyzed by fluorescent microscopy or by flow cytometry. Another level of investigating cellular biocompatibility with bioprinted scaffolds is by evaluating cellular metabolism. The most frequently used assays are: MTT assay, LDH assay, BrdU assay, TUNEL assay, and cell cycle assays. MTT and LDH assays are based on measuring the activity of mitochondrial and cytosolic enzymes that transform the tetrazolium dye MTT (3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide) in formazan, or on measuring lactate dehydrogenase (LDH) level that is rapidly released out of cells when cell membranes are damage. Furthermore, a BrdU (5-bromo-2-deoxyuridine) assay permits the evaluation of cell proliferation. The evaluation of inflammatory markers, such as interleukin (IL)-1, IL-3, IL-6, and IL-12, can be assessed by an enzyme-linked immunosorbent assay, and lipid peroxidation by a TBARS

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assay for signs of oxidative stress. After in vivo animal tests, the ultimate challenge is the use of scaffolds in human clinical trials and obtaining the approval of regulatory bodies, such as the Food and Drug Agency (FDA), EMA (European Medicines Agency), and WHO (World Health Organization) (Keong and Halim, 2009; Wiegand and Hipler, 2009; Holban, 2015).

2.1.8 DEGRADATION RATES Biodegradable scaffolds are the most wanted scaffolds for applications such as drug delivery, implants, and tissue regeneration. The degradation of scaffolds can occur by physical, chemical, and/or biological processes. The degradation rate is important for adapting the number of cells and molecules in order to grow and develop tissues. Controllable degradation rates should match the rate of tissue growth in vitro and in vivo. The biodegradation rate of a polymeric scaffold depends mainly on the intrinsic properties of the polymer, including the chemical structure, the presence of hydrolytically unstable bonds, the level of hydrophilicity/hydrophobicity, crystalline/amorphous morphology, glass transition temperatures, the copolymer ratio, and the molecular weight. Nonbiodegradable scaffolds are also used for replacing parts of hard tissue (hip, knee, and tooth), such as poly-methyl methacrylate and polyethylene (Suri and Schmidt, 2009; Gordon et al., 2004).

2.2 MECHANICAL PROPERTIES Mechanical parameters that are most often evaluated include: (1) Young modulus (elastic modulus, tensile modulus); (2) flexural modulus (bending modulus); (3) tensile strength—maximum stress that the biomaterial can withstand before it breaks; (4) yield strength (yield stress)—the stress level at which a material begins to deform plastically; (5) fatigue—failure under cyclic stress. An important quality for the strength and rigidity of a polysaccharidic scaffold is that it has fibrous proteins which provide integrity and stability. Major factors affecting the mechanical properties and external geometry of scaffolds include: porosity, pore volume, size, orientation, and connectivity (Chang and Buehler, 2014; Manssor et al., 2016) (Table 2.1).

2.2.1 HYDROGEL-DERIVED SCAFFOLDS Hydrogels are a class of crosslinked polymeric substances capable of absorbing and retaining large quantities of water. There are two main types of hydrogels: natural-derived hydrogels (collagen, gelatin, fibrin, chitosan, alginate, agar, hyaluronic acid, maltodextrin) and synthetic-derived hydrogels (polyethylene glycol, poly(lactic-co-glycolic acid), poly-e-caprolactone) (Gasperini et al., 2014).

2.2 Mechanical Properties

Hydrogels offer a wide variety of applications from tissue engineering to drug delivery, contact lenses, and wound dressings (Hospodiuk et al., 2017). Naturalderived hydrogels are generally weak for 3D bioprinting, while synthetic hydrogels are stiffer but lack bioactive molecules. The level of hydration influences the biocompatibility. Hydrogels can absorb up to 1000 times their original weight without dissolving; being permeable to oxygen water-soluble nutrients are required for cell survival (Ahmed, 2013). Hydrogels permit migration of embedded cells in any direction in 3D in contract to polymeric scaffolds. The hydrogel viscosity for extrusion-based bioprinting ranges from 30 to 60 3 107 mPa/s, of course the hydrogel concentration changes the viscosity level. Ouyang et al. (2016) show that there is no significant difference in bioink viscosity when cell density is under 2 3 106 mL21. In inkjet-based bioprinting the hydrogels should have a low viscosity so that it can easily flow through the tubing system and nozzle without clogging. Bioink hydrogels should have rheological properties that permit the viscosity to be increased when a shear is applied. The viscosity level in inkjet-based bioprinting is between 3.5 and 12 mPa/s, which depends on the bioink concentration (Mandrycky et al., 2016). Table 2.1 Main Parameters Analyzed in Bioprinted Scaffolds Main Characteristics of Bioprinted Scaffolds External geometry Surface properties

Porosity and pore size Adherence Biodegradation rates Biocompatibility

Mechanical competence

Features

References

Micro and macrostructure connectivity Surface energy, chemistry, charge, topology, area

Chua and Yeong (2015), Harnett et al. (2007) Harnett et al. (2007), Gentleman and Gentleman (2014), Sarkar (2013), Verma et al. (2011) Murphy et al. (2010), Murphy and O’Brien (2010)

Pore distribution, pore shape, pore volume, pore roughness, pore network Integrins and extracellular and cytoskeletal proteins In cell media, body fluids, FBS, Hanks Buffer, enzyme degradation rate Nontoxic, nonimmunogenic, noncarcinogenic, antithrombotic, compliant, vasoactive Young modulus, tensile/ compressive properties, shear/ torsion properties, bending properties, viscoelastic properties, fatigue

Keong and Halim (2009), Holban (2015) Suri and Schmidt (2009), Gordon et al. (2004) Keong and Halim (2009), Wiegand and Hipler (2009), Holban (2015) Chang and Buehler (2014), Manssor et al. (2016)

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Laser-based bioprinting utilizes laser energy for the fabrication of tissue constructs and can be classified into two groups: •

•

processes based on cell transfer; • laser-guided direct writing, • matrix-assisted pulsed laser evaporation-direct write, • laser-induced forward transfer (LIFT), and processes involving photopolymerization; • stereolithography (STL), • dynamic optical projection stereolithography, • two-photon polymerization.

In laser-based bioprinting the viscosity of hydrogels ranges from 1 to 300 mPa/s. The hydrogel should possess gelation capability to ensure stability and high-mechanical strength for cells (Hospodiuk et al., 2017). Crosslinking methods can be grouped in: 1. Physical crosslinking: ionic, hydrophobic and hydrogen bonding interactions, stereo-complexation, self-assembly of amphiphilic peptides or polymers into micellar structures are some of the well-established mechanisms. This type of crosslinking has the advantage that it is without any exogenous agents, thus, eliminating the risk of chemical contamination or toxicity. In bioprinting, hydrogels that most commonly use physically crosslinking are agarose, alginate, chitosan, collagen, gelatin, Matrigel, and Pluronic (Wang et al., 2015). 2. Chemical crosslinking: covalent bonding between polymer chains confers better mechanical stability compared to physical-crosslinking. Chemical crosslinking involves crosslinking agents (glutaraldehyde, genipin), and the formation of reactive species by photo-irradiation or photo-crosslinking. The degree of crosslinking depends on the concentration of the crosslinker used. A high degree of crosslinking leads to hydrogel with strong mechanical properties which reduces the degradation time of the hydrogel. Growth factors, drugs, and other biological compounds are released more slowly in the crosslinked hydrogel due to stronger encapsulation between chemicallycrosslinked polymer chains. Depending on the concentration of glutaraldehyde, cytotoxicity might be induced, thus, some natural compounds are used, such as genipin. Genipin is used to crosslink gelatin, collagen, chitosan, and fibrin (Hennink and van Nostrum, 2012; Yan et al., 2010). 3. Photopolymerization: is initiated in the presence of a photoinitiator which forms excited molecular species upon irradiation with light, and low molecular weight monomers or oligomers undergo a process called curing to form a crosslinked polymeric network. Frequently pure hydrogels do not crosslink independently unless an external photoinitiator is added, thereby the physical properties of the hydrogel can be controlled by modulation of the rate and degree of crosslinking by photopolymerization. Some photoinitiators

2.2 Mechanical Properties

used in 3D bioprinting are: Irgacure, Biokey, VA-086, Eosin Y. Irgacure is the most widely used photoinitiator, but up to 0.5% (w/v) have toxic effects on cells unless they are leached out of the scaffold. Eosin Y and Biokey are two photoinitiators that crosslink the scaffold with visible light, however, not in UV light such as with Irgacure. Methacrylated gelatin (GelMA) and PEG are two commonly used polymers capable of undergoing photopolymerization (Fairbanks et al., 2009; Fedorovich et al., 2009). 4. Enzymatic crosslinking: applies to hydrogels composed of fibrin. These hydrogels are treated with thrombin, a serine protease that converts fibrinogen into fibrin. Another enzyme used is transglutaminase (for gelatin) which catalyzes the formation of isopeptide bonds. Both enzymes are Ca21dependent enzymes (Benedikt et al., 2000; Go´mez-Guille´n et al., 2011).

2.2.2 AGAROSE HYDROGEL Agarose is a polysaccharide extracted from various seaweeds made up of the repeating unit of agarobiose, a disaccharide composed of D-galactose and 3,6anhydro-L-galactopyranose. Agarose is a neutral copolymer and in aqueous solutions from thermos-reversible gels. The mechanical and thermal properties of agarose hydrogels depend on the polymer concentration, pH, and solvent type (Ferna´ndez et al., 2008). The gelation capacity makes agarose suitable for tissue engineering applications, however its degradation rate is slow and its adhesion properties are low (Zhang et al., 2012). To eliminate these disadvantages, agarose can be blended with other polymers, such as gelatin or chitosan. When agarose is blended with chitosan gels it retains transparency in the hydrated state, suggesting a uniform distribution of chitosan in the agarose matrix without noticeable phase separation. Zamora-Mora et al. (2014) showed that this type of hydrogel has a sponge-like structure and porosity similar to agarose. SEM microscopy showed that the size of agarose and chitosan/agarose gels pores was 73 6 16 and 102 6 39 μm, respectively. Furthermore, no variation in the elastic modulus was observed below the melting temperature of the gel. The elastic modulus of the composite gels increased with agarose concentration, reaching the value of 1 kPa at 20 C for chitosan/agarose (Zamora-Mora et al., 2014). Luo and Shoichet (2004) created an agarose hydrogel with a cysteine compound that containing a sulfydryl protecting group in which immobilized fibronectin peptide fragment, and glycine-arginine-glycine-aspartic acid-serine (GRGDS), promote migration and neurite outgrowth (Luo and Shoichet, 2004).

2.2.3 ALGINATE HYDROGEL Alginate (alginic acid, algin) is a polysaccharide distributed in the cell walls of brown algae and other microorganisms. Alginate is a linear copolymer with homopolymeric blocks of (1-4)-linked β-D-mannuronic acid (M) and its C-5

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epimer α-L-guluronic acid (G) residues, respectively, covalently linked together in different blocks. The copolymer can appear by association of homopolymeric blocks of consecutive G-blocks, consecutive M-blocks, or alternating M-G blocks. The viscosity of alginate decreases with increasing shear rate because in aqueous solution alginate has non-Newtonian characteristics and increase with concentration. Polymer chains can be crosslinked with multivalent cations, such as Ca21 or Ba21, leading to a gel network. The gel stiffness depends on the polymer molecular weight, the distribution of G and M blocks, and the stoichiometry of the chelating cation with alginate (Lee et al., 2000; Zhao et al., 2010). The disadvantages of alginate are: cells do not adhere to alginate and it is not degraded by the mammalian enzyme. The viscosity of gels made of various alginate fractions is extremely similar (1 2 Pa s), and they have a shear moduli variate between 10 kPa (high-MW fraction 0.25) and 24 kPa (high-MW fraction 1.0) (Augst et al., 2006). Alginate hydrogel is not an ideal biomaterial for 3D printing with an extrusion-based printer due to its low ability of viscosity recovery, but this can be improved by mixing it with other materials, such as graphene oxide, gelatin, and cellulose (Li et al., 2016).

2.2.4 CHITOSAN HYDROGEL Chitosan is obtained from insoluble chitin through deacetylation, and is found in the exoskeletons of insects, crustaceans, and fungi. Hydrogen interactions between acetamide groups and hydroxyl groups of chitin gives a rigid crystalline structure. When chitin is partially deacetylated and converted to chitosan, the amount of amino groups and its solubility is increased. There is a proportional increase in chitosan deacetylation and the enhancement of biocompatibility and biodegradability. The structure of chitosan is comprised of glucosamine and N-acetylglucosamine. The charge density of chitosan depends on the degree of deacetylation and the amino group density. The pH of the chitosan solution is linked to the quantity of ionized amino groups. Chitosan is a weak base with pKa 6.5 which can be dissolved in dilute acidic medium. The molecular weight of chitosan lies between 50 and 2000 kDa. In the human body, chitosan can be biodegraded by lysozyme, gastrointestinal enzymes, and colon bacteria. Chitosan hydrogels are used for wound healing. The probable mechanism of healing could be assumed to be in (1) the infiltration and secretion of inflammatory mediators by leukocytes, (2) the migration of macrophages and increases in the amount of collagen, which (3) activates the complement system and cytokines. Chitosan with a deacetylation degree near to 100 is shown to have a high rate of degradation, cell biocompatibility, and high affinity for cell adhesion. The porosity of chitosan-based hydrogels has a huge influence on the properties of hydrogel, such as swelling, cell adhesion, and cell proliferation rate. The most common methods used in forming porous hydrogels for tissue regeneration include (1) freeze drying, (2) gas foaming, and (3) salt leaching. Chitosan scaffolds can be used for

2.2 Mechanical Properties

the regeneration of various tissues, such as bone, cartilage, skin, and nerves (Ahmadi et al., 2015; Cao et al., 2014; Mirahmadi et al., 2013; Miguela et al., 2014; Gnavi et al., 2013).

2.2.5 CELLULOSE HYDROGEL Cellulose is a polysaccharide consisting of a linear chain of several hundred to many thousands of β(1-4) linked D-glucose units. Cellulose has many hydroxyl groups and can form hydrogen bonds easily, but it is difficult to dissolve in common solvents due to its highly extended hydrogen bonded structure. The preparation of cellulose hydrogels requires that cellulose is dissolved in aqueous systems; new solvents being used include N-methylmorpholine-N-oxide (NMMO), ionic liquids (1-butyl-3-methylimidazolium chloride (BMIMCl) and 1-allyl-3methylimidazolium chloride (AMIMCl)), and alkali/urea, thiourea. Pure cellulose hydrogel can be obtained from certain bacterial species (Chang and Zhang, 2011). Although resorption of cellulose in animal and human tissues does not occur; since animal cells are not able to synthesize cellulases, chemical modification, crosslinking, and bounding with bioresorbable moieties can yield resorbablecellulose scaffolds. The most widely used cellulose derivatives are: methylcellulose (MC), ethyl cellulose (EC), hydroxyethyl cellulose (HEC), hydroxypropyl methylcellulose (HPMC), and sodium carboxymethylcellulose (NaCMC). The gelation mechanism involves hydrophobic associations among the methoxy group of macromolecules. At low temperatures, polymer chains in solution are hydrated and simply tangled with one another. As the temperature increases, molecules lose water, until polymer-polymer hydrophobic associations take place, thus, forming a hydrogel network. The sol-gel transition temperature depends on the degree of substitution of cellulose ethers as well as on the addition of salts. The degree of substitution and the salt concentration can be properly adjusted to obtain specific formulations gelling at 37 C. Cellulose based-hydrogel scaffolds are good materials for tissue engineering applications due to their intrinsic properties, like nontoxicity, biocompatibility, tunable and porous microstructure, and good mechanical properties. Cellulose and its derivatives have been used for the treatment of skin burns and in the regeneration of cardiac, vascular, neural, cartilage, and bone tissues (Sannino et al., 2009). Cellulose scaffolds activated in a saturated Ca(OH)2 enable the adherence, proliferation, and vitality of chondrocytes, suggesting the development of cartilage in a way similar to cartilage repair (Mu¨ller et al., 2006). Entcheva et al. (2004) demonstrate that cellulose acetate and regenerated cellulose promote cardiac cell growth and enhance cell connectivity (gap junctions) and electrical functionality. Bacterial cellulose associated with graphene oxide nanoflakes reduced Young’s modulus (  50%) and clear definition of water hydrogel interfaces. Furthermore, investigation of 3D neuronal networks showed that this scaffold accelerated neurite outgrowth (  100 μm/ day), and it supports the formation of well-defined neuronal bilayer networks and synaptic connectivity (Kim et al., 2017).

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2.2.6 FIBRIN HYDROGEL Fibrin is a fibrous protein formed by the action of the protease, thrombin, on fibrinogen and is involved in the clotting of blood. Changes in conformational structure and the exposure of polymerization sites permit fibrin monomers to selfassemble into insoluble fibrin gel. Rheology studies showed that fibrin gel has nonlinear elasticity; fibrin network forming complex that allows a high degree of deformation without breakage. Fibrin gels are used as a bioadhesive in surgeries for hemostasis, for wound closure, and as a sealant. In comparison to synthetic polymeric materials, fibrin gel presents many advantages, such as a controllable degradation rate, nontoxic end-products, and excellent biocompatibility. Moreover, by controlling the precursor concentration and ionic strength, the morphology, mechanical properties, and stability of fibrin hydrogel could be tuned. Unlike collagen-based hydrogel that presents a fast degradation rate, fibrin gel is slowly degraded and permits cell seeding, uniform cell distribution, good adhesive property, and cellular interaction (Li et al., 2015). Christman et al. (2004) showed the feasibility of injecting a cell-scaffold mixture into the damaged hearts, after myocardium infraction, of Sprague-Dawley rats. Five weeks after implantation, a decrease in the thickness of the infarct wall was observed along with preserved cardiac functions based on histological and echocardiography results. Fibrin gel was demonstrated to increase the survival rate of transplanted cells, decrease the infarct size, and increase the blood flow to damaged tissue (Christman et al., 2004). Lee et al. (2012) reported that mesenchymal stem cells (MSCs) mixed into collagen/hyaluronic acid/fibrinogen composite gel and injected into a rabbit model with a knee defect was capable of regenerating and repairing the defect through osteochondral regeneration. Histological analysis showed the presence of glycosaminoglycans and type II collagen, a hyaline-like cartilage construct (Lee et al., 2012). Synthetic polymer channels filled with fibrin improve nerve repair and this is further enhanced if Schwann cells, which bind fibrin through αvβ8 integrins and are required for healthy axon development, are suspended in the fibrin (Tsai et al., 2006). Human umbilical vein endothelial cells (HUVECs) and fibroblast cocultured in fibrin enhance angiogenic behavior, HUVECs being a good model for studying the angiogenic process. In vivo fibrin could cause an immune reaction or be rapidly degraded and it shows weak mechanical properties that do not allow long-term cultures. Due to its weak mechanical properties, fibrin is not suitable for extrusion-based bioprinting, but in combination with thrombin could be printed using droplet-based bioprinting or inkjet bioprinting (Murphy et al., 2010).

2.2.7 GELATIN/COLLAGEN HYDROGEL Collagen is a triple helical protein obtained from natural sources, such as rat tail, bovine tendons, fish skin, and is most frequently used in tissue engineering. Collagen is a highly conserved protein in mammals, and represents about 25% of

2.2 Mechanical Properties

the entire protein mass in most animals. Collagen scaffolds facilitate cell attachment due to rich integrin-binding domains that interact with cell cytoskeleton. Collagen type I is less used in bioprinting because it remains in a liquid state at low temperatures or forms a fibrous structure at high temperatures and neutral pH. Complete gelation can take up to half an hour at 37 C. Because of its low mechanical properties and instability, it is better to blend collagen with other materials. Collagen alone can be bioprinted using extrusion-based bioprinting; better results were obtained in combination with Pluronic (Hospodiuk et al., 2017; Homenick et al., 2011). Collagen has a fibrous microarchitecture not suitable for inkjet bioprinting, but that would be better for microvalve bioprinting. Fibrin-collagen bioink embedded with amniotic fluid stem cells or MSCs, was bioprinted using a valve-based bioprinter as a scaffold that can be used for the treatment for skin burns (Skardal et al., 2012). Chen et al. (2012) examine the role of three collagen-based scaffolds (collagen, collagen-elastin, and collagenchondroitin-4-sulfate) on the microstructures, mechanical properties, and bioactivities in the presence of cardiosphere-derived cells. The pore sizes of all the scaffolds ranged between 100 and 200 μm, being sufficient for oxygen, nutrients, and cell attachment. The presence of elastin increased the pore sizes of the scaffolds because elastin does not form continuous sheets as collagen does, it only embedded in the collagen sheets in the form of short rods. The incorporation of chondroitin-4-sulfate decreased the pore sizes due to the collagen-chondroitin-4sulfate coprecipitation effect in the suspension, which reduced the viscosity. Low viscosity causes a higher freezing rate, producing smaller ice crystals and, therefore, smaller pores in the scaffolds. Elastin and GAGs have a much lower stiffness than collagen, therefore, the moduli values are lower compared to collagen (Chen et al., 2012). Gosline et al. (2002) reported a Young modulus of B1 MPa for elastin and B1 GPa for collagen. Scaffolds formed by 50% collagen and 50% condroitin-4-sulfate had a lower attachment of cells; an explication being that the binding sites of the collagen were saturated by chondroitin. Proliferation in the presence of elastin slowed after 4 days compared with collagen due to its nonintegrin pathway (Chen et al., 2012). Collagen glycosaminoglycan (Col-GAG) scaffolds have demonstrated great potential for skin and bone tissue engineering due to their ability to promote cell growth and tissue development. Pore size is important for regeneration to take place, thus, for skin regeneration and wound healing it was postulated that the range of pore size should be between 20 and 120 mm. Col-GAG scaffolds can be fabricated using a lyophilization process, and with a constant cooling rate technique it is possible to produce scaffolds with a homogenous pore structure (Murphy et al., 2010). Gelatin is a denatured form of collagen protein and is used in food, medicine, and pharmaceutical industries. Fish gelatins have a significantly lower gelling temperature, lower melting point, lower thermal stability, and higher viscosity. These gelatins have a lower content of peptide repetitions, such as prolinamide and hydroxyprolinamide, in their polypeptide chains. The hydrolytic processes of collagen I can be classified into three groups: physical, chemical, and enzymatic.

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Depending on the protocols used for collagen I pretreatment prior to the extraction process, two main types of gelatin can be obtained: Type A and Type B. Type A gelatin has an isoelectric point of 9.0, and is derived from acidic hydrolysis (sulfuric or hydrochloric acid). While type B gelatin has an isoelectric point of 5.0, and is derived from the alkaline hydrolysis of collagen I using alkaline liquid (NaOH). During hydrolytic processes, the triple helix structure of collagen I is partly broken into single-stranded polymeric molecules. The average molecular weight of gelatin is between 15,000 and 400,000 Da. The structures of gelatin are composed of glycine-X-Y peptide triplet repetitions, where X and Y can be any amino acid, but proline for X- and hydroxyproline for Y-positions are the most common. The average length and molecular weight of gelatins depend on the origins of the raw materials, the pretreatment methods, and the hydrolytic processing parameters (e.g., pH, temperature, and time). Strong noncovalent interactions, such as van der Waal forces, hydrogen bonds, and electrostatic and hydrophobic interactions are among the individual gelatin chains (Wang et al., 2017). At low temperatures, gelatin residues self-associate to form helical structures leading to a gel-like form, which reverts back to a random coil conformation as the temperature increases. Gelatin keeps the Arg Gly Asp (RGD) sequence from collagen, is less immunogenic, and promotes cell adhesion, proliferation, migration, and differentiation. Gelatin in gel form can easily liquefy at 37 C, the weight of gelatin gel decreases by 50% after around 10 hours of incubation and completely dissolves within 24 hours. To remove this inconvenience a variety of chemical crosslinking methods have been examined, however, these methods are cytotoxic. The most used methods for crosslinking gelatin are: enzymatic crosslinkers (transglutaminase, horseradish peroxidase), or chemical crosslinkers (hydrogen peroxide, glutaraldehyde) (Murphy et al., 2010). Using extrusion-based bioprinting, gelatin embedded with hepatocytes was bioprinted into spatially-defined 3D structures. The hepatocytes remained viable and led their biological functions for more than two months (Wang et al., 2016).

2.2.8 HYALURONIC ACID HYDROGEL Hyaluronic acid (hyaluronan, HA) is an anionic nonsulfated glycosaminoglycan, present in almost all connective tissues and a major component in the ECM. It is involved in cell proliferation, migration, and the progression of some malignant tumors. The repetitive disaccharide units are D-glucuronic acid and N-acetyl-Dglucosamine linked by alternating β-1,4 and β-1,3 glycosidic linkages. Hyaluronic acid is used in tissue engineering due to its ability to form flexible hydrogels. HA bioink contains chemical modifications of functional groups (glucuronic carboxylic acid group, the secondary hydroxyl group and the N-acetyl group) to enhance its rheological properties. HA has a slow gelation rate and poor mechanical properties. Hyaluronic acid has the advantages that it is cell friendly, with controllable mechanics, architecture, and degradation. Human stem cells encapsulated in HA

2.2 Mechanical Properties

maintained their phenotype, normal karyotype, and full differentiation capability. This is essential for cell migration, nerve regeneration, neuronal and glial development, and wound healing (Murphy et al., 2010; Faroni et al., 2015; Aya and Stern, 2014; Nettles et al., 2004). Hyaluronic acid is used in extrusion-based bioprinting but is generally blended with other biomaterials to enhance its bioprintability and solidification ability. Skardal et al. (2010) crosslinked hyaluronic hydrogels with tetrahedral polyethylene glycol (PEG) for bioprinting blood vessel-like constructs. To form extrudable hydrogels, tetra-PEG molecules were acrylated to form tetraacrylates and then crosslinked with thiolated gelatin and HA (Skardal et al., 2010). Bioink composed of human adipose-derived stem cells or endothelial forming cells and hyaluronic acid and fibrinogen was bioprinted using LIFT laser-based bioprinting. Alternating layers of HA-fibrinogen combined with both cell types permitted interaction between the two cell types (Gruene et al., 2011). Degradation of HA scaffolds in biological environments is made by hyaluronidase. Schante et al. (2012) improved the enzymatic stability of hyaluronic acid by attaching amino acids to HA residues. HA hydrogel can bind to cells through cell surface receptors, such as CD44 RHAMM (hyaluronan-mediated motility receptor) and ICAM-1 (intercellular adhesion molecule 1). HA scaffolds have been widely used since 2001 in burn care and posttraumatic and complicated surgical wounds. HA/agarose and HA/fibronectin scaffolds have shown great potential in wound healing applications (Collinsa and Birkinshaw, 2013). Through the engineering of heart valves using valvular interstitial cells and HA based hydrogels attempts are being made to create artificial valves for the replacement of calcified ones that occur in the case of diabetes and cardiovascular diseases (Burdick and Prestwich, 2011). An in vivo study showed that treatment with HA hydrogels significantly inhibited glial scarring, astrocytic activation, macrophage/microglia infiltration, and chondroitin sulfate proteoglycan deposition (Lin et al., 2009; Khaing et al., 2011). Chen et al. (2014) demonstrated that HA and platelet-rich plasma can restore the downregulation of cartilage gene expression induced by interleukin-1β and tumor necrosis factor-α, including SOX-9, collagen type II, and aggrecan.

2.2.9 MATRIGEL HYDROGEL Matrigel is a commercial product that contains an ECM protein mixture obtained from Engelbreth-Holm-Swarm mouse sarcoma cells. This matrix contains laminin as a major component, collagen type IV, heparin sulfate proteoglycan, entactin, and other minor components. It promotes vasculogenesis by stimulating endothelial cells outgrowth, proliferation, and cells networking formation (Iordache et al., 2017). Matrigel is an expensive material; but it has the ability to create 3D bioprinted constructs with strong mechanic properties in which cells survive better than with popular hydrogels, such as agarose and alginate (Melchels et al., 2012). Thermal gelation of Matrigel and collagen type I is similar; but gelation of

49

Table 2.2 The Most Used Bioprinting Materials to Create Scaffolds for Tissue Engineering Types of Scaffolds Hydrogel-derived scaffolds

Natural

Biomaterial

Main Applications

Agarose Alginate Chitosan Cellulose

Cartilage, bone tissue engineering, drug delivery Cartilage, bone tissue engineering, drug delivery Bone, cartilage, skin, and nerves Skin burns, regeneration of cardiac, vascular, neural, cartilage, and bone tissues Regeneration of cardiac, vascular, neural, and cartilage tissue Skin burns, regeneration of cardiac, vascular, neural, cartilage, and bone tissues Skin burns, regeneration of cardiac, vascular, neural, and cartilage tissues Regeneration of cardiac, vascular, neural, cartilage, and bone tissues Regeneration of cardiac, vascular, neural, cartilage, and bone tissues Regeneration of cardiac, vascular, neural, and cartilage tissues, and drug delivery Regeneration of cardiac, vascular, neural, and cartilage tissues, and drug delivery Including skin, cartilage, ligament bone, skeletal muscle, vascular, and neural tissue, and drug delivery Including skin, cartilage, ligament bone, skeletal muscle, vascular, and neural tissue, and drug delivery Skin, cartilage, ligament bone, vascular, and neural tissue, and drug delivery Skin, cartilage, ligament bone, vascular, and neural tissue, and drug delivery

Fibrin Gelatin/Collagen Hyaluronic acid Matrigel Synthetic

Methacrylated gelatin Pluronic F-127

Fibrous polymerderived scaffolds

Natural Synthetic

Porous polymerderived scaffolds

Natural Synthetic

PEG, PNIAAm, PAA, PMMA, PAam, PDMAEM Collagen, gelatin, elastin, silk fibroin, chitosan, hydroxyapatite (HAP) HAP, TCP, PLAPCL, PLGA, PEVA, PLLA-CL, PU Collagen, gelatin, elastin, silk fibroin, chitosan, HA PLLA, PGA, PLGA, PCL, PDLLA, PEE, PEO, PBT, HAP

HAP, hydroxyapatite; PAA, poly(acrylic acid); PAam, polyacrylamide; PBT, polybutylene terephthalate; PCL, poly(ε-caprolactone); PDLLA, poly(D,L-lactide); PDMAEM, poly(dimethylaminoethylmethacrylate) hydrochloride; PEE, poly(ether ester); PEG, polyethylene glycol; PEVA, poly(ethylene-co-vinylacetate); PEO, poly (ethylene oxide); PGA, polyglycolide; PLA, polylactide; PLGA, poly(L-lactide-co-glycolide); PLLA, poly(L-lactic acid); PLLA-CL, poly(L-lactide-co-ε-caprolactone); PMMA, polymethylmethacrylate; PNIAAm, poly(N-isopropylacrylamide); PU, polyurethane; TCP, tricalcium phosphate.

2.2 Mechanical Properties

Matrigel reversible. Crosslinking occurs between 24 C and 37 C, and the gelation process takes about half an hour beginning at temperatures above 4 C. In extrusion-based bioprinting, Matrigel needs to be bioprinted at a low temperature before becoming fully crosslinked. To accelerate crosslinking, the construct should be printed on a heated plate. For extrusion-based bioprinting Matrigel was used together with osteoblast, endothelial-progenitor cells, and liver cells (Murphy et al., 2010). In a similar study, 3D cellular constructs were bioprinted using human osteosarcoma cells and Matrigel substrates using a LIFT technique (Barron et al., 2004) (Table 2.2)

2.2.10 SYNTHETIC HYDROGELS Synthetic polymers should mimic the native environment similar to ECM, this being possible through chemical modifications, such as crosslinking functional groups and adding domains capable of enhancing the structural and mechanical properties of bioprinted constructs. Gelatin methacrylate (GelMA) is an inexpensive, photocrosslinkable hydrogel tunable for different tissue engineering applications through modifying the polymer concentration, methacrylation degree, or UV light intensity. GelMA has relatively high mechanical strength, low swelling ratio, and permits blending with other hydrogels to increase cell survival. At 5% GelMA, cells elongated, migrated, and formed networks with surrounding cells. At higher concentrations, 10% or 15%, cells do not spread; cells bared limited elongation and few multicellular networks were formed. GelMA has a low viscosity at room temperature and is easy to extrude, crosslinking rate can be manipulated by UV exposure time (Nichol et al., 2010). By extrusion-based bioprinting GelMA together with chondrocyte liver cells and MSCs have been successfully bio printed (Murphy et al., 2010). Kuo and Demirci (2014), by droplets-based bioprinting GelMA embed with human MSCs and bone morphogenetic protein-2 (BMP-2) and transforming growth factor (TGF-β1) it was create a scaffold that mimics the cartilage. This bioink permitted the cells to differentiate toward osteogenic and chondrogenic lineages in a spatial manner, and mimic a native fibro-cartilage microenvironment (Kuo and Demirci, 2014). Poloxamer 407 (trade name Pluronic F127) is a hydrophilic nonionic surfactant from the class of copolymers known as poloxamers. Poloxamer 407 is copolymer consisting of a central hydrophobic block (56 units) of polypropylene glycol flanked by two hydrophilic blocks of polyethylene glycol (100 units). There are more than 11 types of Pluronic polymers which differ in molar mass, functionality, composites, and temperature of crosslinking which range between 10 C and 40 C. Pluronic copolymer quickly degrades, in less than a few hours; thus, to remove this inconvenience it is used in combination with other hydrogels, such as methacrylated hyaluronic acid (Mu¨ller et al., 2015). Pluronic F127 embedded with bone marrow-derived MSCs form a bioink that can be printed by

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CHAPTER 2 Bioprinted scaffolds

extrusion-based bioprinting. The construct permits cells to interact with each other and is capable of adipocyte differentiation under induction stimuli. The bioprintability of Pluronic using extrusion-based bioprinting is achieved at 20 C as it becomes more viscous. In the case of prolonged bioprinting a cooling system might be needed in order to maintain a lower temperature in the bioink reservoir. Pluronic in solid form (at room temperature or higher) can be mixed with other types of hydrogel in order to liquefy it. Laser-based bioprinting has not been attempted because Pluronic is not viscoelastic, and cannot transfer thermal energy to kinetic energy which is essential for jet formation (Wu et al., 2011).

2.3 FIBROUS POLYMER-DERIVED SCAFFOLDS The fabrication of fibrous scaffolds that mimic the architecture of natural human tissue can be done using nanofibers. Common techniques for the synthesis of nanofibers are: electrospinning, phase separation, and self-assembly. Nanofibers favor cell adhesion, migration, proliferation, and differentiation (Ma et al., 2005). Fibrous polymeric scaffolds are mainly used for tissue engineering, including skin, cartilage, ligament bone, skeletal muscle, vascular, and neural tissue, and also for drug delivery. Natural polymers (such as collagen/gelatin, chitosan, HA, silk fibroin) and synthetic polymers (such as PLA, PU, PCL, PLGA, PEVA, PLLA-CL) are used for the formation of fibrous scaffolds with applications in the biomedical field. Usually fibrous scaffolds made with nanofibers do not possess functional groups and will be specifically functionalized depending on the application (Singh et al., 2010). Yao et al. (2016) tested different combinations of PCL/Gelatin composite fibrous scaffolds (4:1, 2:1, 1:1, 1:2, 1:4) for mechanical properties and cellular responses. The results showed that the 2:1 PCL/ Gelatin fibrous scaffolds had the highest tensile strength, 3.7 MPa, and the highest elongation rate, about 90%. Also 2:1 PCL/Gelatin scaffolds showed responses from MSCs in terms of attachment, spreading, and cytoskeleton organization. This result suggests that mechanical property plays an important role in regulating cellular functions (Yao et al., 2016). The morphology of fibrous scaffolds can be controlled by modified parameters, such as viscosity, conductivity, polymer molecular weight, flow rate, applied potential of bioprinter, and ambient conditions (e.g., temperature, humidity). Yang et al. (2005) showed that neural stem cells elongated and their neurites outgrew along the direction of the PLLA fiber orientation. PCL nanofibrous scaffolds seeded with human bone marrowderived MSCs were tested for their ability to support chondrogenesis. The results showed that scaffolds in the presence of TGF-β1 induced differentiation of MSCs to chondrocytes (Chen et al., 2013). Fibrous scaffolds used in tissue engineering are fabricated, in particular, by electrospinning techniques and less so by 3D bioprinting.

2.4 Porous Polymer-Derived Scaffolds

2.4 POROUS POLYMER-DERIVED SCAFFOLDS Porous scaffolds, such as sponge, foam, and mesh scaffolds have high porosity, with pores that form an interconnected network. Porous scaffolds are often used in tissue engineering and regenerative medicine especially in bone and vascular regeneration. The architecture of the scaffolds mimic ECM allowing the cells to establish cell-to-cell and cell-to-polymer interaction. Mesh scaffolds present large open spaces throughout the scaffold so their use is limited to skin regeneration or wound healing. The advantages of foam polymeric scaffolds are: (1) they offer a physical surface for the attachment of cells, (2) they allow for a good circulation of nutrients, and (3) the cluster size is limited according to the pore size of the foam scaffold. The orientation of pore architecture can be controlled depending on the solvent and the phase separating condition (Ma and Zhang, 2001; Freiberg and Zhu, 2004). Designing certain and precise three-dimensional shapes involves refined extrusion technologies. There is not an ideal pore size; the dimensions vary according to cell type (Wei and Ma, 2004). Porous materials for scaffolds include natural polymers, such as collagen/gelatin, chitosan, HA, silk fibroin, cellulose, agarose, alginate, and synthetic biodegradable polymers, such as PLLA, PGA, HAP, PLGA, PBT, PCL, PDLLA, PEE, PEO. Tripathi and Bas (2012) created hydroxyapatite (HAP) scaffolds with pores of 100 300 μm. The compression strength of the 60% HAP scaffold was 1.3 MPa. The cellular response of human osteoblast like SaOS2 was positive, cells were able to adhere, proliferate, and migrate into the pores of the scaffold. Furthermore, cell viability increased on porous scaffolds compared to dense HAPs, and the expression of alkaline phosphate was enhanced as compared to nonporous HAP discs (Tripathi and Bas, 2012). An et al. (2012) investigated porous zirconia/hydroxyapatite ZrO(2)/HAP scaffolds for use in bone reconstruction as an alternative to porous hydroxyapatite which is too fragile for large bone defects. Their results showed that by adjusting the porosity of the zirconia/hydroxyapatite scaffold from 72% to 91%, when the ZrO(2) content increased from 50 to 100 wt%, the compressive strength of the scaffold increased from 2.5 to 13.8 MPa. Cell adhesion and proliferation in the ZrO(2)/HAP scaffold was greatly improved compared to the scaffold made only with ZrO(2). In vivo study showed bone marrow derived stem cell survival and enhanced bone regeneration around the implanted scaffold (An et al., 2012). Microporous tricalcium phosphate (β-TCP) scaffolds can be used as a BMP-2 delivery system in bone lesions. The scaffolds were loaded with 30 and 15 μg of BMP-2 and implanted into the back muscles and femoral defects (condyle and diaphysis) of rabbits for 4 weeks. The results from histological assays showed that BMP-2 was released from the scaffolds and new bone tissue was formed. Porous scaffolds made by β-TCP are frequently used in bone engineering because of their chemical similarity to the inorganic phase of natural bone, as well as their favorable biocompatibility, osteoconductivity, and bioresorbable properties (Sohier et al., 2009). Due to the difficulty of fabricating bioceramic scaffolds

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with smaller pore sizes using the current 3D printing techniques, the effect of smaller pore sizes (below 400 μm) of 3D printed bioceramic scaffolds on bone regeneration and biomechanical behavior has never been studied. Diao et al. (2018) bioprinted β-TCP scaffolds with interconnected smaller pores of three different sizes (100, 250, and 400 μm). Their results showed that scaffolds with the 100 μm pore size were found to present the highest maximum load and stiffness after being implanted for 12 weeks in rats. Tomography and histological analysis indicate that scaffolds with the 100 μm pore size achieved the highest percentage of new bone ingrowth, which correlates to biomechanical properties (Diao et al., 2018). Young modulus should be in the range of 10 1500 MPa for hard tissue formation, respectively 0.4 350 MPa for soft tissue. Using traditional techniques, such as porogen leaching or gas foaming, a maximum compressive modulus of 0.4 MPa can be obtained, below the needed values. 3D bioprinting of D,L--polylactic-polyglycolic acid (PLGA)/L-polylactic acid (L-PLA) in one phase and a L-PLGA/tricalcium phosphate mixture in the second phase form porous scaffolds. Elastic modulus and yield strength were 450 and 13.7 MPa, respectively. For soft-tissue applications poly(ethylene glycol)-terephthalate (PEG/PBT) fibers were bioprinted resulting in a scaffold with orthogonal pore structures ranging between 185 and 1683 μm. Static and dynamic moduli were 0.05 2.5 and 0.16 4.33 MPa, respectively, compared with native cartilage values (0.27 MPa static, 4.10 MPa dynamic) (Hollister, 2005).

2.5 CONCLUSION AND PERSPECTIVES Although there is a wide variety of bioink materials, including hydrogels, decellularized ECM, fibrous polymers, microcarriers, and cell aggregates, to achieve constructs that are similar to the target tissue, several aspects need to be taken into account. First, an important aspect is compatibility with bioprinting modalities. Extrusion-based bioprinting is the most flexible method among the existing bioprinting modalities, due to the mechanism and the larger nozzle diameters. Droplets and laser-based bioprinting permit bioprinting of hydrogels only. The second aspect is the bioprintability of the scaffolds. In this case, the bioprintability of hydrogels is superior to that of other bioink. The third aspect is represented by biomimicry, which is important for good bioprinting. The degradation of the scaffold, and the interaction and proliferation of cells are important for the formation of tissue. The fourth aspect is about the resolution of bioprinting which depends on the bioprinting modality as well as the bioink. Laser-based bioprinting of hydrogels have a resolution of 5.6 6 2.5 μm, while droplets and extrusionbased bioprinting between 50 and 100 μm. The fifth aspect is represented by affordability. Matrigel, fibrin, and collagen hydrogels are expensive comparative to synthetic polymers. In a scaffold hundreds of millions of cells are needed, so obtaining them could be labor-intensive, costly, and time-consuming. Other

References

aspects include scalability, practicality, mechanical and structural integrity, degradability, commercial availability, immunogenicity, and applicability. Despite the multitude of biomaterials that appear daily, relatively little research has been devoted to the development of biomaterials for bioprinting processes. Although a large number of hydrogels present great potential for tissue engineering, just a limited number can be used in bioprinting due to their lack of bioprintability, the toxicity of degradation products, and bioprinter parameters which should be optimized for each specific tissue. Further improvements are essential in bioink technology regarding the synthesis of new materials for bioprinting. In the near future a new field of study will emerge in the development and design of novel materials for bioprinting. The main goals for improving bioprinting are to minimize cell loss, promote cell-cell interactions, and to increase the mechanical properties and biocompatibility of bioink for supporting 3D bioprinted constructs. In the future, new bioink materials compatible with 3D bioprinting processes will be an important technology for tissue engineering and regenerative medicine.

ACKNOWLEDGMENT This chapter was possible thanks to the project “Tissue engineering of blood vessels using three-dimensional bioprinting of endothelial and smooth muscle progenitor cells” Grant PN-III-P1-1.1-PD-2016-1660, Nr. 19/2018.

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Lee, K.Y., Rowley, J.A., Eiselt, P., Moy, E.M., Bouhadir, K.H., Mooney, D.J., 2000. Controlling mechanical and swelling properties of alginate hydrogels independently by cross-linker type and cross-linking density. Macromolecules 33 (11), 4291 4294. Li, H., Liu, S., Li, L., 2016. Rheological study on 3D printability of alginate hydrogel and effect of graphene oxide. Int. J. Bioprint. 2 (2), 54 66. Li, Y., Meng, H., Liu, Y., Lee, B.P., 2015. Fibrin gel as an injectable biodegradable scaffold and cell carrier for tissue engineering. Sci. World J. 685690, 1 10. Lin, C.M., Lin, J.W., Chen, Y.C., Shen, H.H., Wei, L., Yeh, Y.S., et al., 2009. Hyaluronic acid inhibits the glial scar formation after brain damage with tissue loss in rats. Surg. Neurol. Surg. Neurol. 72 (Suppl. 2), S50 S54. Luo, Y., Shoichet, M.S., 2004. A photolabile hydrogel for guided three-dimensional cell growth and migration. Nat. Mater. 3 (4), 249 253. Ma, P.X., Zhang, R., 2001. Microtubular architecture of biodegradable polymer scaffolds. J. Biomed. Mater. Res. 56 (4), 469 477. Ma, Z., Kotaki, M., Inai, R., Ramakrishna, S., 2005. Potential of nanofiber matrix as tissue-engineering scaffolds. Tissue Eng. 11 (1 2), 101 109. Mandrycky, C., Wang, Z., Kim, K., Kim, D.H., 2016. 3D bioprinting for engineering complex tissues. Biotechnol. Adv. 34, 422 434. Manssor, N.A., Radzi, Z., Yahya, N.A., Mohamad Yusof, L., Hariri, F., Khairuddin, N.H., et al., 2016. Characteristics and Young’s modulus of collagen fibrils from expanded skin using anisotropic controlled rate self-inflating tissue expander. Skin Pharmacol. Physiol. 29, 55 62. Melchels, F.P.W., Domingos, M.N., Klein, T.J., Malda, J., Bartolo, P.J., Hutmacher, D.W., 2012. Additive manufacturing of tissues and organs. Prog. Polym. Sci. 37, 1079 1104. Miguela, S., Ribeiroa, M., Brancala, H., Coutinho, P., Correia, I., 2014. Thermoresponsive chitosan agarose hydrogel for skin regeneration. Carbohydr. Polym. 111, 366 373. Mirahmadi, F., Tafazzoli-Shadpoura, M., Shokrgozarb, M., Bonakdar, S., 2013. Enhanced mechanical properties of thermosensitive chitosan hydrogel by silk fibers for cartilage tissue engineering. Mater. Sci. Eng. C Mater. Biol. Appl. 33, 4786 4794. Mu¨ller, F.A., Mu¨ller, L., Hofmann, I., Greil, P., Wenzel, M.M., Staudenmaier, R., 2006. Cellulose-based scaffold materials for cartilage tissue engineering. Biomaterials 27 (21), 3955 3963. Mu¨ller, M., Becher, J., Schnabelrauch, M., Zenobi-Wong, M., 2015. Nanostructure pluronic hydrogels as bioinks for 3D bioprinting. Biofabrication 7, 35006. Murphy, C.M., O’Brien, F.J., 2010. Understanding the effect of mean pore size on cell activity in collagen-glycosaminoglycan scaffolds. Cell Adh. Migr. 4 (3), 377 381. Murphy, S.V., Atala, A., 2014. 3D bioprinting of tissues and organs. Nat. Biotechnol. 32 (8), 773 785. Murphy, C.M., Duffy, G.P., Schindeler, A., O’brien, F.J., 2010. The effect of mean pore size on cell attachment, proliferation and migration in collagen glycosaminoglycan scaffolds for bone tissue engineering. Biomaterials 31, 461 466. Nettles, D.L., Vail, T.P., Morgan, M.T., Grinstaff, M.W., Setton, L.A., 2004. Photocrosslinkable hyaluronan as a scaffold for articular cartilage repair. Ann. Biomed. Eng. 32, 391 397. Nichol, J.W., Koshy, S.T., Bae, H., Hwang, C.M., Yamanlar, S., Khademhosseini, A., 2010. Cell-laden microengineered gelatin methacrylate hydrogels. Biomaterials 31, 5536 5544. Ouyang, L., Yao, R., Zhao, Y., Sun, W., 2016. Effect of bioink properties on printability and cell viability for 3D bioplotting of embryonic stem cells. Biofabrication 8, 35020.

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Sannino, A., Demitri, C., Madaghiele, M., 2009. Biodegradable cellulose-based hydrogels: design and applications. Materials 2, 353 373. Sarkar, S., 2013. Utilizing hydrophobic/hydrophilic interactions in biomaterials: old dog, new tricks!. Biomaterials Forum 1 15. Schante, C.E., Zuber, G., Herlin, C., Vandamme, T.F., 2012. Improvement of hyaluronic acid enzymatic stability by the grafting of amino-acids. Carbohydr. Polym. 87 (3), 2211 2216. Singh, M., Sandhu, B., Scurto, A., Berkland, C., Detamore, M.S., 2010. Microsphere-based scaffolds for cartilage tissue engineering: using subcritical CO2 as a sintering agent. Acta Biomater. 6 (1), 137 143. Skardal, A., Zhang, J., Prestwich, G.D., 2010. Bioprinting vessel-like constructs using hyaluronan hydrogels crosslinked with tetrahedral polyethylene glycol tetracrylates. Biomaterials 31, 6173 6181. Skardal, A., Mack, D., Kapetanovic, E., Atala, A., Jackson, J., Yoo, J., et al., 2012. Bioprinted amniotic fluid-derived stem cells accelerate healing of large skin wounds. Stem Cells Transl. Med. 792 802. Sohier, J., Daculsi, G., Sourice, S., de Groot, K., Layrolle, P., 2009. Porous beta tricalcium phosphate scaffolds used as a BMP-2 delivery system for bone tissue engineering. J. Biomed. Mater. Res. A 92 (3), 1105 1114. Suri, S., Schmidt, C.E., 2009. Photopatterned collagen hyaluronic acid interpenetrating polymer network hydrogels. Acta Biomater. 5 (7), 2385 2397. Thomas, J.D., 2016. Could 3D bioprinted tissues offer future hope for microtia treatment. Int. J. Surg. 32, 43 44. Tripathi, G., Bas, B., 2012. A porous hydroxyapatite scaffold for bone tissue engineering: physico-mechanical and biological evaluations. Ceram. Int. 38 (1), 341 349. Tsai, E.C., Dalton, P.D., Shoichet, M.S., Tator, C.H., 2006. Matrix inclusion within synthetic hydrogel guidance channels improves specific supraspinal and local axonal regeneration after complete spinal cord transection. Biomaterials 27, 519 533. Verma, D., Desai, M.S., Kulkarni, N., Langrana, N., 2011. Characterization of surface charge and mechanical properties of chitosan/alginate based biomaterials. Mater. Sci. Eng. C 31 (8), 1741 1747. Wang, S., Lee, J.M., Yeong, W.Y., 2015. Smart hydrogels for 3D bioprinting. Int. J. Bioprinting 1, 3 14. Wang, X., Ao, Q., Tian, X., Fan, J., Tong, H., Hou, W., et al., 2017. Gelatin-based hydrogels for organ 3D bioprinting. Polymers 9 (401), 1 24. Wang, Z., Jin, X., Dai, R., Holzman, J.F., Kim, K., 2016. An ultrafast hydrogel photocrosslinking method for direct laser bioprinting. RSC Adv. 6, 21099 21104. Wei, G., Ma, P.X., 2004. Structure and properties of nano- hydroxyapatite/polymer composite scaffolds for bone tissue engineering. Biomaterials 25 (19), 4749 4757. Wiegand, C., Hipler, U.C., 2009. Evaluation of biocompatibility and cytotoxicity using keratinocyte and fibroblast cultures. Skin Pharmacol. Physiol. 22, 74 82. Wu, W., DeConinck, A., Lewis, J., 2011. Omnidirectional printing of 3D microvascular networks. Adv. Mater. 23, H178 H183. Yang, F., Murugan, R., Wang, S., Ramakrishna, S., et al., 2010. Electrospinning of nano/ micro scale poly(L[HYPHEN]lactic acid) aligned fibers and their potential in neural tissue engineering. Biomaterials 15, 2603 2610.

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Yan, L.P., Wang, Y.J., Ren, L., Wu, G., Caridade, S.G., Fan, J.B., et al., 2010. Genipincross-linked collagen/chitosanbiomimetic scaffolds for articular cartilage tissue engineering applications. J. Biomed. Mater. Res. A 95 (A), 465 475. Yao, R., He, J., Meng, G., Jiang, B., Wu, F., 2016. Electrospun PCL/Gelatin composite fibrous scaffolds: mechanical properties and cellular responses. J. Biomater. Sci. 27, 9. Zamora-Mora, V., Velasco, D., Herna´ndez, R., Mijangos, C., Kumacheva, E., 2014. Chitosan/agarose hydrogels: cooperative properties and microfluidic preparation. Carbohydr. Polym. 13 (111), 348 355. Zhang, L.M., Wu, C.X., Huang, J.Y., Peng, X.H., Chen, P., Tang, S.Q., 2012. Synthesis and characterization of a degradable composite agarose/HA hydrogel. Carbohydr. Polym. 88 (4), 1445 1452. Zhao, X., Huebsch, N., Mooney, D.J., Suo, Z., 2010. Stress-relaxation behavior in gels with ionic and covalent crosslinks. J. Appl. Phys. 107 (6), 063509.

CHAPTER

Fundamentals of chitosanbased hydrogels: elaboration and characterization techniques

3

Rejane Andrade Batista1, Caio Gomide Otoni2 and Paula J.P. Espitia3 1

Instituto Tecnolo´gico e de Pesquisas do Estado de Sergipe, Rua Campo do Brito, Aracaju, Brazil 2National Nanotechnology Laboratory for Agribusiness, Embrapa Instrumentac¸a˜o, Sa˜o Carlos, Brazil 3Nutrition and Dietetics School, Universidad del Atla´ntico, Atla´ntico, Colombia

3.1 INTRODUCTION Hydrogels can be defined as systems comprising of three-dimensional, physically or chemically bonded polymer networks entrapping water in intermolecular space (Ahmed, 2015). In other words, hydrogels may be referred to as hydrophilic gels or colloidal gels in which the dispersion medium is water (Ahmed et al., 2013). Depending on several factors (e.g., polymer nature, crosslinking nature, and density), hydrogels may be swollen with different amounts of water. The water sorption capacity of hydrogels results from their hydrophilicity, which in turn is provided mainly by capillary, osmotic, and hydration forces (Buwalda et al., 2014). Among the particularities of hydrogels two properties deserve special emphasis, namely: their remarkable ability of absorbing high amounts of liquids—more than 100% in relation to their dry weights—in a rapid fashion as well as their capacity to retain specific compounds without changing their structures in the swollen state when exposed to certain pressures (Bao et al., 2011; Feng et al., 2014; Mahdavinia et al., 2004). Fig. 3.1 illustrates a cellulose-based hydrogel in its original, swollen, and dried stages (Chang et al., 2010). Because of these peculiarities, hydrogels have become essential to numerous businesses as their application potential has been broadened for several areas, including agriculture, waste water treatment, water purification, tissue engineering, sensors, contact lenses, and drug release, among many others (Chang and Zhang, 2011; Feng et al., 2014). Thanks to their applicability potential, absorbent polymers and, particularly, hydrogels have been arousing growing interest from the scientific community. This reflects in a remarkable increase in the number of research and review articles since the 1980’s, as demonstrated in Fig. 3.2. Several classification criteria have been proposed in order to group hydrogels, as illustrated in Table 3.1 (Chang and Zhang, 2011; Zheng and Wang, 2015). Materials for Biomedical Engineering: Hydrogels and Polymer-based Scaffolds. DOI: https://doi.org/10.1016/B978-0-12-816901-8.00003-1 © 2019 Elsevier Inc. All rights reserved.

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FIGURE 3.1 Original (A), swollen (B), and dried (C) hydrogels comprised of cellulose powder and carboxymethylcellulose and crosslinked with epichlorohydrin. Adapted from Chang, C., Duan, B., Cai, J., Zhang, L., 2010. Superabsorbent hydrogels based on cellulose for smart swelling and controllable delivery. Eur. Polym. J. 46, 92
100 with the permission of Elsevier.

FIGURE 3.2 Number of scientific documents (research and review articles) published annually on superabsorbent polymers (black bars; topic: superabsorbent and polymer) and hydrogels (gray bars; topic: superabsorbent and hydrogel) retrieved from Web of Science Core Collection (as of January, 2017).

Hydrogels may be classified according to different properties or characteristics. The most used classifications rely on binding nature, that is, chemical or physical bonds. Physically bonded hydrogels are comprised of polymer chains ordered by intermolecular interactions (e.g., ionic or hydrogen bonds), whereas their chemically

3.1 Introduction

Table 3.1 Classification Criteria of Hydrogels and Main Features Considered Classification Criteria

Main Features Considered

Source

• • • • • • • • • • • • • • • • • • • • • • • • • • • •

Composition (monomers)

Crosslinking process Degradability Physical appearance

Electrical charge

Properties Chemically responsive

Biochemically responsive

Physically responsive

Natural Synthetic Hybrid Homopolymeric nature Copolymeric nature Multicomponent/interpenetrating network Physical Chemical Biodegradable Nonbiodegradable Matrix Film Microsphere (beads) Neutral/nonionic nature Ionic (anionic/cationic nature) Amphiphilic Conventional Smart pH Glucose Oxidant Antigens Enzymes Ligands Temperature Pressure Light Electric/magnetic field

Based on Ahmed, E.M., 2015. Hydrogel: preparation, characterization, and applications: a review. J. Adv. Res. 6, 105
121 and Ullah, F., Othman, M.B.H., Javed, F., Ahmad, Z., Akil, H.M., 2015. Classification, processing and application of hydrogels: a review. Mater. Sci. Eng. C 57, 414
433.

bound counterparts present chains connected by covalent bonds (Chang and Zhang, 2011). A combination of both is also possible (Buwalda et al., 2014). Another widely spread classification criterion takes into account the source of raw materials: natural, synthetic, or hybrid. The list of synthetic polymers used for hydrogel production is extensive and has been reviewed elsewhere (Kabiri et al., 2011; Ullah et al., 2015). Examples include polyacrylamide, poly(sodium acrylate), poly(acrylic acid), and polyvinylpyrrolidone, to mention a few.

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Hydrogels based on naturally occurring polymers, in turn, may be further sorted into two groups: polysaccharides or polypeptides (Zheng and Wang, 2015). The latter includes but is not limited to collagen, gelatin, soy, and fish proteins (Zohuriaan-Mehr et al., 2009), whereas the former encompasses starch (Chen et al., 2015), cellulose, and its derivatives (Bao et al., 2011; Chang and Zhang, 2011; Ma et al., 2015); sodium alginate (Wang and Wang, 2010), guar gum (Thombare et al., 2016), xanthan gum (Feng et al., 2014), and chitosan, and its derivatives, among others (Guilherme et al., 2015; Wang et al., 2013; Yu et al., 2010). Hydrogels made up of chitosan matrix, in particular, have been arousing rising interest because of the unique physicochemical, biological, and mechanical properties of this polymer. Concerning the biomedical industry, these materials present many advantages, such as biocompatibility, low toxicity, high bioactivity, multifunctionality, and biodegradability (Calo´ and Khutoryanskiy, 2015). These characteristics allow them to play their role in direct contact with the human body without any health damages. The aforementioned materials are capable of mimicking natural tissues, in addition to being adaptable and easy to handle as well as having lower cost than the polymers that are conventionally used for this purpose (Chang and Zhang, 2011; Feng et al., 2014). Chitosan-based hydrogels may be presented as fibers, films, gels, membranes, micro- or nanosized particles, and sponges without any functionality impairment (Bansal et al., 2011). When it comes to biomedical engineering, the feasible applications of hydrogels include, but are not limited to, contact lenses, drug release, artificial muscles, bone filling, and dermatology (Calo´ and Khutoryanskiy, 2015).

3.2 CHITOSAN NATURE AND MAIN PROPERTIES Chitosan is a carbohydrate polymer obtained from chitin through a deacetylation process. Chitin ranks second—preceded by cellulose only—amongst the most abundant biopolymers. It is of natural occurrence in marine crustaceans, mollusks, and insects—mainly as part of their exoskeleton—as well as in fungi, in which it plays a structural role (Alves and Mano, 2008). Large-scale chitin extraction from fungal mycelia is preferred over animal sources, chiefly due to the year-long (i.e., nonseasonal) availability of raw material and the completely controlled chitin production, which leads to a final product that features the standard physicochemical characteristics. Chitin extraction from fungi is also advantageous from a safety point of view due to the low allergenic risk, which otherwise would possibly be increased by crustacean components that might remain after the extraction and purification processes from marine sources (Croisier and Je´roˆme, 2013). In its native form, chitin is not suitable for several applications, mainly because of its chemical inertness and poor solubility (Croisier and Je´roˆme, 2013).

3.2 Chitosan Nature and Main Properties

FIGURE 3.3 Representation of chitosan extraction from chitin.

Thus, chitin must be subjected to a deacetylation process in a reaction with a concentrated alkali solution, which results in chitosan (Fig. 3.3). Chitin deacetylation is often carried out in the presence of sodium hydroxide or potassium hydroxide in addition to anhydrous hydrazine and hydrazine sulfate (Dash et al., 2011; Rinaudo, 2006). Unlike chitin, chitosan is highly reactive and is commercially available in powder, paste, fiber, or other forms (Agnihotri et al., 2004). Chemically, chitosan can be defined as a biopolymer with a linear arrangement of glucosamine and N-acetyl-glucosamine as structural and functional units linked by (1-4) glycosidic bonds. The chemical structure of chitosan is determined by the presence of glucosamine units, which affects the degree of deacetylation and, consequently, its reactivity. In this regard, chitosan solubility increases in aqueous-acid solutions when it presents a deacetylation degree higher than 50%, while chitosan biodegradability may be prevented when deacetylation degrees are higher than 69% (Alves and Mano, 2008; Berger et al., 2004). The main physicochemical characteristics of chitosan are presented in Table 3.2. Out of the major biological properties that make chitosan suitable for several applications, its antimicrobial, immunological, wound-healing activities, biodegradability, as well as its nonimmunogenic and noncarcinogenic characteristics may be highlighted (Dash et al., 2011). Regarding its antimicrobial properties, previous research has shown the biological activity of chitosan against Gramnegative and Gram-positive bacteria. One of the proposed mechanisms of action for chitosan antimicrobial activity is the interaction of this biopolymer with

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Table 3.2 Main Physicochemical Characteristics of Chitosan Physicochemical Characteristics Average molecular weight Deacetylation degree Ionic nature Solubility

3.8
20.0 kDa 60%
100% Cationic polysaccharide (at neutral or basic conditions) Insoluble in water, soluble in acid solutions

bacterial cell membranes due to the electrostatic interaction between chitosan— polycationic in nature—and anionic groups located on the bacterial surface. This interaction affects bacterial membrane permeability, resulting in its disruption and, consequently, leakage of intracellular content (Abdelgawad et al., 2014; Shahidi et al., 1999). The biodegradability of chitosan is a result of its biopolymeric nature, constituted of a polysaccharide and structured by means of glycosidic bonds (Croisier and Je´roˆme, 2013). In this regard, the main bonds of chitosan—that is, glucosamine-glucosamine, glucosamine-N-acetyl-glucosamine, and N-acetyl-glucosamine-N-acetyl-glucosamine—are targeted by enzymatic action, leading to chitosan degradation. Lysozyme and bacterial enzymes in the colon have been identified in vertebrates as responsible for chitosan biodegradation. Moreover, higher plants present chitinases, which show enzymatic activity on N-acetyl-glucosamine residues. This indicates a mechanism of self-protection against plant threats, such as microbes and insects, which might have chitin in their structure. On the other hand, the extent of chitosan biodegradation is determined by the deacetylation degree; in this context, high deacetylation degrees result in low biodegradability (Kean and Thanou, 2010). Additionally, chitosan has shown mucoadhesive properties, resulting mostly from its strong electrostatic interactions with the negative charges of sialic acid residues present on mucosal surfaces (Sinha et al., 2004). Therefore, the aforementioned biological properties make chitosan an interesting biopolymer with the potential to be used as a raw material for hydrogel preparation with innumerous biomedical and nonbiomedical applications.

3.3 FUNDAMENTALS OF CHITOSAN HYDROGELS Hydrogels are defined as tridimensional networks comprising of macromolecular compounds with the ability to retain water or biological fluids in their inner side (Berger et al., 2004). Regarding their structures, hydrogels are constituted by two phases: a liquid phase—mostly water or other biological fluids; and a solid phase—polymer chains that confer a gel-like consistency, allowing the structure to entrap water (Croisier and Je´roˆme, 2013). The liquid phase of hydrogels confers biocompatibility and allows for their widespread application in biomedicine,

3.3 Fundamentals of Chitosan Hydrogels

agriculture, food science, and nutrition, to mention a few. Generally, hydrogels are characterized by their particle size, which might range from the nanoscale to several centimeters. Also, they are expected to be highly flexible, allowing them to acquire the shape of the space they are contained in (Bhattarai et al., 2010). The main hydrogel component is a hydrophilic polymer, which can absorb different amounts of water depending on its hydrophilic nature (Bhattarai et al., 2010). As mentioned previously, hydrogels can be classified according to different criteria, such as the nature of the polymeric compound and the method of hydrogel constitution (network nature), among others (Berger et al., 2004). Besides its cationic nature, which allows for the formation of blends with a wide range of polysaccharides, chitosan is among the biopolymers that offer higher flexibility to hydrogel applications. This results from the ease of modification of its molecule through reactions of amino and hydroxyl groups, providing hydrogels with numerous possibilities of structural and morphological profiles as well as absorption potential. As a consequence, this broadens the range of the applications and functions of hydrogels. Such variations, however, are determined by the production techniques since crosslinking is the major cause of the structural, mechanical, and chemical properties of hydrogels, as highlighted in Table 3.3 (Barbucci et al., 2004; Deligkaris et al., 2010; Kalia et al., 2013; Omidian and Park, 2010). Moreover, some features of hydrogels are conditioned to changes in environmental factors, such as temperature, pH, and electric and magnetic properties. These are considered as nonconventional characteristics of chitosan-based hydrogels, which are mainly stimuli-responsive to the aforementioned environmental factors.

3.3.1 PHYSICAL HYDROGELS Hydrogels prepared by physical methods are constituted by noncovalent, reversible links or interactions between polymer chains, such as ionic, electrostatic and hydrophobic interactions, grafting, and entanglement, among others (Berger et al., 2004; Croisier and Je´roˆme, 2013). Factors that can affect hydrogel formation through physical methods include pH, polymer concentration, and temperature. Moreover, hydrogel integrity is determined by the number of interactions that occur between the reacting compounds; therefore, increasing interactions result in stiff hydrogels, while limited interactions result in a soft, weak hydrogel structure (Croisier and Je´roˆme, 2013). Hydrogels prepared by physical methods include polyelectrolyte complexed hydrogels, which are three-dimensional networks resulting from ionic interactions among their constituting polymers. In this regard, a three-dimensional network is formed when two polyelectrolytes with opposite charges react in solution. In the case of chitosan, electrostatic interactions take place between its amino group which features a cationic nature and the anionic group of the other compound present in the solution. This type of hydrogel is considered as an interesting alternative to hydrogels elaborated by chemical methods since no additional molecules

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Table 3.3 Crosslinking Methods Used to Produce Chitosan-Based Hydrogels Main Features

Crosslinking Methods Physical

Chemical

Radiation

Crosslinking agent

Absent

Absent

Interaction nature

Complex coacervation (Magnin et al., 2004) Freeze-thawing (Giannouli and Morris, 2003) H-bonding (Takigami et al., 2007) Heating/cooling a polymer solution (Funami et al., 2007) Ionic interaction (Zhao et al., 2009) Maturation (heat induced aggregation) (AlAssaf et al., 2009) Heat-induced aggregation Polymer solution heating/cooling Drying at high or low temperatures

Agents of different natures depending on the intended applications Covalent Radiation grafting (Cai et al., 2005) Chemical grafting (Spinelli et al., 2008)

Polymerization of the available functional groups Polymer-polymer crosslinking Reticulation triggered by a crosslinking agent Cosmetics

Producing free radicals in the polymer following exposure to high energy source, such as gamma ray, X-ray, or electron beam (Gulrez and Al-Assaf, 2011)

Applied techniques

Applications

Food industry Biomedical products

Aqueous state radiation (Zhao et al., 2003) Radiation in paste (Zhao et al., 2003) Solid state radiation (Kuang et al., 2008)

Biomedicine products Pharmaceutical industry

Based on Gulrez, S. K. H., & Al-Assaf, S. (2011). Hydrogels: methods of preparation, characterisation and applications. In: A. Carpi (Ed.), Progress in Molecular and Environmental Bioengineering-From Analysis and Modeling to Technology Applications (Vol. 1, pp. 117150). London, UK: IntechOpen.

(e.g., catalysts and initiators, among others) are needed for hydrogel formation. Therefore, there is no need for a purification step. Examples of natural polyelectrolytes with anionic nature include polysaccharides with carboxylic groups (COOH), such as xanthan gum, pectin, and alginate. Moreover, collagen, gelatin, keratin, albumin, and fibroin are proteins that have been previously studied for the development

3.4 Characterization Techniques

Table 3.4 Main Advantages and Disadvantages of Physical Hydrogels Advantages

Disadvantages

Additives (catalysts or initiators—mostly toxic) are not needed for hydrogel formation Biocompatibility—essential for biomedical applications Ability to shape the injured tissue as a template

Limited mechanical resistance Easy dissolution should environmental factors not be carefully controlled Limited control of hydrogel pore size

of polyelectrolyte complexed hydrogels (Berger et al., 2004). The major advantages and disadvantages of physical hydrogels are presented in Table 3.4.

3.3.2 CHEMICAL HYDROGELS Chemical hydrogels result from polymeric interactions through covalent bonding. These hydrogels are characterized by the need of chemical modification of chitosan structure and by their nature of having an irreversible structure. Linkages that might take place during the elaboration of chemical hydrogels include amide and ester bonding, as well as Schiff base, among others. During the elaboration process of chemical hydrogels, a crosslinker is required to allow the configuration of the tridimensional network. Most of the time these crosslinkers are small multifunctional molecules, such as tripolyphosphate, ethylene glycol, diglycidyl ether, and others, that react with chitosan and favor interaction with previously activated chitosan active functional groups (Croisier and Je´roˆme, 2013). However, one of the main characteristics of this kind of hydrogel is that a purification step is required after their elaboration, in order to avoid potential toxicity caused by covalent crosslinkers that can remain unreacted after the hydrogel elaboration process (Berger et al., 2004). The main factors that determine the structural characteristics of chemical hydrogels include the crosslinking density and the ratio of crosslinking molecules regarding the polymers used for hydrogel formation (Croisier and Je´roˆme, 2013; Dash et al., 2011).

3.4 CHARACTERIZATION TECHNIQUES Different techniques may be used to characterize the profile of hydrogels. Researchers are encouraged to choose certain ones depending on the desired hydrogel properties, as well as the requirements for specific practical applications. In biomedical engineering, for instance, rheological and cytotoxicological aspects as well as absorption and degradation potentials are of outmost importance.

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Overall, the main aspects that are considered in hydrogel characterization are morphology, swelling, and mechanical resistance. The relevant characterization techniques typically applied to chitosan-based hydrogels may be grouped into structural analysis or property measurements. Structural analyses mainly rely on microscopic and spectroscopic methods. On the other hand, the characterization of performance properties depends on the specific application of hydrogels (Table 3.5).

Table 3.5 Analytical Techniques Typically Used for Hydrogel Characterization Analysis

Technique

Analyzed Factors

References

Structural analyses

X-ray diffraction

• Identification of ions • Mapping and distribution of mineral chemical elements (compositional map) • Phase miscibility and compatibility in blends • Structure length and orientation; Modulation of chemical composition • Topographical analysis • Polymer chain conformation • Morphological, structural, and molecular characteristics of the matrix • Chemical quantification of different compounds (e.g., release of drugs, vitamins, and other active ingredients) • Identification of chemical compounds within the hydrogel matrix • Investigation of possible, important replacements in ingredients

Cullity and Stock (2001)

Scanning electron microscopy Transmission electron microscopy Atomic force microscopy

Ultraviolet
visible spectroscopy

Fourier-transform infrared spectroscopy

Zhou et al. (2008), Zhao et al. (2009) Kamel (2007)

Duran et al. (2006)

Zhou et al. (2008)

Zhao et al. (2009)

(Continued)

3.4 Characterization Techniques

Table 3.5 Analytical Techniques Typically Used for Hydrogel Characterization Continued Analysis

Technique

Analyzed Factors

References

Performance properties measurements

Release of active compounds

• Diffusion of active compounds • Diffusion-controlled mechanism • Relationship of active compound delivery and hydrogel pore size • Mechanical performance • Hydrogel structural integrity • Sol
gel transitions • Determination of hydrogel thermosensibility • Determination of the maximum amount of liquid (water or any biological fluid) retained inside the hydrogel structure • Quantification of hydrophilic degree of developed hydrogels

Bhattarai et al. (2010), Fan et al. (2015, 2016), Zhao et al. (2009), Zhou et al. (2008)

Mechanical resistance

Viscosity

Swelling index

Contact angle

3.4.1 STRUCTURAL ANALYSIS 3.4.1.1 Microstructural and spectroscopic analysis Direct imaging of chitosan hydrogels is important because it allows for the study of microstructural characteristics that might influence hydrogel structural integrity as well as in the loading and release of active compounds. In this regard, microscopic techniques, such as atomic force microscopy (AFM), transmission electron microscope (TEM), and scanning electron microscopy (SEM), allow for the direct imaging of chitosan hydrogels, the formation of new structures, matrix integrity, and porosity. AFM is one of the main tools for studying the surface of chitosan-based hydrogels. Studies done with AFM include the investigation of the threedimensional surface topography as well as the morphological, structural, and molecular properties of hydrogels on a nanoscale (Duran et al., 2006). These techniques allow for the characterization of phase morphology and distribution in blends and composites in addition to the conformation of polymer chains themselves. Such information supports the comprehension of a hydrogel profile.

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TEM allows for a detailed examination of hydrogel samples, including changes in the chemical composition, orientation, and aspect ratio (length-to-diameter or length-to-width ratio) of nanostructures; induction of electronic phase changes; and images based on material absorption. These data support the discussion of the influence of different treatments, as well as the optimization of hydrogel preparation conditions (Kamel, 2007; Williams and Carter, 1996). SEM is useful for determining the morphology of various hydrogel matrices. For SEM imaging, hydrogel samples are usually coated with a thin layer of gold under vacuum prior to sample exposure to electron beam. Zhou et al. (2008) used this technique to analyze thermosensitive chitosan hydrogels and observed that hydrogel morphology was dependent on chitosan characteristics and concentration. In this context, hydrogels prepared with 1% chitosan solution presented a loose and ramified configuration, while increased chitosan concentration resulted in a more compact microstructure. Also, chitosan characteristics, such as low molecular weight, resulted in a loose structure full of holes. Moreover, modifications to the degree of deacetylation resulted in more compact but irregular microstructures when increased from 75.4% to 85.5% (Fig. 3.4). Moreover, thermosensitive chitosan hydrogels combined with αβ-glycerophosphate were developed and characterized using an SEM technique (Zhao et al., 2009). To do so, samples were coated with gold under vacuum, and hydrogel surfaces and cross-sections were subjected to imaging analyses. On the other hand, when coupled with an energy dispersive system (EDS), SEM also determines qualitatively and semiquantitatively the elementary compositions of hydrogel samples by means of emitting X-rays. SEM may indicate miscibility of chitosan hydrogel as well as compatibility of polymer blends with the constituents of the hydrogel matrix. This technique has been previously used by Zhao et al. (2003), who studied the morphological structure, by SEM, as well as the elementary distribution pattern of blend hydrogels based on poly(vinyl alcohol) (PVA) and carboxymethylated chitosan (CM-chitosan) using EDS. In order to determine the pattern distribution of the CM-chitosan components on the surface of the blend hydrogels, developed hydrogels were immersed in a 5% (w/w) CuSO4 aqueous solution for 6 days at room temperature. After this time, the CuSO4-complexed amino groups in the developed hydrogels were determined using an EDS technique. As a result, it was confirmed that CM-chitosan was uniformly distributed in the developed blend hydrogels at low concentrations of CM-chitosan content (,8% w/w); this was probably because CM-chitosan is able to interact more effectively with PVA during hydrogel elaboration due to its higher hydrophilic nature compared to neat chitosan.

3.4.1.2 Ultraviolet
visible spectroscopy and Fourier-transform infrared spectroscopy Spectroscopy techniques are usually employed to determine chemical interactions among functional groups that constitute hydrogel structure. Also, Fourier-transform

3.4 Characterization Techniques

FIGURE 3.4 Scanning electron micrograph of chitosan and αβ-glycerophosphate CS-ab-GP gelation with different formulations (500 3 ). (A) Molecular weight (MW) 5 1360 kDa, degree of deacetylation (DD) 5 75.6%, chitosan concentration 5 1%; (B) MW 5 1360 kDa, DD 5 75.6%, chitosan concentration 5 2%; (C) MW 5 499 kDa, DD 5 75.4%, chitosan concentration 5 2%; and (D) MW 5 1340 kDa, DD 5 85.5%, chitosan concentration 5 2%. Zhou, H.Y., Chen, X.G., Kong, M., Liu, C.S., Cha, D.S., Kennedy, J.F., 2008. Effect of molecular weight and degree of chitosan deacetylation on the preparation and characteristics of chitosan thermosensitive hydrogel as a delivery system. Carbohydr. Polym. 73, 265
273 with permission.

infrared spectroscopy (FTIR) is often used to verify the development of polymerization reactions when hydrogels are formed. The FTIR technique has been used for the characterization of thermosensitive chitosan hydrogels by mixing 2 mg of hydrogel sample with 100 mg of KBr (Zhou et al., 2008). From the obtained results, researchers observed that hydrogel was successfully formed by the interaction of hydrogen bonding between C 5 O of chitosan and
OH of αβ-glycerophosphate and by the junction of N-H of chitosan and
OH of αβ-glycerophosphate. Li et al. (2017) proposed a novel method for preparing a chitosan-based hydrogel and corroborated, through FTIR, that the interaction between NH2 and Ag1 was the major factor leading to crosslinking and that the CTS-Ag1/NH3 hydrogel network was maintained by hydrogen bonds, leading to improved mechanical properties.

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3.4.2 PROPERTY MEASUREMENTS 3.4.2.1 Active compound release assessment The release for active compounds previously incorporated in hydrogels is determined by time and it could be controlled by diffusion, swelling ability of the hydrogel, or by chemical reactions, with the diffusion-controlled mechanism being the most common. However, many active compounds have sizes that are considerably smaller than the pore size of hydrogels, resulting in a burst delivery of these compounds and limiting their controlled release (Bhattarai et al., 2010).

3.4.2.2 Mechanical resistance One of the main characteristics of chitosan hydrogels is their adaptability to the environment where they are used. In this regard, in most cases chitosan hydrogels can mimic the body tissue allowing for their use as scaffolds for biomedical applications, as well as carriers for the delivery of active compounds (Croisier and Je´roˆme, 2013). Thus, chitosan hydrogels must present certain mechanical resistance associated to the maintenance of their structural integrity. For determining the mechanical resistance of chitosan hydrogels, samples are cut in a dumbbell shape. Parameters, such as tension (N) and elongation (mm), are determined using a universal testing machine or a texturometer, while the thickness of hydrogel samples is determined with a caliper (Fan et al., 2015). Finally, the tensile strength and percentage of elongation is determined according to these equations: Tensile strength ðMPaÞ 5 Elongation% 5

F ðW 3 HÞ

S 3 100 L

wherein F is the sample tension at break; S is the elongation at break; W is sample width; H is sample height; and L is sample length. Also, the mechanical resistance of chitosan hydrogel might be determined by puncture test with a texturometer (Gonc¸alves et al., 2017). In this test, a 12.5 mm Teflon probe is forced to puncture the hydrogel samples, allowing for the determination of the maximum force (Bloom index) required to alter a sample’s structural integrity.

3.4.2.3 Viscosity (sol
gel analysis) The measurement of hydrogel viscosity has been used to determine the occurrence of sol
gel transitions. This measurement is quite important when dealing with the development of chitosan thermosensitive hydrogels. Thermosensitive hydrogels are a kind of hydrogel that has the ability to transit from a solution into a gel with temperature modifications. In this regard, Zhou et al. (2008) have developed a thermosensitive hydrogel based on chitosan and αβ-glycerophosphate, which has the ability to transform from solution to gel at

3.4 Characterization Techniques

37 C. Considering viscosity, researchers stablished the sol phase as a physical state of the hydrogel with the main characteristic of flowing liquid, while the gel phase was stablished as a nonflowing physical state of the developed hydrogel. Generally, measurements to determine this property are done in a viscometer, and the hydrogel sample is placed in a water bath at target temperature for sol
gel transition (in this case 37 C 6 0.5 C).

3.4.2.4 Swelling index Swelling index is a quantitative parameter that allows for the determination of the amount of water or biological fluid that can be retained inside the tridimensional structural hydrogel network. Usually, the weight of the dry hydrogel is determined. Following this, the dry hydrogel is placed in distilled water or any other target biological fluid or simulant at room temperature for 48 hours. Also, assay temperatures might vary according to the specific characteristics of the tested hydrogel (e.g., whether it is thermosensible or not). Finally, the tested hydrogel is placed apart from the water, the excess liquid phase is removed from the hydrogel surface, and its final weight is determined (Fan et al., 2015). For swelling index determination this equation is often used: Sw 5

Ws 2 Wd Wd

wherein Sw represents the swelling index; Ws represents the weight of swollen hydrogel; and Wd represents the weight of dry hydrogel (Zhao et al., 2009). If desired the swelling index can be expressed as a percentage (Huber et al., 2017), as: %Sw 5 Sw 3 100

Argin et al. (2014) determined the swelling index of physically crosslinked xanthan-chitosan hydrogels in simulated gastrointestinal conditions. As a result, they observed that xanthan-chitosan hydrogels were sensible to simulated pH media. In this regard, the deionization of hydrogel functional groups affected the transport mechanism inside the tridimensional hydrogel network; thus, highly swellable hydrogels result in large amounts of free water in the hydrogel structure and, consequently, in increased solute release.

3.4.2.5 Contact angle Contact angle measurements allow for the quantification of the hydrophilicity of developed hydrogels. These measurements are usually done at room temperature with a contact angle goniometer, which releases a drop of distilled water onto the surface of the sample. Images of water drop on sample surface are recorded for a specific period of time for angle determination.

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3.4.2.6 Thermal analysis Thermal analysis is done to determine the thermal behavior of developed hydrogels. Hydrogel samples are tested in a thermogravimetric analysis system with temperatures ranging from room temperature to 600 C at a rate of 10 C/min, under an argon or nitrogen atmosphere (Klein et al., 2016; Konwar et al., 2015).

3.4.3 SPECIFIC PROPERTIES FOR BIOMEDICAL ENGINEERING APPLICATIONS 3.4.3.1 Degradability Hydrogels must be both biocompatible and biodegradable in order to be used for biomedical engineering applications (Hoffman, 2002; Shoichet, 2010). Investigating the in vitro degradability of these materials is hence indispensable. To do so, hydrogel samples are put in contact with a simulant medium and organic fluids. Lysozyme is one of the most applied enzymes for analyzing the degradation of chitosan derivatives because it mimics physiological conditions efficiently. Calculations consider sample weight loss throughout immersion in the simulant medium under stirring (80 rpm) at 37 C for approximately 28 days. Hydrogel samples are then removed from the simulant medium at pre-established intervals for weighing. The degradation pattern is plotted with the weight loss data, which is measured with this equation: Weight lossð%Þ 5

Wi 2 Wt 3 100 Wi

wherein Wi and Wt are hydrogel weight at the beginning of the measurements and after each time interval, respectively (Freier et al., 2005; Jin et al., 2009).

3.4.3.2 Cytotoxicity According to the International Standard Organization (ISO), an in vitro cytotoxicity assay is essential for any materials that are intended for direct contact with the human body. In this context, ISO 10993-5 (ISO 10993-5, 2009) described some standard methods for evaluating the toxicity of such materials. Most cytotoxicity tests involve the contact of a sample with a cell culture as well as the further investigation—through the addition of vital dyes as well as the inhibition of cell colonies, for instance—of possible cell disruptions (Nakayama et al., 1997).

3.5 POTENTIAL APPLICATIONS AND FUTURE TRENDS OF CHITOSAN HYDROGELS The development of chitosan hydrogel has wide potential applications in several fields of science and novel products, such as in drug delivery, tissue engineering (scaffolds), and wound-healing applications, among others (Alves and Mano, 2008;

References

Croisier and Je´roˆme, 2013; Dash et al., 2011). Chitosan has been shown as a promising compound for the partial or total substitution of synthetic materials for the elaboration of hydrogels. In this regard, the interest in this compound as a key material for hydrogel development has been notorious, with innumerous studies presenting its potential application in several fields, such as personal care products, agriculture, and pharmaceutics. The main properties of chitosan are its biodegradability, low toxicity, and rheological aspects, which are compatible with the current requirements of biomedical engineering. Thus, these features of chitosan have favored the development of an increasing number of studies toward a better understanding of the relationship between chitosan structure, its properties, and potential applications in new products. Moreover, studies regarding the interaction of chitosan with other biopolymers have been focused on the improvement of performance properties, such as mechanical resistance and swelling ability, aiming at efficient absorbing materials (Ng et al., 2016; Rodrı´guez-Va´zquez et al., 2015). Currently, the challenges regarding the development of hydrogels goes beyond crosslinking techniques and new formulations; these new challenges include further in vitro and in vivo tests, especially regarding drug release, genetic therapy, and others (Rani et al., 2010). Nevertheless, chitosan constitutes an excellent polymeric matrix for the development of hydrogels; some limitations still remain in its structure but they can be easily overcome by implementing new technologies (physical or chemical interactions) that allow for improving its functionality.

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CHAPTER

Bioreabsorbable polymers for tissue engineering: PLA, PGA, and their copolymers

4

Ana Carolina B. Benatti1,2, Ana Fla´via Pattaro1, Ana Ame´lia Rodrigues1, Mariana Vitelo Xavier1,2, Andreas Kaasi1,2,3, Maria Ingrid Rocha Barbosa1, Andre´ Luiz Jardini1, Rubens Maciel Filho1 and Paulo Kharmandayan1,2 1

Institute of Biofabrication, School of Chemical Engineering, University of Campinas, Campinas, Sa˜o Paulo, Brazil 2School of Medical Sciences, Surgery Department, University of Campinas, Campinas, Sa˜o Paulo, Brazil 3Eva Scientific Ltda, Sa˜o Paulo, Sa˜o Paulo, Brazil

4.1 TISSUE ENGINEERING Current advances in modern medicine have enabled the development and improvement of biomaterials and biofabrication techniques that improve the quality of life of people who suffer from certain limitations caused by disease or trauma (Barbanti et al., 2005; Karande and Agrawal, 2008; Liu et al., 2012). The technique known as tissue engineering seeks to enable the functional repair or regeneration of living organs and tissues, and may be defined as a multidisciplinary science that applies principles and methods from life science and engineering to develop biological substitutes in order to restore, maintain, or improve function (He´lary and Desimone, 2015; Badylak, 2005). In this manner, the development of biomaterials for tissue engineering and regenerative medicine is of fundamental importance, since these may be utilized in clinical practice in different areas of medicine, such as orthopedics, dermatology, plastic surgery, and others, and likewise for use in odontology and veterinary medicine. Tissue engineering aims to combine the in vitro culture of cells suitable for a given tissue, along with a biocompatible, biodegradable, three-dimensional structure: a support (scaffold). This will provide a physical structure that will also act as a basis for cell development, differentiation, and cellular migration, as well as for the development of extracellular matrix (ECM) components necessary for the creation and maintenance of tissue. Fig. 4.1 identifies the main steps in tissue engineering. Most structures in tissue engineering are composed of at least two important components: a group of cells and a material in which they can grow. These elements play an important role in the development of new tissue. The cellular component is required for the generation of new tissues through ECM production and is responsible for its long duration and maintenance. The scaffold material Materials for Biomedical Engineering: Hydrogels and Polymer-based Scaffolds. DOI: https://doi.org/10.1016/B978-0-12-816901-8.00004-3 © 2019 Elsevier Inc. All rights reserved.

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FIGURE 4.1 Steps for tissue engineering. (Revati et al., 2015).

provides short-term mechanical stability and provides a three-dimensional organization model for tissue development. The interaction of these two components, such as the coordination of degradation rates of polymers with the rate of cellular biosynthesis and cellular inoculation characteristics of the polymer, are critical for the success of the construction in tissue engineering (Bonassar and Vacanti, 1998). Normally, the preparation of the tissue engineering products are obtained following these steps (Barbanti et al., 2005): selection and processing of the support; inoculation of the cell population on the support; premature tissue growth; growth of mature tissue in a physiological system (bioreactor); and assimilation of the product. For the reconstruction of an organ or tissue it is necessary to select a support (scaffold) in which cell growth will occur. The selection of this support considers type, site of the lesion, and its extension, and basically follows two strategies of application when prepared with bioresorbable polymers, as is the case of poly (α-hydroxy acids) (Barbanti et al., 2005). The first strategy consists of the development of the material so that physical and mechanical support of the cells occur from inoculation to the reimplantation of the material into the host. The function of the implant, consisting of the polymer and the cell, is to provide support for cell growth and also to serve as a mechanical/structural substitute for the original tissue until the formation of new tissue and the complete bioabsorption of the material (Barbanti et al., 2005). In the second strategy, the implant consists of formed mature tissue. The scaffold, which is a polymeric support, is designed with mechanical properties and degradation time suitable for inoculating the cells to their insertion into a bioreactor, where the formation of mature tissue will occur. Only after tissue formation, is the implant inserted into the body (Barbanti et al., 2005).

4.3 Biomaterials

4.2 SCAFFOLDS Scaffolds are critical for tissue recovery, and have functions essentials to the success of tissue regeneration. Those functions are: to serve as spacers to prevent tissue growth in the vicinity of the affected site; to provide a temporary support structure for the tissue to be replaced; to serve as a substrate for cell growth, proliferation, migration, and differentiation; to provide spaces for vascularization, for the formation of new tissues and remodeling; and to allow for efficient transport of nutrients, growth factors, blood vessels, and the removal of materials (Karande and Agrawal, 2008). To take on these functions, scaffolds need to meet some basic requirements. The main requirement for any type of scaffold is biocompatibility. The biocompatibility is vital for a beneficial response of the cells, as well for an adequate immune response of the host to the tissue or implant to be provided. This means that the interactions that occur between the scaffold, cells, and the tissue of the host should be favorable, without any potential damage due to induced cytotoxicity, the generation of an adverse immune response, or the activation of blood coagulation pathways. A number of factors contributes to the tissue generated by the biomaterial, including implant shape and size, its chemical reactivity, degradation rate and its by-products, as well the implant location and the host species. Scaffolds may be of natural or synthetic origin. However, all scaffolds must have mechanical properties, such as resistance, tenacity, and ductility corresponding to the host tissue; so that scaffolding degradation does not occur before the generation of new tissue. The size, geometry, interconnectivity, branching, and chemistry of the pore and canal surfaces can directly affect the extent and nature of nutrient diffusion and internal tissue growth. Generally, living tissue is observed in the outside regions of the scaffold while the inner part does not support viable tissue due to a lack of adequate diffusion (Mirtchi et al., 1989). Currently, scaffolds manufactured with poly(α-hydroxy acids) are some of the most used, due to their high biocompatibility and degradability, which facilitate tissue regeneration. Some of these polymers are: poly(lactic acid) (PLA); poly (ε-caprolactone) (PCL); and poly(glycolic acid) (PGA).

4.3 BIOMATERIALS Biomaterials are considered substances of natural or synthetic origin that may be tolerated in a temporary or permanent manner by tissues of a living body. Biomaterials may be used as individual units or as part of a more complex system that offers treatment, repair, or substitution of a tissue, organ, or function (Williams, 1987a,b,c). Biomaterials may also be used as medical devices (Piotrowska and Sobczak, 2015), for example, through implants (sutures, bone plates and pins, joint replacement, heart valves, dental implants, intraocular

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lenses, etc.) or e medical apparatuses (pacemakers, artificial hearts, etc.), with a view of improving the quality of life of a patient (Ramakrishna et al., 2001; Wiebe et al., 2014; Piotrowska and Sobczak, 2015). Fig. 4.2 shows different uses of biomaterials in the human body. Biomaterials of synthetic origin represent the largest group of biodegradable polymers. These may be produced through controlled conditions in a laboratory using reproducible mechanisms and with constant physicochemical characteristics and properties, such as mechanical strength, elasticity, and degradation rate (Dhandayuthapani et al., 2011). On the other side, synthetic polymers present, as

FIGURE 4.2 Different uses of biomaterials in the human body. PE, Polyethylene; PMMA, Poly(methyl methacrylate); PLA, poly(lactic acid); PGA, poly(glycolic acid) (Marques, 2011).

4.3 Biomaterials

a disadvantage, limitations in their bioactivity due to their hydrophobic nature (Nooeaid et al., 2012). The group of poly(α-hydroxy acids) are some of the most widely used synthetic polymers at present, such as PLA, PCL, and PGA (Carson and Bostrom, 2007; Cima et al., 1991; Vacanti et al., 1995). Table 4.1 contains some examples of synthetic polymers and their applications. On the other hand, polymers such as agarose, alginate, collagen, and hyaluronic acid are materials of natural origin. These present advantageous characteristics, such as the capacity to adapt to different formats with great flexibility, in addition to having molecular domains that aid cell and tissue development and increase the biological interaction between biomaterials and adjacent tissues (Carson and Bostrom, 2007). Table 4.2 contains natural polymers and their biomedical applications. In general, biomaterials must adhere to strict requirements, since in addition to repairing the function of injured tissue by way of compatible mechanical properties and anatomy-specific compatible size and shape, in order to effectively carry out their functions, they should not, under any circumstance, cause biological disturbances in the organism over time. In this way, biomaterials should present compatibility with the receptor organism, that is, they should not cause toxic effects to the organism (Jahno, 2005). As described previously, biomaterials may be composed from synthetic polymers, metals, ceramics, and natural macromolecules, manufactured to comply with the site and function they are to carry out of the tissue with which they will be contacting. Moreover, they may be classified according to their physiological comportment as: biotolerable, bioinert, bioactive, and absorbable (Hench and Andersson, 1993). Biotolerable biomaterials are isolated from adjacent tissue by the formation of a fibrous tissue around the material, where the thickness of this tissue is inversely proportional to the tolerance of the tissue adjacent to the biomaterial. This is the case for most synthetic polymers and metals (Hench and Andersson, 1993). Bioinert biomaterials are those that the body tolerates the presence of, with minimal formation of fibrous tissue around the material, due to an almost Table 4.1 Description of Biodegradable Synthetic Polymers and Their Applications in Medicine (Vozzi et al., 2003; Nakamatsu et al., 2006; Serra et al., 2013; Vroman and Tighzert, 2009; Rentsch et al., 2012) Polymers

Biomedical Applications

PLA, PGA, PLGA PLA/HAP PLGA/HA PCL

Surgical suture; scaffolds; orthopedic medical device for bone fixation Bone tissue engineering Growth factor delivery; tissue vascularization Bone and cartilage repair; orthopedic surgery sutures

PLA, Poly(lactic acid); PGA, poly(glycolic acid); PLGA, poly(lactic-co-glycolic acid); HAP, hydroxyapatite; HA, hyaluronic acid; PCL, Poly(ε-caprolactone).

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Table 4.2 Natural Biodegradable Biomaterials (Lee and Mooney, 2001; Vinatier et al., 2006; Brandl et al., 2007; Pescosolido et al., 2011; Collins and Birkinshaw, 2013) Polymers

Biomedical Applications

Agarose Alginate Collagen Hyaluronic acid

Bacterial culture; electrophoresis Scaffolds; drug delivery Artificial skin; scaffolds Angiogenesis

inexistent release of material substances into the tissue. This is the case with materials such as titanium, zirconium, titanium alloys, and carbon (Park and Lakes, 2007). Bioactive biomaterials create a chemical bond between the material and the tissue, due to chemical similarity between the two, as is the case with glass and glass-ceramic with bone tissue, due their composition of calcium phosphate and hydroxyapatite (HA) (Park and Lakes, 2007). Finally, absorbable biomaterials are those that after a certain period of time, end up degraded or absorbed by the organism, in a manner whereby they do not cause an inflammatory process or any type of toxicity in the tissue or organ, since these biomaterials, upon degradation, form nontoxic waste products that are eliminated by the organism through natural metabolic pathways, such as the Krebs cycle or through urine excretion (Ciccone et al., 2001).

4.3.1 POLYMERIC BIOMATERIALS The usage of polymers as biomaterials has generated great attention in the scientific community in the past decades. Different chemical structures and functional groups present in polymers govern the morphology and properties of these materials, allowing for molecular architectures suiting different applications in regenerative medicine (Kwon et al., 1991; Shi, 2006; Khan et al., 2015). Macromolecules, which form part of polymeric materials, are formed from smaller fundamental and repetitive structures called monomers, giving rise to long chains formed through polymerization reactions. Among polymeric polymers, some biopolymers are produced from raw material from renewable sources, such as sugar cane, corn, and cellulose (Bhardwaj and Kundu, 2010). The mechanical properties of a polymer are dependent on certain factors, such as its molecular weight, chemical structure, orientation and molecular crystallinity, and purity, among others. Mechanical properties, such as Young’s modulus and maximum tensile strength, increase with increasing molecular weight, which varies from material to material. The crystallinity of the material also directly affects its physical characteristics and degradability, where increasing polymer

4.3 Biomaterials

crystallinity and orientation improve the mechanical properties of the polymer, such as the mechanical strength (Vert et al., 1992). A fundamental feature that a biopolymer should present, in order to be used as a medical device or for use in tissue engineering, is biocompatibility. In sum, this would be the absence of toxicity, that is, it should not cause any negative effect to the cell or tissues with which the material will be in contact. The use of biopolymers in organisms to accelerate tissue recovery, has been studied by various groups using in vitro and in vivo approaches, where the former is a preassessment of the latter (Lucchesi et al., 2008). The success of the tissue development is based on three factors: adequate selection of the type of cell that ought to be used and implanted; selection of a material with adequate structural and physicochemical properties for the implantation site, in addition to its function as a three-dimensional support (scaffold); and interaction of the material with the cells of the tissue (Malafaya et al., 2007; Cramer, 2013). Among the most promising biocompatible biomaterials is the group of poly (α-hydroxy acids). The main polymers within this group are: PLA, PCL, PGA, and poly(p-dioxanone) (PDO) (Nampoothiri et al., 2010). However, these polymers present some deficiencies, such as poor flexibility, slow biodegradation, and high crystallinity.

4.3.2 BIOREABSORBABLE BIOPOLYMERS Bioreabsorbable biopolymers are a class of extremely important biopolymers, due to their capacity to break down both in vitro and in vivo. This capacity is known as biodegradability, and represents the capacity of a biomaterial to degrade, accompanied by the formation of nontoxic substances, without the need for surgical intervention to remove the material after the tissue recovery period. A bioreabsorbable biomaterial must possess suitable mechanical properties for the site of application and a suitable degradation time for the recovery of the tissue in question (Middleton and Tipton, 2000). Also, it must not provoke any toxic or inflammatory response in the tissue where it was implanted; it must be metabolized into nontoxic substances, and then eliminated through the organism’s natural metabolic pathways; and it must be easily processed and sterilized. These materials are used in tissues that require a temporary supportive structure for tissue reconstruction. In the medical field, temporary bioreabsorbable implants have been used since 1960, in surgical suture based on lactic acid and glycolic acid (Gilding and Reed, 1979), in orthopedic medical devices, stents, and more recently as controlled drug delivery systems. Currently, the desire of tissue engineering of being able to reconstruct tissues and organs and returning the quality of life to a patient, without the risk of rejection, is a concept embraced by various research centers around the world (Lasprilla, 2011). Materials such as PCL, PGA, and PLA are of substantial interest, since they are hydrolytically unstable and degrade naturally due to ester bond cleavages; a process that takes place over time and which may vary from a few days to years,

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and because of this they have been targeted for studies for the production of biodegradable sutures, reabsorbable prostheses, and artificial organs (Nampoothiri et al., 2010).

4.4 POLY(α-HYDROXY ACIDS) Poly(α-hydroxy acids) are widely used polymers in biomedical research, and the most common polymers in this category are: PLA, PGA, and the copolymer poly (lactic-co-glycolic acid) (PLGA) (Basu et al., 2016). The application field of these polymers is vast and they are employed as medical devices in implants, suture materials, prostheses, orthopedic repair materials, intermedullary nails, and for controlled drug delivery (Basu et al., 2016). PLA has excellent biocompatibility and mechanical properties, with these being pretexts for it becoming one of the most extensively studied and utilized materials for tissue engineering, in addition to being one of the most widely used biopolymers for medical devices, such as screws, plates, valves, and others (Auras et al., 2010). This diversity of different PLA uses comes down to the fact that small changes in its physicochemical structure give rise to changes in its characteristics, making it useful for different areas. It is one of the most promising and widely used biopolymers. Beyond its excellent physicochemical characteristics and biocompatibility, it may be formed from its basic organic building block using materials from renewable sources, by the fermentation of sugars obtained from natural sources, such as sugar cane or corn (Drumright et al., 2000; Moon et al., 2001; Gupta et al., 2007; Cheng et al., 2009; Auras et al., 2010). PLA may be prepared by two main polymerization types: polymerization by direct condensation of lactic acid and polymerization by the opening of the cyclic lactic dimer ring (Basu et al., 2016). The different properties of PLA, such as melting point, mechanical strength, and crystallinity, are given by the structure of the polymer itself (due to different proportions of L-, D-, or meso-lactide) and also by its molecular weight (Auras et al., 2010). PLA contains two stereoisomers: poly(L-lactic acid) (PLLA) and poly(D-lactic acid) (PDLA) and its racemic mixture leads to poly(D, L-lactic acid) (PDLLA) (Vert et al., 1992). PDLA and PLLA (levorotary and dextrorotary) are mirror images of each other, both being optically pure and semi-crystalline. PDLLA is racemic, amorphous, and optically an inactive polymer (Tsuji and Ikada, 1997). PLA is a semi-crystalline polymer, with a glass transition temperature of 57
C and a melting temperature between 174
C and 184
C; the methyl group of PLA is responsible for its hydrophobic nature and for it being more resistant to hydrolysis. High molecular weight is a parameter that has a direct effect on the quality of the produced polymer medical devices, since increasing molecular weights lead to

4.5 Poly(α-Hydroxy Acids) Synthesis

medical devices with better mechanical properties, this being a crucial factor in the search for a high-quality biomaterial (Gogolewski, 2000). PLLA has excellent physical characteristics, such as mechanical strength and thermal plasticity, in addition to presenting good processability and the property of material hydrolysis giving rise to cleavage of polymeric bonds. From these oligomers are obtained, and after that monomer units of lactic acid, and these may suffer from enzymatic attack and are then naturally eliminated by the human body through metabolic pathways (Proikakis et al., 2002). Initially, the hydrolysis of the polymer leads to a decrease in molecular weight, particularly in amorphous regions, in a way that water diffuses into the polymer and causes its fragmentation, contributing to the loss of mechanical strength and potentially a reduction of mass (Middleton and Tipton, 2000). PGA, on the other hand, is a biocompatible polymer known since 1954. In 1962, PGA was developed as the first absorbable synthetic suture material, and received the name Dexon by the manufacturer, American Cyanamide DuPont Dexon sutures are rapidly reabsorbed by the body, losing their mechanical characteristics after a period of implantation that varies between 2 and 4 weeks. Due to its simple structure and stereochemistry, PGA can present different degrees of crystallinity, ranging from completely amorphous to a maximum of 52% crystallinity. Due to its great crystallinity, it is insoluble in most organic solvents and, for the same reason, presents the best mechanical properties among biodegradable polyesters (Singh and Tiwari, 2010). PGA may be polymerized in a direct or indirect manner, from glycolic acid. Direct polycondensation produces a polymer with a molecular weight below 10,000 Da, due to the requirement of a high degree of dehydration (99.28%) and the absence of monofunctional impurities. To obtain PGA with a molecular weight above 10,000 Da, it is necessary for its synthesis to take place by polymerization through opening of cyclic dimer rings in the glycolic acid (Ching-Chang, 2000; Gunatillake and Adhikari, 2003; Huang, 2004; Zhao, 2007; Singh and Tiwari, 2010).

4.5 POLY(α-HYDROXY ACIDS) SYNTHESIS The main method of PGA and PLA synthesis, to obtain a low molecular weight polymer, is direct polycondensation of glycolic or lactic acid. To produce PGA and PLA of high molecular weight, different techniques have been developed, but the preferred method is that of polymerization through the opening of cyclic dimer rings. Factors that affect the mechanical performance of polymers include the selected monomer, the selection of initiator, process conditions, and the presence of additives. These factors affect polymer hydrophilicity, crystallinity, melting and glass transition temperatures, molecular weight, distribution of molecular weight, terminal end groups, and the presence of residual monomers and additives (Middleton and Tipton, 2000).

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With direct polymerization by polycondensation, the presence of hydroxyl and carboxyl groups in the molecule allows its conversion to a polyester through a polycondensation reaction, with this being considered a low-cost solution (Auras et al., 2010). The output is a polymer of low molecular weight and limited mechanical properties, and may be used for controlled drug delivery devices, surgical suture, and scaffolds for cell cultures, while not being indicated for other applications, such as screws, plates, and others (Drumright et al., 2000). The procedure of polymerization by condensation consists of the heating of the monomer at atmospheric pressure and a temperature within the range of 175
C185
C until the point where there is no more water distillation taking place. Subsequently, the pressure is reduced while maintaining the temperature for a certain period of time, giving rise to PGA or PLA of low molecular weight (Gunatillake and Adhikari, 2003; Zhao, 2007; Singh and Tiwari, 2010). The first attempt to prepare poly(α-hydroxy acids) using the ring opening method was performed by Carothers in 1932 (Carothers et al., 1932). The process of polymerization by cyclic dimer ring opening may follow the process of melt polymerization, also known as bulk polymerization, or it may follow the process of solution polymerization, suspension polymerization, or emulsion polymerization. The mechanism of ring opening polymerization may be divided into three reaction classes: cationic, anionic, and coordination-insertion. The former two are considered alternative methods, with the latter being the most widely used industrially (Gunatillake and Adhikari, 2003; Zhao, 2007; Singh and Tiwari, 2010). There have been many suggestions of polymerization of lactide using tin octoate as a catalyzer, however no consensus exists as to what mechanism will faithfully reproduce the reaction. A plausible theory was presented by Nijenhuis et al. (1992), where the tin octoate catalyzer (Sn(Oct)2) coordinates with the lactide through its free p- or d-orbitals with a CQO bond. Just as with the solvation of Lewis’ acid catalyst, the complex presents a cationic feature, as seen in Fig. 4.3. The complete mechanism is presented in Fig. 4.4. Since the electron density in the carbon is diminished by resonance, there is a facilitation of nucleophilic attacks mediated by OH-containing compounds. These groups may come from the catalyst itself, which contains small quantities of water as a result of its production method. Polymerization is initiated when the OHcontaining compound reacts with the lactone/Sn12(Oct) complex, through a carbon R

R

R O

Sn Oct

O

O

O O

Sn Oct

O

Sn Oct

FIGURE 4.3 Resonance structures of the lactone/Sn 1 (Oct) complex (Nijenhuis et al., 1992).

4.5 Poly(α-Hydroxy Acids) Synthesis

Oct

Oct Sn

Sn

R' O H

O

O

O O

O

Oct

Oct

Sn

R' H

O

R

R

R

1

2

Oct

Oct

Sn O H R 6

H O O R'

H O R'

O

O

O R'

H O

R

–H

3

Sn

O 5

O +H

Oct

Sn

R

Sn

O R'

O H

O 4

O R' R

FIGURE 4.4 Mechanism of reaction for polymerization of lactones by Sn(Oct)2 (Nijenhuis et al., 1992).

nucleophilic attack as may be seen in resonances 1 and 3 in Fig. 4.2. After coordination, the complex in structure 6 (Fig. 4.4) is obtained, which may be used to coordinate with another lactone, giving rise to species such as the type in structure I, where R represents chain growth. In this mechanism, the catalyst is not chemically bound to the end of the chain and this allows the molecule to “jump off” the end of the chain to another, so that the number of polymerized chains to be produced may be lower than the number of catalyst molecules (Nijenhuis et al., 1992). Bendix (Bendix, 1998) also suggested a mechanism. The difference between this mechanism and the one presented by Nijenhuis et al. (1992) is the absence of the OH group (considered a result of catalyst impurities). The suggested mechanism consists of the lactide ring opening by the catalyst through acyl ester bond cleavages. By this mechanism, the chain extremities will be formed in this order: the beginning and the end of the chains will have ester and carboxyl acid terminations, respectively, as per Figs. 4.4 and 4.5. Studies indicate that the reaction of polymerization using Sn(Oct)2 happens in a faster manner and with better control, when there is a protic reagent present, such as with alcohol (Dechk-Cabaret et al., 2004). Recent investigations have enabled the characterization of different intermediary tin complexes that heavily support the mechanism of coordination-insertion. However, the most debated matter concerns the very nature of the initiation complex. While it is still accepted that protic reagents, such as alcohols, react with Sn (Oct)2 to form covalent tin(II) alkoxides, the coordination stage may occur with a retention of covalent ligands (Eq. 4.1) or with the release of octanoic acid (Eq. 4.2), and the conditions of reaction (temperature, ratio alcohol/tin, solvent, and others) have a strong influence on the process (Dechk-Cabaret et al., 2004).

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CHAPTER 4 Bioreabsorbable polymers for tissue engineering

Equações juntas Sn (Oct)2 + ROH Sn(Oct)2 + ROH (lactide)n

RO

Sn(Oct) + ROH

(ROH)Sn (Oct)2 (RO)Sn(Oct) + OctH (lactide)nH

(RO)Sn(Oct) + RO

Equações separadas Sn (Oct)2 + ROH

(lactide)n

RO O H3C

Sn(Oct)2 + ROH

(RO)Sn (Oct) + OctH

Sn(Oct) + ROH

(RO)Sn(Oct) + RO

Acyl bond cleavege

O O

+

CH3

O L-Lactide

Sn

(ROH)Sn (Oct)2

O

O 2+

R1

O 2

Stannous (II) ethylhexanoate

O

CH3 O

CH3

O

(lactide)n

CH3

O O n

Polymer backbone

O CH3

H

O Sn 2+

O

Growing chain end

R2

octoate (Oct)

CH3

O R1 =

CH3

O O

O

O

O

O O

CH3

O

O O

CH3

Ester terminated start group (L-Lactide attacks the octoate)

Terminated chain end (Octoate attacks the terminal carbonyl group)

O O

CH3 O

CH 3

OH

O

Terminated chain end (Water attacks the terminal carbonyl group)

FIGURE 4.5 Mechanism of polymerization of lactide using Sn(Oct)2 proposed by Bendix (1998).

Finally, it should not be underestimated that other than the initiation of the polymerization reaction, protic agents may be involved in the reversible transfer of chains, as per the polymeric chain growth (Eq. 4.3), thus, making essential the optimization of the ROH/Sn(Oct)2 rate (Dechk-Cabaret et al., 2004). SnðOctÞ2 1 ROH2ðROHÞSnðOctÞ2

(4.1)

SnðOctÞ2 1 ROH2ðROÞSnðOctÞ 1 OctH

(4.2)

RO 2 ðLactideÞn 2 SnðOctÞ 1 ROH2ðROÞSnðOctÞ 1 RO 2 ðLactideÞn 2 H

(4.3)

4.5 Poly(α-Hydroxy Acids) Synthesis

R CH3 O

O Sn

O

H3C

O O

2 CH3OH

H3 C O

O

H

O H

H 3C

H3C O H

O

O

R

CH3

O CH 3 O

O

O

O

O

O

O

O CH 3

Sn

H 3C

O

H

R

O H3C

O

O CH3 R Sn O O H3 C O O O O H CH3 O R

O H

H3C

H3 C O

O O

CH3

O

O

R

Sn

H

O

H

O

Sn

O O

CH3 R O O

O R

CH3

FIGURE 4.6 Mechanism of reaction of Lactide polymerization with Sn(Oct)2, in the presence of methanol (Dechk-Cabaret et al., 2004).

Fig. 4.6 shows the mechanism of the reaction of Lactide polymerization by ring opening using Sn(Oct)2 as the catalyst and methanol as the initiator. Two molecules of methanol are coordinated with Sn(Oct)2. As with the model of Sn(Oct)2, both coordinations are favorable and occur in an associative manner, that is, with the retention of two halves of octoanate (hydrogen bonds are formed between alcohol and octoanate ligands). A weak lactide complexation is expected at this stage. The last step of the coordination induces the migration of methanol protons to the neighboring octoanate ligand, in a way that the alcohol is converted into alkoxide. Subsequently, the insertion happens in two stages, known as the coordinated nucleophilic attack of alkoxide on lactide, followed by a ring opening, resulting in the insertion of lactide in the OH bond of the coordinated methanol. This route suggests that the octanoic acid remains coordinated with the tin during the propagation, but if both entropy and reaction temperature are considered, it may be that the octanoic acid dissociates from the tin alkoxide during the ring opening. In coordination-insertion polymerizations, the efficiency of control of molar mass depends on the rate of kpropagation/kinitiation, but it is also dependent on lateral transesterification reactions. These transesterification reactions may occur in two ways: intramolecularly (giving rise to macrocyclic structures and shorter chains) and intermolecularly (redistribution of chains), as may be seen in Fig. 4.6. The equilibrium between polymerization and depolymerization may be considered as a particular case of the reaction of intramolecular transesterification (Dechk-Cabaret et al., 2004) (Fig. 4.7).

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CHAPTER 4 Bioreabsorbable polymers for tissue engineering

Intramolecular n

O R

2

R 4

O

3

1O

5

O

n

6

O

M

O

R

O

O 4

3 2

O +

M

6

O

1

R OR'

5

R

O

O

OR' R

Intermolecular

R

M

O O

O O

O

R O

OR' R

O

O OR'

M

R

O

R O

R

OR' R

O O

M

O O

+

O

OR'

O

M

R

FIGURE 4.7 Representation of intramolecular and intermolecular lateral transesterification reactions (Dechk-Cabaret et al., 2004).

Oct Oct

Sn

ROH

Oct Oct

H Oct Sn O Oct R Glycolide O

H Sn O

R

O 1 5 O

Oct

O Sn 5 Oct

O

1 O

O

O

n

O O O

R

O O

HO

O O

H 2n+2

FIGURE 4.8 Reactional mechanism for the synthesis of PGA (Mohammadi-Rovshandeh and Sarbolouki, 2001).

All these lateral reactions give rise to a wide distribution of molar mass, and sometimes to polymers with nonreproducible molar masses (Dechk-Cabaret et al., 2004). The mechanism of the proposed PGA polymerization is similar to the polymerization of PLLA, as both polymers are polyesters. Fig. 4.8 presents a proposed mechanism for the process of PGA polymerization by ring opening, using tin octoanate as the catalyst and an alcohol as the initiator.

4.6 Copolymerization of Poly(α-Hydroxy Acids)

4.6 COPOLYMERIZATION OF POLY(α-HYDROXY ACIDS) One of the ways of modifying the properties of a polymer is by copolymerization with other monomers. Mechanical properties, degradation rate, hydrophilic/hydrophobic balance, and other parameters, may be altered resulting in a great variety of materials. Even though PLA and PGA present similar structures, their chemical, physical, and mechanical properties are extremely different, and this is due to the presence of a methyl group bound to the alpha carbon, and, as such, the copolymerization of these materials results in a polymer with distinct features (Gentile et al., 2014). The copolymer PLGA may be prepared from different proportions of its constituent monomers, and depending on the quantity of lactide and glycolide used in the polymerization, different copolymer forms may be obtained (Erbetta et al., 2012). Different synthesis mechanisms are used in the production of copolymers, but the process parameters strongly affect the physicochemical characteristics of the end-product. The process of PLGA copolymerization may be carried out by two main approaches: the first is by the reaction of the direct polycondensation of lactic acid and of glycolic acid, giving rise to copolymers of low molecular weight; and the second is by the reaction of ring opening of cyclic lactide and glycolic dimers, giving rise to copolymers of high molecular weight, and, consequently, possessing better mechanical properties (Gilding and Reed, 1979; Gentile et al., 2014). As lactide is much more hydrophobic than glycolide is, a polymer rich in lactide will be less hydrophilic and will absorb less water, resulting in a slower polymer chain degradation process. It is important to emphasize that the relationship between the copolymer physicochemical properties and the quantities of glycolide and lactide is nonlinear (Gilding and Reed, 1979). Fig. 4.9 presents the reaction of the formation of the PLGA copolymer using cyclic dimers as monomers, whereas Fig. 4.10 presents the polymerization reaction using lactic acid and glycolic acid as monomers.

O m

O

O

O O

+

n

O

O O

O

O

Lactide

Glycolide

O O

O

O

m O

O

n Polyglycolide PGA

Polylactide PLA PLGA

FIGURE 4.9 Reaction of PLGA copolymerization using cyclic dimers as monomers (Erbetta et al., 2012).

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CHAPTER 4 Bioreabsorbable polymers for tissue engineering

m

O HO

OH + n

Lactic acid

O

O HO

OH

Glycoli cacid

O

HO

O O

m

H n

Poly-(lactic-co-glycolicacid)

FIGURE 4.10 Reaction of PLGA copolymerization using lactic acid and glycolic acid as monomers (Erbetta et al., 2012).

Since glycolide has higher reactivity when compared to lactide, the proportion of glycolide in the final polymer is, in general, higher than the initial quantity at the beginning of the reaction (Gilding and Reed, 1979; Erbetta et al., 2012). Gilding and Reed (1979) determined that the rate of reactivity of glycolide for a reaction at 200
C was 2.8, whilst for lactide it was 0.2; therefore, glycolide will be preferentially copolymerized at low conversions and lactide will be incorporated more and more as the glycolide polymerization reaches completion (Erbetta et al., 2012). The degradation of PLGA copolymer occurs by hydrolysis of ester bonds in the presence of water. This process happens over four stages: hydration, initial degradation (cleavage of covalent bonds with accompanying decrease of molar mass), constant degradation, and solubilization (Gentile et al., 2014).

4.7 MECHANISMS OF DEGRADATION OF POLY(α-HYDROXY ACIDS) Many factors are related to the degradation rate of bioabsorbable polymers, such as the site of implant; molar mass; molar mass distribution; mechanical properties; chemical and stereoisometric composition; crystallinity; morphology; surface roughness; porosity; surface charge; surface free energy; pH; presence of additives; and others (Lam et al., 2008). Bioabsorbable polymers may, as such, have different degradation rates, leading to different degradation times varying from weeks to years. Water, when present, diffuses into amorphous regions, which allows for an easier penetration of water than into crystalline regions. Therefore, hydrolytic degradation takes place in the amorphous regions, where there is a reduction of mean molar mass, yet without causing a loss of mechanical properties in the material. Only after an almost total loss of the amorphous regions does water begin to slowly penetrate into the crystalline regions, followed by hydrolysis and accompanied by the degradation of the mechanical properties of the material (Middleton and Tipton, 2000).

4.8 Biocompatibility

Completely amorphous materials tend to undergo hydrolysis more easily than the partially crystalline ones. However, factors such as hydrophilic monomeric residues, more hydrophilic terminal acid groups, more reactive hydrolytic groups of the main chain, and a medical device of smaller size can also accelerate the process of hydrolysis (Middleton and Tipton, 2000). The degradation of biodegradable polymeric materials happens through random scissions of the main chain, in which there is a gradual decrease of the mean molar mass, or by scissions of terminal groups of the polymeric chain, in which terminal carboxyl groups formed during the degradation are subject to the effect of autocatalysis. PLA and its derivatives suffer from hydrolytic degradation due to the mechanism of random scissions of the main chain. PLA is known for having a longer degradation rate that can vary between 6 months and 2 years, driving many scientists to search for blends of PLA with other materials, such as nano-hydroxyapatite, with the view of maintaining the characteristics of biocompatibility and mechanical capacity yet with a shorter degradation period (Zhao et al., 2006; Mansourizadeh et al., 2014). PGA has a hydrolytic instability due to the presence of ester bonds in its molecule. Its process of degradation is erosive, occurring over two stages: first the polymer is, once again, converted into its glycolic acid monomer, in which water diffuses into the amorphous part of the molecule, cleaving the ester bonds; in the second stage, after the amorphous regions have been eroded, the crystalline portion of the molecule becomes exposed, leaving it susceptible to hydrolysis and leading to the dissolution of the polymeric chain (Gunatillake and Adhikari, 2003).

4.8 BIOCOMPATIBILITY A crucial factor for the success of biomaterials in science is the biocompatibility of these with the adjacent body of tissues and fluids. Initially, this term was defined as the capacity of a given material to exert an appropriate tissue response in the host in a specific application (Williams, 1987c). Since then, the general concept of biocompatibility has been extended over the years by tissue engineering and defined in terms of the performance or success of these materials in a specific tissue. In this context, materials are considered biocompatible if they, in their clinical applications, do not provoke harmful, toxic, or carcinogenic reactions in the patient (Ratner, 2004). With the view of confining the assessment of novel materials, this must be done in a manner where requirements, previously established through standards and directives with global reach, have to be satisfied. For this, metabolic and pathophysiologic processes that are triggered by the investigated material are assessed by experimental models in vitro and in vivo, respectively. The in vitro assays are preliminary studies that are strongly recommended for the assessment

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of basic cell function and structure. The subsequent studies (in vivo), on the other hand, assess tissue reactions to the material of study. In both assays, directives are adopted (e.g., ISO 10993-1: General Principles; Biological of Medical Devices, 2009a,b) that regulate the analysis, comprising of the characterization stage, selection of cell culture tests, and animal models. In the latter case, ethical concerns are also considered with regards to the animals used in the experiments. This standard, alongside other standards that guide the development of in vitro assays for the assessment of materials and metabolic conditions, are described in Sections 4.9.14.9.3.

4.9 TOXICITY OF POLY(α-HYDROXY ACIDS) 4.9.1 IN VITRO CYTOTOXICITY TESTS The increasing production of materials for use in the human body, be it with surgical materials or drug delivery materials, has led to an intensification of the testing and safety standards governing these materials. As such, every material with intended medical purposes must be studied prior to use, initially in vitro and later in vivo. The use of animals for biological tests is a sensitive subject and is widely discussed in society, with the existence of standards and laws that govern their utilization, and their use is restricted in many countries. As such, in the past 30 years, there has been an increase in the search for more efficient in vitro tests, with the aim of reducing animal testing. Several standards have been published as references for cell culture tests, so as to determine in vitro cytotoxicity of materials for medical use (Cannas et al., 1995; ISO 10993-5 Part 1; ISO 10993-5 Part 5). Agencies, such as the International Organization for Standardization (ISO) and the American Society for Testing and Materials (ASTM), are international entities that have established suggested rules and tests that should be carried out for the biological analysis of medical devices. The standards demand, depending on the type of material, its function, and implantation site, that different tests be carried out to scrutinize whether or not the material presents a toxic effect on the cells or tissue in question. Controlled experiments with cells are important, since they exhibit a high sensitivity, and allow for the characterization of abnormal cellular events and likewise of genetic bases and phenotypical processes that are involved in cell transformation. It is also possible to compare the behavior of normal and abnormal cell lines, in addition to the possibility of the analysis of acute and chronic toxicity, which are much more pronounced in vitro than in vivo (Goldberg, 1987; Vilela et al., 2003). With the advancement of cell biology, new techniques for the assessment of biomaterials have been proposed (Chiellini, 2006). These promising advances

4.9 Toxicity of Poly(α-Hydroxy Acids)

may assist in the reduction of the use of animal experimentation and simplify the analysis process. Standard ISO 10993-5 (ISO 10993-5 Part 1; ISO 10993-5 Part 5), describes the general principles that govern biological assessments of medical devices. This standard has as its foundation, the categorization, the nature of the material, its site and period of application in the body, in addition to the selection of suitable tests. The standard is divided into 20 parts, in which there are directives for characterization, assay selection, and technical requirements that are to be followed. Part 5 is responsible for the initial assessment of biocompatibility using a cytotoxicity assessment. The procedures for this assessment are described in international standardization format, and come with regular revisions, with the first having been published in 1993 and the the most recent in 2009 (ISO 10993-5 Part 1; ISO 10993-5 Part 5). Among the biocompatibility tests, there are general and specific in vitro tests. The former uses established cell lines that are easily cultured and reproduced, such as fibroblasts (e.g., L929 rat fibroblasts, Vero cells from african green monkey, BALB 3T3 clone 31), and the results of these tests constitute an initial sorting of the materials. The specific tests, on the other hand, use primary cells, such as the use of osteoblasts to examine the toxicity of a material that is to contact bone. These results lead to a better understanding of the in vivo performance of the material in question. The cytotoxity caused by toxic substances/materials may cause damage to cell membranes; reduction of metabolic activity; and damage to genetic material (genotoxicidade). These toxic effects can lead to a decline in energy production and protein synthesis, resulting in alterations to cell membrane mediated active transport, reduced cell proliferation, and, as a consequence, cell death. As such, the methods of assessment of in vitro cytotoxicity involve the analysis of changes in cell behavior, for example, reduction of cell adhesion; reduced cell proliferation; alteration of cell morphology; and cell death, manifested by cells rounding up and becoming detached from the substrate, in addition to an absence or change in metabolic activities as well as cell lysis (Wataha et al., 1993). For a test to be efficient, it is necessary that it simulates the in vivo situation, in a way that nutrient transport, coordination of multicellular processes, cell stability, and cell microenvironment all be present in the testing conditions, which helps to preserve the phenotypic condition of the cells. For that, phenotype and cell function are extremely dependent on the interactions with the ECM and with neighboring cells, which organize in a tridimensional manner. To reproduce the in vivo state on in vitro conditions, the use of 3D cell cultures has become common, where a scaffold with defined 3D architecture will enhance the development of the tissue (Tsang and Bhatia, 2007). However, most in vitro studies commonly carried out use simple cell lines, which do not reflect the true interaction between the different cell types that

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make up in vivo tissue. It, thus, becomes indispensable to carry out an in vivo study for the safety assessment of different materials. In in vivo tests, it is possible to analyze the interaction between different cell types, hormonal factor effects, interactions with the ECM, and likewise with proteins and molecules. Cell culture tests do present important information on the contact between these and the medical device, but for a comprehensive assessment, it still becomes necessary to complement these tests with animal experimentation tests (Cannas et al., 1995).

4.9.2 IN VITRO HEMOCOMPATIBILITY TEST Upon coming into contact with blood, a biomaterial leads to the emergence of complex phenomena that may trigger the destruction of red blood cells, denaturation of proteins, and thrombus formation. Of these, blood clot formation is the most important, as its presence may trigger a sudden loss of blood flow to vital organs (Williams, 1987a,b,c). Blood-surface interactions taking place in the polymeric material are governed by different factors: surface texture, electrostatic effects, minimum interfacial free energy, ratio of polar to nonpolar phases, and others (Arranz and Chaves, 1989). For a better understanding of the interaction of blood with biomaterial, it is necessary to recapitulate the composition of blood. Blood is considered a specialized connective tissue, composed of a suspension of cells in an aqueous medium called plasma. Within this latter component there are proteins, such as albumin and hemoglobin, electrolytes, micronutrients, and hormones and substances responsible for coagulation. Blood cells are divided into white and red blood cells, where it is the red blood cells, or erythrocytes, that are responsible for the oxygenation of tissues and elimination of carbon dioxide. Of the cell elements, platelets carry out an important role in the coagulation process. When they are not activated, these platelets present themselves in the shape of a disc and with a diameter of aproximately 23 μm. They possess a complex platelet membrane with numerous receptors for interaction with key proteins contained in plasma. In their interior, a series of granules (mitochondria, dense granules, A granules, and others) are found, alongside a multitude of proteins that control aggregation capacity and the interaction of platelets with other structures (Hanson and Ratner, 1996). The hemolysis test determines the degree of red blood cell (erythrocyte) lysis in blood. The main resulting effect of the interaction of a material with red blood cells, is the accelerated aging or premature mechanical destruction of these cells, culminating in the release of hemoglobin (Almeida, 2000). As for the kinetics of the formation of thrombi on the polymeric surface, an assessment is taken to determine the time it takes for the material to provoke blood coagulation, resulting in the formation of thrombi (composites of fibrin and adhered platelets). Short or very long coagulation times suggest that the material

4.9 Toxicity of Poly(α-Hydroxy Acids)

is activating or inactivating the coagulation system, which serves as evidence for lack of blood compatibility (Almeida, 2000).

4.9.3 IN VIVO BIOCOMPATIBILITY TESTS As previously described, biomaterials have several applications, and for this reason there are a number of in vivo biocompatibility tests available, each one tied to the site and function of the different biomaterials.

4.9.3.1 General tests for bone implants The results encountered in studies related to osseointegration, have enabled the development of new techniques that seek the perfection of rehabilitation with osseointegrated implants (Zanetti, 2008). Osseointegration is defined as an anatomical and functional amalgamation between remodeled bone and the implant surface, where the efficacy of the implant essentially depends on the bioactivity of the material, that is, on its capacity to establish a mechanically solid interface without the presence of fibrous tissue and with fully unified material surface and osseous tissue (Machado, 2008). Osteoinduction is the capacity of inducing the host to produce new bone tissue. Most studies in the area of biomaterials, aim to find bone substitutes and are directed toward a material that may present a good osteoinductive potential, or, alternatively, a good osteoconductive potential, whereby the material undergoes bioabsorption in a short amount of time (Zanetti, 2008). For that purpose, lactic acid and glycolic acid-based biomaterials are highlighted here, which are used by themselves or, more frequently, are associated to become copolymers, such as those of PLA/PGA. These substances are widely used in the make-up of surgical devices (Zanetti, 2008). Factors such as the low density of the PLA/PGA copolymer cause its biodegradation to occur slowly, and simultaneous bone deposition may occur (Queiroz et al., 2008). The biodegradation of some of these copolymers occurs by hydrolysis and the length of time this takes is density dependent, leaving a high percentage of tissue in place (Serino et al., 2003). In general, in vivo biocompatibility tests aiming at the analysis of the mentioned factors require a protocol for the use of animals, the majority with rats or rabbits, in which, through the preparation of bone defects and insertion of the implant, the interaction of the biomaterial can be measured using laboratory assays. The most used tests for these evaluations are histocompatibility studies with immunohistochemistry using antibodies against the proteins RANK L, osteoprotegerin (OPG), and osteocalcin (OC). These testes have the objective of obtaining markings for qualitative and semiquantitative analysis that would represent the expression of these proteins characterizing reabsorption and bone neoformation. The identification of the mechanisms of regulation and cellular growth of osteoblasts, osteoclasts, and osteocytes, among other cells involved in the process of

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bone remodeling, have made the use of these techniques useful for the evaluation of events during the bone repair process (Zanetti, 2008).

4.9.3.2 General tests for stents In 2002 and 2003, in the United States, a number of pharmacological coronary artery stents made of biopolymers were approved for clinical use and fueled a great boost in the percutanous treatment of coronary stenoses, by the pronounced reduction of the incidence of intra-stent restenosis and the need for a new revascularization of the target vessel (Morice et al., 2002; Stone et al., 2004). However, the manifestation of late and very late thrombosis events, after the implantation of these prostheses, was a danger signal as to the safety of these pharmacological stents. With the view of overcoming the limitations and adverse events related to these first generation stents, new pharmacological stents were developed. These new stents presented variations in their metallic alloy composition, the thickness of the rods, the design of the meshes, the drug carrier polymer, and the drug class and dose or the site of its release at the stent surface. Tests done for the in vivo assessment of biopolymer stents, including stents of PLA, PGA, and composites of these, are generally carried out in pig models and require histomorphological studies for the analysis of inflammation, fibrin deposition and degree of injury, intracoronary ultrasound with planar analysis of luminal areas, stent areas, neointimal hyperplasia areas, percentage of neointimal hyperplasia areas and quantitative coronary angiography for the assessment of minimum luminal diameters (MLD), reference diameters (RD), and percentage of angiographic stenosis and late luminal losses.

4.10 APPLICATIONS OF POLY(α-HYDROXY ACIDS)—PLA AND PGA The market for biomaterials is significant, considering both the number of annually commercialized units and the commercial turnover, and may be readily segmented based on two distinct criteria. The first relates to the types of composites used in the biomaterial composition, such as metals, ceramics, polymers, and natural origin materials. The second criterion is based on the application area of the biomaterial, for example, for orthopedic, cardiovascular, odontological or ophtalmic use, or in plastic surgery, tissue engineering, injury treatment, or neurological and central nervous system disorders. In addition, this also includes medical devices with other applications, such as gastrointestinal and urinary uses, or as drug delivery systems and for bariatric surgery (Pires et al., 2015).

4.10 Applications of Poly(α-Hydroxy Acids)—PLA and PGA

Specifically in the area of tissue engineering, the most widely used synthetic polymers are biodegradable polymers, such as saturated aliphatic polyesters, including PLA and PGA, as well as PLGA (Serino et al., 2003). These composites undergo hydrolytic degradation through ester hydrolysis. Upon degradation, the monomeric components of each polymer are removed through natural pathways, and the organism presents highly regulated mechanisms to fully remove, for example, monomeric lactic acid and glycolic acid components (Wataha et al., 1993). Due to these properties, in addition to use in tridimensional scaffold fabrication for cell proliferation, these polymers have been utilized as biodegradable suture, absorbable bone fixation devices, and matrices for drug delivery (Pires et al., 2015).

4.10.1 NONMEDICAL APPLICATIONS OF POLY(α-HYDROXY ACIDS)—PLA AND PGA PLA is a biodegradable and bioactive plastic derived from renewable sources, such as sugar cane and corn, and, in 2010 it became the second most used biodegradable plastic in terms of global volume (Gupta et al., 2007). Its characteristics of biodegradability, where the time of degradation is 6 months to 2 years, and its natural origin have spurred the use of this biomaterial for different industrial areas. Its varied processing methods, such as extrusion, 3D printing, film formation, and others, give rise to different uses of PLA as a biodegradable plastic material. Some of the most common uses of PLA are for the production of plastic films, glasses, bottles, kitchen utensils such as plates and forks, as well as for feminine hygiene products and diapers. PLA may also be used as a biomaterial for 3D printing for the production of a multitude of objects, just as it may be used for the creation of scaffolds for use in tissue engineering, by providing a temporary base for tissue regeneration whilst being absorbed by the tissue without causing toxic effects in the tissue. Not only may PLA be used as a scaffold, it may also be used as a medical device, in the form of pins, screws, plates, and suture thread. High molecular weight PGA, on the other hand, is commercialized for use as a food packaging material, providing a protective barrier for perishable foodstuff that would lose their freshness upon air exposure. Bottles of thin wall thickness may also be made using high molecular weight PGA, but that still maintain their barrier properties (Schmitt, 1973).

4.10.2 MEDICAL APPLICATIONS OF POLY(α-HYDROXY ACIDS)—PLA AND PGA Reabsorbable biopolymers, such as PLA, have been used not only as biodegradable plastics, but also as biomedical and pharmaceutical materials since the 1960s, thanks to their excellent mechanical properties and the fact that they are

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naturally hydrolysable in the human body (Reneker and Chun, 1996). Initially, they were used for surgical suture, where Kulkarni and co-workers in 1966 used PLLA in the form of plates and screws for bone fracture fixation. Since then, due to its biocompatibility and degradation rate, they have been widely studied and utilized in the orthopedic field and for oral fixation devices, especially due to the high mechanical strength present in a high molecular weight PLLA (Jahno, 2005). Commercial polymers are available between the range of 2,000 and 300,000 Mw (Jahno, 2005), where high molecular weight polymers are used for orthopedic surgery and low molecular weight polymers for drug delivery carriers. In pharmaceutical applications, low molecular weight lactide copolymers are preferred, since they are more rapidly degraded in the body than high molecular weight PLLA. Their application as scaffolds and other medical devices, such as rods, plates, screws, fibers, membranes, sponges, granules, and microspheres has also become more widespread, especially in the area of tissue engineering. With the advent of 3D printing, PLA gained importance as a raw material for material printing both within and outside the medical sector. Since the 1980s, a decade that saw the first industrial 3D printing production systems become available, this technique has evolved to the point of having a prominent position also for the health sector. Plastics and metals are now being used to create personalized replicas of organs or skeleton parts, which assist in precision surgery, surgical guides to indicate incision and insertion sites, bone substitution implants, or scaffolds for the formation of organs, as well as some types of prostheses. The natural glass transition temperature (Tg) of PLA is around 65
C, but normally the polymer is mixed with other plastics to make it more suitable for 3D printing. For low-cost printing applications that do not require biocompatibility, PLA is highly recommended for its low cost compared to other raw materials and for the environmental protection aspect, as PLA possesses a much shorter degradation time than other polymers. For the medical sector, many studies are still being carried out for the verification of the adaptation of cells on the surface of 3D printed PLA scaffolds, as well as for the hemocompatibility and bioactivity of these scaffolds. PGA was developed as the first absorbable synthetic suture, receiving the name Dexon, by the company American Cyanamide DuPont in 1962 (Gunatillake and Adhikari, 2003). Suture made of PGA is naturally degraded in the human body through hydrolysis over a period of 6090 days. Its properties include high mechanical strength, ease of use, ease of slipping through tissue, and a great ability to make knots, facilitating its use. It is commonly used for subcutaneous sutures and for thoracic and abdominal surgeries. The use of PGA as a suture material has opened up space for its use in other medical fields, such as for implants, plates, pins, screws, and anastomosis rings. It has also been used for controlled drug delivery (Middleton and Tipton, 2000). PGA has also been widely studied by tissue engineering, for use as a biodegradable scaffold, aiding the tissue recovery process.

4.11 Future Trends in Biofabrication

4.11 FUTURE TRENDS IN BIOFABRICATION 4.11.1 ELECTROSPINNING Electrospinning is a technique commonly used in tissue engineering for the manufacture of products of micrometric and nanometric origin. Nanofibers are considered as fibers with diameters ranging from 10 to 2000 nm, in which different alignments of the electrophilic fibers lead to the formation of unique functional structures, such as nanotubes and nanowires, as well as blankets with high porosity and connectivity for use as scaffolds (Li and Xia, 2004; Lannutti et al., 2007; Pham et al., 2006). The principle of the process is relatively simple: a polymer solution is injected through a capillary, forming a droplet that is induced in an electrostatic field, thus, forming a jet (Taylor cone). The jet moves through the air leading to the evaporation of the solvent, and upon reaching a metal target forms nanofibers which have an adjustable diameter (Reneker and Chun, 1996; Shin et al., 2005). This technique is extremely interesting for tissue engineering, since nanofibers having a high area to volume ratio, a highly interconnected pore network, and fibers with diameters that mimic the dimensions of ECM proteins, such as collagen, lead to increased cell proliferation, migration, and differentiation. The control of the diameter of the nanofibers is achieved through the variation of the concentration of the dissolved polymer, as well as the applied load between the capillary and the metallic plate collector. Other factors influencing the formed nanofibers include the molecular weight of the polymer that has been dissolved, as well as the distribution of their polymer chains; the viscosity and conductivity of the polymer solution; the difference of the electric potential applied during the electrospinning process; the distance between the capillary and the collector plate (allowing solvent volatility); and the time of the electrospinning process. Another parameter is the ability to control the deposition of the fibers in the collector plate, leading to different uses of the manufactured nanofibers (Smith and Ma, 2004; Yang et al., 2006). The conventional electrospinning method leads to the formation of microfibers and nanofibers that are randomly oriented. In order to control the deposition pattern of the fibers, a cylindrical collector with controlled rotation can be used to guide the deposition of the fibers in a single direction. Another way of controlling the deposition of the fibers is through use of a collector plate with a specific shape in order to control the areas of fiber deposition, which can create regions of higher and lower density of fiber deposition (Vaquette and Cooper-White, 2011). The creation of a complex fiber deposition design, with a pattern to guide cell growth, is fundamental to obtaining an interaction between the cultured cells and their ECM proteins, thus, increasing the exchange of nutrients and leading to a greater proliferation. The manufacture of a three-dimensional scaffold constituted of nanofibers with a functional biodesign,

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is possible by modifying the traditional metal collecting plate on a plate with the desired design that the fibers should form. Thus, the fibers will use the plate pattern to guide their deposition, both in morphology and in density (Macedo et al., 2015). One of the major challenges of traditional electrospinning is the creation of a 3D scaffold, which allows a fiber deposition with a homogeneous interconnectivity between structures, which when absent stops uniform tissue growth. Variations in the diameter of electrophilic fibers is also a problematic factor, since it can either help or prevent cell growth depending on the diameter. Controlling these parameters, as well as a customized design for each type of tissue, is fundamental to the efficiency of the scaffold for tissue regeneration. The development of a porous, customized scaffold with an interconnected pore network and a micrometric spacing between fibers that allows a more efficient proliferation of cells, is one of the objectives of the use of electrospinning for tissue engineering.

4.11.2 3D BIOPRINTING RAPID PROTOTYPING 3D bioprinting rapid prototyping (RP) is a newly developed technique for fabricating well-designed tissue engineering scaffolds. RP uses a computer-aided design model to create a 3D project with detailed control of morphological characteristics, chemical composition, and mechanical properties. Then, the project is built through an additive process where successive layers of the material are printed. This method allows for the manufacture of extremely reproducible scaffolds, which allows for the customization of size and shape according to specific requirements (Yang et al., 2002; Yeong et al., 2004; Yang et al., 2002; Peltola et al., 2008). There are different additive processes for 3D printing available, such as selective laser sintering, stereolithography, and fused deposition modeling (Li et al., 2014). The 3D printing technique is used for the manufacture of biopolymer-based scaffolds for use in tissue engineering, and recently, living cells and growth factors were embed into hydrogel-type scaffolds during fabrication, to enhance tissue regeneration because of its similarity with the ECM and the in vivo environment (Murphy and Atala, 2014).

4.11.3 BIORESPONSIVE HYDROGELS Hydrogels are polymeric materials used in various medical applications with a high versatility degree of chemical and physical properties. They have a 3D network with crosslinked form containing hydrophilic polymer chains, which provide swelling in water. Hydrogels are also known for being biocompatible with low immunogenicity and allow a fast mass transfer between cells and the surroundings (Kuo and Chang, 2013). They are used as scaffolds in tissue engineering (Chang et al., 2013), controlled drug delivery, (Casolaro et al., 2012) and it may provide the initial structural support required to retain cells in the defective area

4.11 Future Trends in Biofabrication

for cell metabolism, growth, differentiation, and new matrix synthesis (Remya and Nair, 2013). Bioresponsive hydrogels are active systems with controllable physical and biochemical properties that can respond to stimulus through natural biological processes (Wilson and Giuseppi-Elie, 2013). When exposed to specific conditions in vivo, such as changes in pH, temperature, or a biological target (enzyme, growth factor, antibody, etc.) molecular recognition events trigger changes in molecular interactions that translate into macroscopic responses, such as swelling/ collapse, degradation, mechanical deformation, or solution-to-gel transitions (Ebara et al., 2014). Bioresponsive hydrogels have been used in drug delivery, for drugs like insulin for instance (Albin et al., 1985; Rehor et al., 2005); as biosensors, such as a glucose sensor for patients with diabetes mellitus (Holtz and Asher, 1997); and in tissue engineering, to create a hydrogel scaffold that permits cell migration by mimicking the ECM (Lutolf and Hubbel, 2005).

4.11.4 BIOPOLYMER COMPOSITES IN TISSUE ENGINEERING Biopolymer composites are important in tissue engineering since they provide an advantageous environment for the growth, migration, and differentiation of cells. They play an essential role in tissue engineering through cell seeding and proliferation and the development of new tissue in three-dimensions, therefore, they present great potential in tissue engineering research (Ninan et al., 2015). Biopolymer composites are constituted of two or more materials and may have significant improvements of mechanical properties, biocompatibility, and biodegradation (Park et al., 2016; Swetha et al., 2010). Scaffolds made of biopolymers composites can be obtained using several manufacturing techniques, like solvent casting and particulate leaching, thermally induced phase separation, emulsion freeze-drying, electrospinning, and RP (Park et al., 2016). These materials have several applications in the medical field. They can be used in the regeneration of skin, bones, cartilage, vascular grafts, and other organs (Park et al., 2016). One area of biocomposites that has gained great attention in past years is nanocomposites (Okamoto and John, 2013). In this type of material, at least one of its constituent phases (usually the load) must have, at least one of its dimensions, in a nanoscale size (Raquez et al., 2013). Researchers had been working to obtain PLLA nanocomposites using montmorillonite. The main routes of synthesis to obtain these materials are: solvent intercalation, melt intercalation, and in situ intercalation (Raquez et al., 2013). In the case of PLLA nanocomposites, studies indicate improvements in mechanical properties, such as the gas barrier and an increase in the biodegradation rate of the polymer (Raquez et al., 2013).

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4.12 CONCLUSIONS In the past few years, there has been significant growth in the development of materials that could be used for tissue engineering and in regenerative medicine. Bioabsorbable polymers, such as PLLA, PLGA, and its copolymers, can be used in different areas and present satisfactory results in the improvement of the quality of life of patients. The development of composites as well as new biofabrication techniques can result in significant advances for the use of biomaterials and be a gain to society, however, this will require a lot of effort and investment in research in the future.

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FURTHER READING ISO 10993 Biological evaluation of medical devices Part 1-20. Shalaby, W.S., Burg, K.J.L., 2003. Absorbable/biodegradable polymers: technology evolution. CRC Press. Tokiwa, Y., Calabia, B.P., Ugwu, C.U., Aiba, S., 2009. Biodegradability of plastics. Int. J. Mol. Sci. 9 (9), 37223742.

CHAPTER

Technological challenges and advances: from lactic acid to polylactate and copolymers

5

Luciana Fontes Coelho1, Susan Michelz Beitel1 and Jonas Contiero1,2 1

Department of Biochemistry and Microbiology, Institute Bioscience, Sa˜o Paulo State University (UNESP), Sa˜o Paulo, Sa˜o Paulo, Brazil 2Associate Laboratory IPBEN-UNESP, Rio Claro, Sa˜o Paulo, Brazil

5.1 LACTIC ACID Lactic acid (2-hydroxypropanoic acid), CH3CHOHCOOH, is an organic acid that has been extensively used worldwide in a variety of industrial and biotechnological applications. The existence of an asymmetric carbon in the alpha position of the acid function is the source of two enantiomeric forms of this molecule, called L(1) and D( ) (Fig. 5.1) (Ghaffar et al., 2014; Sodegard and Stolt, 2002). Lactic acid may be obtained chemically or by microbial fermentation. Production by fermentation results in the formation of D( ) or L(1) lactic acid, or a racemic mixture, depending on the microorganism used. Homofermentative methods are preferred for use in industrial production, since this pathway results in high product yield and low byproduct formation (Mehta et al., 2007). Lactic acid synthesized by most bacteria occurs in two forms of stereoisomers, however, only a few have the homofermentative character. In the last stage of lactic fermentation, pyruvic acid is converted to lactic acid by the enzyme lactate dehydrogenase (LDH) which is NAD 1 dependent. Each isomer consists of a specific LDH; L-lactate dehydrogenase (L-LDH) is related to the conversion of the L(1) isomer, whereas D( ) lactic acid is produced by D-lactate dehydrogenase (Stock et al., 1997). Some bacteria present the lactate racemase (LR), responsible for the conversion of isomer D( ) to L(1) and vice versa. In addition, the enzymes responsible for determining the lactic acid isomerism produced are expressed at different levels depending on the genetic characteristics of the microorganisms (Garvie, 1980). The purity of the isomers L(1) and D( ) lactic acid is the main factor determining the physical properties of PLA, different types of lactic acid polymers are formed. The properties of PLA depend on the proportion of enantiomers, which allows the production of polymers with different characteristics aimed at specific Materials for Biomedical Engineering: Hydrogels and Polymer-based Scaffolds. DOI: https://doi.org/10.1016/B978-0-12-816901-8.00005-5 © 2019 Elsevier Inc. All rights reserved.

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FIGURE 5.1 Chemical structures of lactic acid enantiomers (Shi et al., 2013).

industrial applications (Auras et al., 2003). PLA with high crystallinity and a high melting point that is suitable for fiber production can be produced from pure isomers of L(1) lactic acid and D( ) lactic acid, but not from racemic isomers (mixture of L(1) and D( ) isomers). Therefore, many studies aim to obtain the pure form of D( ) (Tanaka et al., 2006; Lu et al., 2009; Beitel et al., 2017) or L(1) isomers (Wee et al., 2006b; Lima et al., 2010; Coelho et al., 2018). The isomeric purity of lactic acid is influenced by the culture medium of the microorganism (Madhavan Nampoothiri et al., 2010). Some studies involving metabolic engineering work on directing the production of only one isomer by silencing a specific LDH in order to obtain high purity lactic acid production (Okano et al., 2010; Assavasirijinda et al., 2016; Awasthi et al., 2018). To solve this problem, bacteria have been engineered to increase the chemical and optical purity of lactic acid. Native Lb. helveticus produces a racemic mixture of L(1) and D( ) lactic acid. By removing the promoter region of the gene for D ( ) lactate dehydrogenase, Lb. helveticus produced only the L(1) isomer (KylaNikkila et al., 2000). In Escherichia coli (Mazumdar et al., 2010), genes encoding enzymes that catalyze the conversion of pyruvate to succinate, acetate, and ethanol were inactivated, resulting in increased lactate production and chemical purity. Lactic acid bacteria are able to ferment sugar through different routes resulting in either a homo- or heterofermentative process. The homofermentative pathway only results in lactic acid as the end product of glucose metabolism via the Embden Meyerhof Parnas pathway (Fig. 5.2A), whereas in the heterofermentative route equimolar amounts of lactic acid, carbon dioxide, and ethanol or acetate are formed from glucose via pentose-phosphate (Fig. 5.2). Although large-scale fermentation technology for the production of L(1) lactic acid has been well established, the processes related to economical and efficient production of D( ) lactic acid on an industrial scale still have to be improved. According to Liu et al. (2014), to obtain relevant information related to largescale production processes of D( ), it is necessary to demonstrate an efficient

5.1 Lactic Acid

(A)

Glucose

(B)

Glucose

1 ATP

Fructose-6P

Glucose-6P

1 ATP

Fructose-1,6DP

4 ATP 2 NADH

2 Pyruvate

2 NADH

2 Lactate

Xylulose-5P

Ethanol

2 NADH Dihydroxyacetone-P Glyceraldehyde-P AcetyI-P

Glyceraldehyde-P

LDH

1 NADH

CO2

2 ATP 1 NADH Pyruvate 1 ATP 1 NADH LDH Acetate Lactate

FIGURE 5.2 Catabolic pathways of lactic acid bacteria: (A) Homofermentative; (B) Heterofermentative. Adapted from Hofvendahl, K., Hahn Ha¨gerdal, B., 2000. Factors affecting the fermentative lactic acid production from renewable resources. Enzyme Microb. Technol., 26, 87 107.

pilot scale fermentation using a promising microbial strain with efficient culture medium conversion and low cost pH neutralizers. In the early 1960s, chemical synthesis of lactic acid was developed, considering the interest in the baking industry. Manufacturing of synthetic lactic acid on a commercial scale began in 1963 in Japan and the United States (Holten and Mu¨ller, 1971; Vickroy, 1985). In chemical synthesis, lactonitrile is produced by the combination of hydrogen cyanide and acetaldehyde, in the presence of a catalyst, in the liquid phase. Then, lactonitrile is hydrolyzed to lactic acid by sulfuric or hydrochloric acid. In this process, ammonium chloride is generated as a byproduct (Pal and Dey, 2012). However, chemical synthesis results in a mixture of D( ) and L(1) lactic acid, moreover, it is not an economically viable process (Wee et al., 2006a). Compared to chemical synthesis, the biotechnological process for the production of lactic acid offers several advantages, such as low substrate cost, and low temperature and energy consumption (Datta and Henry, 2006). Fig. 5.3 shows a diagram with the main steps in the production of lactic acid by chemical and fermentative routes. The number of investigations involving the production of lactic acid by fermentation has increased exponentially since the 1990s (Fig. 5.4). Lactic acid occupies the third position among the 30 most studied molecules of 2016, following C6 sugars and polyhydroxyalkanoates (Biofuelsdigest, 2016).

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FIGURE 5.3 Lactic acid production methods: (A) Chemical synthesis and (B) Microbial fermentation. Adapted from Wee, Y.J., Kim, J.N., Ryu, H.W., 2006a. Biotechnological production of lactic acid and its recent applications. Food Technol. Biotechnol., 44, 163 172.

800 700

Number of publications

120

600 500 400

July 2016 300 200 100 0

1990

1992

1994

1996

1998

2000

2002

2004

2006

2008

2010

2012

2014

2016

Year

FIGURE 5.4 Number of publications related to the production of lactic acid via fermentation. With permission from Biofuelsdigest. The 30 Hottest Molecules of 2016: sneak preview. Available , http://www.biofuelsdigest.com/bdigest/2016/01/04/the-30-hottest-molecules-of-2016-Sneak . (accessed July 2016).

5.1 Lactic Acid

FIGURE 5.5 Global Market Forecast for Lactic Acid, 2014 22 (US$ Million).

The world production of lactic acid in 2006 was estimated to be between 130,000 and 150,000 metric tons/year, with a 19% annual growth, mainly due to its use in the production of biodegradable polymers such as PLA (Wee et al., 2006a). In 2009, world production was 258,000 metric tons (Global Industry Analysts INC, 2011). The largest producer, NatureWorks, presented a production of 140,000 metric tons of PLA in 2011. By 2013, the world production of lactic acid (D, L, and DL) was 284,000 metric tons (Dammer et al., 2013). A 329,000tonne growth was projected for 2015 (Global Industry Analysts INC, 2011), and the estimated lactic acid production for 2017 is 367,300 metric tons (AbdelRahman et al., 2013). The market forecast for lactic acid by Credence Research (2016) presented an annual estimate of growth from 2014 to 2022, based on public domain information reported by some producer companies (Fig. 5.5).

5.1.1 FACTORS THAT INFLUENCE LACTIC ACID PRODUCTION The use of a microorganism in a biotechnological process generally involves the production of several molecules. However, species found in nature normally do not produce large amounts of metabolites, only the necessary amount for survival. In this way, several works have been developed in order to improve on this level of production, aiming at industrial scale application (Parekh et al., 2000). Operational parameters are controlled in order to optimize the fermentation process. Various factors influence lactic acid production by acid-lactic bacteria, such as pH (Mussatto et al., 2008; Andersen et al., 2009), temperature (Tango and

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CHAPTER 5 Technological challenges and advances

Ghaly, 1999), aeration (Okino et al., 2008), carbon source (Bulut et al., 2004), nitrogen source (Nancib et al., 2001; Altaf et al., 2007), mineral salts (Hauly, 2001), and vitamins (Xu et al., 2008), among others. Microorganisms present determined optimum pH and temperature ranges for their growth, which vary according to species or strain. Studies have shown that when pH is controlled during fermentation, lactic acid production is improved (Mussatto et al., 2008). Extracellular pH is one of the main factors related to the production of lactic acid, presenting an influence on the catalytic activity of enzymes. The optimal pH for lactic acid production varies between 5.0 and 7.0; Lactobacillus strains present an optimum pH of around 5.7 and are known as tolerant acid. Fermentation is strongly inhibited by low pH and ceases at a pH below 4.5 (Kashket, 1987; Silva and Mancilha, 1991; Panesar et al., 2007). Many bacteria are sensitive to low pH, and this sensitivity has contributed to food preservation and industrial fermentation manipulation (Miller and Wolin, 1981). pH even has an effect on the fermentation pathway of microorganisms, as reported by Russell and Hino (1985), who, using Streptococcus bovis, showed a production of lactic acid, acetate, and ethanol at neutral pH, but at low extracellular pH showed homolactic fermentation. The same was reported by Stokes (1949) while studying the behavior of E. coli, where it was found that a low pH decreased the production of acetate, formate, and ethanol, however, it increased the production of lactic acid. A pH change may also influence cell growth, as reported by Fiedler et al. (2011); in their research they observed that a change in pH is related to the yield of cell mass produced per mole of ATP in Streptococcus pyogenes and Lactobacillus lactis, while higher yields were verified at pH close to the natural habitat of the microorganisms. Thus, lactic acid produced during fermentation has to be constantly neutralized. To control pH during fermentation, some neutralizing agents are added to the fermentation medium, such as calcium carbonate, which is the most commonly used with agitated bottles and in reactors (Yen et al., 2010). However, there are some problems associated with the addition of agents and neutralization toward pH control, because lactate instead of lactic acid is formed at a high pH value, which will result in an increase in purification costs due to the need for the recovery of lactic acid. In addition, when calcium carbonate or calcium hydroxide are used, calcium sulfate can be produced in the process of converting lactate to lactic acid, which can cause considerable environmental problems due to the formation of gypsum, as well as to extra recovery costs. The recovery and purification of lactic acid accounts for 50% of production costs (Eyal and Bressler, 1993). In order to avoid or minimize the use of neutralizing agents, genetically improved strains that are resistant to acidic environments and are lactic acid producers at low pH may be useful. Engineering strains to increase the growth and production of lactic acid at low pH will reduce the formation of residue and reduce the cost with lactic acid production and purification

5.1 Lactic Acid

Acid stress and end-product inhibition are among the main challenges in lactic acid production. Therefore, more studies on metabolic engineering for the identification of genes and proteins associated with stress responses and tolerance could improve biosynthetic pathways. Mainly using yeast and fungi, once these microorganisms are more acid tolerant than lactic bacterium has been an option (Upadhyaya et al., 2014). Zheng et al. (2010) improved the acid tolerance, as well as the production of D( ) lactic acid of Sporolactobacillus inulinus ATCC 15538, using ultraviolet irradiation, diethyl sulfate mutation, and protoplast fusion. They obtained the recombinant F3-4, which produced 119% more D( ) lactic acid (93.4 g/L) at pH 5.0 compared to the original strain. Temperature is an environmental factor, of which any variation has an effect on the course of fermentation; reactions catalyzed by enzymes are especially sensitive to small changes in temperature (Merritt, 1966). Different species of microorganisms present different optimum temperatures, which are related to cell growth and lactic acid production, so it is convenient to study the appropriate temperature in the fermentation process for each case (Peleg, 1995). When a microorganism is submitted to temperatures different than those considered ideal, inhibition of growth occurs and in some cases destruction of products (Scheper, 2000). The agitation is related to the homogenization of the medium, dissipation of the heat produced by the metabolic reactions, heat transfer to the temperature control, as well as to minimize cell death resulting from the addition of concentrated acid and base for pH control (Charles, 1985; Scheper, 2000).

5.1.2 CULTURE MEDIUM FOR LACTIC FERMENTATION: ALTERNATIVE SOURCES OF CARBON AND NITROGEN Social behaviors and consumption habits culminate in social, environmental, and economic problems, which make sustainability a topic often discussed, drawing the attention of governments, organizations, and the academic community (Malhotra et al., 2013). Concern for the environment has stimulated researchers worldwide to develop methods to produce a wide range of molecules and materials using sustainable technology (Abdel-Rahman et al., 2013). Commercial culture media used for the growth of fastidious microorganisms, such as lactic acid bacteria, are not economically attractive as high nutrient-rich culture media with high-cost carbon source (glucose), nitrogen (yeast extract), amino acids, and mineral salts are needed (Kyla-Nikkila et al., 2000). The use of refined sugars (glucose) as feedstock increases the cost of lactic acid production. The cost of raw materials is between 40% and 70% of the total cost of production and yeast extract, used as a source of nitrogen in the fermentation medium, accounts for 38% of the total cost of production (Tejayadi and Cheryan, 1995). The cost of purification is directly related to the complexity of the culture medium.

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CHAPTER 5 Technological challenges and advances

An alternative is the use of alternative substrates, such as molasses (Wee et al., 2004), sugarcane juice (Calabia and Tokiwa, 2007; Farooq et al., 2012), sugar beet (Kotzamanidis et al., 2002), hydrolyzed soybean oil (Kwon et al., 2000), starch (Yumoto and Ikeda, 2004, Ohkouchi and E Inoue, 2006), rice starch (Fukushima et al., 2004), soybean oil waste (Yumoto and Ikeda, 2004), and the water of cassava (Coelho, et al., 2010). As well as several agricultural products and wastes (Narayanan et al., 2004, Zhang et al., 2007, Ezejiofor et al., 2014). The complex alternative sources used to reduce production costs have substances or physical characteristics that may influence the production of lactic acid and even the growth of the microorganism used (Ohkouchi and E Inoue, 2006). Studies have reported the production of lactic acid through biomass conversion (Tokuhiro et al., 2008). However, to make sugars from biomass accessible, the chemical pretreatment and enzymatic hydrolysis of lignocellulosic biomass, or enzymatic saccharification of amylaceous biomass (Li and Cui, 2009) are required. In traditional pretreatments of lignocellulosic biomass, chemicals such as acids and bases are added, which will influence the final stage of separation and purification of the product, besides increasing the cost of the process (Hofvendahl and Hahn Ha¨gerdal, 2000). Due to the high demand for the use of more economical and renewable resources, metabolic engineering appears to be a revolutionary tool for the development of strains that can use the residual lignocellulosic biomass (pentose) and that are resistant to the inhibitory compounds formed during the pretreatment of biomass. However, there are some preferred wastes that do not need biomass pretreatment, and that, therefore, do not form inhibitory compounds (Upadhyaya et al., 2014). The cost of producing lactic acid can be significantly reduced by using the byproducts of other processes, such as whey or sugarcane molasses, which contain readily fermentable sugars (Dumbrepatil et al., 2008). Whey is a byproduct of cheese production, and is rich in proteins, fats, and lactose with excellent functional, nutritional, and technological properties (Pelegrine and Carrasqueira, 2008). Molasses is a viscous material considered as a byproduct of sugar production, composed of sucrose, glucose, and fructose, with a total carbohydrate concentration of 45% 60% (Mariam, et al., 2009). Among the alternative sources of nitrogen reported are: peanut flour (Wang et al., 2011), soybean meal (Kwon et al., 2000), corn steep liquor (Yu et al., 2008), and yeast autolyzed (Lima et al., 2009, Coelho et al., 2011), among others (Tang et al., 2016). Proflo is a fine yellow powder produced from cotton seeds. It is rich in protein (56% w/v), and contains 24% carbohydrate, 5% oil, and 4% ash, the latter being rich in calcium, iron, and phosphorous chloride (Okafor, 2007). Corn steep liquor is an extremely economical alternative to nitrogen sources such as yeast and peptone extract. This residue is a byproduct from corn starch processing (Wee et al., 2006b), with nitrogen being available as amino acids, such as alanine, arginine, aspartic acid, cystine, glutamic acid, histidine, isoleucine, leucine, lysine, methionine, phenylalanine, proline, threonine, tyrosine, and valine, with some B vitamins also present (Cardinal and Hedrick, 1948).

5.1 Lactic Acid

To solve the problem of high costs due to the high nutritional requirement of lactic bacteria, Corynebacterium glutamicum, which can produce lactic acid in minimal medium, was engineered to express the D( ) lactate dehydrogenase gene from L. bulgaricus with .99.9% optical purity (Jia et al., 2011). Due to its high availability and reduced cost, glycerol could be an ideal carbon source for the production of lactic acid. With this in mind, some authors have been investigating alternatives for the production of D( ) lactic acid and L(1) lactic acid from glycerol. Mazumdar et al. (2013) performed metabolic engineering on E. coli for the microaerobic production of L(1) lactic acid from glycerol, through knockout of the genes responsible for the synthesis of succinate, acetate, and ethanol, and overexpression of the glycerol dissimilation respiratory pathway (GlpK/GlpD). The modified strain produced 50 g/L of L(1) lactate from 56 g/L of crude glycerol. Similar experimentation was conducted by Mazumdar et al. (2010), but the metabolic pathway was directed toward the production of D( ) lactic acid, in which case E. coli produced 32 g/L of D-lactate from 40 g/L of glycerol. Excess substrate in fermentation, such as carbon and nitrogen source, can inhibit the production of lactic acid by osmotic stress and catabolic repression. In general, the commercial production of lactic acid requires the use of a robust and nonfastidious microorganism that has the ability to ferment both pentose and hexose, does not present inhibition by the substrate and the final product, is resistant to inhibitory compounds formed during the pretreatment of lignocellulosic biomass, and minimizes catabolic repression (Hofvendahl and Hahn Ha¨gerdal, 2000).

5.1.3 PRODUCTION OF LACTIC ACID BY FERMENTATION Lactic acid can be produced by different methods of fermentation, such as submerged or solid-state fermentation, continuous fermentation with simultaneous saccharification, continuous processes with cellular recirculation, batch fermentation, or fed batch fermentation, among others (Marques et al., 2008, Coelho et al., 2011). Fed batch fermentation is a strategy to prevent inhibition by excess substrate and catabolic repression, resulting in higher production, productivity, and cellular growth rate (Bernardo et al., 2016; Marques et al., 2016; Hu et al., 2015). Among the different types of bioreactors used in biotechnological processes, the most used is the stirred tank in submerged processes, corresponding to 90% of the total of bioreactors used industrially. The first application of this process was during the Second World War for the production of penicillin (Baily, 1980). Submerged fermentation allows for a better dissolution of the nutrients of the medium, facilitating their contact with microorganisms, as well as for temperature and pH control during fermentation (Schmidell et al., 2001). The batch process is the most used, but it presents low productivity due to the long fermentation time, as well as low cell concentrations due to inhibition by the

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CHAPTER 5 Technological challenges and advances

final product (Abdel-Rahman et al., 2013). An important factor to be considered in batch fermentative processes is catabolic repression by the substrate. Despite the fact that the rate of conversion of the substrate into product by the microorganism is high, the initial substrate is consumed in a short period of time, ceasing the fermentation. However, a high concentration of the starting substrate may lead to inhibition of production. Catabolic repression can be avoided by maintaining a low concentration of substrate within the reactor, providing noninhibitory sugar amounts throughout the process as required, which characterizes the process as fed batch. Several studies have used this technique in order to avoid inhibition of production by the substrate and, thus, to raise production and productivity levels (Abdel-Rahman et al., 2015). Different feeding strategies can be used in fed batch fermentation, such as pulse feed, continuous, pH Stat, and exponential feeding. Several studies have reported improvements in lactic acid production when using the pulse-feeding strategy (Son and Kwon, 2013; Wang et al., 2011; Meng et al., 2012). In this technique, certain concentrations of substrate are inserted once or more into the reactor, at predetermined intervals, as required. Ding and Tan (2006) studied the effects of different feeding strategies for L(1) lactic acid production by Lactobacillus casei, and according to their results, exponential feeding was the most efficient method; increasing lactic acid production by 56.5% compared to the batch process. For exponential feeding, different software are used with specific formulas that are set with predetermined data containing kinetic fermentation values. These programs are able to adjust the feed flow as requested, taking into account the conversion rates as well as the speed at which they need to occur. In the pH stat method, the feed is based on pH control by considering the linear relationship between the consumption of the base used for the control of pH and substrate utilization during lactic fermentation. The base and the substrate are mixed proportionally and then the substrate concentration is controlled through pH adjustment automatically. In continuous batch fermentation, the necessary nutrients are provided continuously, while the cells, products, and process residues are removed. The bioreactor volume and the nutrient concentrations are kept constant. It is not a preferred process in the industrial sector due to the high risk of contamination and mutation of the producing microorganism (Schmidell et al., 2001). Lu et al. (2012) studied a fermentative process in a pilot scale reactor equipped with a microfiltration system for the recycling of cells and for pulse feeding. In this process the cells were recycled 12 times, and according to the authors the method was proven to be efficient and promising for use on an industrial scale. Some studies report lactic acid production using immobilized cells in order to increase the cell density of the process. However, promising results related to increased production and productivity were not reported, and better results were reported using free cells (Hofvendahl and Hahn Ha¨gerdal, 2000).

5.1 Lactic Acid

In order to optimize lactic acid production, metabolic engineering is one alternative to solve the problems faced in fermentation processes, like acidity, end product and substrate tolerance, as well as the high costs of substrates.

5.1.4 MICROORGANISMS INVOLVED IN THE PRODUCTION OF LACTIC ACID The production of lactic acid has been reported by bacteria, fungi, and yeasts (Ilme´nemail et al., 2013, Sun et al., 2012, Ding and Tan, 2006), but the largest number of studies report such production by bacteria. Osawa et al., (2009) reported a production of 85 g/L of L(1) lactic acid by Candida boidinii genetically modified with alterations in the pyruvate carboxylase gene in order to reduce ethanol production. The most reported filamentous fungus for lactic acid production is Rhizopus oryzae, usually presenting production from byproducts such as glycerol (Vodnar et al., 2013), residues such as manure (Sun et al., 2012), or residues rich in starch (Taskin et al., 2012, Yen et al., 2010). Lactic acid bacteria have traditionally been used to produce lactic acid and are the predominant candidates for industrial production (Okano et al., 2010). This group of bacteria is characterized as being: Gram-positive; aerobic or facultative anaerobes; present in the form of cocci or bacilli, which produce lactic acid by fermentation of carbohydrates; incapable of using lactate; and nonpathogenic (recognized as “GRAS”—Generally Recognized As Safe) (Axelsson, 2004; Hofvendahl and Hahn Ha¨gerdal, 2000). They belong to several genera, including Aerococcus, Carnobacterium, Enterococcus, Lactobacillus, Lactococcus, Leuconostoc, Pediococcus, Streptococcus, Tetragenococcus, Vagococcus, and Weissella (Rattanachaikunsopon and Phumkhachorn, 2010; Carr et al., 2002). According to Garrity et al. (2004), the bacteria of the genus Lactobacillus belong to the family Lactobacillaceae, order Lactobacillales, class Bacilli, and phylum Firmicutes. Lactobacilli are found in substrates rich in carbohydrates, fermented foods, mammalian mucosa, and plants (Hammes and Vogel, 1995). Some species of this genus, such as Lactobacillus rhamnosus, are able to produce L(1) lactic acid, while others have the potential to produce the D( ) isomer, such as Lactobacillus delbrueckii (Hofvendahl and Hahn Ha¨gerdal, 2000). The bacteria of the species delbrueckii are characterized as being: Grampositive, present in the form of bacilli, facultative anaerobic, without motility, and without the formation of spores. L. delbrueckii, is part of the group of lactic acid bacteria, therefore, it is acid tolerant and has strictly fermentative metabolism and lactic acid is the main end product (Axelsson, 2004; Kandler and Weiss, 1986). Studies have shown lactic acid production by several L. delbrueckii strains from different substrates as well as by other species of the genus Lactobacillus (Table 5.2). Another important genus of D( ) lactic acid producing bacteria is Sporolactobacillus, belonging to the phylum Firmicutes, class Bacilli, order Bacillales, and family Sporolactobacillaceae (NCBI- National Center for

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CHAPTER 5 Technological challenges and advances

Table 5.1 Microorganisms of Genus Sporolactobacillus Involved in D( ) Lactic Acid Production Microorganism

D(

) Lactic Acid Production (g/L)

Nitrogen and Carbon Sources

Reference

Sporolactobacillus sp. CASD Sporolactobacillus nakayamae Sporolactobacillus inulinus ATCC 15538 Sporolactobacillus sp. strain CASD

207

Glucose and peanut flour Crystallized sugar and peanut flour Glucose and yeast extract Glucose and yeast extract

Wang et al. (2011) Beitel et al. (2016) Zheng et al. (2010) Zhao et al. (2010)

112.93 93.4 90.00

Biotechnology Information, 2013). It is a Gram-positive bacterium producing spores, microaerophilic and mesophilic, which exclusively produces D( ) lactic acid via homofermentation. Found mainly in wild plant rhizospheres (Holzapfel and Botha, 1988). This genus was named by Oki Nakayama, a Japanese microbiologist who isolated a large number of Sporolactobacillus strains (LPSN, 2013), being known for its ability to tolerate high concentrations of lactic acid as well as its production (Zheng et al., 2010). Few studies involving the production of lactic acid by this genus have been reported (Table 5.1). Some species and strains of Bacillus sp. are commercially promising for the production of lactic acid (Pleissner et al., 2016; Zhou et al., 2013; Ma et al., 2014; Su et al., 2011). The advantages of the use this microorganism include the possibility of producing lactic acid using lignocellulose substrates because of their ability to utilize both C5 and C6 sugars. In addition, they are thermotolerant and nonfastidious. The advantages of using high temperatures in fermentation on an industrial scale are: the increase in enantiomeric purity; lower risk of contamination by mesophiles; faster reactions, therefore, higher productivity; and greater solubility of the fermentation broth. However, there are some bottlenecks for the accomplishment of this, no thermotolerant microorganism producing D( ) lactic acid has been found in nature, it is only possible to find L(1) lactic acid-producing thermotolerant Bacillus. According to Kranenburg et al. (2013) the genus Bacillus is composed of more than 200 different species, however few of them can be modified genetically. For the production of D( ) lactic acid, some strains of Bacillus coagulans have been genetically modified using specific genetic tools. The native ldhL gene from B. coagulans, DSM1 (a producer of L-lactic acid), was deleted and the D-lactate dehydrogenase (ldhD) gene was overexpressed to generate lactic acid (Kranenburg et al., 2013); the engineered strain B. coagulans RDSM1 produced 28 g/L of lactic acid, (99.5% D-lactic acid) from 50 g/L of glucose. The wild strain of B. coagulans, DSM1, produced the same concentration, 28 g/L, of lactic acid, but 99.8% L isomer. Wang et al. (2014) engineered a thermotolerant strain of B. coagulans P4102B (a potent producer of optically pure L-lactic acid .99%), by directing the

5.1 Lactic Acid

Table 5.2 Microorganisms of Genus Lactobacillus involved in D( ) Lactic Acid Production Microorganism

Isomer

Lactic Acid Production ( g/L)

Nitrogen and Carbon Sources

References

Lactobacillus casei

L(

180.0

Glucose and yeast extract Glucose, yeast extract, and peptone Sugar cane juice and yeast extract

Ding and Tan (2006) Yang et al. (2015)

Calabia and Tokiwa (2007) França et al. (2009)

)

Lactobacillus lactis

143

Lactobacillus delbrueckii NCIM 2365 Lactobacillus delbrueckii Lactobacillus delbrueckii ATCC6949 Lactobacillus plantarum LMISM6

D(

)

135

D(

)

120.0

Sugar cane juice

D(

)

101.0

Sugar cane molasses

94.8

Coelho et al. (2011)

Lactobacillus delbrueckii NCIMB 8130 Lactobacillus delbrueckii subsp. lactis QU 41 Lactobacillus manihotivorans LMG 18010

D(

)

88.0

Sugar cane molasses and corn steep liquor Beet molasses and yeast extract

D(

)

20.1

Glucose and yeast extract

Tashiro et al. (2011)

L(1)

12.6

Starch

Guyot et al. (2000)

DL

Kadam et al. (2006)

Kotzamanidis et al. (2002)

gene coding for enzyme D-lactate dehydrogenase to produce high D( ) lactic acid by the deletion of the L-lactate dehydrogenase gene (LdhL) and the acetolactate synthase gene (alsS). The engineered strain produced 90 g/L of D-lactic acid with optical purity greater than 99% from glucose at 50 C. The CRISPR (Clustered Regularly Interspaced Short Palindromic Repeats) technology associated with the cas gene forms a bacterial immune system against strange DNA, such as phage or plasmids, and has revolutionized the field of genetic engineering due to its ability to perform the genome quickly and efficiently. It has been used for gene deletion, knockout, and genetic recombination of a variety of microorganism species (Hidalgo-Cantabrana et al., 2017). Technologies based on CRISPR have been successfully used to increase the phage resistance of industrial strains (Stefanovic et al., 2017) The CRISPR-Cas9 system was used by Ozaki et al., 2017 in order to genetically modify Schizosaccharomyces pombe to produce D( ) lactic acid from glucose and cellobiose, by the insertion of a gene encoding D-lactate dehydrogenase

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CHAPTER 5 Technological challenges and advances

of Lactobacillus plantarum and the deletion of pyruvate dehydrogenase (ADH) and glycerol-3-phosphate dehydrogenase (GPD). The new strain produced 25.2 g/ L of D( ) lactic acid from 35.5 g/L of glucose. The expression of betaglucosidase was also made to facilitated production through cellobiose and the resulting strain produced 24.4 g/L of D( ) lactic acid from 30 g/L of cellobiose in minimal medium. Mougiakos et al. (2017) have constructed a genomic editing system for thermophilic microorganisms using CRISPR/Cas9 technology, in which it is possible to perform, in only 4 days, the elimination, knockout, and insertion of genes, with an efficiency of 90%, 100%, and 20% respectively, with the great advantage of using only one plasmid and one promoter. In other words, a lot of genetic tools are not required. Thus, homologous recombination was performed using plasmids at between 45 C and 55 C in thermophilic Bacillus smithii as spCas9 is inactive in vivo above 42 C, then the transfer was carried out at 37 C, which allowed for counter-selection because at this temperature spCas9 is active and it is possible to introduce lethal DNA that breaks into unedited cells. In this way, this model can be used as an important tool in genomic editing of thermophilic bacteria to increase the production of D( ) lactic acid and L(1) lactic acid.

5.1.5 EXTRACTION AND PURIFICATION OF LACTIC ACID The technological barriers to the production of low cost lactic acid are mainly in the processes of separation and purification of lactic acid from the fermentative medium. In addition, the efficiency of this process is essential for the subsequent synthesis of PLA, since the presence of sugar, proteins, and other organic acids must be minimal in order to obtain optimum polymerization results. In the conventional method of lactic acid purification, the precipitation of calcium lactate occurs with esterification and hydrolysis through reactive distillation. It is an economical process, simple and reliable, but generates large amounts of calcium sulfate, which is not environmentally friendly. In order to solve this problem, some recovery methods have been developed in the literature to remove the lactic acid from a fermentation broth, such as cross flow filtration with cell recycling (Sikder et al., 2012), electrodialysis (Wang et al., 2013), and ion exchange resins (Boonmee et al., 2016). It is also possible to perform an integrated membrane separation process composed of ultrafiltration, nanofiltration (NF), ion exchange, and vacuum evaporation. Lee et al. (2017) used this method and obtained lactic acid with high purity ( . 99.5%). To recover and purify the L(1) lactic acid produced from microbial fermentation media economically and efficiently, ion exchange chromatography is used among a variety of downstream operations(Tayyba Ghaffar et al., 2014). The purification of lactic acid from ion exchange columns has been shown to be a technological option for the process (Gonza´lez et al., 2008), since the necessary equipment is relatively simple and cheap, however its use is recommended when the solution of lactic acid has low salt concentrations (Quintero et al., 2012).

5.2 Poly(lactic Acid)

Many studies focus on the recovery of lactic acid using polymeric anionic adsorbents. This method presents an advantage: there is no need to acidify the fermentation broth prior to adsorption. The adsorption of ion exchange resin is a practical method in the industry due to its economy, ease of manipulation, reduction in chemical consumption, and low waste production. The NF process has also been shown as an option for the process of separating lactic acid from a fermentation medium. Its selectivity is due both to steric hindrance and to the effect of electrostatic repulsion. Most NF membranes can firmly hold compounds of molecular weight up to 150 250 g/mol, charged species, and polyvalent ions. NF is a purification step that can be performed before and/or after the ion exchange process, since NF membranes have low rates of rejection to lactic acid and high retention of mono- and disaccharides and divalent ions (e.g., Ca12 and Mg12) (Kang et al., 2004), it is also possible to obtain a high recovery rate of lactic acid. With the participation of NF membrane modules and microfiltration in a stable production system, 76 77 L/m2 h was obtained for L(1)lactic acid with a purity greater than 95% (Pal and Dey, 2012). Advances in separation and purification techniques based on membrane technologies, particularly in microfiltration, ultrafiltration, and electrodialysis, has led to the creation of new processes for the production of lactic acid without the generation of residues. The literature shows that the production of lactic acid is viable from lactate salts through the use of bipolar membrane electrodialysis which can convert solutions of salts into acids. First, the monopolar electrodialysis membrane that will be responsible for concentrating and purifying the broth is used, and then the biopolar electrodialysis membrane is used to convert lactate salts into lactic acid (Habova et al., 2004). Through ionic exchanges, the membrane is capable of supplying hydrogen ions to the broth and recovering the cations present, so that the base used during the fermentation can also be recovered. In other words, at the end of the process there is the conversion of the lactate salts into acid form and also the recovery of the base used in the fermentation process. The study of procedures for the purification of lactic acid involves unit operations, and the critical steps are evaluated in relation to lactic acid recovery and the removal of contaminants, such as proteins, sugars, and color.

5.2 POLY(LACTIC ACID) The main challenges for the commercialization and use of PLA include: cost production; physical structure of polymer; PLAs are not produced directly by microorganisms; biodegradation; and biocompatibility. Therefore, the reduction of costs, as well as the development of efficient fermentation processes that result in high productivity are current study objectives (Abdel-Rahman et al., 2013).

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Cost reductions can be achieved using cheap substrates, optimizing lactic acid production processes, substituting energy sources used in PLA production with more sustainable energy sources, such as wind or solar energy, optimizing PLA production processes, and increasing PLA demands (Jamshidian et al., 2010).

5.2.1 PLA CHEMICAL AND PHYSICAL PROPERTIES There are several kinds of degradable polymers, however, PLA is one of the most popular materials as it is considered as a biocompatible, compostable, recyclable, and renewable thermoplastic polyester, with the brightest development prospects, and is denominated as an ecofriendly material (Madhavan Nampoothiri et al., 2010; Maharana et al., 2015). The production of PLA also consumes carbon dioxide (Dorgan et al., 2001) and uses 25% 55% less fossil energy sources than petroleum-based polymers. Cargill Dow developed process improvements, for the near future, where the use of fossil energy sources can be reduced by more than 90% compared to any of the petroleum-based polymers (Vink et al., 2003). Considering the environmental concerns around petroleum-based polymers and the associated “white pollution” that results in severe urban environmental consequences, the low carbon economy has begun receiving the relevant attention. In this way, PLA is an important environmentally friendly plastic derived from renewable sources, with performance comparable to many petroleum-based plastics (Ren, 2010). PLA presents some properties ideal for use in packaging, textiles, and other consumer products, such as its high wicking performance, light weight, good dyeability, antibacterial feature, good ultraviolet resistance, high water vapor transmission rates, good printability, low process temperature, and ease of conversion into different forms (Chen et al., 2016; Lim et al., 2008). On the other hand, its potential market applications could be limited by poor thermal stability, mechanical properties, and processability (Cheng et al., 2015). Unoriented PLA is brittle, however, it displays good strength and stiffness (Auras et al., 2005). It is a semicrystalline polymer with good transparency, mechanical strength, and melt processability; however, the physical properties of PLA can vary widely, depending on its stereochemical composition, processing temperature, annealing time, and molecular weight, which have an effect on its melting point and on the rate and extent of polymer crystallization (Chen and Patel, 2012; Frone et al., 2016). PLA even exhibits mechanical properties comparable to those of poly(ethylene terephthalate) and better than those of polystyrene (Auras et al., 2004). This polymer is considered as a biodegradable polyester due to its potentially hydrolysable ester bonds, and for this reason is considered as a sustainable alternative to petrochemical plastics. PLA exists in three stereochemical forms: poly (L-lactide) (PLLA), poly(D-lactide) (PDLA), and poly(DL-lactide) (PDLLA) (Madhavan Nampoothiri et al., 2010). Savaris et al. (2016) studied the modifications to PLA films after exposure to sterilization methods, and verified that PLA sterilized by saturated steam showed

5.2 Poly(lactic Acid)

morphological, chemical, thermal, and physical changes, thus, this process is not recommended. However, under other sterilization processes, such as ethylene oxide, hydrogen peroxide plasma, electron beam radiation, and gamma radiation, the samples exhibited only thermal and physical changes, therefore, these processes can be used for PLA sterilization.

5.2.2 PLA SYNTHESIS 5.2.2.1 Chemical polymerization Considering that PLA is the best candidate polymer to replace nonbiodegradable synthetic plastics, several studies related to synthesis techniques have been performed in order to increase the optical or mechanical properties of this polymer (Sungyeap Hong, 2014). In the beginning of PLA technology development, its applications were limited to the medical field due the high costs involved (Chen et al., 2016). Attaining a PLA with good physical properties and high molecular weight is desirable (Liu et al., 2013). Over the past decade, new technologies of polymerization were developed, promoting the economical production of polymers with high molecular weight (Lim et al., 2008). There are different polymerization techniques for achieving a high molecular weight PLA, including direct condensation polymerization, azeotropic dehydrative condensation, and polymerization through lactide formation or ring opening polymerization (ROP) (Fig. 5.6). ROP is a common polymerization route used to build hydrolytic polymers, including PLA, and is the most convenient technique for controlling the molecular weight of this polymer (Nair and Laurencin, 2007, Masutani and Kimura, 2014), however, this method is relatively complicated and expensive (Liu et al., 2013). This route, starts with low molecular weight lactic acid oligomers from catalytic depolymerization of internal transesterification using tin catalysis and stereoselective initiators in order to enhance the rate and selectivity of the intramolecular cyclization reaction. Then the ring of lactide opens to form high molecular weight PLA (Auras et al., 2004). Some companies, like NatureWorks LLC (USA) among other manufacturers of PLA, use the ROP route for their production. Currently, much effort is being made to establish the direct polycondensation method, since this technique has been less studied compared to the ROP method (Masutani and Kimura, 2014). However, chemical catalysts have drawbacks, for example, they often require high operating temperatures and metal catalysts may cause problems for certain uses (Xiao et al., 2006). Microwave irradiation has been used in PLA synthesis, and in this technique the polymerization rate is much faster than with conventional heating, also different catalysts can be used. Bakibaev et al. (2015) to get the PLA polymer 100 times faster than the conventional heating method of synthesis, obtained polymer

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FIGURE 5.6 Synthesis of poly(lactic acid) (Lunt, 1998).

with same optical characteristics. Similar results reported by other authors explain that microwave irradiation gives rapid energy transfer and high-energy efficiency, conducive to a faster reaction rate (Singla et al., 2012; Nagahata et al., 2007). Nikolic et al. (2010) compared poly(D, L-lactide) synthesis using ROP to poly (D, L-lactide) synthesis using microwave irradiation, and reported that polymerization by ROP takes over 30 hours in bulk at 120 C, in contrast, microwave irradiation performs the same bulk polymerization process much faster and with less energy consumption, with a reaction time of about 30 minutes at 100 C.

5.2.2.2 Enzymatic polymerization: production of PLA directly by genetically modified microorganism Enzyme catalyzed polycondensation is a green alternative approach that present some advantages compared to chemical synthesis, since mild conditions can be achieved, because the reactions are usually metal free and are carried out at lower temperatures. In this technique, high selectivity, high efficiency, and recyclability of enzymes have also been reported. Therefore, this process has aroused interest for its medical applications, in order to avoid toxic solvents and metal residues (Sen and Puskas, 2015). Researchers have developed a strategy for the efficient production of PLA and its copolymers using microorganisms, in only one step, through metabolic

5.2 Poly(lactic Acid)

engineering. One of the strategies was the use of polyesters containing lactate with different comonomers by using the PHA biosynthetic pathway in recombinant E. coli (Park, Kim, et al., 2012a; Park, Lee, et al., 2012b) Jung et al., (2010) reported the production of a homopolymer PLA and its copolymers, poly(3-hydroxybutyrate-co-lactate), P(3HB-co-LA), by direct fermentation of metabolically engineered E. coli. The introduction of heterologous metabolic routes involving CoA transferase and polyhydroxyalkanoate (PHA) synthase for efficient generation of lactyl-CoA and incorporation of lactyl-CoA into the polymer, respectively, allowed for the synthesis of PLA and P(3HB-co-LA) in E. coli, but the efficiency was low. Jung et al., (2010), in their study, engineered the metabolic pathway of E. coli by knocking out the ackA, ppc, and adhE genes and exchanged the ldhA promoters and acs for trc promoters. Lassalle et al., (2008) reported effective PLA production using lipases. Candida antarctica lipase B (CAL-B) is an effective biocatalyst, presenting 60% lactic acid (LA) conversion and 55% recovered solid polymer. Based on the fact that PHA monomeric constituents are structurally analogous to LA, Taguchi et al., 2008 explored the capacity of PHA synthase to exhibit polymerizing activity toward LA-coenzyme A (CoA), in order to establish a biological process for the synthesis of LA-based polyesters. An LA-CoA producing E. coli strain with a CoA transferase gene was constructed and an engineered PHA synthase gene was introduced into the resultant recombinant strain. As a result, a LA-incorporated copolyester, P(6 mol% LA-co-94 mol% 3HB), was reported.

5.2.3 KINDS OF POLYMERS, COPOLYMERS, AND THEIR FEATURES The bioplastics industry has made dramatic changes over the past decades in order to achieve durable bioplastics with high biobased content. In the automotive and electronics industries, for example, the requirements of this material are that they are tough, strong, and durable. The high strength and low toughness characteristics of PLA are presented as limitations especially in applications where mechanical toughness such as plastic deformation at high impact rates or elongation is required (Nagarajan et al., 2016; Kfoury et al., 2013). The need to improve the chemical and physical properties of PLA to meet consumer and biological applications, brings forth several technologies based on PLA (Rasal et al., 2010). In this way, in order to expand these potential properties, copolymerization or blending with other polymers are been developed to driving to new materials, that presents high performance, low cost and easy processing (Hamad et al., 2014; Schreck and Hillmyer, 2007). Another example of modifications needed relates to the permeability of PLA to oxygen, gas, and moisture, which is much higher than in most other plastics, such as PE, PP, and even PET. The penetration of moisture and consequently the hydrolytic degradation of PLA, as well as its low glass transition temperature,

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poor thermal stability and low toughness and ductility can significantly interfere in the manufacturing, storage, and transport processes of PLA, reducing, in this way, the range of its applications. To extend the applications of PLA, properties such as impact strength or flexibility, stiffness, barrier properties, thermal stability, and production costs should be improved (Jamshidian et al., 2010). Copolymer composition, functionalities, as well molecular weight and terminal groups, need to be controlled in order to drive PLA polymer applications (Masutani and Kimura, 2014). Blends of PLA with other flexible polymers have been previously studied, like poly(lactide-co-glycolide) (PLGA), and applied in various biomedical applications, such as vaccination, cancer treatment, inflammation medication, tissue engineering, and regenerative medicine, among others. PLGA is considered one of the most promising polymers in design, development, and optimization for medical applications (Pan and Ding, 2012; Szle˛k et al., 2016; Danhier et al., 2012). PLGA-based have been used also to prepare microspheres to multiparticulate dosage applications under different sizes, depending on the route of administration (Szle˛k et al., 2016). PLGA nanoparticles have attracted attention due some interesting properties, such as biodegradability and biocompatibility. The Food and Drug Administration and the European Medicine Agency approval in drug delivery systems for parenteral administration, protection of drug from degradation as well as possibility of sustained release (Danhier et al., 2012). In turn, poly(lactic acid)/polycaprolactone (PLA/PCL) blends (80/20) showed the greatest elongation and impact strength compared to that of neat PLA (Chavalitpanya and Phattanarudee, 2013; Rao et al., 2011). The synthesis of poly (D, L-lactic acid)/poly(L-lactic acid) bentonite nanocomposites increases the tensile strength of this polymer as well as its biodegradability and barrier properties (Sitompul et al., 2016). The system of triblock copolymers poly(lactic acid)-poly(ethylene oxide)-poly (lactic acid) has been widely explored for several applications related with controlled and sustained release of drugs and in tissue engineering devices (Saffer et al., 2011). Xu and Guo (2010) reported that blending poly(butylene succinate) (PBS) with PLA improves the tensile strength and elastic modulus of the polymer without much loss of ductility. On the other hand, Hassan et al., (2013) related in their study that the thermal stability of the blends (PLA/PBS) was higher than that of pure PLA and the tensile strength and modulus of the blends decreased with increasing PBS content, however, the impact strength was improved by about twofold compared to pure PLA and PBS also increased the storage modulus. The low heat resistance and slow crystallization rate characteristics of PLA limit its application in food packaging. The combined effect of annealing and cellulose nanofibers on the crystalline structure of PLA was studied by Frone et al. (2016). The transformation of the grained structure specific to amorphous PLA into a crystalline lamellar structure after annealing was detected, showing an improvement of the surface modulus and hardness for the new composites compared to PLA.

5.2 Poly(lactic Acid)

Hashima et al., (2010) developed a ductile polymer with high glass transition temperature by blending PLA with a hydrogenated styrene-butadiene-styrene block copolymer with the aid of reactive compatibilizer, poly(ethylene-co-glycidyl methacrylate). The aging resistance was also improved by incorporating polycarbonate. The architecture of polymers can be controlled by the addition of hydroxylic compounds, which permits precise control over the speed of crystallization, the mechanical properties, and the processing temperatures of the material (Auras et al., 2004). In order to increase the crystalline content under typical polymer processing conditions, Li and Huneault (2007) studied different strategies to promote PLA crystallization. Talc used as a nucleating PLA agent is highly effective in the 80 C 120 C temperature range and the combination of nucleant and plasticizer developed significant crystallinity at high cooling rates. Charles et al., (2010) in their study, prepared poly(L-lactic acid)/hydroxyapatite/poly(ecaprolactone), in order to prepare a bone-repair material that matches with the modulus of bone. They reported the production of composites with flexural moduli near the lower range of bone. There has been much research based on PLA improvements considering the medical field. It was verified through biological assays that PLA scaffolds containing low loads of diamond nanoparticles were not cytotoxic using L929 cells, presenting bioactive properties and favored cell adhesion. The potential range of applicability of electrospun polyesters diamond nanoparticles loaded for tissue engineering purposes was also related (Pereira et al., 2016). In Brazil, BASF offers the Ecovio, which consists of the biodegradable plastic ecoflex (synthetic polyester obtained from the condensation of 1,4-butanediol with terephthalic and adipic acid) and PLA, which is obtained from renewable raw materials based on sugar. This plastic can be adapted according to customer requirements (Araujo et al., 2007). PLA and PLGA have been used commercially as membranes according to their properties (Table 5.3).

5.2.4 PLA APPLICATIONS The actual needs of society drives the progressively increasing plastic production, consequently resulting in high plastic waste generation (Balaguer, et al., 2016). In this context, the concept of ecofriendly materials has attracted attention by several sectors of society all over the world (Sukan et al., 2015). Considering the interest in green alternatives, greater efforts have been made to develop degradable biological materials without any environmental pollution to replace plastics derived from petroleum, increasing the manufacturing of biodegradable PLA materials (Abdel-Rahman et al., 2013; Madhavan Nampoothiri et al., 2010). This biopolymer has bacterial origins, presents a similar level of quality to traditional plastics, and has been considered a candidate to replace unsustainable products, like polystyrene and polyethylene terephthalate, however, the cost of its production is an inconvenience yet (Sukan et al., 2015).

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Table 5.3 The Most Commonly used Commercially Available Reabsorbable Synthetic Polymeric Membranes Commercial Name (Manufacturer)

Materials

Properties

Function Time

Resorption Rate

Guidor (Sunstar Americas, Inc. mnear Chicago, IL, USA) Resolut Adapt (W. L. Gore and ASSOC, Flagstaff, AZ, USA) Resolut Adapt LT (W.L. Gore and ASSOC, Flagstaff, AZ, USA) Epi-Guide (Curasan, Inc., Kleinostheim, Germany) Vivosorb (Polyganics, Groningen, The Netherlands)

Poly-D, L-lactide and Poly-L-lactide, blended with acetyl tri-n-butyl citrate Poly-D, L-lactide/Coglycolide

2-layer

$6 weeks

13 months

Good space maintainer

8 10 weeks

5 6 months

Poly-D, L-lactide/Coglycolide

Good space maintainer

16 24 weeks

5 6 months

Poly-D, L-lactic acid

3-layer Selfsupporting

20 weeks

6 12 months

Poly(D, L-lactideε-caprolactone)

Can also be used as a nerve guide

10 weeks

24 months

Studies have been carried out on the utilization of biodegradable materials for plastic applications, such as packaging, paper coatings, sustained release systems for pesticides/fertilizers, compost bags, and textile applications, among others (Madhavan Nampoothiri et al., 2010; Hamad et al., 2014; Chen et al., 2016). Efforts have been made to develop nanotechnology approaches to packaging material science for improving performances and decreasing prices (Jamshidian et al., 2010). Considering that the major drawback of PLA for packaging applications is its brittleness, Javidi et al., (2016) verified that PLA-essential oil composite films were more flexible than neat PLA films, leading to modification of the tensile behavior. The potential antimicrobial performance of Origanum vulgare L. essential oil (OEO) in PLA-based matrices was evaluated too and the results indicated that antimicrobial PLA film was effective against Staphylococcus aureus and E. coli, suggesting that developed PLA films with active substances could be used in designing antimicrobial packaging materials. Another way to solve the problem of the PLA stiffness limitation, is adding a plasticizer in order to enhance flexibility and improve the electromechanical properties of this polymer. Therefore, several types of plasticizer could be used for this purpose, such as poly (ethylene glycol), glyceryl triacetate, citrate esters, lactide monomer (Lemmouchi et al., 2009; Thummarungsan et al., 2016). Lemmouchi et al., (2009) reported that blending PLA (80 wt.%) with plasticizers (20 wt.% tributyl citrate) improves the

5.2 Poly(lactic Acid)

thermomechanical properties. Besides that, it was evidenced that the plasticizers investigated in their work enhanced the degradation of the PLA matrix in compost conditions. PLA has been used in the medical field, including drug delivery systems, tissue engineering, wound management, orthopedic devices (Hamad et al., 2015), degradable sutures, nanoparticles, porous scaffolds for cellular applications (Lasprilla et al., 2012), implantable composites and bone fixation parts (Madhavan Nampoothiri et al., 2010), and plain membranes for guided tissue regeneration (Vert, 2004), among others. These applications require materials with specific properties to provide efficient therapy (Gupta et al., 2007). PLA had attracted great attention in medical applications due its advantageous features over nonbiodegradable polymers, such as biocompatibility and bioabsorption, eliminating the subsequent need to remove implants (Hamad et al., 2015). The expected trend for future years, is that permanent implants will be replaced by biodegradable devices, which could auxiliate the body to regenerate damaged tissues (Nair and Laurencin, 2007). Bioabsorbable polymers receive significant attention for biomedical applications because their degradation occurs by simple hydrolysis to metabolizable products by the human body (Savioli Lopes et al., 2012). Deepthi et al., (2016) in their study, developed a tendon construct of electrospun aligned PLLA nanofibers in order to mimic aligned collagen fiber bundles, and they layered PLLA fibers with chitosan-collagen hydrogel to mimic the glycosaminoglycans of sheath extracellular matrix for tendon regeneration. Their results indicated that this newly developed scaffold would provide a proper construct for flexor tendon regeneration under immobilized conditions. Proanthocyanidins extracted from grapes have several bioactive properties, giving them potential medical uses. In this way, nanoencapsulation with poly-D,Llactide polymer was accomplished, and in vitro release studies, through stomach and intestinal simulation, showed a sustained release of proanthocyanidins (Ferna´ndez et al., 2016). Considering the packaging industry, PLA is an alternative “green” food packaging polymer, ideal for fresh products, such as fruits, vegetables, salads, and those whose quality is not damaged by PLA oxygen permeability (Jamshidian et al., 2010). Jamshidian et al., (2010) reported in their research an overview of the capabilities of PLA, discussing the antimicrobial and antioxidant trends of this polymer, required in different areas when used in packaging. Some applications of nanomaterials in combination with PLA structures were also reported for creating new PLA nanocomposites with promising abilities. Chang et al., 1996 studied PLA for use as a matrix for controlled release of herbicides as well growth stimulation and yield improvement potential, applied in preplant soil incorporation with soybeans. This research demonstrated that both lactide and PLA were able to increase soybean leaf area, pod number, bean number, and plant dry weight. Also PLA used as an encapsulation matrix for herbicides could provide reduced environmental impacts and improved weed control.

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PLA is already used in cosmetic packaging, and it is well-known that polymeric packaging can interact with components, such as active ingredients, excipients, and solvents, used in a variety of cosmetic formulas. In their study, Capra et al. (2014) evaluated the mechanical, physicochemical, and organoleptic properties of PLA bottles present in the cosmetic market. The results indicated that the polymer was modified under a heating process in combination with chemical treatment.

5.2.5 PLA MARKET DEVELOPMENT On a global scale, about 275,000 tons of lactic acid are produced per year, and the application of lactic acid is directed to several industrial departments, such as food and beverages, personal care, and solvents, however the greater application of lactic acid is directed to the production of polymers (Fig. 5.7) (The essential chemical industry, 2016). Among several applications of PLA, a large part is intended for the packaging industry, followed by applications in agriculture, as shown in Fig. 5.8. Some PLA producer companies have developed PLA that can withstand high temperatures making it suitable for packaging and other applications where this quality is required, such as catering and for serving hot beverages. PLA can be also used for the manufacture of electronic appliances, automotive parts, and many consumer goods like diapers, toys, etc. (Occams Business Research, 2016).

FIGURE 5.7 Biotechnological lactic acid applications. From The Essential Chemical Industry, 2016. essentialchemicalindustry-online, Ed. J.N. Lazonby and D.J. Waddington, Published by the Department of Chemistry, University of York, York, UK. Used with permission from Professor David Waddington, Department of Chemistry.

5.2 Poly(lactic Acid)

FIGURE 5.8 PLA application market, 2013 (Occams Business Research, 2016).

According to a study by Grand View Research (2016), the increasing demand for PLA is expected to reach US$2,169,6 million by 2020 and will increase the lactic acid market, which is expected to reach US$4,312,2 million in the same period.

5.2.6 PLA BIODEGRADATION, BIOCOMPATIBILITY, AND TOXICITY It is desired that biodegradable biomaterials do not cause an inflammatory response; do not produce toxic degradation products, have appropriate mechanical properties, appropriate permeability and processability, as well as, an ideal degradation time depending on the purpose of use. It is also important to be sterilizable (Ikada and Tsuji, 2000). Biodegradable polymers do not require excellent biocompatibility, since they do not stay in the body for a long time but disappear within hours, days, or weeks without leaving any trace of residue. It has not yet been fabricated biocompatible implants for permanent use (Ikada and Tsuji, 2000). Many of factors affect the degradation rate of PLA, such as the local environment (temperature, water, pH, salinity, oxygen, nutrients); and the chemical and physical characteristics of the polymer, such as molecular weight, crystallinity, purity, permeability, porosity, volume, size, and terminal carboxyl or hydroxyl groups (Madhavan Nampoothiri et al., 2010). Low molecular weight polymers can be degraded more easily than high molecular weight as well as amorphous polymers, for example, PDLLA can be degraded more easily than semicrystalline PLLA and PDLA and then scPLA (stereocomplex of PLLA/PDLA with higher crystallinity) (Xu et al., 2006). The decomposition of PLA in the body varies between 6 months to 1 year, depending on the type used (Pietrzak et al., 1997). According to Li et al. (1990), the weight of PLLA did not reduce until 5 weeks in saline solution. Inflammatory responses were observed in the cells of human and animal bodies due to the products generated by the degradation of PLA in the form of implants, drug delivery materials, and sutures (Vainionpa¨a¨ et al., 1989). In general, PLA is a biocompatible polymer with low toxicity. Although some experiments in vitro have shown low cell multiplication (van Sliedregt et al.,

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1992) and high acid concentrations generated due to degradation; hydrolysis of the polymer (Taylor et al., 1994). Because of the pH reduction, some undesired reactions in the tissue were observed, such as osteolysis, bone resorption, as well as irritation at the implant site (Claes and Ignatius, 2002, Suganuma and Alexander, 1993). One way to solve this problem would be the incorporation of hydroxylapatite (HA) or tricalcium phosphate, which controls the rate of acid formation (Hile et al., 2004). Another problem of PLA is its vulnerability, with the risk of breaking during surgery. The structure of the material can be enhanced by the addition of trimethylene carbonate (TMC) (Zhang et al., 2004). Despite the disadvantages reported in the literature, some researchers have found innovative ways for biopolymers indicating that they show great potential for use in tissue engineering. Yoon et al., 2017 verified, in vivo and in vitro, that the weight and microstructure of PLLA were not degraded over time. There was no reduction in PLLA mesh tension in vitro for 180 days. In addition, they did not observe PLLA induction of inflammation in the subcutaneous tissue. A new technology that is nontoxic and biocompatible, NanoMatrix3D (NM3D), was suggested by Pogorielov et al., 2018.

5.3 CONCLUSION PLA have been showed an extraordinary increasing of applications, and, for this reason this demand will increase the lactic acid market for the next years, however the cost of this production is an inconvenience yet. New technologies for polymerization have been developed for promoting the economical production and high molecular weight of PLA. Chemical catalysts are being gradually replaced by different polymerization processes that present easy conditions of operating, and advances have been developed in genetic engineering toward the same purpose. Cost reductions can be achieved using cheap carbon and nitrogen sources in lactic acid fermentation, optimization of fermentation parameters, and the replacement of electric energy with a cheaper sustainable energy source, such as solar or wind energy. To attend to the necessities of consumers, chemical and physical PLA properties are emerging. By boosting copolymerization or blending technologies with other polymers, new materials with interesting characteristics for diverse applications will emerge. Further studies are needed to assess the compatibility of the material in the body and the toxicity of the degradation products that can cause inflammatory responses, as well as the appropriate mechanical properties, permeability and processability, to have an optimal degradation time depending on the purpose of use. Optimizations of lactic acid production processes, polymerization, and the related technologies, as well as biodegradation researches present continuous progress and novelties.

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PLGA scaffolds: building blocks for new age therapeutics

6

Hafsa Ahmad1, Abhishek Arya2, Satish Agrawal2 and Anil Kumar Dwivedi2 1

Pharmacognosy & Ethnopharmacology Division, CSIR-National Botanical Research Institute, Lucknow, India 2Division of Pharmaceutics & Pharmacokinetics, CSIR-Central Drug Research Institute, Lucknow, India

6.1 CHALLENGES IN NEW AGE THERAPEUTIC STRATEGIES The field of therapeutics via drug delivery or biomedical engineering has undergone several transformations since its inception and is still continuously evolving. The emergence and use of biomaterials and scaffolds could be observed as a major breakthrough in this area. The advent of nanotechnology gave a new impetus to the utilization of nanoscale materials in the fields of drug delivery, engineering, diagnostics, and imaging (Wang et al., 2015,b; Shi et al., 2010; Mulder et al., 2009; Caruthers et al., 2007; Farokhzad and Langer, 2006; Lanone and Boczkowski, 2006; Sahoo and Labhasetwar, 2003). However, the challenges associated with the use of these materials are manifold. An understanding of the mechanisms of activities occurring in the cellular environment, like cellular uptake, transport, and trafficking, and the eventual fate of these biomaterials across complex biological networks is imperative to formulate an efficient therapeutic strategy. To achieve the ability to control and manipulate therapeutic strategies that can overcome obstacles, like rapid clearance by the immune system, poor targeting efficiency, and inability or difficulty crossing biological barriers; a good understanding of the cellular transport of these biomaterials becomes essential. Another challenge faced by new age therapeutics is the rising concern about the safety and toxicity of these biomaterials. This has led to the emergence of nanotoxicology; a field parallel to nanomedicine. Drug delivery strategies lead to the inhalation, ingestion, and absorption of these nanomaterials into human systems, hence the potential negative impact their interaction with biological systems might cause cannot be ruled out. Thus, checks and standards in manufacturing and disposal processes that ensure the safety of these therapeutic agents must be

Materials for Biomedical Engineering: Hydrogels and Polymer-based Scaffolds. DOI: https://doi.org/10.1016/B978-0-12-816901-8.00006-7 © 2019 Elsevier Inc. All rights reserved.

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enforced. Guidelines and protocols that establish the safety of these novel materials must be put into place. Large-scale production of these novel materials and scaffolds is yet another challenge. Achieving scale up laboratory or pilot technologies that ensure reproducible production and commercialization is a complex task. Due to the nature of the method of preparation and the huge material costs involved these technologies are often rendered incompatible with large-scale production. Since economy of production is practically unattainable, the factor of cost is a major challenge affecting new age therapeutics. Besides these factors, other significant factors include solubility and structural stability issues, biocompatibility, attainment of tunable release profiles, and restricting the therapeutic action to the targeted site (Andresen and Berg, 2013; Lammers, 2013; Park, 2013; Bamrungsap et al., 2012; Ruenraroengsak et al., 2010; Jeong and Borm, 2008; Jayagopal and Shastri, 2007). There are several other practical problems associated with treatment modalities that utilize nanomaterials. One such problem is the lack of correlation between the transport and behavior of these particles from in vitro or in vivo studies and actual 3D solid human tumors given their heterogeneity in chemotherapeutic modalities. Also, there has been too much importance being given to passive targeting approaches based on enhanced permeation and retention effects in order to improve the accumulation of chemotherapeutic agents, which has also become a matter of concern. These problems taken together pose a challenge in extrapolating and translating experimental leads into human clinical trials (Satalkar et al., 2016; Bregoli et al., 2015; Eetezadi et al., 2014; Sengupta, 2014). If these challenges are left unaddressed they could tremendously restrict the true potential of new age therapeutics to provide better outcomes in healthcare. A strategy that addresses most, if not all, of these concerns would be greatly sought after in a scenario like this. Nanoengineering of delivery systems is one such strategy which can confer unique characteristics to a drug and can be achieved at different levels of the formulation process. These approaches have the capacity to provide favorable endpoints, like enhanced efficacy and residence, and reduced dosage requirements, toxicity, and costs, thus, leading to overall enhanced patient compliance (Jayagopal and Shastri, 2007). Tissue engineering is yet another strategy that aims to fabricate scaffolds that can mimic the environment of the extracellular matrix. These scaffolds (nanofiber and microfiber based scaffolds) help in the cellular scale reproduction of tissue and provide templates for the macroscopic shape of the target tissue (Sperling et al., 2016). Saturated poly(α-hydroxy esters), like poly(glycolic acid) (PGA), poly (lactic acid) (PLA), and poly(lactic-co-glycolide) (PLGA) copolymers, are the most commonly utilized biodegradable synthetic polymers for the design of 3D scaffolds in tissue engineering (Mano et al., 2004; Lin et al., 2002). PLGA has been used extensively for tissue engineering and biomedical applications due to its mechanical strength, biodegradability, and biocompatibility (Gentile et al., 2014; Vasita et al., 2010).

6.2 Poly(Lactide-co-Glycolide): General Introduction

6.2 POLY(LACTIDE-CO-GLYCOLIDE): GENERAL INTRODUCTION Biodegradable materials can either be natural or synthetic in origin with a tendency to get degraded in vivo by enzymatic or nonenzymatic means, and get transformed into relatively safe byproducts that are biocompatible and can be eliminated by normal metabolic pathways. The use of biomaterials for therapeutic strategies has observed a dramatic surge in the past few decades. The most popularly used biomaterials in drug delivery and therapeutics can be categorized as: (1) synthetic biodegradable polymers, which are relatively hydrophobic in nature and include α-hydroxy acids (e.g., PLGA), polyanhydrides, and several others. (2) Naturally occurring polymers like complex sugars (chitosan, hyaluronan) and inorganics like hydroxyapatite (Makadia and Siegel, 2011). Across all categories of biodegradable biomaterials, PLGA is a popular choice as a carrier for the delivery of drugs, biologicals, proteins, and macromolecules (DNA, RNA, peptides) and as scaffolds for tissue engineering. It is a biodegradable FDAapproved copolymer of PLA and PGA. PLA has an asymmetric α-carbon referred to as either D or L; the enantiomeric forms being PDLA and PLLA. PLGA is a general acronym for poly(D, L-lactic-co-glycolic acid); the D- and L-lactic acid forms being in equal ratio. It has been an attractive choice in drug delivery and tissue engineering and is often referred to as a “smart polymer” owing to these merits: biodegradability and biocompatibility, controlled or sustained release characteristics, favorable degradation properties, suitability for parenteral usage, long clinical use and well documented preparation methods for formulations utilizing a wide variety of hydrophilic as well as hydrophobic drugs and molecules, modifiable surface properties, possibility of formulating a given targeted therapy, protection of therapeutic agent against degradation, and tunable physical properties (Kapoor et al., 2015; Danhier et al., 2012; Bouissou et al., 2006; Ruhe et al., 2003; Jain et al., 2000). These multiple properties render it useful in a wide spectrum of applications (Fig. 6.1). PLGA can be obtained in different forms based on the ratio of lactide to glycolide used in polymerization, which can be identified in terms of monomer ratio (e.g., PLGA 75:25 meaning 75% lactide and 25% glycolide). Degradation of PLGA copolymer is affected through hydrolysis or biodegradation via cleavage of its ester linkages into oligomers and subsequently into monomers. This takes place by bulk degradation of the matrix when penetration of water into the matrix is greater than the rate at which the polymer degrades. PLGA degradation is a cumulative effect involving various processes, like bulk and surface diffusion and bulk and surface erosion collectively. Since the rate of PLGA copolymer degradation is governed by a number of factors, like chain compositions, the degree of crystallinity, molecular weight and glass transition temperature of the polymer, and hydrophobic/hydrophilic balance, its release rate patterns are seldom predictable. The time in which PLGA copolymer degrades into its monomers depends on the ratio of monomer used in its production. PLGA copolymer composed of

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CHAPTER 6 PLGA scaffolds: building blocks for new age therapeutics

FIGURE 6.1 Properties and biomedical applications of PLGA.

greater glycolide content has shown a lesser degradation time (Gentile et al., 2014; Danhier et al., 2012; Haripal, 2012). Studies have shown that PLGA follows a dose-dependent and nonlinear biodistribution and pharmacokinetics pattern. Furthermore, its dose and composition also have a significant influence on its clearance from systemic circulation and its uptake by cells of the mononuclear phagocyte system. Organ distribution studies of several nanoparticulate PLGA formulations via autoradiography and quantitative distribution experiments have demonstrated its rapid accumulation in the bone marrow, spleen, lymph nodes, peritoneal macrophages, and liver. PLGA carrier has exhibited rapid early degradation which later slows down and observes subsequent respiratory clearance by the lungs. Several studies have advocated for the use of surface modification in order to overcome these demerits and to enhance its time in circulation (Esmaeili et al., 2008; Yang et al., 2001; Panagi et al., 2001; Bazile et al., 1992).

6.3 POLY(LACTIDE-CO-GLYCOLIDE) SYNTHESIS The process parameters used in the synthetic scheme for the preparation of PLGA has a strong influence on the physicochemical properties of the resultant product.

6.3 Poly(Lactide-co-Glycolide) Synthesis

PLGA could be prepared by a direct polycondensation reaction of lactic acid and glycolic acid (Scheme 1 in Fig. 6.2), which can be affected in solution as well as in the melt or solid state. Low molecular weight PLGA (MW , 10 kDa) could be produced by polycondensation in solution at temperatures higher than 120
C and under conditions of water removal. A high degree of dehydration was a crucial requirement for preparing high molecular weight PLGA via polycondensation, thus, this method was deemed unsuitable for the same (Gentile et al., 2014; Avgoustakis, 2005). However, Ajioka et al., reported a single step preparation of high molecular weight PLGA (MW .160 kDa) by direct polycondensation of lactic acid and glycolic acid in azeotropic solvent diphenyl ether (2040 hours at 130
C) using powdered tin as a catalytic agent. However, the use of this solvent generated high complexities in process control as well as in product purification, thus, leading to increased costs. Thus, it was thought that polycondensation should preferably be carried out in the melt/solid state (Scheme 2 in Fig. 6.2) to overcome the Scheme 1 CH3 CH

nHO

+

CH2

nHO

H

CH

O

Glycolic acid

Lactic acid

H2 C

O

OH + 2nH2O n

CO

CH3

COOH

COOH

CO

Low MW PLGA

Scheme 2 O

O H

O

H

H H3C O

H

O

H

O CH3

O

H

H3C

H

O

H

H

O

O

D-Lactide

CH3

CH3 H3C

O

O

O Glycolide

O

O

H

O

L-Lactide

Meso-lactide

Scheme 3 H3C

R Melt state HO

CH

H

O

CH

H3C O C

OH m

Solid state

H

O

CH

Oligomer

FIGURE 6.2 Routes of chemical synthesis of PLGA copolymer.

OH n

COOH R = H or CH3

O C

Polymer

159

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CHAPTER 6 PLGA scaffolds: building blocks for new age therapeutics

limitations arising from the use of azeotropic solvents (Moon et al., 2001; Takahashi et al., 2000; Ajioka et al., 1995, 1998). Ring opening polymerization of lactide, glycolide, and cyclic diesters (Scheme 3 in Fig. 6.2) was observed as another method for the preparation of high molecular weight PLGA in a short reaction time under conditions of metal catalysis and high temperature in the range of 130
C220
C. The various catalysts used here included aluminum isopropoxide, tin (II) 2-ethylhexanoate, tin (II) alkoxides, stannous octoate, powdered zinc, and Lewis acids (zinc chloride, antimony trifluoride). Among these, stannous octoate is the most widely accepted choice due to its acceptance as a food stabilizer by the USFDA (Kowalski et al., 2000; Kricheldorf et al., 1992). This method resulted in a shorter reaction time but the possibility of contamination from metallic residues could not be ruled out. Thus, toxicity concerns in biomedical applications were relatively high with this method. Enzymatic ring opening polymerization was an alternative approach for the preparation of low molecular weight PLGA uncontaminated by any toxic residues. However, this synthetic route utilized a milder temperature, pH, and pressure (Duval et al., 2014). It has been demonstrated that analog PLGA prepared via ring opening polymerization takes a longer time to degrade as compared to random PLGA. A new method has been reported by Li et al., for the preparation of repeated sequence PLGA copolymers having different activities and utilizing 1,3 diisopropyl carbodiimide and 4-(dimethylamino)pyridinium p-toluenesulfonate as agents for catalysis. The PLGA produced by this method presented high control over sequence and stereochemistry and could be modified and manipulated to reduce the rate of hydrolysis. Thus, having positive implications in drug delivery and therapeutics as the release rate of drugs is significantly affected by the degradation rate of the polymer used (Li et al., 2011).

6.4 POLY(LACTIDE-CO-GLYCOLIDE) PROPERTIES The physicochemical as well as biological properties of PLGA significantly depend on various factors, these include the initial molecular weight of the monomers, composition (lactide-glycolide ratio), lactide stereoisomeric composition (L- or DL-lactide), exposure time to water, and storage temperature. PLA exists as two optical isomers of lactide; D and L depending on the position of the pendant methyl group on its alpha carbon. PLGA is available as D-, L- or DL-isomeric forms. PGA, unlike PLA, does not have the methyl side group, thus, it is crystalline in nature; however, PLGA copolymers are largely amorphous and exhibit a glass transition temperature (Tg) in the range of 40
C60
C. It shows glassy behavior and possesses a fairly rigid chain structure. It has also been demonstrated that the Tg of PLGA copolymer tends to decrease with a subsequent decline in lactide content and in its molecular

6.4 Poly(Lactide-co-Glycolide) Properties

weight. PLGA can be processed into any shape and size and is capable of encapsulating molecules of varied sizes (Gentile et al., 2014; Houchin and Topp, 2009; Park, 2006; Avgoustakis, 2005). PLGA is readily soluble in most common solvents, including chlorinated solvents, tetrahydrofuran, acetone, or ethyl acetate. PLGA with a less than 50% glycolide content is easily soluble in common organic solvents, like chloroform, dichloromethane, acetone, ethyl acetate, tetrahydrofuran, and dioxin. PLGA with a glycolide content of 50% or greater is insoluble in most organic solvents. Uncommon solvents, like hexafluoroisopropanol, have been used for characterizing PLGA rich in glycolide content (Danhier et al., 2012; Makadia and Siegel, 2011; Avgoustakis, 2005; Wu and Wang, 2001; Uhrich et al., 1999). The degree of crystallinity of PLGA impacts its degradation rate as well as mechanical properties. Its crystallinity, in turn, depends on its composition (lactide-glycolide ratio) and the stereoisomeric composition of lactide units in the copolymer. Both PLGA composed of 0%75% glycosyl units as well as that with 25%75% glycolide content are found to be amorphous (Vert et al., 1981; Gilding and Reed, 1979). Molecular weight, composition (in terms of lactide/glycolide ratio), the stereochemistry of the lactide, and processing are important factors that govern the mechanical properties of PLGA, such as strength, toughness, and elasticity. PLGA being less crystalline or relatively amorphous has lower mechanical strength than the more crystalline polymers like PLA and PGA (Daniels et al., 1990; Vert et al., 1981). Molecular weight and polydispersity index are key factors that influence the mechanical strength of PLGA as well as its ability to be formulated as a drug delivery carrier and controlling its degradation rate and hydrolysis. However, the type of drug being used also affects the release rate. Crystallinity also affects swelling behavior, mechanical strength, and capacity to undergo hydrolysis, as well as biodegradation rate. When the more crystalline PGA is copolymerized with PLA, the resultant PLGA has a lower degree of crystallinity and a higher hydration rate and hydrolysis. Assuming generalization, the higher the PGA content, the faster the degradation. However, PLGA (50:50) does not comply with this rule and exhibits faster degradation. The degree of crystallinity and melting point are directly related to molecular weight. Commercially available PLGA polymers are usually characterized by their intrinsic viscosity, which is directly related to their molecular weight (Danhier et al., 2012; Houchin and Topp, 2009; Avgoustakis, 2005). PLGA copolymers are categorized as thermoplastic materials and they exhibit sufficient heat stability when protected from moisture. Hence it can be melt processed to produce sutures, orthopedic fixation devices, and drug delivery systems. Prolonged heating at temperatures above 200
C under a vacuum or nitrogen environment results in the degradation of PLGA into lactide and glycolide units. Thermal degradation is a function of time at lower temperatures and is considerably accelerated by impurities, residual monomers, and humidity (Avgoustakis, 2005; Gilding and Reed, 1979).

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PLGA is known to be a biocompatible and nontoxic polymer. These properties were first established by the production of biodegradable sutures. Dexon (PGA) and Vicryl L-PLGA (8:92) are biodegradable sutures that have a history of clinical use older than 30 years. The physicochemical properties of a polymer affect the host response to a polymeric implant. PLGA implantation studies in the bone or soft tissues of animals have shown either no or very mild inflammatory responses that ease out with time. Thus, no toxicity or allergic responses were reported with PLGA (Avgoustakis, 2005; Tiainen et al., 2004; Anderson and Shive, 1997; Visscher et al., 1985; Vert et al., 1984).

6.5 POLY(LACTIDE-CO-GLYCOLIDE) SCAFFOLDS FOR BONE TISSUE ENGINEERING The tissue engineering approach is based on the utilization of a variety of combinations of scaffolds with viable cell types or biomolecules derived from these cell types that facilitate repair and regeneration or restoration of tissues. Here the key is to let the body recognize these electrospun scaffolds introduced into it as its own and then use them for tissue regeneration, thus, resulting in healing (Rezvani et al., 2016). Bone tissue engineering is an emerging field that encompasses several clinical applications, like replacement of bone in orthopedic defects, stabilizing spinal segments, treatment of pseudoarthrosis, bone neoplasms, and tumors. Other applications include craniofacial, maxillofacial, orthopedic, trauma, reconstructive, and neck and head surgery (Amini et al., 2012; Ferrone and Raut, 2012; Dimitriou et al., 2011; Martou and Antonyshyn, 2011). Novel nanofiber scaffolds have gained considerable interest for bone tissue engineering in the past decade. These scaffolds serve as an excellent substitute for the use of allografts, autografts, and xenografts, and are based on biomaterials with the properties of bioactivity, biocompatibility, biodegradation, osteoconduction, and osteoinduction (Li et al., 2013; Stevens, 2008). Electrospinning is a technique that allows for the fabrication of suitable scaffolds from a wide range of nanobiomaterials having the potential to control drug release characteristics in defective tissues (Rezvani et al., 2016). These materials might be metals (superior mechanical properties), ceramics (excellent biocompatibility), and biodegradable polymers (superior physicochemical as well as biological properties making them suitable for tissue engineering) (Gentile et al., 2014; Cordonnier et al., 2011; Salgado et al., 2004). Saturated poly(α-hydroxy esters), like PLA, PGA, and PLGA, are the most commonly employed synthetic biodegradable polymers for the preparation of 3D scaffolds for bone tissue engineering. The chemical properties of these polymers enable their hydrolytic degradation via deesterification. The monomeric components of these polymers formed upon degradation are removed by the body’s natural pathways. Out of these, PLGA is the most preferred polymer for the

6.5 Poly(Lactide-co-Glycolide) Scaffolds for Bone Tissue Engineering

fabrication of bone constructs as it offers better control over degradation properties by modifying or varying its monomeric composition (Mano et al., 2004; Lin et al., 2002). Many innovative approaches and techniques with the aim of preparing improved bone constructs have come to the fore in the recent past, some of them have been discussed here (Table 6.1). PLGA is most commonly used for porous scaffolds, films, spun fibers, and hydrogels or injectables for bone tissue applications (Fig. 6.3). Their brief descriptions and methods of preparation have been discussed here.

6.5.1 POROUS SCAFFOLDS Porous scaffolds fabricated from PLGA matrices are popular choices for tissue engineering and regenerative medicine because of their mechanical strength, biocompatible and biodegradable nature, tunable rate of degradation, and processability (Gentile et al., 2014; Pan and Ding, 2012). The past decade has observed some facile fabrication techniques at room temperature being put to use for the production of 3D porous scaffolds. Some of these techniques include gas foaming (Harris et al., 1998), freeze drying (Whang et al., 1995), phase separation (Zhang and Ma, 1999), porogen leaching (Mikos et al., 1994), fiber bonding (Mikos et al., 1993), electrospinning (Zeng et al., 2003), 3D printing, and fused deposition modeling (Zein et al., 2002). Amongst these, porogen leaching has been popularly adopted by several researchers due to its ease of operation and the efficient control of porosity and pore size on account of variations in the size and amount of porogen used. The use of porogen leaching in combination with a molding technique still remains a challenge (Gentile et al., 2014; Pan and Ding, 2012). Several researchers have also reported the use of particulate leaching for scaffold fabrication. However, it is associated with numerous disadvantages, like incomplete solvent removal by evaporation, greater suitability to produce thin scaffolds (up to 2 mm thickness), and absence of interconnectivity and open-pore structure in scaffolds that require low porosity (McMahon et al., 2013; Zhang et al., 2011). In order to avoid or reduce degradation in the processing of biodegradable polymers, the use of high temperatures must be avoided. Several improvements of these major fabrication techniques at moderate or even room temperature have been reported. Some of these are discussed here. 1. Modified thermal compression molding-particulate leaching method at modified temperature: Biomedical and clinical applications of tissue engineering require scaffolds that resemble the anatomical shapes of given organs or tissues with a defect. A new and feasible fabrication method was developed by the Fudan group which led to the simultaneous formation of an internal interconnected pore structure and an external complicated anatomical shape of porous scaffolds (Wu et al., 2005). This method was called the modified thermal compression molding-particulate leaching approach

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Table 6.1 PLGA Intervention in Bone Tissue Applications S. No.

Scaffold System

Purpose/Findings

Reference

1

Hydroxyapatite based biodegradable mPEGPLGA nanoparticles of risedronate prepared using a water miscible dialysis method Thin layer PLGA coated beta-tricalcium phosphate (β-TCP) scaffold with sustained release of vascular endothelial growth factor (VEGF) 3D PLGA/ nanohydroxyapatite/β-TCP scaffolds prepared by 3D prototyping technique using 3D Bio-Extruder equipment Simvastatin loaded PLGA microparticles entrapped in porous freeze-dried chitosan-gelatin scaffolds

Sustained delivery of risedronate for the treatment and prevention of osteoporosis and to minimize its adverse effects upon oral administration In vitro bone regeneration

Rawat et al. (2016)

Improving preosteoblast cell (MC3T3-E1) adhesion, proliferation, and differentiation

Roh et al. (2016)

Localized and controlled release of simvastatin from scaffolds that promotes hFOB cell proliferation and osteoblast differentiation Composite scaffolds with sufficient mechanical and physicochemical properties to support defect regeneration and to maintain their stability during the formation of new tissue PLGA incorporation enhanced the mechanical properties of brittle CPC and demonstrated enhanced structural and bone formation parameters (osteoid volume, osteoid surface); enhanced bone mineral density and biomechanical compression strength; and decreased bone erosion PCL-PLGA-duck beak scaffolds showed potential for fracture healing by promoting new bone formation in rabbit radius by inducing repair induction

Gentile et al. (2016)

2

3

4

5

Hydroxyapatite-gelatin scaffolds incorporated with dexamethasone-loaded PLGA microspheres fabricated using a freeze casting technique

6

Brushite forming calcium phosphate cement reinforced with PLGA fibers in a minimally invasive sheep lumbar vertebroplasty model

7

Poly(E-caprolactone)-PLGAduck beak scaffolds

Khojasteh et al. (2016)

Ghorbani et al. (2016)

Maenz et al. (2017)

Lee et al. (2016)

(Continued)

6.5 Poly(Lactide-co-Glycolide) Scaffolds for Bone Tissue Engineering

Table 6.1 PLGA Intervention in Bone Tissue Applications Continued S. No.

Scaffold System

Purpose/Findings

Reference

8

Hypoxia-mimetic agent deferoxamine (DFO) loaded PLGA scaffolds

Jia et al. (2016)

9

Simvastatin loaded PLGA microspheres combined with a rapidly absorbable calcium sulfate bone substitute; the effect on bone healing was studied

10

PLGA nanoparticleschitosan bioactive glass scaffold prepared by lyophilization Metallic magnesium particle/PLGA composite scaffold as dental bone grafting material prepared using a solvent casting, salt leaching method

DFO dramatically stimulated bone formation and angiogenesis in critically sized osteoporotic femur defects in ovariectomized rats. DFO promoted healing of osteoporotic bone defects via enhanced angiogenesis and osteogenesis This system showed good osteoconductive and osteoinductive properties and showed promise for bone regeneration. It promoted healing in critical-sized calvarial bone defects in rats. The production of bone morphogenetic protein-2 and neovascularization were enhanced in the area of the defect Controlled release reservoirs for favorable biodistribution of model drugs to implanted bone Potential for dental socket preservation and orthopedic bone regeneration. These scaffolds could reduce inflammation, unlike clinically used PLGA devices, and enhanced osteogenesis, unlike previous magnesium devices These composites possessed good biocompatibility, and demonstrated the ability to promote cartilage formation and reconstruction Enhanced bone regeneration and in vivo healing capacity demonstrated upon implantation in critical-sized calvarial defects in a rat model

11

12

PLGA-Hydroxyapatite-Zein composite scaffolds fabricated by electrospinning

13

Pro-osteogenic factor rhBMP-2-PLGA-alginate microparticles in collagenhydroxyapatite scaffolds, which had been previously optimized for bone regeneration, as delivery platforms to produce a device with enhanced capacity for bone repair

Fu et al. (2015)

Nazemi et al. (2015)

Brown et al. (2015)

Lin et al. (2015)

Quinlan et al. (2015)

(Continued)

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Table 6.1 PLGA Intervention in Bone Tissue Applications Continued S. No.

Scaffold System

Purpose/Findings

Reference

14

Tetracycline based PLGA nanoparticles (TC-PLGA NPs)

Wang et al. (2015a)

15

Triple layered PLGA based composite membrane prepared in a single step using a combination of techniques, like solvent casting and thermally induced phase separationsolvent leaching Biomorphic PLGA/ nanohydroxyapatite composite scaffolds were successfully prepared using cane as a template Boron incorporated PLGA scaffolds

Improved curative effects of Simvastatin on recovery of bone mineral density and better bone targeting when delivered through TC-PLGA NPs as compared to either Simvastatin loaded PLGA NPs or Simvastatin alone Periodontal guided bone regeneration

Support cell attachment, proliferation, and differentiation (MC3T3-E1)

Qian et al. (2014a)

Improved in vitro proliferation, attachment, and calcium mineralization of rADSCs. Enhanced bone regeneration via enhanced osteocalcin, VEGF, and collagen type I protein levels in a femur defect model. Enhanced rate of bone healing and high potential in functional bone tissue engineering Scaffold for mineralized and unmineralized tissues Enhanced targeting ability of zoledronate and its enhanced endocytosis demonstrated by in vitro and in vivo studies for bone metastasis Suitable scaffold for engineering bone tissue in vitro

˘ Dogan et al. (2014)

16

17

18 19

20

Nanofibrous PLGAhydroxyapatite system Zoledronate anchored PLGA-PEG nanoparticles

Poly(b-hydroxybutyrate-cob-hydroxyvalerate) microspheres embedded in PLGA matrix

Jamuna-Thevi et al. (2014)

Kolluru et al. (2013) Chaudhari et al. (2012)

Huang et al. (2010)

6.5 Poly(Lactide-co-Glycolide) Scaffolds for Bone Tissue Engineering

FIGURE 6.3 PLGA interventions in bone tissue engineering.

(MTCM-PL). In this method, conventional solvent casting was used for the initial preparation of a polymer-particulate mixture (PLGA and NaCl particles) which were then compression molded in a specially designed combined flexible-rigid mold. This method allowed for PLGA scaffolding to be carried out at moderate temperatures (i.e., a temperature higher than the Tg of the PLGA but lesser than its flow temperature). Particulate leaching in the final step resulted in a strongly interconnected porous scaffold having excellent mechanical properties (Wu et al., 2005; Wu and Wang, 2001). 2. Room temperature compression molding-particulate leaching method: The use of room temperature provides for convenience in scaffolding and prevents the degradation of the polymer. A new method was reported that utilized compression molding-particulate leaching at room temperature (RTCM-PL). In this technique, a highly concentrated polymer solution was used, unlike polymer melt in other techniques. The mixture of polymer and porogen was molded. Two solvents were used in this method; the organic solvent resolves the polymer but not the porogen and water resolves the porogen particles and not the polymer. RTCM-PL was found to be a suitable technique for fabricating complex scaffolds at low pressure and temperature conditions (Jing et al., 2006; Wu and Ding, 2004). 3. Room temperature injection molding-particulate leaching method: An approach based on injection molding and particulate leaching at room

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temperature (RTIM-PL) with the use of a solvent of PLGA was developed for the successful fabrication of highly porous scaffolds with complicated shapes (tubular, etc.,) avoiding the use of high pressure and preventing any thermal degradation. This was a precise and repeatable method for producing PLGA scaffolds (Wu et al., 2006). 4. Design of new porogen by control of pore shape and inter-pore connectivity: Inter-pore connectivity is an important parameter for cell loading and cell induction and it usually increases with an increase in porosity. The design of a unique porogen could be sought after as a strategy to improve the interpore connectivity of PLGA scaffolds under defined conditions of porosity. The porogen technique based on spheric paraffin particles reported by Ma et al., was combined with compression molding at room temperature to successfully fabricate polyester scaffolds of large size. The scaffolds so developed had ordered macroporous structures, improved inter-pore connectivity, and porosity in the range of 77%97% based on porogen content. Gelatin particles have been used as a porogen in some cases (Gong et al., 2007; Zhang et al., 2006; Ma and Choi, 2001; Mikos et al., 1994; Chen et al., 2001). 5. Thermally induced phase separation (TIPS): This technique is based on utilizing changes in thermal energy to facilitate demixing of a homogenous solution of polymer into a biphasic or multiphasic system domain. PLGA biomimetic scaffolds based on nano-biphasic components and consisting of hydroxyapatite and β-tricalcium phosphate powders as reinforcement material were reported to be fabricated by TIPS (Okamoto and John, 2013; Forgacs and Sun, 2013; Ebrahimian-Hosseinabadi et al., 2011).

6.5.2 FIBROUS SCAFFOLDS Scaffolds in the form of fibers have been endowed with superior mechanical strength and biocompatibility and have good suitability in bone tissue engineering. Transforming PLGA into a textile structure, like fibers, is a complicated process and is dependent on various factors, like changes in the structure of copolymer while being processed. Copolymer extrusion into monofilament and multifilament for fabrication of micro and mono fiber composite scaffolds can be carried out using different fiber-forming techniques, such as melt spinning, wet spinning, and electrospinning. The technique adopted for fiber formation has an effect on the properties of the fiber produced (Azimi et al., 2014; Puppi et al., 2011). Some of these techniques include: 1. Melt Spinning: This was one of the first techniques used to prepare fibers and various researchers have utilized it for fabricating PLGA fibers under different conditions. In this method, the polymer is initially melted and filtered followed by extrusion via spinneret through the spinneret hole at melt temperature. There is a draw zone where the extruded filaments are

6.5 Poly(Lactide-co-Glycolide) Scaffolds for Bone Tissue Engineering

cooled to solidification temperature and further to below Tg. In the final step, these filaments come to take-up bobbins, and their temperature is less than Tg (Cicero et al., 2002a,b; Cicero and Dorgan, 2001; Yuan et al., 2001; Fambri et al., 1977). In a study, a fiber matrix of PLGA (85:15) was fabricated as scaffolds for cell culture by centrifugal melt spinning method (Wang et al., 2011). In another study, the melt spinning technique was used for the preparation of sutures in combination with an antitumorigenic drug by melt extrusion of a PLGA (75:25) mixture (Intra et al., 2011). 2. Wet Spinning: Solution spinning, dry spinning, and wet spinning techniques are utilized for nonmelting polymers. Melt spinning becomes unfeasible when the copolymer degrades during melting or due to the thermal instability of the melt. Solution spinning techniques both dry and wet involve filtration and deaeration of the polymer solution which is finally pumped via spinneret. Solvents are evaporated off in the dry spinning method; however, wet spinning involves polymer coagulation in a fluid compatible with the spinning solvent but that is itself not the solvent for the polymer. When the viscose polymer solution reaches the coagulation bath then phase separation occurs due to solvent outflow and nonsolvent inflow and it leads to the polymer being precipitated as fibrils (Azimi et al., 2014; Arbab et al., 2008, 2011). Many investigators have employed a wet spinning technique for fabricating PLGA filaments. A study reported the preparation of hollow fibrous scaffolds combined with human bone marrow stromal cells for bone repair and regeneration using the wet spinning technique (Morgan et al., 2007). Another simple wet spinning methodology for PLGA monofilaments was reported (Nelson et al., 2003). A wet inversion method was adopted by Wen et al., for fabricating permeable biodegradable hollow PLGA fiber membranes. Similar fiber membranes were prepared using dry-wet and wet spinning techniques in another study where 1,4-dioxane and 1-methyl-2-pyrrolidinone were used as solvents and water as a nonsolvent (Ellis and Chaudhuri, 2007; Wen and Tresco, 2006). 3. Electrospinning: This is a simple and versatile method for the fabrication of ultrathin nonwoven fibers of PLGA and other polymers with sizes in the submicron range. The method usually involves subjecting a polymeric solution or melt to strong electric fields followed by ejection of the liquid phase polymer through a nozzle. These ejected fibers are significantly reduced in size as they move toward a collector (Azimi et al., 2014; Huang et al., 2003). Fibers produced by electrospinning have several advantages over those produced by other techniques. These have small inter-fibrous pore size with tunable porosity, high surface to volume ratio, and the capability of achieving desirable characteristics. Electrospun fibers are thought to have a significant use in bone regeneration by virtue of their biodegradability, mechanical properties, and osteoconductivity (Peng et al., 2011; Mouthuy et al., 2010; Nie and Wang, 2007; Wang et al., 2002).

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6.5.3 HYDROGELS Hydrogels can be described as products composed of hydrophilic polymers which hold water in large amounts in their 3D network chains making them two- or multicomponent systems. These hydrogels have enormous water content holding capacities, especially in the swollen state where the mass fraction of water is greater than the mass fraction of polymer depending on the type of polymer in use. Hydrogels can be prepared using several methods, including single step methods, such as polymerization and parallel crosslinking of multifunctional monomers; and multistep methods that involve the synthesis of polymer molecules with reactive groups and their subsequent crosslinking, as well as by reaction of polymer molecules with suitable crosslinking agents (Ahmed, 2015). Hydrogels are widely used scaffolds for tissue engineering and drug delivery applications. A temperature sensitive scaffold system with a Tg below 37
C for in vivo bone repair was reported to be prepared by blending PLGA with a plasticizer, like poly(ethylene glycol) (PEG) (Rahman et al., 2014; Dhillon et al., 2011). Composite scaffolds composed of HA and PEG copolymer named PLGAG-PEG were fabricated by Lin et al. (2012). Multiblock amphiphiles, like triblock (ABA) copolymers, containing biodegradable PLA, PGA, or their copolymers have surfaced as the most popular substitutes for poloxamer-based copolymers. PEG-PLGA-PEG is an example of where the hydrophobic block is at the center (B) block, while PLGA-PEG-PLGA fabrications contain the hydrophobic block in the arm (A) block (Hoare and Kohane, 2008). PLGA-PEG-PLGA type scaffolds afforded a diffusion controlled release for hydrophilic compounds, while hydrophobic compounds followed an initial diffusion controlled release followed by a prolonged polymer degradation controlled release pattern (Qiao et al., 2005).

6.5.4 INJECTABLE MICROPARTICLES Injectable scaffolds for regenerative medicine should possess mechanical properties, a pore diameter, and porosity that support tissue formation and offer a minimally invasive therapy. Porosity and pore diameter assume significance as they influence cell attachment, cellular proliferation, and migration, in addition to affecting the delivery of therapeutic agents and the removal of waste. Such scaffolds should ideally have sufficient strength to retain their structure and exhibit multiscale porosity (microporosity and macroporosity), with pore diameters in the range of ,20 to .100 μm. The versatile properties of PLGA make it a popular candidate for the fabrication of such scaffolds owing to its wide-ranging porosity and mechanical properties. Discreet polymer microspheres are popularly used as injectable scaffolds for tissue engineering applications. Microspheres can be fabricated using different biodegradable polymers (chitosan, gelatin, and PLGA) through different techniques and are most commonly reported to encapsulate cells and growth factors for tissue repair in bone, skin, and the brain. Qutachi et al., reported the fabrication

6.6 Poly(Lactide-co-Glycolide) Scaffolds in Anticancer Therapy

of porous PLGA microspheres (average size 84 6 24 μm) that could form solid porous scaffolds at body temperature and function as successful injectable carriers for in vitro NIH-3T3 cell attachment and proliferation (Qutachi et al., 2014). Jain et al., reported an in situ fabrication of injectable PLGA microspheres by yet another method. Here a stable dispersion of PLGA pre-microspheres in a vehicle mixture was initially prepared, which upon injection, came into contact with water from the aqueous buffer or physiological fluid; thus, the pre-microspheres were hardened into solid matrix-type microparticles with a drug entrapped. Drug release from these microspheres was controlled (Jain, 2000). PLGA microspheres have the advantage of modifiable biodegradation properties; where manipulation of biodegradation can be achieved from some weeks to several months-time by bringing about variations in its monomer composition and, thus, making it possible to influence the rate of release and degradation of the incorporated drugs and biomolecules (Gentile et al., 2014). PLGA microspheres have been prepared using a traditional O/W emulsification method and several modifications have been made in order to attain injectable microspheres with biomimetic properties through the addition of hydroxyapatite. In another study apatite-coated injectable PLGA microspheres containing osteoblasts were prepared by W/O/W, following immersion in simulated body fluid for 5 days at 37
C. These were successful in significantly enhancing new bone formation in mice. Another study demonstrated the use of injectable hydroxyapatite PLGA microspheres for delivery of the antiosteoporotic drug alendronate which was fabricated using a solid/O/ W or W/O/W technique. These microspheres improved osteoblast proliferation and upregulated ALP (Wang et al., 2013,b; Shi et al., 2009; Kang et al., 2008).

6.6 POLY(LACTIDE-CO-GLYCOLIDE) SCAFFOLDS IN ANTICANCER THERAPY Cancer; a major cause of world mortality is a condition characterized by the uncontrolled growth and spread of abnormal cells. Current treatment modalities include chemotherapy, surgical procedures, radiation, and hormone therapy where chemotherapy is one of the most obvious choices for cancer treatment. The major drawback of conventional chemotherapy is its highly nonspecific targeting, which leads to high vulnerability of healthy cells toward the untoward effects of chemotherapeutic drugs, thus, imposing restrictions on the maximum allowable dosage of these drugs. Nanoparticle based therapeutics has emerged as a novel strategy for preferential targeting of chemotherapeutic agents in large doses and therapeutic genes into malignant cells while sparing healthy cells. This treatment strategy has great potential in the field of oncology as it minimizes or overcomes the limitations of conventional therapy, like severe adverse effects and toxicity, and cancer cell drug resistance and undesirable biodistribution (Prabhu et al., 2015). Polymeric nanoparticulate therapy has gained attention in the recent past as

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attractive vectors for tumor targeting. These polymeric nanoparticles have defined morphology and composition in their core as well as periphery. The chemotherapeutic moiety may either be surface conjugated to them or encapsulated within their polymeric core. Functionalization of these polymeric nanoparticles further extends their systemic residence and reduces their nonspecific distribution. Functionalization also enables tissue targeting with ligands, like aptamers, peptides, antibodyantibody fragments, and small molecules. Various types of polymeric nanoparticles, like polymer conjugates, polymer hybrids, solid polymeric particles, polymeric micelles, dendrimers, polyplexes, and polymersomes, might be used for the treatment of cancer. Polymeric nanoparticles enabled therapy is capable of sustained delivery and targeting of a wide range of chemotherapeutic drugs to the tumorigenic site with enhanced efficacy and minimized adverse effects. This strategy also provides protection to the drug against rapid metabolism and clearance, thus, improving its stability and site specificity. Polymers used for fabricating these nanocarrier systems may be natural or synthetic in origin. Such nanocarrier systems provide protection to drugs against fast metabolism during circulation and clearance by the liver, kidney, and reticuloendothelial system (RES), which further enhances a given drug’s stability and target specificity. In these scaffolds, the active principle or drug might be encapsulated, entrapped, or adsorbed onto the polymer matrix as in solid colloidal systems. Their structure depends on the method of their preparation. They might be either nanosphere (therapeutic agent dispersed throughout the particle) or nanocapsules (vesicle based reservoirs where the therapeutic agent is entrapped in a cavity (aqueous/ oily) enveloped by a singular polymer membrane). Stimuli-sensitive polymeric nanocarriers are a new advancement in the field of cancer treatment; these can move across intracellular delivery barriers to release a drug in response to the microenvironmental trigger of the tumor cell. Polymers used for preparing these nanocarriers might be natural or of synthetic origin. The most commonly employed polymers are PLGA, PLA, PGA, PCL, poly(D, L-lactide), chitosan, and PLGA-PEG, which can be used for passive and ligand-targeted delivery of therapeutic moieties (Swain et al., 2016; Gupta et al., 2013; Chan et al., 2010; Alexis et al., 2008; Parveen and Sahoo, 2008; Letchford and Burt, 2007; Sinha et al., 2006). PLGA has been in safe clinical use since its approval by the FDA in 1969. PLGA particles for drug release were approved in 1989 (Lupron Depot) and since then have been commonly used due to their controlled release properties for eliciting a persistent therapeutic effect (Zhou et al., 2012). PLGA nanoparticles can be prepared using a nanoprecipitation/solvent diffusion method, where during the particle preparation process, target drugs can be effectively entrapped in the precipitated rigid core of nanoparticles. PLGA nanoparticles are attractive candidates for tumor targeted therapy and imaging. Incorporating and encapsulating drugs in PLGA matrices; enables their slow release over a prolonged period of time, which leads to a reduced frequency of dose administrations, reduced discomfort and peak associated adverse effects. This results in protection of the drug or bioactive

6.6 Poly(Lactide-co-Glycolide) Scaffolds in Anticancer Therapy

principle within the body, and maintenance of its more constant blood levels. These benefits eventually increase overall patient compliance. In order to overcome extracellular and intracellular barriers; nanoparticles require specific properties which can be achieved through functional modifications. These PLGA scaffolds can be surface functionalized to improve their therapeutic benefits. These are stabilized with a surface hydrogel layer of PEG which can be functionalized with numerous therapeutic moieties and ligands for site-specific targeting. Stabilization with PEG minimizes opsonization by the mononuclear phagocytic system and prolongs systemic residence (Dias et al., 2015; Chung et al., 2010). In a study, the surface of PLGA nanoparticles were decorated with PEI utilizing a cetyl derivative to improve surface functionalization and siRNA delivery. PEI is a functional polycation that presents a strategy to bind nucleotide based drugs and facilitate endosomal escape (Andersen et al., 2010). PLGA nanoparticles being solid, unlike lipid polyplexes, provide better stability and protection to nucleic acids against degradation while in systemic circulation and also allows for their long-term storage and convenient usage in clinical settings. These advantages make them promising vectors for the delivery of oligonucleotides, like siRNA, for tumor targeting. Zhou et al., reported the fabrication of PLGA nanoparticles for siRNA delivery with multiple functionalities. These octafunctional PLGA particles demonstrated efficient tumor targeting and significant antitumor efficacy on human lung cancer cell line A549 in a xenograft mice model (Zhou et al., 2012). Murata et al., reported the development of anti-VEGF siRNA loaded PLGA microspheres with a transfection agent (arginine or branched PEI). These sustained release PLGA systems were prepared using a W/O/W drying method and were investigated for their antitumor efficacy in mice bearing S-180 tumors. siRNA release from PLGA microspheres could be sustained for over 1 month and an improved sustained suppressive effect on VEGF gene expression was demonstrated (Murata et al., 2008). PLGA nanoparticles have a rigid core which sometimes lowers their tumor targeting efficacy as they are highly localized in the liver. In a study investigating cellular distribution of PLGA nanoparticles upon injection, it was found that these PLGA particles were majorly taken up by Kupffer cells followed by liver sinusoidal endothelial cells and hepatic stellate cells. About a meager 7% hepatocytes were found to take up these PLGA nanoparticles. These particles showed increased retention in liver sinusoidal endothelial cells and hepatic stellate cells when Kupffer cells were found to be depleted by clodronate liposomes, however, the same was not true for hepatocytes (Park et al., 2016; Chung et al., 2010). Block copolymers based on PEG and poly(α-hydroxy acid), like PLGA, have shown potential for localized cancers, like brain, ovarian, and esophageal cancers. PLGA-b-PEG-b-PLGA (a thermosensitive ABA block copolymer) used for drug delivery applications reversibly transitions into a gel at body temperature. Both hydrophilic as well as hydrophobic drugs can be incorporated into these scaffolds in a solution state and injected at a desired site where it

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subsequently changes to a gel state and provides sustained release of the drug for a prolonged period of time and eventually biodegrades into nontoxic byproducts (Cho et al., 2016). Their most significant clinical application has been in the treatment of localized cancers as an adjunct or alternative to surgical oncology. These gel scaffolds can be surgically resected into tumor cavities and also be used in a neoadjuvant setting to reduce tumor burden and enhance tumor resection. PLGA-b-PEG-b-PLGA sol-gel injections guided by endoscopic and ultrasonographic procedures have been attempted for the treatment of pancreatic cancer. Gliadel wafer has been used a successful strategy to treat brain tumors. PLGAb-PEG-b-PLGA based Regel has been extensively studied for the local treatment of esophageal cancer. Oncogel is Regel at 23% weight polymer/weight water containing paclitaxel at 6.0 mg/mL. It has also been used for endoscopy guided treatment of esophageal cancer (Cho et al., 2016; Wolinsky et al., 2012; Elstad and Fowers et al., 2009; Matthes et al., 2007). Various types of PLGA interventions for the treatment of cancer have been enumerated in Table 6.2.

6.7 POLY(LACTIDE-CO-GLYCOLIDE) INTERVENTIONS IN CENTRAL NERVOUS SYSTEM DELIVERY An increase in the aging population has led to an upsurge in the incidence of many neurodegenerative diseases. In a scenario like this, drug delivery for brainrelated disorders, like neurodegeneration and cerebral tumors, assumes great significance. The current treatment methodologies for the central nervous system (CNS) have rather remained underdeveloped despite extensive efforts in the past two decades largely due to restrictions imposed by the bloodbrain barrier (BBB) for the penetration of drugs. Novel drug delivery has emerged as a popular treatment strategy for brain-related diseases. Polymeric nanoparticles, being biocompatible and biodegradable, ensure safer therapies for brain disorders. PLGA based nanoparticulate systems provide a new platform for the development of potential targeted therapies directed to cross the BBB (Fornaguera et al., 2015; Grover et al., 2013; Krol, 2012; Danhier et al., 2012). PLGA nanoparticles can be fabricated using an emulsification-solvent evaporation technique. Here the hydrophobic drug and the polymer are dissolved in an organic solvent like dichloromethane. An aqueous solution of the surfactants, like polysorbate-80 or poloxamer-188, is added to the polymer solution to produce an O/W emulsion which upon sonication or homogenization changes into nanosized droplets. The solvent is evaporated out or extracted and nanoparticles are collected postcentrifugation. Another technique called the double emulsion W/O/W method was developed to incorporate hydrophilic drugs, like proteins, peptides, and nucleic acids. The nanoprecipitation method, also known as interfacial deposition method, is yet

6.7 Poly(Lactide-co-Glycolide) Interventions

Table 6.2 PLGA Interventions in Cancer Treatment Breast Cancer • Folic acid decorated and PEGylated PLGA nanoparticles for delivery of 5-Fluorouracil against breast cancer in MCF-7 cells and colon cancer in HT-29 cells (El-Hammadi et al., 2017). • Irinotecan loaded PLGA nanoparticles with hyaluronic acid moieties against CD44overexpressing breast carcinoma cells (HS578T) (Giarra et al., 2016). • Disulfiram loaded (mPEG-PLGA/PCL) mixed nanoparticles for antitumor efficacy on MCF-7 cells (Song et al., 2016). • Chlorambucil loaded PLGA nanoparticles for antitumor efficacy on MCF-7 cells (Dias et al., 2015). • Doxorubicin loaded mPEG-PLGA nanopolymersomes against murine breast tumors (Alibolandi et al., 2015). • Doxorubicin loaded, low molecular weight cell-penetrating peptide (protamine) assisted PLGA nanoparticles against multidrug resistance breast cancer (Wang et al., 2014b). • Mitoxantrone (DHAQ) or fluorescence agent Rhodamine B (Rb) loaded cyclic peptide (arginine-glycine-aspartic-glutamic-valine acid, cRGD)-modified monomethoxy (PEG)PLGA-poly (L-lysine) nanoparticles (Liu et al., 2012). • Antihuman epidermal growth factor receptor (HER-2, ErbB2) antibody anchored PLGAPEG immuno-nanoparticles (Dhankar et al., 2011). • Monoclonal antibody bound PLGA immuno-nanoparticles (Kocbek et al., 2007). Ovarian Cancer • APRPG (Ala-Pro-Arg-Pro-Gly) peptide-modified PEG-PLGA nanoparticles encapsulating inhibitors of angiogenesis TNP-470 with tumor suppressing effects in SKOV3 cells (Wang et al., 2014a). • Surface smooth and porous PLGA microspheres original and collagen I coated as carriers for 3D culture of ovarian cancer HO8910 cells (Zhang et al., 2014). • Epidermal growth factor modified mPEG-PLGA-polylysine (mPEG-PLGA-PLL) encapsulated cisplatin nanoparticles for antitumor efficacy against SKOV3 cells (Wang et al., 2013d). Prostate Cancer • PLGA-PEG nanoparticles conjugated to anti-CD24 for targeted delivery of docetaxel in luciferase expressing PC3M prostate cancer tumor (Bharali et al., 2017). • Bicalutamide loaded folic acid conjugated chitosan functionalized PLGA against prostate cancer (Dhas et al., 2015). • Small PLGA nanoparticles (4995 nm) loaded with Paclitaxel (Broc-Ryckewaert et al., 2013). • PLGA-b-PEGCOOH nanoparticles conjugated to the A10 RNA aptamer that binds to the prostate specific membrane antigen (PSMA) (Cheng et al., 2007). Colon Cancer • Folic acid decorated and PEGylated PLGA nanoparticles for delivery of 5-Fluorouracil against colon cancer in HT-29 cells (El-Hammadi et al., 2017). • Paclitaxel loaded PLGA nanoparticles after surface conjugation with wheat germ agglutinin against colon cancer cells (Wang et al., 2010). • Cisplatin loaded biodegradable PLGA nanoparticles for anticancer effects against DHD/ K12PROb adenocarcinoma colon cell line (Moreno et al., 2010). (Continued)

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Table 6.2 PLGA Interventions in Cancer Treatment Continued Lung Cancer • Disulfiram loaded PLGA nanoparticles against NSCLC cells (Najlah et al., 2016). • LFC131 peptide conjugated sodium carboxylmethyl cellulose coated PLGA nanoparticles loaded with doxorubicin delivered to CXCR4 expressing lung cancer cells (Chittasupho et al., 2014). • Doxorubicin loaded inhalable highly porous large PLGA microparticles surface-attached with TRAIL (Kim et al., 2012, 2013). • Paclitaxel loaded TPGS-functionalized PLGA nanoparticles (Wang et al., 2013a). • Paclitaxel and isopropyl myristate loaded PLGA nanoparticles surface conjugated with wheat germ agglutinin against malignant (A549 and H1299) and normal (CCL-186) pulmonary cells (Mo and Lim, 2005). Liver Cancer • Genistein loaded star shaped copolymer mannitol-functionalized PLGATPGS nanoparticles for liver cancer treatment in liver cancer cell line and hepatoma-tumorbearing nude mice (Wu et al., 2016). • Disulfiram loaded polysorbate 80-stabilized PLGA nanoparticles against hepatocellular carcinoma on Hep3B cells (Hoda et al., 2016). • Oleanolic acid loaded PLGATPGS nanoparticles against liver cancer using HepG2 cells (Gao et al., 2016). Pancreatic Cancer • Anthothecol (a limonoid isolated from Khaya anthotheca) loaded PLGA nanoparticles against pancreatic cancer stem cells (Verma et al., 2015). • Gemcitabine loaded PLGA nanospheres against MiaPaCa-2 and ASPC-1 pancreatic cancer cells (Jaidev et al., 2015). Others • Bromelain loaded hyaluronic acid-grafted-PLGA copolymer for antitumor effects against Ehrlich ascites carcinoma via targeting CD44 receptors (Bhatnagar et al., 2016). • Curcumin loaded PLGA/poloxamer nanoparticles for antitumor effects in mesothelioma cells (Mayol et al., 2015).

another method. Here the solution of polymer and drug in an organic solvent, like acetone, is added dropwise to water. The organic solvent is evaporated off and nanoparticles are collected as pellets after centrifugation. PLGA nanoparticles are sometimes also prepared using a spray drying technique. Drug loading into nanoparticles can be achieved either during their preparation or later by adsorption onto the prepared particles (Danhier et al., 2012). BBB is an anatomic barrier that aims to limit the penetration of harmful substances into the brain. It is presented by the cerebrovascular endothelium sealed with tight junctions. There are other additional supportive cells which include pericytes, astrocyte end-feet, and a basal lamina. The basal lamina behaves like an extracellular matrix and provides a scaffold for cell migration, mechanical support for cell attachment, and separation of adjacent tissue. It imposes a mesh, creating an obstacle for the entry of microorganisms, hydrophilic molecules, and nanoparticles into the brain. The major challenge for PLGA based

6.7 Poly(Lactide-co-Glycolide) Interventions

therapies is to overcome the highly controlled access to the brain. This restricted access is majorly based on these factors (Grover et al., 2013; Krol, 2012): 1. 2. 3. 4. 5.

The presence of tight junctions that seal the intercellular gap Decreased pinocytosis rate from the luminal side No fenestration which blocks the intercellular passage of the endothelium The presence of an enzymatic barrier (second line protection) The presence of an efflux transporter system (e.g., P-glycoproteins).

Polymeric nanoparticles are subjected to low brain delivery due to various reasons. The major limiting factors include the modifications imposed on nanoparticles surface by bodily fluids and in transit across cellular compartments after administration of these systems. Physical parameters including blood flow, the ratio of diseased to healthy cells, and residence time at the desired site are other major limiting factors. Biological factors that restrict brain delivery include highly selective physiological barriers, lack of specific cellular biomarkers for disease cells, and limited transporters for specific tissues. Differences in pH and the presence of perineuronal nets surrounding specific neurons further cause restricted delivery of drugs or molecules to the brain (Krol, 2012). Cellular internalization of PLGA nanoparticles is effected via fluid phase pinocytosis along with clathrin-mediated endocytosis. These PLGA based nanocarriers are capable of entering the cytoplasm rapidly within 10 minutes of incubation by escaping the endolysosomes which facilitate their interaction with vesicular membranes resulting in transient and localized membrane destabilization and eventually the escape of nanoparticles into the cytosol (Danhier et al., 2012). Most hydrophobic nanocarriers are identified by the RES as foreign bodies and are, therefore, eliminated. These nanocarriers when coated with molecules that provide a hydrophilic layer on their surface make them less vulnerable to recognition by the RES. This process is referred to as “surface modification” and the most commonly used moieties used for this include PEG, poloxamer, poloxamine, and chitosan. Amongst these usage of PEG increases the blood residence of nanoparticles in addition to providing excellent biocompatibility. Surface modification of these polymeric PLGA nanoparticles inhibits electrostatic and hydrophobic interactions which help opsonins to bind to particles. It also helps in site-specific targeting (tumor/organs) by enhanced selective cellular binding and internalization via receptor-mediated endocytosis. Targeting moieties (ligands) can be grafted to nanoparticle surfaces via PEG linkage. Surface modification also enables charge reversals as surface charge plays an important role in cellular interactions and the uptake of nanoparticles. Negatively charged PLGA nanoparticles can be surface modified by PEGylation or by chitosan coating to produce neutral or positively charged particles respectively (Grover et al., 2013; Danhier et al., 2010, 2012; Tahara et al., 2009; Owens and Peppas, 2006). Surface modification of certain proteins and other specific agents is also a popular method that aids in brain targeting. Transferrin receptors are expressed by the cellular lining of the BBB, hence

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nanoparticles coated with the same aid in drug delivery across the BBB. Doxorubicin and paclitaxel-loaded transferrin-coated PLGA nanoparticles exhibited enhanced tumor inhibition in in vivo-glioma models. It has also been reported that glutathione; an antioxidant, when used as a coating improves the therapeutic performance of the delivery system. Glutathionecoated PLGA containing paclitaxel could cross the BBB and elicit tumor inhibition in in vitro and in vivo rat glioma models. Once they were incorporated into the cell they released the drug inducing cytotoxicity (Cui et al., 2013; Geldenhuys et al., 2011). Systemically administered drugs exhibit low bioavailability in the brain, and poor drug efficacy within the CNS due to peripheral toxicity. PLGA interventions have been successfully used to address these issues. Several reports of PLGA based therapeutic strategies are available for a wide range of CNS conditions and brain-related disorders (Fig. 6.4). Fornaguera et al., reported that surface functionalized loperamide loaded PLGA nanoparticles prepared by phase inversion could successfully cross the BBB and, thus, provide efficient therapies for neurological diseases (Fornaguera et al., 2015). In a study, it was reported that camptothecin showed good tolerability and safe administration (up to 20 mg/kg) when encapsulated in PLGA nanoparticles. It demonstrated enhanced tumor accumulation and suppressed GL261 intracranial tumor growth in immune competent C57 albino mice (Householder et al., 2015).

FIGURE 6.4 PLGA interventions in CNS therapy.

6.7 Poly(Lactide-co-Glycolide) Interventions

Stroke is a complex pathophysiological condition that occurs due to occlusion or rupturing of cerebral blood vessels and is one of the major reasons of global mortality and morbidity. The treatment options remain largely limited to tissue plasminogen activator or mechanical recanalization (Ahmad et al., 2015, 2016). Cyclosporin A promotes the stimulation of endogenous neural stem/progenitor cells beneficial to recovery after stroke, however, its systemic delivery immunecompromises the subject. Thus, PLGA encapsulated Cyclosporin A hydrogel system was fabricated which showed equivalent therapeutic activity to drug alone and exhibited a constant concentration of the drug when implanted. This was a promising study for the therapy of stroke (Caicco et al., 2013). In another study, the stimulation of endogenous neural stem/progenitor cells was sought by sequential delivery of two growth factors. PEGylated EGF loaded PLGA nanoparticles and erythropoietin encapsulated in a biphasic microparticle with a PLGA core and poly(sebacic acid) coating were successfully fabricated and delivered via HAMC hydrogel which spatially confined the particles and attenuated the inflammatory response of brain tissue. This sequential delivery exhibited tissue repair in a mouse stroke model preventing any damages due to ICV infusion (Wang et al., 2013c). In another study, it was reported that PLGA microparticles provided an efficient system for cytokine delivery for ischemia. VEGF165 loaded PLGA microparticles were evaluated in an ischemia-reperfusion model in rats and it was found that they could be detected in the infarcted myocardium for more than 1 month posttransplant and provided sustained delivery of active protein in vitro and in vivo (Formiga et al., 2010). In another study using rats, it was found that PLGA microparticles of VEGF and Coenzyme Q10 could elicit benefits in managing myocardial ischemia (Simo´n-Yarza et al., 2013). Another study described a PLGA based scaffold system for brain tissue restoration following stroke. Here a (ppAAm)-treated PLGA scaffold was prepared for direct injection into the lesion cavity. It was successful in forming primitive neural tissue when implanted, thus, promoting the repair and regeneration of disintegrated tissues due to stroke (Bible et al., 2009). Another investigation on neural repair using PLGA based grafts showed that an artificial PLGA construct allowed for differentiation of modified neural stem cells into neurons and for establishing connections and exhibiting synaptic activities (Xiong et al., 2009). Alzheimer’s disease (AD); a complex degenerative disorder is marked by progressive memory loss and deterioration of other cognitive functions. PLGA nanospheres encapsulating VEGF were investigated as a possible therapy for AD. Results showed that this system improved behavioral deficits, reduced Aβ plaques, and promoted angiogenesis resulting in reduced neuronal loss and cerebrovascular abnormalities (Herra´n et al., 2013). PLGA based therapies have also been reported to be successful for spinal cord injury (SCI). Methylprednisolone loaded PLGA nanoparticles and GDNF loaded PLGA nanoparticles have shown improved outcomes in SCI models. PLGA based hydrogels loaded with proregenerative factors were also found to be useful as therapies for SCI treatment (Kabu et al., 2015).

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6.8 POLY(LACTIDE-CO-GLYCOLIDE) STRATEGIES FOR GENE THERAPY AND VACCINE DELIVERY Gene therapy involves the transfer of genetic material to specific cells to produce a given intended therapeutic response. Gene delivery systems commonly use viral and nonviral vectors. Viral vectors have certain drawbacks on account of their possible toxicity, immunogenicity, and inflammatory potential, whereas nonviral vectors are preferred systems. These nonviral systems are composed of biocompatible polymers and provide benefits like lack of intrinsic immunogenicity, ease of large-scale production, can be characterized, and are safe and stable systems. PLGA nanoparticles have been investigated as nonviral vectors for gene therapy and particularly for RNA and DNA delivery. However, PLGA nanoparticles bear a negative charge that limits their interaction with DNA bearing a similar negative charge. Besides this, PLGA systems encapsulating DNA suffer poor transport across cellular membranes. These demerits can be overcome by a cationic surface modification which can be achieved using polycations, like PEI, poly(2-dimethyl-amino)ethyl methacrylate, cetyltrimethylammonium bromide, dodecyl dimethyl ammonium bromide, and chitosan. Cationic modification makes them readily bind and condense the DNA. In a study; PLGA nanoparticles which could bind to antisense oligonucleotides, 20 -O-methyl-RNA with improved cellular uptake were fabricated (Nafee et al., 2007). Benfer et al., reported a PLGA based system composed of PVA with diamine modification for siRNA delivery. These siRNA loaded nanoparticles showed high and rapid uptake and a clathrinmediated internalization process. Specific green fluorescence protein (GFP) knockdown demonstrated that these systems could successfully be used for gene silencing (Benfer and Kissel, 2012). In a study, Frede et al., prepared PEI coated PLGA nanoparticles encapsulating siRNA and a calcium phosphate core for local treatment of colon inflammation. These multishell PLGA nanoparticles exhibited minimal toxicity and rapid cellular uptake in culture studies. Intrarectal application of this PLGA based therapy in a dextran sulfate sodium-induced colonic inflammation mice model showed promising results for the treatment of intestinal inflammation (Frede et al., 2016). In another study, siRNA and pDNA were codelivered to hMSCs by complexing two different genes with PEI and then coating them onto PLGA nanoparticles (Jeon et al., 2012). PLGA therapeutics has also been successfully utilized in stem cell therapy. Kim et al., reported a PLGA based nonviral vector system for enhanced hMSC differentiation. PLGA nanoparticles with pDNA complexed to it in high amounts resulted in robust gene expression in hMSCs, and induced chondrogenesis. PEI polyplexing augmented the in vitro and in vivo cellular uptake of SOX9 DNA complexed with PLGA nanoparticles (Kim et al., 2011). Park et al., demonstrated a method for polymeric matrix functionalization by polyplexing SOX9 genes and heparinized TGF-b 3 followed by coating them onto PLGA microspheres loaded

6.8 Poly(Lactide-co-Glycolide) Strategies for Gene Therapy

with dexamethasone. This system was used as a gene carrier as well as a cell delivery vehicle (Park et al., 2012). In another investigation led by Park et al., the enhancement of chondrogenic differentiation and chondrogenesis of hMSCs was proposed through the use of a nonviral gene carrier system. PEI polyplexes with a SOX trio (SOX 5, 6, and 9) fused with GFP, yellow fluorescence protein (YFP), or red fluorescence protein (RFP) was coated onto PLGA nanoparticles. This PLGA based system used as a safe and stable gene carrier showed enhanced transfection efficiency and chondrogenesis in hMSCs (Park et al., 2011). Vaccine delivery utilizing a new generation of prophylactic and therapeutic vaccines to promote effective immunization has gained impetus in recent times for the prevention and control of diseases. Carrier-mediated vaccine delivery has several advantages as it enables control on the spatial and temporal presentation of antigens to the immune system and results in efficient targeting and sustained release. Immune responses can be efficiently mediated by weak immunogens in lower doses via carrier mediated vaccination unlike the prime and booster doses required in conventional vaccination (Saroja et al., 2011). These carrier systems can augment antigen or adjuvant uptake by antigen presenting cells (APCs) and evoke improved immune responses. PLGA nanocarriers have been popularly used for vaccines and cancer immunotherapeutics and various antigens, like peptides, lipopeptides, proteins, cell lysates, viruses, and pDNA, have been formulated as PLGA based vaccine therapies. PLGA nanoparticles can encapsulate antigens, their combinations, or even combinations of an antigen and adjuvant in low doses and induce strong T-cell responses. They have demonstrated continuous and prolonged release of entrapped antigens. PLGA nanoparticles are used for the delivery of exogenous antigens that can be cross-presented through MHC-I complexes to CD8 1 cells in cancer immunotherapy. These PLGA nanocarriers get internalized by dendritic cells and reach the MHC-I pathway (Danhier et al., 2012). Location and method of antigen entrapment might affect the modulation of immune response; hence, Liu et al., investigated this theory by preparing PLGA based carriers through three different antigen loading methods. PLGA hybrid nanoparticles were prepared by antigen adsorption, entrapment, and adsorptionentrapment techniques. It was found that PLGA carriers formed by entrapment and adsorption-entrapment techniques exhibited significantly better lysosomal escape and cross-presentation of antigens from dendritic cells. Based on in vivo data, it was found that PLGA carriers based on an adsorption-entrapment method of antigen loading showed sufficient initial antigen exposure along with long term antigen persistence at the site of injection (Liu et al., 2016). Studies utilizing PLGA based therapeutics for vaccine delivery against different diseases are discussed here. sLiAg and MPLA adjuvant loaded PLGA nanoparticles surface modified with TNFα mimicking peptide were fabricated as an effective therapy against leishmaniasis (Margaroni et al., 2016). PLGA microspheres were also evaluated as efficient carriers for vaccination against acute toxoplasmosis. rSAG1 and rGRA2 proteins (derived from different stages of Toxoplasma gondii life cycle) were adsorbed onto PLGA microspheres fabricated

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via double emulsion solvent evaporation and subcutaneously injected into BALB/ c mice. The vaccinated mice exhibited a significant partial protection and longer survival time associated with increased IFN-g/IL-10 ratio and greater amounts of toxoplasma-specific IgG antibodies in comparison to the control groups. The mice that were vaccinated with a mixture of rSAG1 and rGRA2 showed more potent cellular and humoral responses as compared to the mice dosed with only rSAG1 or only rGRA2 (Allahyari et al., 2016). Tzeng et al., described PLGA microspheres encapsulating inactivated polio vaccine (IPV) along with stabilizing excipients. It was observed that this PLGA formulation enabled two separate burst releases of IPV, that is, delivery of two boluses at an interval of 1 month; and showed more robust and long-lasting humoral immune response compared to a single bolus injection. Thus, this PLGA based intervention could be successfully used to minimize the number of vaccine administrations and could be useful in polio eradication (Tzeng et al., 2016). Varypataki et al., investigated the use of PLGA nanoparticles loaded with SLP along with SLP loaded TLR ligand-adjuvant cationic liposomes as an alternative cancer vaccine to clinically used montanide ISA-51 and squalene-based emulsions. These systems enhanced dendritic antigen uptake and activation of T-cells in in vitro studies. These preparations showed enhanced potential to induce cellmediated immune responses in vivo upon subcutaneous injection in mice (Varypataki et al., 2016). In another study, HBsAg loaded chitosan and glycol chitosan coated PLGA nanoparticles were fabricated for nasal vaccine delivery. These formulations were characterized and evaluated for systemic uptake, biodistribution, and their immune-adjuvant capacity. It was found that the positively charged glycol chitosan coated PLGA particles induced higher systemic as well as mucosal immune responses as compared to the positively charged chitosan coated PLGA particles and the negatively charged PLGA particles (Pawar et al., 2013). Oral vaccination is largely weak and nonimmunogenic due to limitations of insufficient antigen uptake resulting from enzymolysis and hydrolysis in the gastrointestinal tract. A study reported the preparation of acid resistant PLGA nanoparticles loaded with HP55 for oral delivery. Vaccinated mice showed complete protection after 1 month of Helicobacter pylori challenge. This delivery system ensured antigen release in an acidic environment (pH ,5.5) and exhibited potent T-cell immune responses. Thus, this PLGA based strategy could be suitably used for oral vaccination against gastrointestinal infection (Tan et al., 2017). In another study on oral vaccination by Ma et al.; ovalbumin loaded PLGA nanoparticles coated with a lipid monolayer were fabricated using a membrane emulsification method. Release studies showed that lipid-coated PLGA particles could protect encapsulated ovalbumin from an acidic environment for a prolonged period of time and exhibited lower initial burst as compared to noncoated PLGA particles loaded with ovalbumin. These lipid-coated PLGA nanoparticles can be used as an efficient strategy for inducing mucosal and humoral immune responses by oral vaccination (Ma et al., 2014).

6.9 Miscellaneous Poly(Lactide-co-Glycolide) Therapeutics

6.9 MISCELLANEOUS POLY(LACTIDE-CO-GLYCOLIDE) THERAPEUTICS In addition to wide usage in strategies for bone tissue engineering, cancer chemotherapy, cerebral diseases, immunology, and vaccination; PLGA based interventions have also been reported for the treatment of inflammatory diseases, infectious diseases, cardiovascular disorders, skin ailments, ocular conditions, imaging, and diagnostics, as well as others (Danhier et al., 2012). In a study, PLGA nanoparticles were used for improving the oral bioavailability of atorvastatin; an antihyperlipidaemic and antiatherosclerotic drug (Li et al., 2016). PLGA mediated protein therapy with growth factor; neuregulin, for the treatment of myocardial ischemia has also been reported in a rat model (Pascual-Gil et al., 2015). One earlier study investigating thrombolysis has shown PLGA based nanoparticles loaded with t-PA to have efficient thrombolytic capabilities in a blood clot occluded tube model (Chung et al., 2008). Biodegradable PLGA has also been reported to be used in drug-releasing stents in patients with coronary artery lesions (Qian et al., 2014b). The lactate from of PLGA is known to accelerate neovascularization and to, thus, promote wound healing. In a study by Chereddy et al., PLGA nanoparticles encapsulating LL37 (an endogenous human host defense peptide) were fabricated. The prepared system was found to promote angiogenesis, wound closure, and exhibited antimicrobial activity (Chereddy et al., 2014). PLGA has been gaining attention of late for the selective treatment of infectious diseases. African trypanosomiasis is a disease with devastating effects in the sub-Saharan Africa caused by a protozoan pathogen: Trypanosoma brucei. Arias et al., developed PEGylated PLGA nanoparticles coupled to a single domain heavy chain antibody fragment capable of recognizing the surface of the pathogen T. brucei and loaded with pentamidine; a first line drug for trypanosomiasis treatment (Arias et al., 2015). Results from an in vivo murine model of African trypanosomiasis were suggestive of the fact that the fabricated system could successfully cure all infected mice at a dose 10-times lower than that of free pentamidine, and reportedly cured 60% of the treated mice at a 100-times lower dose. Malaria; another endemic parasitic disease that claims numerous lives globally. PLGA interventions for malaria treatment can reduce the dose requirements of antimalarial drugs and, thus, lower the related toxicity. In a study by Surolia et al., monensin loaded PLGA nanoparticles were fabricated through an emulsionsolvent evaporation method and were found to be 10 times more effective in inhibiting Plasmodium falciparum growth in vitro as compared to free monensin (Surolia et al., 2012). PLGA based delivery systems have also been implicated in the management of skin diseases. Quercetin loaded PLGATPGS nanoparticles have been reported to improve the poor hydrophilicity of quercetin and to augment its antiUVB effect. These nanocarriers were successful in reaching the dermis through

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the epidermis. Macroscopic and histopathological changes associated with UVB radiation were also attenuated in mice skin (Zhu et al., 2016). PLGA interventions have also been used for the ocular delivery of drugs. Pranoprofen loaded PLGA nanoparticles were fabricated and dispersed in hydrogels prepared from either carbomer or azone for ocular administration. Azone based formulations were more effective in the treatment of edema on the ocular surface as compared to carbomer hydrogels (Abrego et al., 2015). Another study suggested the use of thermosensitive PLGA-PEG-PLGA triblock as a matrix material for ocular delivery of dexamethasone. Studies in a rabbit model demonstrated that PLGA-PEG-PLGA copolymer was a suitable in situ gel-forming material for dexamethasone administration to the eyes and could also increase its bioavailability and afforded a Cmax which was about sevenfold greater than that of dexamethasone eye drops (Gao et al., 2010). PLGA based applications have also been found to be extremely useful for imaging and diagnostics. Ultrasound imaging utilizes PLGA microcapsules as contrast agents since PLGA gradually degrades into nontoxic byproducts (lactic and glycolic acid) in vivo, which further undergo degradation into carbon dioxide and water via the tricarboxylic acid cycle. Absorbable PLGA microbubbles have demonstrated potential use in left ventricle opacification and myocardium imaging. These microbubbles could detect myocardial perfusion defects (Lu¨ et al., 2009).

6.10 CONCLUSIONS AND FUTURE TRENDS The landscape of drug delivery and therapeutics is one which is complex and continually evolving. Numerous challenges associated with biomedical applications, whether through drug delivery or engineering, include but are not limited to an understanding of the cellular environment, safety and toxicity issues, scale up for large scale manufacturing, and cost burdens. A strategy that addresses most of these concerns would be a boon to drug delivery and biomedical engineering. PLGA has been one of the most popular polymers since its approval by the USFDA and EMA. It represents a unique class of polymers useful in medicine and holds great promise for future biomedical applications. It provides the benefits of a minimally invasive strategic regenerative medication. It can be invariably tailored to meet the demands of specific therapies and controlled drug deliveries by modifying its lactide/glycolide ratio and presents the possibility of surface modifications. Its biodegradability, biocompatibility, and minimal toxicity make it a safe choice for applications in bone tissue engineering, cancer therapeutics, vaccine and gene delivery, brain-related disorders, and several other biomedical uses. PLGA scaffolds have been extensively used in bone tissue engineering. Porous scaffolds, injectable microspheres, hydrogels, and fibers have gained attention for treating various bone related disorders. Futuristic strategies aimed at bone repair

6.10 Conclusions and Future Trends

based on combinations of PLGA with hydroxyapatite along with surface functionalization are expected to create an osteoconductive and osteoinductive gradient, allowing for an increased success of bone tissue regeneration. PLGA nanocarriers result in improved treatment efficacy due to sustained release of the therapeutic agent, improved pharmacokinetic and pharmacodynamic profiles, and reduced dosing frequency. As far as PLGA based nanocarriers are concerned; futuristic strategies aim at improving loading capabilities, controlling release profiles, and overcoming the instability of the incorporated drug or biologicals; intended for various therapeutic actions. These strategies might be designed by improvising techniques for the preparation of these PLGA carriers, some of these might include the modification of a solvent evaporation technique that is classically used or the use of newer methods like microfluidics, supercritical fluid technology, hydrogel templating or coaxial electrospray. Surface modification or complexations with PEG, chitosan, hyaluronic acid, or SiO2 appear to be promising approaches. PEGylated PLGA systems have been shown to act as stealth carriers by increasing systemic residence and controlling the release of encapsulated therapeutic agents. The use of these systems can enhance internalization and receptor-mediated endocytosis. Functionalization with specific proteins and receptors can give new dimensions to targeting in cancer therapy. Surface functionalized PLGA nanocarriers are future strategies directed toward brain targeting, skin permeation, and cancer chemotherapeutics. PLGA interventions are promising therapeutic strategies in the treatment of AD, stroke, gliomas, and regenerative medication for brain-related disorders. PLGA nanocarriers have emerged as alternatives to surgical oncology and radiation therapy for cancer treatment. PLGA nanoparticles can reduce the dose requirement and related toxicity of chemotherapeutic agents. These are safe systems that provide selective targeting, thus, sparing healthy cells from the deleterious effects of a given drug. PLGA carriers set the ground for successful future applications in lung, colon, pancreas, breast, ovarian, and prostate cancers. Cationically modified PLGA nanocarriers represent a future trend in the delivery of RNA and DNA. PLGA nanocarriers modified with PEI have shown the potential to be successfully used for siRNA delivery for numerous conditions. PLGA interventions have also shown promise in stem cell therapy. PLGA nanocarriers encapsulating various peptides, lipopeptides, proteins, cell lysates, viruses, and pDNA have provided a novel platform for immunology and vaccine delivery. PLGA based vaccine delivery can be achieved for polio, leishmaniasis, toxoplasmosis, and H. pylori infections, etc. The use of PLGA could serve as a beneficial strategy in oral vaccination by providing protection to the encapsulated antigen or protein from the acidic gut environment. PLGA based systems have increasingly become feasible candidates to act as drug delivery vehicles for a variety of biomedical uses as previously discussed. However; newer methodologies need to be adopted in the design and fabrication of these delivery systems in order to cater to the diagnosis and treatment of a variety of health conditions based on an improved understanding of the

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underlying pathophysiologies in various diseases. In depth studies on the mechanisms of action of these PLGA based strategies, their functionalities, in vivo evaluation, pharmacokinetic modeling, biodistribution, and toxicity assessment must be accomplished. These findings should be validated and reproducible in clinical trials. More laboratory leads should find their way to successful clinical trials and demonstrate therapeutic benefit and proof of efficacy in these trials.

ACKNOWLEDGMENTS The authors are thankful to knowledge resource center, CSIR-CDRI, Lucknow, and CSIR, New Delhi (CDRI Communication No. 9783).

LIST OF SYMBOLS AND ABBREVIATIONS % , .
C kDa mm μm mg/kg Aβ AD APCs BBB CD8 1 Cmax CNS DNA EGF EMA FDA GDNF GFP HAMC HBsAg hMSCs ICV IFN-g/IL-10 IgG IPV MHC-I MPLA

percentage less than greater than degree celsius kilodalton millimeter micrometer milligram per kilogram amyloid beta Alzheimer’s disease antigen presenting cells bloodbrain barrier cluster of differentiation 8 maximum or peak serum concentration central nervous system deoxyribonucleic acid epidermal growth factor European medical agency food and drug administration glial cell line-derived neutrophic factor green fluorescent protein hyaluronan methylcellulose hepatitis B surface antigen human mesenchymal stem cells intracerebroventricular interferon gamma-interleukin 10 immunoglobulin G inactivated polio vaccine major histocompatibility complex I monophosphoryl lipid A

References

MTCM-PL MW NaCl O/W PCL PDLA pDNA PEG PEI PGA pH PLA PLGA PLLA ppAAm PVA RES RFP RNA RTCM-PL RTIM-PL SCI SiO2 siRNA sLiAg SLP SOX9 Tg TGF-b 3 TIPS TLR TNFα TPGS USFDA UVB VEGF W/O/W YFP

modified thermal compression molding-particulate leaching approach molecular weight sodium chloride oil in water polycaprolactone poly(D-lactic acid) plasmid DNA poly(ethylene glycol) poly(ethyleneimine) poly(glycolic acid) “pouvoir hydrogen” or hydrogen power poly(lactic acid) poly(lactic-co-glycolide) poly(L-lactic acid) plasma allylamine polyvinyl alcohol reticuloendothelial system red fluorescence protein ribonucleic acid room temperature compression molding-particulate leaching method room temperature injection molding-particulate leaching method spinal cord injury silicon dioxide small interfering RNA soluble leishmania antigens synthetic long peptide sex determining region Y-box 9 glass transition temperature transforming growth factor beta 3 thermally induced phase separation toll like receptors tumor necrotic factor-alpha tocopheryl polyethylene glycol succinate United States Food and Drug Administration ultraviolet B vascular endothelial growth factor water in oil in water yellow fluorescence protein

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Electrospun biomimetic scaffolds of biosynthesized poly(β-hydroxybutyrate) from Azotobacter vinelandii strains. cell viability and bone tissue engineering

7 Angel Romo-Uribe

Johnson & Johnson Vision Care, Inc., Advanced Science & Technology, Jacksonville, FL, United States

7.1 INTRODUCTION 7.1.1 POLYMERS AS MEDICAL DEVICES The development of biocompatible polymeric materials with tuned physical properties and added functionality has recently surged due to their significantly potential biomedical applications. Furthermore, adding biodegradability broadens the scope of these applications (Lendlein and Langer, 2002; Mather et al., 2009; Dumitriu, 2011; Domb et al., 2011). Although materials are an essential component for a given application, the processing route can also define the properties and behavior of these materials. Moreover, for any material to have commercial viability it must perform as designed, must withstand processing, handling and sterilization procedures, and must have a lifetime appropriate to particular applications and needs of users. Furthermore, during shelf life storage the polymers must retain their physical properties. Therefore, the development of polymeric materials for potential biomedical applications goes from the synthesis to processing and understanding the influence of sterilization and storage time on their physical properties. Examples of close coupling between synthesis and processing to produce smart, functional polymers are: shape memory polymers, smart coatings, and electrospun membranes. The production and applications of these sorts of polymeric materials are reviewed in this chapter. This revision is by no means exhaustive as the important subject of polymer hydrogels, a foundation of J&J Vision Care technologies, is not addressed. It is rather intended to provide a flavor of the Materials for Biomedical Engineering: Hydrogels and Polymer-based Scaffolds. DOI: https://doi.org/10.1016/B978-0-12-816901-8.00007-9 © 2019 Elsevier Inc. All rights reserved.

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varied applications and advantages arising from smart choices of polymers and processing routes.

7.1.2 SHAPE MEMORY POLYMERS Shape memory polymers (SMPs) are one class of smart materials which can “remember” their original shape and are finding more applications into the medical field since the pioneering work of Lendlein and Langer (2002). SMPs respond dynamically to external stimuli like heat, electric field, magnetic field, infrared radiation, and UV radiation. Because SMPs can store a temporal shape and return to their original (permanent) shape simply by applying the proper stimulus, these materials have acquired relevance in medicine. A shape memory phenomenon is not an intrinsic property of polymeric materials. Appropriate synthetic routes (e.g., crosslinking) and thermo-mechanical programming, enable polymeric materials to exhibit more or less temporal shape stability—so-called shape “fixing”— and shape recovery to the permanent shape (Mather et al., 2009; AlvaradoTenorio et al., 2011, 2015). Lendlein and Langer (2002) demonstrated the possible use of biodegradable SMPs as smart sutures. Xu and Song (2010) synthesized a high performance SMP for possible tissue engineering applications, utilizing hybrid organic-inorganic nanoparticles, polyhedral oligomeric silsesquioxane (POSS), and polylactides (PLAs). Luo and Mather (2010) reported a triple shape memory effect in polycaprolactone-epoxy composite. Triple shape memory consists in a material capable of fixing two temporary shapes and sequential recovery to the permanent shape. A combination of an electrospun nonwoven web of polycaprolactone (PCL) and epoxy matrix was applied, thus, clearly exemplifying the close coupling between materials and processing to endow the final product with new properties. Luo and Mather (2013) further took advantage of this approach to produce self-healing coatings. Saatchi et al. (2015) reported polymeric networks with reversible bidirectional shape memory effect which can be activated at human body temperature, thus, opening up applications in reversible closure systems and folding guides for nerve electrodes. Gu et al. (2016) reported the utilization of the shape memory effect for the removal of bacterial biofilms. Liang et al. (2016) combined crystallizable polymers with solid-state fluorescent dyes to produce thermoresponsive shape memory fluorescent polymers. These materials could find applications as biosensors and recording materials.

7.1.3 SMART POLYMERIC COATINGS Another approach where the processing method plays a major role in the final properties is the production of smart coatings (Baghdachi, 2009). Coatings have traditionally been a passive layer utilized for decorative purposes and as

7.1 Introduction

protective layer of surfaces. For instance, protection against corrosion, humidity, and weather conditions, among others. However, in biomedical applications, the aim is to produce smart surfaces. These surfaces will promote growth (of cells) or will prevent growth (of biofilm, i.e., antifouling surfaces) (Pavlukhina et al., 2010; Engel et al., 2012). Likibi et al. (2008) produced biomimetic nanocoatings utilizing electrostatic self-assembly to polypeptides. These nanocoatings were in vitro tested on human osteoblast cells, opening up potential dental and orthopedic applications. Gerber et al. (2012) described the effectiveness of utilizing polymeric layers to produce functional living materials. “Old,” well-known polymers for over 50 years like polyvinyl chloride (PVC) and polycarbonate (PC) were arranged in a three-layer system to produce a smart artificial two-dimensional biological habitat. Biocompatible and biodegradable smart polysaccharide coatings against bacteria and yeasts were reported by Cado et al. (2013). These authors reported that antimicrobial peptide release was triggered by enzymatic degradation of the coating provoked by the bacteria and yeasts themselves. A truly “smart” response of this coating. Zhuk et al. (2014) reported the production of smart antibacterial coatings utilizing layer-by-layer assembly. These authors also demonstrated that the antibiotic is not released in infection-free physiological conditions and that the release of the antibiotic is triggered in response to a bacterial challenge. Another fine example of a “smart” polymer. Finally, Libera and coworkers (Krsko and Libera, 2005; Krsko et al., 2009; Wang et al., 2012; Wang and Libera, 2014; Wang et al., 2014) have developed antifouling coatings and scaffolds by micro patterning and inducing biospecificity to polymeric surfaces. That is, controlling the topography of polymeric surfaces at the micron and nanometer scale, characteristic of individual cells and proteins, has been effective to promote adhesion (spatial controlled adhesion) and/or antifouling behavior. Libera’s research points out that “cell-surface interactions are regulated by the spatial distribution of adhesive sites on a surface, therefore suggesting a simple method to influence cellular processes associated with healing after the implantation of a tissue-contacting biomedical device” (Wang and Libera, 2014; Wang et al., 2014).

7.1.4 ELECTROSPUN FIBROUS SCAFFOLDS The processing of polymers by electrospinning is a technique that can produce micrometer- to nanometer-scale fibrous morphologies and greatly modify the physical properties of polymers. These morphologies are ideal to construct biomimetic scaffolds useful for diverse biomedical applications. This usefulness arises from closely mimicking the fibrillar components (e.g., collagen) of the extracellular matrix (ECM). The ECM surrounds the cells of any biological tissue (Sell et al., 2007). Thus, the fibrous scaffolds with high porosity and specific surface

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area and topography are ideal candidates for engineering ECM promoting naturally the cell functions. However, in order to assess the feasibility of an electrospun polymer as scaffold it is essential to perform biocompatibility tests and functionality evaluation in vitro and in vivo by seeding cells and determining the cell viability. Electrospinning occurs when the electrical forces at the surface of a polymer solution or melt overcome the surface tension and cause an electrically charged jet to be ejected. When the jet dries or solidifies, an electrically charged fiber remains. This charged fiber can be directed or accelerated by electrical forces and then collected in sheets or other useful geometrical forms. Important factors in electrospinning processing are the applied voltage (V), the solution flow rate (Q), polymer concentration (c), polymer molecular weight (M w ), and nozzle-tocollector distance (H). (Doshi and Reneker, 1995; Alban˜il-Sanchez et al., 2012). The proper choice of electrospinning parameters enables the production of a thin fiber web constituted by continuous smooth fibers of homogenous diameter to be deposited onto a substrate. The filament diameters can range from nanometer to micrometer scale (Doshi and Reneker, 1995; Romo-Uribe et al., 2009; Tumbic et al., 2016). Since the pioneering work of Reneker at Akron, polymer electrospinning has proved to be effective to modify the properties of bulk polymers and is finding diverse industrial, medical, automotive, and aerospace applications (Doshi and Reneker, 1995). Electrospun polymeric fibers are being utilized to produce polymer nanocomposites (Huang et al., 2003; Alban˜il-Sanchez et al., 2014). Polymer electrospinning has been utilized to incorporate ferromagnetic nanoparticles into polymer filaments opening up opportunities in catalysis, anodic materials in lithium ion batteries, and in electronic nanodevices (Maensiri et al., 2009). This approach was utilized to immobilize iron nanoparticles in polymeric nanofibers of polyacrylic acid (PAA) and polyvinyl alcohol (PVOH) reinforced with multiwall carbon nanotubes (MWCNT) (Xiao et al., 2010). Electrospun nylon membranes composed of nanometer scale filaments were effective to reinforce polyaniline, an electrically conductive polymer, opening up opportunities in separation membranes and actuators (Romo-Uribe et al., 2009). Electrospinning was also utilized to incorporate graphene nanosheets into nylon nanofibers (Alban˜il-Sanchez et al., 2012). Therefore, the application of electrospinning to biodegradable and/or biocompatible polymers is elucidating novel applications in tissue engineering, antibacterial mats and hydrogels, metal and biomolecules immobilization systems, and shape memory composites. As far as biomedical applications are concerned, Song et al. (2012) electrospun PVOH-collagen-hydroxyapatite to produce biomimetic bone-like ECM for the modification of orthopedic prosthetic surfaces. A similar approach was utilized to incorporate hydroxyapatite and alumina into a PCL scaffold (ReyesLopez et al., 2013). Furthermore, crosslinked electrospun collagen scaffolds proved effective for osteoblast viability (Torres-Giner et al., 2009)

7.1 Introduction

Electrospun fibrous scaffolds of PCL were investigated for biomolecules immobilization and tissue engineering (Mattanavee et al., 2009). On the other hand, electrospinning PLA and amino functionalized POSS was effective to immobilize palladium nanoclusters to produce a novel catalytic system (Gardella et al., 2013). Wound-dressing materials with antibacterial properties have also been produced by electrospinning polymer/silver nitride solutions (Rujitanaroj et al., 2008) as well as PLA-silver nanoparticle solutions (Vargas-Villagran et al., 2012, 2014). On the other hand, polyethylene glycol (PEG)/POSS-based polyurethanes were electrospun and silver nitride incorporated to produce hydrogel mats with antibacterial properties (Wu et al., 2009a,b). Robertson et al. (2015) applied dual electrospinning of PCL and a polyetherbased thermoplastic polyurethane to produce shape memory elastomeric composites. Furthermore, Tumbic et al. (2016) described nanoscale blending of PCL and PLA by utilizing dual electrospinning of the respective polymer solutions. The close interaction between blend components proved essential to tune the micro/ nanostructure and thermo-mechanical properties of the blends. Therefore, bio-mimicking electrospun fibrous polymeric structures are of great potential for scaffold-assisted tissue and bone repair, and regeneration applications. Adding functionalities, like shape memory and biodegradability, further broadens the scope of these applications.

7.1.5 POLY-β-HYDROXYBUTYRATE Poly(β-hydroxybutyrate) (PHB) is a polyester produced by numerous bacteria, from the polyhydroxyalkanoate family (PHAs), (Holmes et al., 1982; Sudesh et al., 2000; Pen˜a et al., 2014a, b) which are biodegradable. Recently and based on their properties of biocompatibility and biodegradability, new applications for PHB have been proposed in the medical field, where the chemical composition and product purity are critical. The degradation properties of PHB and PHAs in general are desirable for the absorption of these materials in biomedical applications. Furthermore, the degradation product of PHB, D-(
)-3-hydroxybutyrate, is a common intermediate metabolite present in animal cells, and was also detected in relatively large amounts in human blood plasma (Reusch and Sadoff, 1985). Therefore, there are high expectations for PHAs in the biomedical field, for instance, in surgical sutures, drug delivery, coatings for cardiovascular implants, and scaffolds in tissue engineering among others (Cheng and Wu, 2005; Valappil et al., 2006; Wu et al., 2009a,b). Wang et al. (2004) reported the suitability of poly(3-hydroxybutyrate-co-3hydroxyhexanoate) (PHBHHx) scaffolds to promote in vitro adhesion, proliferation and differentiation of bone marrow cells. Wang et al. (2005) extended this investigation into different compositions of the copolyester PHBHHx and determined the cell viability of fibroblast and osteoblast cells. The utilization of PHBHHx in tissue engineering has been recently reviewed by Chang et al. (2014). PHAs have also

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been utilized in neural regeneration and nerve tissue engineering thanks to their high biocompatibility and low cytotoxicity (Lu et al., 2013). Sevastianov et al. (2003) studied the hemocompatibility properties of the PHAs. On the other hand, Shishatskaya et al. (2006) reported the growth and differentiation of murine marrow osteoblast cells in poly(3-hydroxybutyrate)hydroxyapatite. All these studies aimed to establish the biocompatibility of such materials for different applications of tissue engineering as well as in other medical applications. PHB is an aliphatic polyester composed of 3-hydroxybutyrate monomers in which the carboxyl group of one monomer forms an ester bond with the hydroxyl group of the neighboring monomer (Madison and Huisman, 1999). PHB has mechanical properties similar to conventional plastics like polypropylene, although it exhibits higher rate of crystallization and higher degree of crystallinity, leading to brittleness and low elongation to break (Barham et al., 1984; Barham and Keller, 1986; Dominguez-Diaz et al., 2011, 2015a,b). It has been reported, however, that when hydroxyvalerate, HV, is copolymerized with PHB the resultant copolymer is much less fragile, due to a disruption in degree of crystallinity (Holmes et al., 1982; Holmes, 1988). Recent results have shown that when HV is copolymerized with PHB, it induced higher thermal stability and modified the rheological behavior, thus, enabling melt processing of the PHB/PHV copolymers (Dominguez-Diaz and Romo-Uribe, 2012). Pen˜a et al. (2014a, b) have recently reviewed the fermentation and scaling strategies for microbial PHB production and for tuning the mean molecular weight. Fibrous scaffolds based on PHB produced from bacterial source (Pen˜a et al., 2014a; Dominguez-Diaz et al., 2015a,b) were manufactured using electrospinning and the cell viability of normal human osteoblast cells was investigated. Firstly, the correlation between electrospinning conditions and morphology of PHB fibrous scaffolds is discussed. The aging behavior of PHB scaffolds held at room temperature up to 11 months was characterized by changes in crystallinity, and the influence of sterilization procedures (ultraviolet radiation vs autoclave) on physical properties was investigated. Finally, the cell viability of normal human osteoblast cells cultured on PHB fibrous scaffolds and cast films is evaluated. It will be shown that cells remained viable and proliferated, with higher proliferation in fibrous electrospun scaffolds. The results of this investigation may be valuable for tissue engineering approaches aimed at designing and implementing pre-vascularized scaffolds and functional blood vessel networks.

7.2 METHODS OF CHARACTERIZATION 7.2.1 MATERIALS A commercial poly(3-hydroxybutyrate) (PHB) from Sigma-Aldrich (St. Louis MO, USA), and a specially PHB produced from a mutant strain of Azotobacter

7.2 Methods of Characterization

Table 7.1 Average Molecular Weight, M v , Melting (Tm), Crystallization (Tc), and Decomposition (Tdec) Temperatures for PHB Homopolymers Mv a

Tmb

ΔHmb

Tcb

ΔHcb

Tdecc

Sample

(KDa)

( C)

(J/g)

( C)

(J/g)

( C)

PHB50 PHB230

50 230

159 178

60 67

103 106

61 66

237 238

a

Mw determined by dilute solution viscosimetry, using Mark-Houwink equation (Marchessault et al., 1970). b Tm and Tc determined by differential scanning calorimetry from the second heating and first cooling, respectively, scanning a 20 C/min. c Determined by thermogravimetric analysis under dried nitrogen, scanning at 20 C/min (DominguezDiaz and Romo-Uribe, 2012; Dominguez-Diaz et al., 2015a,b).

vinelandii (OP) (Pen˜a et al., 2014a; Dominguez-Diaz et al., 2015a,b) were used in this investigation. The biosynthesis procedure consisted of the bacteria being cultured in Burk’s medium, and the strain was cultured in 500 mL Erlenmeyer flasks during 72 hours in an orbital incubator shaker with a shaking diameter of 2.5 cm and a shaking frequency of 200 rpm at 29 C and pH 7.2. The aeration conditions were monitored through changes in the oxygen transfer rate (OTR). The specimens physical chemical properties are summarized in Table 7.1. The average molecular weight was determined by capillary viscometry using an Ubbelohde-type viscometer (manufactured by Cannon Instrument Company Inc., State College PA, USA). Relative viscosity measurements were carried out on diluted PHB/chloroform solutions at 30 C. Viscosity averaged molecular weight, Mv, was determined from intrinsic viscosity values, [η], using the Mark
Houwink equation (Eq. 7.1) and the coefficients k 5 0.77 3 1024 and α 5 0.82 reported by Marchessault et al. (1970) α

½η 5 kUM v

(7.1)

As-cast PHB films were prepared dissolving the polymers in chloroform, pouring in a Petri dish of 25 mm diameter and drying first at room temperature overnight in a hood, and then under vacuum at 40 C for 6 hours. The final films’ average thickness was c. 35 μm.

7.2.2 SCAFFOLD FABRICATION The fibrous scaffolds were prepared by electrospinning from chloroform solution utilizing an in-house built instrument (see Alban˜il-Sanchez et al., 2012). The polymers were dissolved in chloroform (analytical grade, purchased from SigmaAldrich Chemicals) at 60 C under continuous gentle magnetic stirring. Solution concentration varied from 12% to 14% w/v and the applied voltages ranged from 8 to 12 kV. The distance between needle and target (aluminum foil screen) was

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kept constant at 6 cm. Mats were electrospun for 2 hours at a rate of 0.2 mL/h utilizing a NE-1000 Programmable Single Syringe Pump (NEW ERA Pump Systems Inc., Farmingdale NY, USA). A high voltage source model CZE 1000R manufactured by Spellman High Voltage Electronics Corp. (Hauppauge NY, USA) was utilized.

7.2.3 FOURIER-TRANSFORMED INFRARED SPECTROSCOPY Fourier-transformed infrared spectroscopy, FTIR, was used to characterize the polymers. IR spectra in attenuated total reflectance (ATR) mode were obtained with the FTIR Nicolet iS10 of Thermo Scientific (USA). For each spectrum, sixteen scans between 4000 and 400 cm
1 were averaged, at a resolution of 4 cm21.

7.2.4 THERMAL ANALYSIS The decomposition temperatures, Tdec, of the specimens were determined by thermogravimetric analysis (TGA), at 10 C/min, using the TGA Q500 analyzer (manufactured by TA Instruments, Newcastle DE, USA). These experiments were carried out under nitrogen atmosphere scanning from 0oC up to 500oC. On the % % other hand, the thermal transitions were determined by differential scanning calorimetry (DSC) using the DSC6000 calorimeter (manufactured by Perkin Elmer, Waltham MA, USA). Temperature and enthalpy calibration were carried out using analytical grade indium (Tm 5 156.6 C). Samples of c. 4 mg were prepared, and data from the second heating and first cooling scans are reported. The thermal scans were carried out at a rate of 10 C/min under dry nitrogen atmosphere.

7.2.5 X-RAY SCATTERING Two-dimensional wide-angle X-ray scattering (WAXS) patterns were obtained using a Micro Star rotating anode generator with copper target manufactured by Bruker (Germany). WAXS patterns were recorded using a Mar345 Image Plate with a resolution of 3450 3 3450 pixels, and 100 μm/pixel; a sample-to-detector distance of 20 cm was used. The patterns were analyzed using the X-ray scattering software POLAR v2.6 (Stonybrook Technology and Applied Research Inc., New York, USA). The degree of crystallinity, αc, was calculated from the ratio of the area under the crystalline peaks to that of the total diffraction curve. The amorphous halo was approximated to a Gaussian curve, and the crystalline peaks were fitted to Gaussian functions, utilizing the software PowderX (Dong, 1999). The expression for the degree of crystallinity is (Alexander, 1969): ÐN Ic ðqÞ q2 dq αc 5 Ð0N 2 0 IðqÞ q dq

(7.2)

7.2 Methods of Characterization

where Ic is the part of the intensity associated to the crystalline phase, I is the total diffracted intensity, and q is the magnitude of the scattering vector defined as: q5

4 πsinθ λ

(7.3)

On the other hand, the crystallite size, Lhkl, was determined from the fullwidth at half maximum (FWHM) of the 020 reflection, using Scherrer’s equation (Alexander, 1969) Lhkl 5

Kλ β o cosθ

(7.4)

where Lhkl is the mean dimension of the crystallites perpendicular to the planes hkl, λ is the radiation wavelength, β o is the FWHM in radians of the hkl reflection at 2θ, and K is a constant that is commonly assigned a value of unity (Alexander, 1969).

7.2.6 SMALL-ANGLE LIGHT SCATTERING Small-angle light-scattering (SALS) patterns were obtained using an in-house instrument (Romo-Uribe et al., 2010). A He-Ne laser (wavelength λ 5 632.8 nm) model 1500 manufactured by JDS Uniphase Corporation (Santa Rosa CA, USA) was utilized and the laserbeam was collimated using three circular pinholes. SALS patterns in HV polarization condition were imaged on a sand blasted glass (Edmunds Optics Co.) placed 126 mm downstream from the sample. Patterns were recorded with a charge coupled device (CCD) model PC-23C (Super Circuits, Taiwan) with C-mount ring and wide-angle lens (Computar, Japan). Images at a resolution of 200 μm/pixel were acquired.

7.2.7 CONTACT ANGLE The contact angle of a water droplet on the scaffolds was determined by the sessile drop technique. A distilled water drop (obtained by purification) was placed on the substrates using a Pasteur pipette and acquired by a digital optical microscope MicroView 800X. The contact angles were determined using the ImageJ software (software freely available from the National Institute of Health, USA, https://imagej.nih.gov/ij/index.html ).

7.2.8 POLARIZED OPTICAL MICROSCOPY Films cast from chloroform solution were prepared and heated up to 200 C, holding for 3 min and then cooled to 30 C at 10 C/min using the hot-stage FP82HT and a controller FP90 manufactured by Mettler Toledo (Columbus OH, USA).

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The morphology of as-cast films was investigated by polarized optical microscopy (POM) using an Eclipse E600 microscope manufactured by Nikon Instruments Inc. (Melville NY, USA) equipped with Motic 2.0 digital camera (Motic, China).

7.2.9 SCANNING ELECTRON MICROSCOPY The samples were mounted on a sample holder using carbon tape and coated with gold for 60 s using an Evex gold sputter coater to yield coating thickness of ˚ , suitable for scanning electron microscopy (SEM) observaapproximately 200 A tion without charge accumulation. The SEM micrographs were obtained with a Hitachi SU1510 scanning electron microscope with secondary electrons at 15 kV under vacuum conditions. The filaments diameter of electropun scaffolds was measured from SEM micrographs using ImageJ software.

7.3 PHB ELECTROSPUN FIBROUS SCAFFOLDS Firstly, the influence of electrospinning parameters, solution concentration, and applied voltage on the morphology of the scaffolds was studied. The flow rate and sample to collector distance were kept constant. The average filament diameter, degree of porosity, and water contact angle were determined to assess the influence of the processing parameters on final morphology and physical properties of the scaffolds.

7.3.1 SCAFFOLDS MORPHOLOGY Fibrous nonwoven scaffolds with micrometer scale pores and random orientation of filaments were obtained via electrospinning, as demonstrated by SEM analysis. Furthermore, the results demonstrated that solution concentration and applied voltage significantly influenced the morphology of the electrospun scaffolds. Firstly, the influence of solution concentration and applied voltage was extensively investigated for the smaller molecular weight specimen, PHB50. The findings were then applied to the higher molecular weight specimen, PHB230. Fig. 7.1A corresponds to a scaffold of PHB50 produced from 12% w/v solution concentration and applying 9 kV. It can be seen that the morphology consists of cylindrical filaments, relatively smooth and with no beads present. However, at relatively larger scale the morphology is relatively inhomogeneous as the filaments are arranged into large, c. 100 μm agglomerates. The mean diameter of the filaments is 3.7 6 0.7 μm. On the other hand, a slight increase of the solution concentration of PHB50 to 13% w/v while maintaining the applied voltage at 9 kV produced very homogeneous, randomly oriented, cylindrical filament morphology with uniform filament

7.3 PHB Electrospun Fibrous Scaffolds

FIGURE 7.1 (A) Nonwoven scaffold of PHB50 electrospun from 12% w/v solution in chloroform applying a voltage of 9 kV. (B) Nonwoven scaffold of PHB50 electrospun from 13% w/v solution in chloroform applying a voltage of 9 kV; (C) enlargement of micrograph in (B).

diameter, as shown in Fig. 7.1B. Note the smoothness of the filaments and the absence of bead morphology as well as uniform spatial distribution of filaments. Fig. 7.1C shows a SEM micrograph at larger scale evidencing the homogenous fibrous morphology of the scaffold. The mean diameter of the filaments was found to be 2.8 6 0.4 μm. That is, an increase in solution concentration produced more homogeneous fibrous scaffolds, and with reduced filament diameter. However, keeping the solution concentration of PHB50 at 13% w/v, but increasing the voltage to 12 kV produced scaffolds with nonhomogeneous morphologies. The SEM micrograph showed a morphology consisting mostly of thicker filaments and large aggregates of filaments. There is no evidence of bead morphology, and the large aggregates appear to consist of fused filaments. It is

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believed that when electrospinning at higher acceleration voltage there is not enough time, at this given needle-to-collector distance, for complete solvent evaporation. Therefore, there is propensity for the filaments to fuse with each other. The mean diameter of the filaments was found to be 6.7 6 1.7 μm. Solution concentration of PHB50 was further increased up to 14% w/v and, in order to enable solvent evaporation, the acceleration voltage was reduced to 10 kV. The morphology now consists of relatively few thick filaments (c. 10 6 1.1 μm thick) and predominantly thinner, fused, filaments. The mean diameter of the thinner filaments was found to be 4.8 6 0.7 μm. It is noted that the fused filaments exhibit a honeycomb-like morphology. It is suggested that the higher viscosity of the polymer solution slowed down the rate of evaporation of the solvent, thus, plasticizing the filaments. Therefore, after deposition on the collector the filaments in close contact would weld to each other giving rise to a honeycomb-like morphology. These results then suggest that electrospinning of the higher molecular weight PHB230 requires lower solution concentration and smaller acceleration voltage. Indeed, scaffolds electrospun from a 12.5% w/v solution at 8 kV consisted of relatively thick and randomly oriented filaments as well as fused filaments forming a porous network. Note that the scaffold morphology resembled that obtained with PHB50 solutions of 14% w/v and electrospinning at 10 kV. The mean diameter of the filaments was found to be 4.5 6 0.7 μm. In this case, the larger molecular weight produces relatively viscous solutions at this concentration. Therefore, solvent evaporation was slowed down and electrospun filaments were partially plasticized, thus, promoting filament fusion. A detailed investigation was carried out on the agglomerates found in the fibrous electrospun scaffolds. SEM micrographs of PHB50 scaffold electrospun from 13% w/v solution at 12 kV exhibited randomly oriented and entangled cylindrical filaments and relatively large aggregates/bundles of filaments. Examination of these micro-aggregates at higher magnifications revealed a porous morphology in the filaments. These micrographs clearly exhibited highly porous filaments fused on each other. It is suggested that the porosity is due to incomplete evaporation of the solvent when the filaments were ejected from the needle and before depositing on the ground collector. These results also suggest that the incomplete evaporation of solvent gave rise to the aggregation/fusion of filaments. This hypothesis is supported by closely examining the scaffolds with homogeneous fibrous morphology. SEM micrographs of a scaffold of PHB50 electrospun from 13% w/v solution, but lowering the applied voltage down to 9 kV consisted of smooth, cylindrical, and randomly oriented filaments. Furthermore, higher magnification micrographs showed no sign of porosity in the filaments. The porosity was determined using image analysis of the corresponding SEM micrographs. The values at 0% porosity correspond to solution cast films. Table 7.2 summarizes the variety of morphologies and characteristics (average diameter of filaments and percent of porosity) of the nonwoven electrospun PHB scaffolds of different molecular weight. Note that scaffolds with homogenous

7.3 PHB Electrospun Fibrous Scaffolds

Table 7.2 Filament Diameter, Bead Diameter, Porosity Percent and Contact Angle, in As-Cast Films and Electrospun Nonwoven Mats of Polyhydroxyalkanoates Materials

PHB 50 KDa

PHB5V 70 KDa

PHB12V 530 KDa

Processing method Water contact angleb ( ) Electrospinning conditions SEM image

Cast film 61 6 0.2 12 % w/v & 9 kV

Cast film 70 6 0.2 12 % w/v & 9 kV

Cast film 59 6 0.3 12 % w/v & 9 kV

Filament diametera (μm) Bead diametera (μm) Porositya (%) Water contact angleb ( ) Electrospinning conditions SEM image

3.5 6 0.7
57.8 89 6 0.4 13 % w/v & 9 kV

3.2 6 0.3
15.2 90 6 0.2 13 % w/v & 8 kV

0.5 6 0.2 25 6 8.3 6.5 69 6 0.3 13 % w/v & 9 kV

Filament diametera (μm) Bead diametera (μm) Porositya (%) Water contact angleb ( ) Electrospinning conditions SEM image

2.8 6 0.4
22.5 75 6 0.3 13 % w/v & 12 kV

2.3 6 0.3
3.6 75 6 0.2 13 % w/v & 9 kV

1.3 6 0.3 18.5 6 6.4 6.4 84 6 0.3 14 % w/v & 10 kV

Filament diametera (μm) Bead diametera (μm) Porositya (%) Water contact angleb ( )

6.7 6 1.7
80.6 96 6 0.3

1.8 6 0.1
31.8 80 6 0.2

2.6 6 0.3 13.2 6 3.5 21 114 6 0.4

Scale Bar Corresponds to 10 μm. Electrospinning distance between needle and target was 6 cm in all cases. a Determined by SEM images. b Determined by sessile drop technique.

fibrillary morphology the degree of porosity is smaller than scaffolds with exhibit nonhomogeneous fibrillary structure. The morphology as well as the topography of polymer scaffolds influence adhesion and cell viability as these closely mimics the fibrillary components (e.g., collagen) of the ECM (Bao et al., 2014). The topography and roughness of electrospun scaffolds and solution cast film was investigated via atomic force microscopy (AFM). Fig. 7.2 shows AFM micrographs of scaffolds electrospun from (A) PHB50, 13% w/v and 9 kV, (B) PHB230, 12.5% w/v and 8 kV, and

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FIGURE 7.2 AFM micrographs of (A) PHB50 scaffold electrospun from 12% & 9 kV, (B) PHB230, electrospun from 12.5% & 8 kV, and (C) PHB230 solution cast film. Images of water drops on surfaces of (D) PHB50 scaffold electrospun from 13% & 12 kV, (E) PHB230, electrospun from 12.5% & 8 kV, and (F) PHB230 solution cast film.

(C) PHB230 film cast from chloroform solution. The uniform filamentous morphology of PHB50 is clearly appreciated in the AFM micrograph, Fig. 7.2A. Note that these micrographs show that the topography of well-formed fibrous morphology is similar to the topography of cast films, as shown by Fig. 7.2A and C. On the other hand, the topography of PHB230 reflects the fibrous morphology and aggregates and it is highly nonhomogeneous with pronounced tops and valleys. The scaffold of PHB50 with smooth cylindrical filaments (Fig. 7.2A) exhibited a roughness of 2.3 nm. On the other hand, the scaffold of PHB230 with relatively large aggregates of fused filaments (Fig. 7.2B) increased significantly, it exhibited a roughness of 44 nm. Not surprisingly the solution cast film exhibits a relatively uniform thickness and the roughness (Fig. 7.2C) was determined to be 1.8 nm. As stated above, the degree of roughness is an important property that influences cell adhesion (Moffa et al., 2013; Zhang et al., 2014). However, it will also be shown in the next section that roughness also influences the water contact angle.

7.3 PHB Electrospun Fibrous Scaffolds

It is well documented that PHB exhibits fast crystallization kinetics, even under rapid quenching (Barham et al., 1984; Dominguez-Diaz et al., 2011). POM showed that solution cast films as well as electrospun scaffolds exhibit considerable birefringence due to crystallization of PHB (Romo-Uribe et al., 2017). The pronounced degree of crystallinity already present in as-cast films and as-spun scaffolds was found to be a good indication of the influence of storage time (aging) on their physical properties, that is, higher degree of crystallinity produces more brittleness. Aging was, therefore, correlated with degree of crystallinity, and this will be discussed in Section 3.3.

7.3.2 WETTING BEHAVIOR The wetting ability of the electrospun nonwoven scaffolds was determined by measuring the contact angle of water drops, and the behavior was compared to that obtained from solution cast films. Moffa et al. (2013) and Zhang et al. (2014) have pointed out on the critical role that wettability plays in the adhesion, proliferation and viability of cells. Fig. 7.2D
F shows images of water drops on electrospun scaffolds of (D) PHB50 electrospun from 13% w/v, 12 kV, (E) shows a water drop on a scaffold of PHB230 electrospun from 12.5% w/v, 8 kV, and (F) water drop on as-cast PHB film. The PHB50 scaffold with uniform filamentous morphology (i.e., no agglomerates), electrospun from a 13% w/v solution concentration and 9 kV, exhibited larger contact angle than on as-cast films. Strikingly, water contact angles were significantly larger for scaffolds with inhomogeneous fibrous morphologies as shown in Fig. 7.2D and E. At first sight, this data suggest that the contact angle is independent of the molecular weight of PHB. However, more research needs to be carried out before conclusions can be drawn. It is interesting to note that these results agree with those reported by Zhu et al. (2006) where super-hydrophobic behavior was attained in quite inhomogeneous morphologies of electrospun polyhydroxyalkanoate scaffolds which exhibited a mixture of filaments, beads and aggregates. The inhomogeneous morphology of electrospun scaffolds and the associated porosity appear to be the driving force for high water contact angles, regardless of the polymer utilized, as shown by Vargas-Villagran et al. for PLA scaffolds (2012, 2014). Table 7.2 summarizes the correlation between solution concentration, applied voltage, and water contact angle for scaffolds of PHB of varying molecular weight. Data from solution cast films have been included for the sake of comparison. Although these results suggest a trend toward higher contact angles for the scaffolds, there is not a clear correlation between processing parameters and contact angle. In order to rationalize these results, it is suggested that air trapped in the rough morphology is the cause for the high contact angle (i.e., hydrophobic behavior) observed in the electrospun scaffolds (Zhu et al., 2006). Therefore, the results prompted us to correlate the degree of porosity (an intrinsic property of the

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morphology) in the scaffolds with the water contact angle. As stated in the experimental section, the porosity was determined using image analysis of the corresponding SEM micrographs. The values at 0% porosity correspond to the solution cast films. The results are reported in Table 7.2; it was found that contact angle increased linearly with the degree of porosity. These results also showed that the scaffold with the most uniform filament morphology, that is, smooth cylindrical filaments with nearly constant diameters, exhibit low degree of porosity and therefore the contact angle is quite close to that obtained for cast films. Furthermore, the results showed that for similar degree of porosity the hydrophobic behavior of PHB230 is only slightly larger than PHB50, thus, reinforcing the notion that scaffold morphology, that is, degree of porosity, drives the contact angle on the scaffolds. These findings are important in the context of scaffolds for tissue engineering because they suggest that the conditions of electrospinning that control the scaffold morphology determine the adhesion to the surface of the scaffolds.

7.3.3 AGING The influence of storage time, that is, aging, of materials needs to be addressed as physical or chemical changes may occur affecting their performance. PHB is prone to become brittle due to increases in crystallinity (Barham et al., 1984; Dominguez-Diaz et al., 2011). In order to address the influence of storage time, the electrospun scaffolds were stored at room temperature under dry conditions for up to 11 months. The degree of crystallinity, α, and crystal size, Lhkl, were monitored during the storage time. Fig. 7.3A, inset, shows a typical WAXS pattern of a PHB50 electrospun scaffold. The processing condition were 13% w/v concentration and at 9 kV. The WAXS pattern was obtained with the incident beam perpendicular to the scaffold surface. The crystalline isotropic rings indicate a random orientation within the plane of the scaffold denoting lack of preferred macromolecular orientation (Alexander, 1969). Similar results were obtained for all scaffolds investigated, that is, no macromolecular alignment was detected regardless of electrospinning conditions. The intensity peaks are in agreement with the α-form of the orthorhombic unit cell of PHB (Hurrell and Cameron, 1998a,b). Fig. 7.3A shows the evolution of the degree of crystallinity α within the scaffolds and Fig. 7.3B shows the lateral crystal size values along the 020 direction, L020, as a function of storage time at room temperature for all the scaffolds investigated. It is well-known that PHB becomes brittle on storage time at room temperature (De Koning and Lemstra, 1993; Biddlestone et al., 1996; Hurrell and Cameron, 1998a,b). The deterioration of the mechanical properties has been associated to the progressive development of secondary crystallization that has been found to extend over months; physical ageing of the amorphous regions seems to be subsidiary in this phenomenon. The data of Fig. 7.3 show that for all the PHB scaffolds, only PHB13&9 exhibits a clear increase of the degree of

7.3 PHB Electrospun Fibrous Scaffolds

FIGURE 7.3 (A) Degree of crystallinity, α, and (B) lateral crystal size along the (020) direction, L020, for the electrospun PHB50 scaffolds, as a function of storage time at room temperature. Concentration and voltage for electrospinning conditions: (K) 12% and 9 kV, (’) 13% and 9 kV, and (V) 13% and 12 kV.

crystallinity with increasing storage time. The concurrent variation of α and L020 for the PHB13&9 scaffold suggests that the observed α-enhancement is associated to an improvement in the lamellar morphology due to secondary crystallization and not, for example, to the destruction of amorphous regions during an incipient degradation process. The fact that the crystallinity α (and crystal size L020) remains constant with storage time for the other PHB scaffolds indicates that the morphology plays a significant role on the development of crystallinity upon storage.

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It is noteworthy that the PHB12&9 and PHB13&12 scaffolds exhibit high levels of porosity (above 50%) with respect to PHB13&9 (  22%). The fact that secondary crystallization cannot be detected for PHB12&9 and PHB13&12 could be associated to the large surface area characteristic of porous morphologies that would favor a more efficient degradation under ambient conditions that could result in destruction of lamellar stacks. On the other hand, it may be that some solvent from the electrospinning process could still be occluded within the polymer filaments preventing the proper development of interlamellar stacks.

7.3.4 STERILIZATION METHODS AND INFLUENCE ON PHYSICAL PROPERTIES Biomaterials intended for medical applications need to go through a process of sterilization in order to avoid contamination and/or biofouling. This issue is hardly addressed as if the sterilization procedure has not consequences on the biomaterials’ physical properties. Sterilization procedures like ultraviolet (UV) radiation and autoclaving are widely used methods as they are simple and can be easily adapted to mass production. However, it is well-documented that exposure to UV radiation may degrade polymeric materials by bond cleavage. On the other hand, autoclaving subjects the polymers to high temperatures (c. 100 C). This limits the application of autoclaving to polymers with Tg and/or Tm . 100 C. Even so autoclaving may trigger cold crystallization, like in the case of the slow crystallizing biopolymer polylactic acid, PLA, which leads to shrinkage and whitening and these effects need to be taken into account when designing medical devices (Reyes-Mayer and Romo-Uribe, 2014). The influence of both, UV and autoclaving sterilization was investigated for PHB electros-spun scaffolds and reported in detail by Romo-Uribe et al., 2017. Autoclave was carried out at 120 C for 20 minutes. Infrared spectra of the (1) as-received PHB produced from OP strain (230 kDa) as well as corresponding spectra after (2) UV and (3) autoclave sterilization did not show changes in the resonance bonds of PHB suggesting that the sterilization procedures did not degrade the polymer. Furthermore, differential scanning calorimetry scans were applied to assess any possible changes in thermal phase transitions of PHB230. Investigation on a series of PHBs with broad range of molecular weights (50
2050 kDa) showed that the thermal transitions are strongly influenced by molecular weight, and reduction of molecular weight by degradation would be reflected in lower melting and crystallization temperatures (Dominguez-Diaz et al., 2015a,b). The results on specimens before and after sterilization confirmed that the thermal transitions were not influenced by any of the sterilization procedures, the melting (Tm) and crystallization (Tc) temperatures remained constant at about 180 C and 106 C, respectively. Therefore, these results suggested that any of these sterilization procedures did not degrade PHB (Romo-Uribe et al., 2017). The autoclave

7.4 Cell Viability and Bone Tissue Regeneration

sterilization procedure was chosen for cell viability and tissue engineering studies. It is important to determine the influence of sterilization procedures on the physical properties of biomaterials as biomedical applications or tissue engineering requires sterile materials to guarantee the purity of the constructs.

7.4 CELL VIABILITY AND BONE TISSUE REGENERATION First, recent results on the suitability of a biosynthesized PHB elecrospun scaffold to support cell growth, that is, cell viability, will be presented. The PHB was biosynthesized utilizing a modified mutant strain of A. vinelandii (OPN), as described by Pen˜a et al. (2014a) and Dominguez-Diaz et al. (2015a,b). Cell viability and cytotoxicity was evaluated using human embryonic kidney 293 cells (HEK293). HEK293 is a transformed, noncarcinogenic cell line. The results will show that PHB scaffolds indeed support cell growth and exhibit higher cell viability than empty dish cultures. Then, recent results on the application of PHB scaffolds to bone tissue regeneration will be presented. For this, and in order to closely mimic the in vivo environment normal human osteoblast (NHOst) cells were utilized.

7.4.1 CELL VIABILITY AND HEK293 CELLS Human embryonic kidney 293 cells (HEK293) were seeded on electrospun scaffolds of PHB230. The PHB scaffolds (about 0.5 3 0.5 cm) were electrospun from 8.5% w/v solution at 8 kV. Prior to sterilization and seeding the scaffold was stored under dry conditions at room temperature, and storage time did not exceed 1 month. The culture medium was DMEM-F12 (Dulbecco’s Modified Eagle Medium, GIBCO Invitrogen) supplemented with 10% of Bovine Fetal Serum (BFS) and 1.2 g of sodium bicarbonate (NaHCO3), and streptomycin/penicillin antibiotic (GIBCO Invitrogen); pH was adjusted to 7.2 6 0.03. The PHB scaffolds were sterilized by autoclave, and placed into a Petri dish of 3.5 cm in diameter where HEK293 cells were inoculated. The initial concentration of cells was kept constant at 1.425 3 106 cells/mL and cultivation was monitored for 144 hours. Cell growth was monitored using a phase contrast microscope Olympus CK40 (objective 40 3 ). In order to determine cell concentration and viability cells were removed from the Petri dish after washing with 1 mL of sterile phosphatebuffered saline (PBS) and 2 mL trypsin/Ethylenediaminetetraacetic acid (EDTA) 1 3 (Gibco). Then, cell counting using the trypan blue exclusion technique was performed in triplicate using a hemocytometer and observed in an optical microscope Axiostar (objective 10 3 ). The HEK293 cells were seeded and the morphology of the cells in Petri dishes was monitored during cell cultivation for up to 6 days (144 hours), as shown by

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the micrographs in Fig. 7.4. The results show that upon seeding (day 0) the cells were fairly well dispersed all over the plate. After 24 hours culture, the cells exhibited the typical elongated morphology. Moreover, it was observed that cell growth was homogeneous throughout the plates, indicating no rejection of PHB by the cells and no toxic effect that could be detected by cell death, as compared to a control specimen (i.e., dish without PHB scaffold). After 72 hours of culture the cells exhibited normal epithelial cell morphology on and into the PHB scaffolds as well as in the plate surface. Finally, after 144 hours the cells were arranged in layers which covered entirely the scaffold. It is noted that the control also exhibited a layer of cells but it did not fill entirely the surface of the plate as open spaces could be easily appreciated, as shown in Fig. 7.4. It is noted that the cell culture in all cases started with 1.425 3 106 cells/mL, and after 144 hours the viable cells in the PHB scaffold reached 1.8 3 107 cells/ mL. On the other hand, the control, empty dish exhibited viable cells of only 1.6 3 107 cells/mL. The results showed that the cell viability for the scaffold, after 6 days culture, was nearly 96% whereas for the control was only about 92%, as shown in Fig. 7.5. Recent results on scaffolds of PHB obtained by solution cast film also showed effectiveness to support cell proliferation of HEK293 (Dominguez-Diaz et al., 2015a,b). The fact that HEK293 cells maintained their growth for 6 days in both electrospun and cast film scaffolds is indicative that no cytotoxic effects were induced by the biosynthesized PHB. In summary, these results demonstrated the biocompatibility of bacterial PHB scaffolds biosynthesized from a modified A. vinelandii strain. That is, cell growth was enabled even at higher densities than Petri dish cultures, and without any

FIGURE 7.4 Bright field optical micrographs of Human embryonic kidney 293 cells (HEK293) seeded on Petri dish (control) and PHB electrospun scaffold for up to 6 days.

7.4 Cell Viability and Bone Tissue Regeneration

FIGURE 7.5 Cell concentration and viability of HEK293 cells after 6 days seeded on Petri dish (control) and PHB electrospun scaffold.

cytotoxic effects. Furthermore, HEK293 cells were attached to PHB scaffolds but were also easily released by trypsin without any damage to the constructs.

7.4.2 BONE TISSUE REGENERATION AND HUMAN OSTEOBLAST CELLS The osteoblasts are responsible for bone formation in bone tissue by secreting and mineralizing the bone matrix, which is composed mainly of type I collagen and the inorganic mineral hydroxyapatite (HAp). In order to test the efficacy of PHB scaffolds on supporting osteoblast differentiation and bone formation, a first step is to investigate osteoblast adhesion, spreading, and proliferation by in vitro culture. This is achieved by investigating the cytocompatibility and suitability of a biomaterial for a specific application. A high degree of proliferation of cells is generally required as the degree of proliferation determines the rate of ECM formation. Osteoblast proliferation will then be followed by matrix maturation and matrix mineralization (Stein and Lian, 1993). These last two stages are not discussed here. Furthermore, in order to closely mimic the in vivo environment in this research Normal Human Osteoblast (NHOst) cells were utilized.

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The NHOst cells from a male donor (CAS: CC-2538), cell medium BulletKit, a ReagentPack, Trypsin/EDTA, trypsin neutralizing solution (TNS), and HEPES buffered saline solution were purchased from Lonza (NJ, USA). The NHOst were cultured in 5 cm culture flasks dishes with 5 mL of culture medium and amplified for obtain enough cells for the experiment. The sub culturing and maintenance of the cells were carried out according to Lonza recommended protocol for this cell line, namely: “when the cells are 80% confluent and for each 25 cm2 of cells to be sub cultured wash cells with 5 mL of HEPES-BSS, cover with 2 mL of trypsin/EDTA to detach the cells (2
6 minutes), after the cells are detached neutralize the trypsin with 4 mL of TNS and rinse the flask with 2 mL of HEPED-BSS to collect residual cells and centrifuge the harvested cells at 220 g for 5 minutes to obtain the cell pellet.” For more details see Lonza Manual. The NHOst cells were used in the fourth pass and were obtained from two 25 cm flask dishes (10 mL of medium for each). The scaffolds were cut into squares of 0.5 3 0.5 cm and sterilized with a Hinotek Boxum Pressure gauge YXQ-LS-SII Vertical-Type autoclave for 20 minutes at 120 C in wet mode. The experiments were carried out in two 24-well culture Corning polystyrene plates using control (wells only with cells, no scaffolds), and scaffolds. First the scaffolds were placed in wells and the cells were seeded at 0.875 3 105 cells/mL. The cell culture was monitored for up to 168 hours using a phase contrast microscope Olympus CK40 (objective 40 3 ). The cell viability was measured at 0, 24, 96 and 168 hours for cells detached, in triplicate for each sample (control and scaffold), using trypan blue technique and a Neubauer chamber cell, and observations were carried out with an optical microscope Axiostar equipped with a 10 3 objective. The results showed (Fig. 7.6) that the attachment ability of osteoblasts cultured onto PHB fibrous scaffolds and culture dishes are comparable at day 1. However, after 6 days, the proliferation of cells cultured onto all fibrous scaffolds of PHB was significantly higher than that onto culture dishes. Moreover, analysis of the area covered by the cells, which reflects the cellular spreading, adhesion and growth abilities demonstrated similar tendency as that of the cell proliferation results in Fig. 7.6. Interestingly, the cell area on PHB scaffolds was nearly twofold larger than that on culture dishes. The morphology of osteoblasts attached and proliferated on the fibrous PHB scaffolds was observed by optical microscopy. After 3 days of culture, there were more cells attached on the scaffolds as compared to those grown on culture dishes. After 4 days of culture, there was higher cell numbers and an extensive network of cell lamellipodia and filopodia woven into and integrated within the PHB network of microfibers, as shown in Fig. 7.7. Note that the term lamellipodia and filopodia are used here to describe the morphological features of the NHOst cells. Lamellipodia are thin, sheet-like membrane protrusions found at the leading edge (front) of motile cells such as endothelial cells, neurons, immune cells, and epithelial cells. These structures are generally devoid of major organelles and are instead composed of a dense and dynamic

7.4 Cell Viability and Bone Tissue Regeneration

FIGURE 7.6 Bright field optical micrographs of NHOst cells seeded on Petri dish (control) and PHB electrospun scaffold for up to 7 days.

FIGURE 7.7 Proliferation and morphology of osteoblast cultured onto PHB electrospun scaffold for 4 days.

network of actin filaments. On the other hand, filopodia are thin, actin-rich plasma-membrane protrusions that function as antennae for cells to probe their environment. Consequently, filopodia have an important role in cell migration, neurite outgrowth, and wound healing and serve as precursors for dendritic spines in neurons. Furthermore, it was observed that the overall NHOst cell population was still noticeably less dense on culture dishes than on fibrous scaffolds.

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FIGURE 7.8 Cell concentration and viability of NHOst cells after 6 days seeded on Petri dish (control) and PHB electrospun scaffold.

Trypan blue was utilized for cell counting and the results showed that growth of cells was slightly more pronounced on the electrospun scaffold. This can be seen in Fig. 7.8 which shows a plot of cell concentration and viability after 6 days of culture. The overall results showed that the cell viability of the NHOst cells after 168 hours of cell culture was c. 90% for the control and c. 95% for the scaffold. Finally, the integrity of the NHOst cells was assessed by staining with acridine orange (AO) and propidium iodide (PI) using an epifluorescence microscope. Such a technique is helpful to determine the type of death exhibited by cells and it is based on the use of a dye that diffuses inside the cells when they have lost the permeability membrane. Cell death could occur by a passive or active process, and death cells are stained blue while white cells are viable. Fig. 7.9 shows that the nuclei of cells detached from the culture dishes and PHB scaffold did not present fragmentation or signal of apoptosis (i.e., a process of programmed cell death that occurs in multicellular organisms), nor necrosis (i.e., the death of most or all of the cells in an organ or tissue due to disease, injury, or failure of the blood supply). After 168 hours, the NHOst cells remained green (viable) and the integrity of nuclei was maintained showing biocompatibility of the biosynthesized PHB scaffold to promote cell growth. Such results in conjunction with kinetic results also showed that cell attachment to the PHB scaffold is promoted and does not inhibit cell growth. Furthermore, it was observed that after 24 hours the time to detach the cells from the cast PHB film and scaffold was twofold than for the control, and after

7.5 Concluding Remarks

FIGURE 7.9 Micrographs obtained by epifluorescent microscopy of NHOst cells stained with Acridine Orange and seeded on Petri dish (control) and PHB electrospun scaffold.

168 hours the time to detach the cells from the PHB scaffolds was threefold. This difficulty to detach the cells may be responsible that cells exhibited irregular membrane, and suggests that PHB promotes adhesion of NHOst cells to its surface. Overall, these results indicate that osteoblasts seeded on fibrous scaffolds spread, migrate, adhere, and communicate much faster compared with that seeded on the smooth surface of culture dishes. This can be attributed to high specific surface area with these biomimetic microfibers. Osteoblasts are anchoragedependent cells, therefore, the high specific surface area and porosity of electrospun fibrous templates could promote attachment and proliferation of cells on the scaffold. Therefore, it has been shown that the microbial PHB produced from a mutant strain of A. vinelandii is cytocompatible. Furthermore, the electrospun fibrous scaffolds mimic the ECM architecture and have a good capability to induce more efficient cell proliferation in vitro opening up opportunities for tissue engineering, and membrane guided bone repair.

7.5 CONCLUDING REMARKS This investigation focused on the electrospinning fabrication for the production of fibrous PHB scaffolds utilizing a biosynthesized PHB from a modified mutant strain of A. vinelandii. It was also assessed the cellular responsive behavior pertaining to bone tissue engineering. The hypothesis is that biomimetic scaffolds

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could ultimately enhance the bone regeneration efficacy upon transplantation. On the one hand, the overall topographical resemblance of electrospun fibrous scaffolds to the ECM architecture could provide a favorable milieu to naturally regulate cell adhesion, proliferation, and differentiation. The results showed that solution concentration and applied voltage determined the morphology of electrospun scaffolds, including fiber diameter, and porosity. The fibrous and porous morphology, thus, produced mimics natural constructs and induced slightly hydrophobic behavior in an otherwise relatively hydrophilic material. Aging of the fibrous scaffolds for up to 11 months showed that nonhomogeneous morphologies were relatively stable whereas uniform filamentous morphologies exhibited significant increase of crystallinity leading to brittleness. The biosynthesized PHB stand the sterilization processes commonly used in biological laboratories or commercial applications for materials in contact with cells: UV and autoclave. Moreover, the PHB scaffolds were cytocompatible and promoted cell growth of HEK293. The cell morphology in the culture dishes and scaffold was normal. The cell viability of the HEK293 cells after 6 days of cell culture was 92% in dish and 95.7% in the electrospun scaffold. That is, the scaffolds not only enabled cell growth, but were more efficient. Furthermore, the results showed that after 18 days of culture the scaffold lost c. 90% of mass. This is a favorable outcome for the use of PHB in tissue engineering where it is expected that the material will degrade as the tissue of the body will regenerate. Finally, the (PHB) electrospun scaffolds are a viable alternative to bone regeneration. It was shown that the scaffolds were effective to promote the cell adhesion and viability of Normal Human Osteoblast (NHOst) cells. The results showed that after 168 hours of cell culture NHOst cells were still growing and cell viability was higher for the electrospun scaffold relative to culture dishes. No significant changes were observed in the cells by epiflourescence microscopy, demonstrating that PHB scaffolds did not induce death by apoptosis or necrosis. Therefore, the biomimetic scaffold would enable the possibility of implantation for tissue remodeling in a well-controlled manner. These results provide strong evidence that the fibrous architecture of PHB for osteoblasts scaffolding is beneficial for promoting bone formation, cell adhesion, and proliferation in vitro. These results on cytocompatibility, cell growth and proliferation (i.e., bone formation ability) warrant future work aimed at in vivo performance evaluation.

GLOSSARY OF TERMS AFM Cu Kα DSC ECM HEK293 kDa

atomic force microscopy copper radiation source differential scanning calorimetry extracellular matrix human embryonic kidney 293 cells kilo daltons (103 g/mol)

References

NHOst Mv OP PHA PHB PLA POM SEM Tc Tm Tdec TGA WAXS

normal human osteoblast cells viscosity-average molecular weight Azotobacter nivelandii strain poly(hydroxyalkanoates) poly(β-hydroxybutyrate) polylactic acid polarized optical microscopy scanning electron microscopy crystallization transition temperature melting transition temperature onset of thermal decomposition temperature thermogravimetric analysis wide-angle X-ray scattering

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53. Maensiri, S., Sangmanee, M., Wiengmoon, A., 2009. Magnesium ferrite (MgFe2O4) nanostructures fabricated by electrospinning. Nanoscale Res. Lett. 4, 221
228. Marchessault, R.H., Okamura, K., Su, C.J., 1970. Physical properties of poly(3-hydroxybutyrate). II. Conformational aspects in solution. Macromolecules 3, 735
740. Mather, P.T., Luo, X., Rousseau, I.A., 2009. Shape memory polymer research. Annu. Rev. Mater. Res. 39, 445
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1085. Moffa, M., Polini, A., Sciancalepore, A.G., Persano, L., Mele, El, Passione, L.G., et al., 2013. Microvascular endothelial cell spreading and proliferation on nanofibrous scaffolds by polymer blends with enhanced wettability. Soft Matter. 9, 5529
5539. Pavlukhina, S., Lu, Y., Patimetha, A., Libera, M., Sukhishvili, S., 2010. Polymer multilayers with pH-triggered release of antibacterial agents. Biomacromolecules 11, 3448
3456. Pen˜a, C., Lopez, S., Garcia, A., Espin, G., Romo-Uribe, A., Segura, D., 2014a. Biosynthesis of poly-β-hydroxybutyrate (PHB) with a high molecular weight mass by a mutant strain of Azotobacter vinelandi (OPN). Ann. Microbiol. 64, 39
47. Pen˜a, C., Castillo, T., Garcia, A., Millan, M., Segura, D., 2014b. Biotechnological strategies to improve production of microbial poly-(3-hydroxybutyrate): a review of recent research work. Microbial. Biotechnol. 7, 278
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4180. Reyes-Lo´pez, Y., Alvarado-Tenorio, B., Romo-Uribe, A., 2013. Polycaprolactone/ α-alumina and hydroxyapatite-based micro- and nano- structured hybrid fibers. MRS Symp. Proc 1569. Available from: https://doi.org/10.1557/opl.2013.798. Romo-Uribe, A., Alvarado-Tenorio, B., Romero-Guzma´n, M.E., 2010. A small-angle light scattering instrument to study soft condensed matter. Rev. LatinAm. Metal. Mat 30, 190
200. Romo-Uribe, A., Arizmendi, L., Romero-Guzma´n, M.E., Sepu´lveda-Guzma´n, S., CruzSilva, R., 2009. Electrospun nylon nanofibers as effective reinforcement to polyaniline membranes. ACS Appl. Mater. Interfaces 1, 2502
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References

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138. Valappil, S.P., Misra, S.K., Boccaccini, A.R., Roy, I., 2006. Biomedical applications of polyhydroxyalkanoates: an overview of animal testing and in vivo responses. Exp. Rev. Med. Dev. 3, 853
868. Vargas-Villagra´n, H., Tera´n-Salgado, E., Domı´nguez-Dı´az, M., Flores, O., Campillo, B., Flores, A., et al., 2012. Non-woven membranes electrospun from polylactic acid incorporating silver nanoparticles as biocide. MRS Symp. Proc. 1376. Available from: https://doi.org/10.1557/opl.2012.285. Vargas-Villagra´n, H., Romo-Uribe, A., Tera´n-Salgado, E., Dominguez-Diaz, M., Flores, A., 2014. Electrospun polylactic acid non-woven mats incorporating silver nanoparticles. Polym. Bull. 71, 2437
2452. Wang, Q., Uzunoglu, E., Wu, Y., Libera, M., 2012. Self-assembled poly(ethylene glycol)co-acrylic acid microgels to inhibit bacterial colonization of synthetic surfaces. ACS Appl. Mater. Interfaces 4, 2498
2506. Wang, Q., Libera, M., 2014. Microgel-modified surfaces enhance short-term osteoblast response. Coll. Surf. B. Biointerfaces 118, 202
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5452. Wang, Y.W., Wu, Q., Chen, G.Q., 2004. Attachment, proliferation and differentiation of osteoblasts on random biopolyester poly(3-hydroxybutyrate-co-3-hydroxyhexanoate) scaffolds. Biomaterials 25, 669
675. Wang, Y.W., Yang, F., Wu, Q., Cheng, Y.C., Yu, P.H., Chen, J., et al., 2005. Effect of composition of poly(3-hydroxybutyrate-co-3-hydroxyhexanoate) on growth of fibroblast and osteoblast. Biomaterials 26, 755
761. Wu, J., Hou, S., Ren, D., Mather, P.T., 2009a. Antimicrobial properties of nanostructured hydrogel webs containing silver. Biomacromolecules 10, 2686
2693. Wu, Q., Wang, Y., Chen, G.Q., 2009b. Medical application of microbial biopolyesters polyhydroxyalkanoates. Artif. Cells Blood Substitutes Immobilization Biotechnol. 37, 1
12. Xiao, S., Shen, M., Guo, R., Huang, Q., Wang, S., Shi, X., 2010. Fabrication of multiwalled carbon nanotube-reinforced electrospun polymer nanofibers containing zerovalent iron nanoparticles for environmental applications. J. Mater. Chem. 20, 5700
5708. Xu, J., Song, J., 2010. High performance shape memory polymer networks based on rigid nanoparticle cores. PNAS 107, 7625
7657. Zhang, D., George, O.J., Petersen, K.M., Jimenez-Vergara, A.C., Hahn, M.S., Grunlan, M. A., 2014. A bioactive “self-fitting” shape memory polymer scaffold for potential to treat cranio-maxillo facial bone defects. Acta Biomater. 10, 4597
4605. Zhu, M., Zuo, W., Yu, H., Yang, W., Chen, Y., 2006. Superhydrophobic surface directly created by electrospinning based on hydrophilic material. J. Mater. Sci. 41, 3793
3797. Zhuk, I., Jariwala, F., Attygalle, A.B., Wu, Y., Libera, M.R., Sukhishvili, S.A., 2014. Selfdefensive layer-by-layer films with bacteria-triggered antibiotic release. ACS Nano 8, 7733
7745.

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FURTHER READING Chung, T., Romo-Uribe, A., Mather, P.T., 2008. Two-way reversible shape memory in a semicrystalline network. Macromolecules 41, 184
192. Dominguez-Diaz, M., Flores, A., Cruz-Silva, R., Romo-Uribe, A., 2012. Morphology induced hydrophobic behavior of electrospun polyhydroxyalkanoate membranes. MRS Symp. Proc. 1466. Available from: https://doi.org/10.1557/opl.2012.1256.

CHAPTER

Polyurethane-based structures obtained by additive manufacturing technologies

8

Pablo C. Caracciolo, Nayla J. Lores and Gustavo A. Abraham Research Institute for Materials Science and Technology, INTEMA (UNMdP-CONICET), Mar del Plata, Argentina

8.1 INTRODUCTION Additive manufacturing (AM) or three-dimensional printing is a fabrication technology where 3D solid objects are built by a layer-by-layer manufacturing process through computer-aided design (CAD) techniques (Chen et al., 2007). This technology is highly versatile and suitable for biomedical applications due to its ability to fabricate complex 3D scaffolds with high precision, geometric control at both macro and micro-scales, customized production, and energy and time consumption lower than that of conventional techniques. These advantages are attractive for biomedical applications, particularly in the tissue engineering field, where scaffolds, cells, and mechanobiochemical factors play a key role in the tissue/organ regeneration process. The design, fabrication, and implantation of both body external and internal biomedical devices present a high complexity. The development of permanent or temporary implants requires a combination of biocompatible materials, processing techniques, and cell cultures to produce 3D biomimetic functional structures (Moroni and Elisseeff, 2008). Moreover, soft- and hard-tissue engineering constructs involve bioresorbable highly-porous materials with pore interconnectivity, suitable surface properties and degradation time, and biomechanical properties matching that of native tissues (Bettinger et al., 2007). Several traditional processing techniques have been explored to prepare tissueengineered scaffolds, including fiber processing, solvent casting, particle leaching, freeze-drying, thermally-induced phase separation (TIPS), chemical/gas forming, hydrocarbon templating, self-assembly, and others (Ma and Elisseeff, 2006). In the past decade, electrospinning emerged as an attractive and versatile electrohydrodynamic technique to produce micro/nanofibrous nonwoven scaffolds for tissue engineering and many bio/nanotechnological applications (Jiang et al., 2015). On the other hand, CAD-based AM technologies have also received significant attention, allowing for the fabrication of acellular, cellular, or hybrid cell/scaffold Materials for Biomedical Engineering: Hydrogels and Polymer-based Scaffolds. DOI: https://doi.org/10.1016/B978-0-12-816901-8.00008-0 © 2019 Elsevier Inc. All rights reserved.

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CHAPTER 8 Polyurethane-based structures

Drop-on-demand Fused deposition inkjetprinting modeling (FDM) Continuous inkjet printing

1950

1960

Laser printing

1970

3D printing

Stereolithography (SLA)

1980

Chondrocytes redifferentiate in 3D environment First cell sorter

Microcontact printing

1990

2000

Tissue enfineering coined

Decade of additive tissue manufacture • Intern. Soc. for Biomanufacturing • Strong and bioactive hydrogels • Combinations of techniques • Digital voxel fabricators • Upscaling • Vascularised constructs

2010

Organ printing coined

2020

Robotic dispensing of cell-laden hydrogels Biolaserprinting (BLP) of cells

Protein printing Hydrogels with RGDsequence

First in vivo use of AM (BLP) Cell-containing gels by SLA specialized journal Biofabrication Inkjet printing of cells Laser direct writing of cells Tissue engineering scaffolds by FDM

FIGURE 8.1 Historical evolution of AM: applications in tissue engineering, new technologies, and relevant findings. Reprinted from Melchels, F.P.W., Domingos, M.A.N., Kleina, T.J., Malda, J., Bartolo, P.J., Hutmacher, D.W., 2012. Additive manufacturing of tissues and organs. Prog. Polym. Sci. 37, 1079
1104, with permission from Elsevier.

constructs (Hollister, 2005; Chen et al., 2007; Li et al., 2015). AM techniques can be classified following different criteria, such as initial physical condition of the processed material and the physical or chemical process used for the solidification of each layer (Gurr and Mu¨lhaupt, 2016). Regarding the latter criterion, techniques include inkjet-based systems (three-dimensional printing or inkjet gluing: 3D Printing, inkjet curing: Polyjet), laser-based systems (stereolithography: SLA systems, selective laser sintering: SLS systems), and extrusion-based manufacturing processes (direct ink writing: DIW, fused deposition modeling: FDM, lowtemperature deposition manufacturing: LDM, integrated organ printing: IOP, 3D bioplotting or biocell printing). Excellent reviews describing the basis and fundamentals of each technique have been reported (Bartolo et al., 2012; Melchels et al., 2012; Bajaj et al., 2014; Gurr and Mu¨lhaupt, 2016; Guvendiren et al., 2016). Fig. 8.1 illustrates the historical evolution of AM showing applications in tissue engineering, introduction of new technologies, and relevant scientific findings. Biofabrication is emerging as a rapidly growing research field closely related to bioprinting and biomanufacturing. Biofabrication for tissue engineering and regenerative medicine was defined as “the automated generation of biologically functional products with structural organization from living cells, bioactive molecules, biomaterials, cell aggregates such as micro-tissues, or hybrid cell-material constructs, through bioprinting or bioassembly and subsequent tissue maturation processes” (Groll et al., 2016). Thus, biofabrication within tissue engineering is not restricted to AM approaches, as it exploits automated processes to generate

8.2 Bioresorbable Polyurethanes in Biomedical Devices

constructs that may mature into functional tissues. As an example, biological selfassembly and bioprinting was proposed as a scaffold-free approach to engineer vascular and nerve grafts (Marga et al., 2012). AM can rapidly produce complex 3D customized models and constructs using a wide range of raw materials. Hybrid processes utilizing polymer melt deposition and electrospinning techniques were employed to fabricate micro/nanostructured architectures in view of tissue-engineered functionalized scaffolds. Moreover, another technique unifies FDM and electrospinning to obtain the ability of an AM technique to create controllable extruded patterns and of electrospinning to form nanofibers (Chanthakulchan et al., 2015). In this chapter, an overview of the latest advances in AM, specifically addressing bioresorbable polyurethanes and their composites, is presented and discussed.

8.2 BIORESORBABLE POLYURETHANES IN BIOMEDICAL DEVICES Biostable polyurethanes have been used in medical devices since 1965 with the introduction of Biomer into the market of cardiovascular devices (Boretos and Pierce, 1967). During many decades, the improvement of in vivo hydrolytic biostability of biomedical polyurethanes has been extensively investigated (Anderson et al., 1998). With the development of scaffolds for tissue engineering applications, the main efforts have focused on the design of biodegradable scaffolds with controlled biodegradation in vivo to support the ingrowth of cells and tissues. Thus, the formulation of bioresorbable polyurethanes needs to accomplish several design requirements that include biocompatible and noncytotoxic degradation products, tissue-mimetic mechanical properties, and degradation rate (Guelcher, 2008). Polyurethanes are a family of polymers that contain the carbamate or urethane (
NHCOO
) group in their chemical structure. Linear segmented polyurethanes (SPUs) are synthesized by step growth or condensation polymerization. In this reaction, bifunctional monomers react in a stepwise manner to produce macromolecules without the production of small molecules as byproducts (Szycher, 2012). Fig. 8.2 shows the general chemical structure of a SPU. The main reactions of

FIGURE 8.2 Schematic of poly(ester urethane) block copolymer structure formed from diisocyanate component (OCN-R-NCO), diol chain extender (HO-R0 -OH), and macrodiol building blocks.

237

238

CHAPTER 8 Polyurethane-based structures

polyurethane synthesis are: the formation of urethane linkage that occurs when an isocyanate functional group reacts with an alcohol; and the urea linkage that occurs when an isocyanate reacts with an amine. Two methods are used to synthesize SPU. In the one-shot method, a diisocyanate, a long chain oligomeric diol (macrodiol), and a chain extender are mixed and allowed to react at once. In a typical SPU synthesis, a prepolymer or two-step method is used. In this route, the macrodiol is reacted with a slight excess of diisocyanate to form a diisocyanateend capped prepolymer. The prepolymer is then reacted with a chain extender to form a linear block copolymer with alternating blocks of hard and soft segments. This method is the preferred procedure since it yields more ordered structures and controlled properties (Szycher, 2012). In general, macrodiols have low glass-transition temperatures (Tg) that are in the rubbery regime at physiological temperatures. The length and chemical structure of macrodiols play a key role in degradation rate, and physical and mechanical properties, ranging from low-modulus thermoplastic elastomers to rigid polyurethanes. The hard segment typically consists of a low molecularweight isocyanate reacted with a low molecular-weight diol or diamine chain extender. Hard segments are semicrystalline or highly ordered domains with high Tg that act as rigid fillers and thermally reversible physical crosslinks. The high elastic recovery and fatigue resistance of SPU have attracted significant interest in the biomedical field, particularly in applications where mechanical properties are key design criteria. Isocyanates and chain extenders also strongly influence the physical properties of SPU. The final properties are affected by several factors, including phase mixing, soft and hard segment chemistry, hard-to-soft segment ratio, molecular weight, surface chemistry, and processing techniques used for scaffolding (Santerre et al., 2005; Kro´l, 2007; Szycher, 2012). Polyurethane networks can be prepared by the incorporation of at least one reactant with functionality higher than 2. The formulation of networks by reactive two-component liquid molding make them suitable as injectable formulations for noninvasive therapies. Aromatic diisocyanates have high reactivity and the resulting SPU display excellent mechanical properties. However, aliphatic diisocyanates are the preferred option for bioresorbable SPU due to the fact that they degrade into nontoxic byproducts. Among these compounds, hexamethylene diisocyanate (HDI), butane diisocyanate (BDI), L-lysine methyl ester diisocyanate (LDI), isophorone diisocyanate (IPDI), and dicyclohexylmethane diisocyanate (H12MDI) are the most commonly used. Macrodiols ranging from 400 to 5000 g/mol include polyesters, polyethers, and combinations of these in the form of diblocks or triblocks. Polyesters are hydrolytically unstable and polyethers are susceptible to oxidation mediated by adherent inflammatory cells. Both types of macrodiols are commonly chosen for the formulation of bioresorbable SPU. Poly(glycolide) diol (PGA diol), poly

8.2 Bioresorbable Polyurethanes in Biomedical Devices

(L-lactic acid) diol (PLLA diol), poly(D, L-lactic acid) diol (PDLLA diol), poly (ε-caprolactone) diol (PCL diol), and poly((R)-3-hydroxybutyric acid-co-(R)-3hydroxyvaleric acid) diol (PHBV diol), and their copolymers are the most commonly used polyester diols. Among polyether diols, poly(tetramethylene) oxide (PTMO), poly(hexamethylene oxide) (PHMO), poly(propylene oxide) (PPO), and poly(ethylene) oxide (PEO) or the same polyethers with low molecular weight (PTMG, PHMG, PPG, and PEG) are selected when hydrophilicity and flexibility are required. Triblock copolymers, such as PCL-b-PEG-b-PCL, PLLA-b-PEG-bPLLA, and PEG-peptide-PEG are also used (Abraham et al., 2006; Guelcher, 2008). Multicomponent SPU based on PEG, PLLA, and poly(trimethylene carbonate) (PTMC) macrodiols are also reported (Bergamo Trinca et al., 2015). On the other hand, chain extenders include low molecular-weight diols or diamines. 1,4-Butanediol (BDO), 1,4-butanediamine (BDA, putrescine), ethylene glycol (EG), ethylenediamine (EDA), and 1,3-diaminopropane (DAP) have been extensively used in biomedical polyurethanes. BDO-BDI-BDO triblocks have also been used as urethane-based chain extenders to produce polyurethanes and polyurethane ureas with enhanced mechanical properties. Phosphatidylcholine diols have been used to suppress protein adsorption on SPU surfaces. The use of enzyme sensitive linkages or cell-responsive moieties are often incorporated into the classical polyurethane structure to improve the biological performance of these materials. Thus, degradable chain extenders, such as amino acids, biological peptides such as Arg
Gly
Asp
Ser (RGDS), diester-diols, and diamine terminated amino acid diesters are also incorporated into hard segment structures to encourage degradation of the material in vivo (Skarja and Woodhouse, 2000; Marcos-Ferna´ndez et al., 2006). The preparation, characterization, and properties of novel elastomeric bioresorbable SPU scaffolds, were reported based on PCL, HDI or LDI, and chain extenders with diurea diol groups (Caracciolo et al., 2008) or diphenol terminated amino acid diester derived from desaminotyrosine (Caracciolo et al., 2009a). These polymers presented a combination of mechanical properties and promising in vitro biological properties (Caracciolo et al., 2008, 2009b). Degradation rate depended on the chemical structure of the chain extenders, as well as the morphology and crystallinity of the materials (Caracciolo et al., 2011). Moreover, SPU/PLLA blends allowed for the preparation of bilayered small-diameter nanofibrous tubular structures (Montini Ballarin et al., 2014) with interesting biomimetic mechanical behavior (Montini-Ballarin et al., 2016a) and degradation properties (Montini-Ballarin et al., 2016b). Waterborne or water-based polyurethanes were developed as an alternative to the traditional solvent-borne SPU in response to environmental issues. These SPU can be synthesized as nanoparticles and dispersed in water by incorporating ionic hydrophilic groups onto hydrophobic backbones, without using external emulsifiers. The low viscosity of waterborne SPU dispersions complicates their processing. Therefore, to enhance processability, water-soluble polymers, such as PEO and polyvinyl alcohol (PVA), are added to increase the viscosity of the dispersion.

239

240

CHAPTER 8 Polyurethane-based structures

8.3 ADDITIVE MANUFACTURING FOR BIOMEDICAL POLYURETHANE PROCESSING Biodegradable SPU can be processed into various products, such as freeze-dried foams, electrospun fibrous scaffolds, and 3D-printed scaffolds, using solvents, radiation, or heat. There are many available AM technologies for processing polyurethanes, but each method has its own disadvantages and applicability. SLA often requires the use of toxic photoinitiators that cannot be completely removed after processing. Therefore, SLA is not suitable for manufacturing implantable biomedical devices. Inkjet printing and SLS need a large amount of polymer powder which produces waste material. In FDM, polymeric filaments may undergo thermal or chemical degradation and LDM often uses toxic solvents. Water-based 3D printing is limited to only some natural polymers, such as chitosan and alginate, while others require crosslinking treatments to improve their mechanical properties. 3D polyurethane structures are obtained from almost every one of them, however, the main printing technologies employed to obtain polyurethane-based scaffolds for biomedical applications are mainly inkjet printing, extrusion-based methods, and particle binding.

8.3.1 INKJET PRINTING Inkjet printing technology is based on the deposition of small droplets (1
100 pL) of an ink from a nozzle that solidify upon contacting a collecting plate, obtaining coatings or 3D structures at the end of the process. Ceramic or polymer suspensions, polymer solutions, and solutions containing cells or bioactive agents have been employed. The main advantage of inkjet printing is its spatial resolution (B10 μm in x
y-axes), and the process can be accelerated using multinozzle systems.

8.3.2 EXTRUSION-BASED METHODS In these methods, a viscous liquid or melt is extruded through a nozzle and solidifies on a collecting platform. The method of DIW entails the use of hydrogels that increase their viscosity when extruded, or polymer solutions from watermiscible low-boiling point solvents that rapidly evaporate upon extrusion, thereby obtaining 3D scaffolds. Low-temperature or liquid-frozen deposition manufacturing (LFDM) is a DIW technique in which ink is extruded and immediately freezes when it makes contact with a low-temperature collecting plate. Thus, the structure does not collapse during printing. Bioactive agents can be employed as limited heating is needed at the nozzle. Cells can also be included in printing inks but the use of cryoprotectants is needed to maintain cell viability. The polymer system goes through a TIPS during deposition and can be freeze-dried to eliminate

8.4 Additive Manufacturing of Composite Polyurethanes

residual solvent. Double-nozzle low-temperature deposition manufacturing (DLDM) is another DIW technique, basically with the same concept as LFDM, but with the possibility of the coextrusion of another biomaterial with different properties. This can improve, for example, hydrophilicity and cell adhesion of the final scaffold. Another multinozzle ink writing technology proposed for tissueengineered constructs is the three-dimensional IOP technique, where different polymer solutions and cell suspensions can be coprinted into a 3D-complex structure. Finally, FDM involves the use of a solid filament that is rolled and extruded through a heated nozzle. This methodology employs filaments of materials that can be melted without thermal degradation, this being one of the main drawbacks of FDM. The spatial resolution of these techniques is in the order of B25 μm in x
y-axes.

8.3.3 PARTICLE BINDING Particle binding technology uses a solution that is extruded through a nozzle onto a bed of powder and binds particles together in a layer-by-layer process. Spatial resolution is in the order of B700
1000 μm in x
y-axes, and B100 μm in z-axis, being strongly affected by particle size.

8.4 ADDITIVE MANUFACTURING OF COMPOSITE POLYURETHANES Several polyurethane-based scaffolds have been obtained using different 3Dprinting techniques. Table 8.1 summarizes the systems found in the literature.

8.4.1 INKJET PRINTING Studies comprising of biomedical polyurethanes processed by inkjet printing were conducted. A SPU synthesized from 4,40 -methylenebis(phenyl isocyanate) (MDI), PCL diol (Mn 5 530 Da), and N,N-bis(2-hydroxyethyl)-2-aminoethanesulfonic acid (BES) as the chain extender, to yield a pH-sensitive polymer, was employed to obtain single-layer patterns (Zhang et al., 2008a). First, SPU were dissolved in an alkaline solution and dropped onto glass slides to form liquid layers. The slides were then placed under the print head of a HP desktop 3900 printer with a HP 21 black ink cartridge containing an acetic acid solution. The acetic acid solution was printed over the slides, thereby obtaining patterns due to polyurethane precipitation because of a local pH decrease. This group carried out another work with the same concept, employing a pH-sensitive SPU synthesized from HDI, PCL diol (Mn 5 530 Da), and N,N-bis(2-hydroxyethyl)glycine (bicine) as a chain extender (Zhang et al., 2008b). The cartridge was filled with an acidic calcium

241

Table 8.1 Polyurethane-Based Scaffolds Obtained by Additive Manufacturing Techniques SPU Scaffold

Fabrication Technique

PCL:HDI:bicine/Ca12 PEG:IPDI:BDO:glycerol/silica

Inkjet printing DIW

PCL:MDI:BDO and Texin/PEGDGE-Jeffamine hydrogels PCL:BDI:BDA/paclitaxel PCL/PDLLA:IPDI:DMPA:EDA/NSCs PCL/PLLA-co-PEO:IPDI: DMPA:EDA/hMSCs PCL/PDLLA:IPDI:DMPA:EDA/SPI/NSCs PCL/PEBA:IPDI:DMPA:EDA/PEO (72/28 to 80/20 ratio) PCL/PEBA:IPDI:DMPA:EDA/hyaluronan (HA), and TGFβ3 or Y27632 PCL/PEG:HDI:BDO/gelatin-lysine PCL/PEG:HDI:BDO/type I collagen

Mechanical Performance

Cell Type

Reference


Osteoblast-like cells


Zhang et al. (2008b) Pfister et al. (2004)

DIW


Edry 5 740.7/εdry 5 7.6Ewet 5 386.4/ εwet 5 12 E 5 3.68/ε 5 247

DIW DIW DIW DIW LFDM





ESPU/PEO 5 0.48/ESPU/PEO 5 0.40

Perkins et al. (2014) Hsieh et al. (2015) Tsai et al. (2015) Lin et al. (2016) Hung et al. (2014)

LFDM

ESPU/HA 5 0.33

Rat SMCs NSCs hMSCs NSCs Rat chondrocytes hMSCs

DLDM DLDM





Xu et al. (2008a,b) Cui et al. (2009a)

In vivo rat study RSC96 Schwann cells Rat ADSCs

Wang et al. (2009)

PCL/PEG:HDI:BDO/type I collagen

DLDM

PCL/PEG:HDI:BDO/type I collagen

DLDM

PCL/PEG:HDI:BDO/ADSC/gelatin/alginate/ fibrinogen PCL/PEG:HDI:BDO/ADSC/gelatin/alginate/ fibrinogen Tecoflex/C2C12PCL/NIH3T3

DLDM

EPU 5 30.8/εSPU . 500/σSPU 5 3.64 ESPU 12%B12.5/εSPU 12%B530/Ecoll 12%B12.5/εcoll 12%B530 ESPU 12%B12.5/εSPU 12%B530/Ecoll 12%B12.5/εcoll 12%B530 ESPU 12%B12.5/εSPU 12%B530/Ecoll 12%B12.5/εcoll 12%B530


DLDM

ESPU

IOP

ESPU/C2C12 5 0.39/Einter 5 1.03/EPCL/ NIH3T3 5 46.67 ESPUB250
300 Edry 5 580/εdry 5 1.8Ewet 5 2.5/εwet 5 29

SPU/PVA (Gel-Lay) ZP11 powder (starch, cellulose, dextrose)/2propanol/glycerol/LDI

FDM Particle binding

coat 5 67/εSPU coat 5 169

Rat ADSCs C2C12 and NIH3T3 hFOB Osteoblast-like cells

Agrawal et al. (2013)

Hung et al. (2016)

Cui et al. (2009b) He and Wang (2011) Wang et al. (2013) Merceron et al. (2015) Castro et al. (2016) Pfister et al. (2004)

E, Young’s modulus (MPa); σ, tensile strength (MPa); ε, elongation at break (%), DIW, Direct ink writing; DLDM, double-nozzle low-temperature deposition manufacturing; FDM, fused deposition modeling; IOP, integrated organ printing; LFDM, liquid-frozen deposition manufacturing.

8.4 Additive Manufacturing of Composite Polyurethanes

chloride water solution (pH 4.0, 0.6 M), and patterns were obtained due to solution printing over basic polyurethane liquid layers (pH 8.6). Although 2D patterns were obtained for all SPU, this technique can potentially be adapted to 3D patterns. Furthermore, drug-release scaffolds could also be produced through the incorporation of bioactive molecules into the cartridge, leading to encapsulation during printing, such as is performed with calcium chloride.

8.4.2 EXTRUSION-BASED METHODS 8.4.2.1 Direct ink writing DIW techniques have also been used to obtain polyurethane-based scaffolds. Pfister et al., produced a composite polyurethane scaffold from liquid polyurethane precursors (Pfister et al., 2004). By adjusting the viscosity of the mixture through the addition of amorphous pyrogenic silica as a thixotropic agent to prevent structural collapse, low viscosity prepolymers composed of IPDI, PEG (Mn 5 600 Da) as a soft segment, BDO as a chain extender, glycerol as a curing agent, and dibutyl tin dilaurate (DBTDL) as a catalyst, were prepared. The paste was printed at 25 C in silicone oil with the aid of a compressed air line, employing a 3D Bioplotter (Envision Technologies, Marl, Germany). The scaffolds were then cured at 60 C. Porous scaffolds with smooth surfaces and homogeneous crosslinking were obtained. Experiments with an osteoblast-like cell line were also performed, suggesting that the scaffold is suitable for cell seeding. Agrawal et al., employed two aromatic SPU to obtain blend fibers for hydrogel-reinforcement (Agrawal et al., 2013). SPU-A was synthesized from PCL diol, MDI, and BDO with propylene glycol as chain extenders and SPU-B (Texin DP7-1205, Bayer Material Science LLC) were employed. A 3:1 ratio of SPU-A to SPU-B was employed to prepare a 20% (w/v) solution in N,N-dimethylformamide (DMF). This solution was loaded in a syringe with a 100 μm needle and placed in an Asymtek Automove 402 dispensing system. The process to obtain 3D fibrous scaffolds was carried out under water, and fibers were formed due to solvent exchange. The density of the scaffolds was dependent on the distance between fibers in each layer, while the fiber diameter was a function of the syringe back pressure and writing speed. Fiber-reinforced hydrogels prepared from poly(ethylene glycol), diglycidyl ether, and Jeffamine (poly(oxyalkylene)amine) presented higher fracture toughness values than that of natural cartilage. Perkins et al., employed DIW technology to surface modify intricate implant geometries depositing drug loaded thin-film coatings (Perkins et al., 2014). To enhance the biocompatibility of surgical implants and devices, multilayer coatings of a bioresorbable poly(ester urethane urea) synthesized from PCL diol (Mn 5 2000 Da), BDI, and BDA, loaded with paclitaxel as an antiproliferation drug were deposited on metallic substrates (Ti6Al4V). Monodisperse droplets of solutions with 0.5% (w/v) SPU in 1,1,1,3,3,3-hexafluoro-2-propanol (HFIP) and 5%
10% (w/w) drug/SPU were deposited on multiple layers using a 50μm

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nozzle. Coatings led to the reduction of platelet adhesion and rat smooth muscle cells proliferation, being promising for 3D contours of cardiovascular devices. Shan-hui Hsu’s group published several works on DIW of water-based biodegradable polyurethane dispersions for tissue engineering applications. This methodology avoids the use of heat, toxic organic solvents, or toxic photoinitiators to produce scaffolds. Hsieh et al., produced a cell printing ink from aqueous dispersions of thermoresponsive SPU nanoparticles synthesized from a green, water-based process (Hsu et al., 2014) that gels near 37 C. The soft segment was based on PCL diol (Mn 5 2000 Da) and PDLLA diol (Mn 5 2000 Da) mixed in a 4:1 ratio. IPDI, 2,2-bis(hydroxymethyl) propionic acid (DMPA), and EDA were employed as chain extenders. The stoichiometric ratio of IPDI/oligodiols/DMPA/ EDA was 3.52:1:1:1.52, comprising of .65% of soft segments (Hsieh et al., 2015). Water solutions contained SPU nanoparticles (B30
40 nm) with a solid content of about 30%, this being the main factor influencing the stiffness of the hydrogel. The SPU dispersions incorporated neural stem cells (NSCs) at a density of 4 3 106 cells/mL before gelation, and were then printed at 37 C using selfdeveloped FDM equipment. This equipment included a nitrogen injection system and a movable heated-platform. Fibers were printed through a 250 μm nozzle and a constant nitrogen pressure of 55 kPa. NSCs presented excellent proliferation and differentiation in the hydrogel, and constructs could rescue the function of impaired nervous system in a zebrafish embryo neural injury model. Moreover, the function of adult zebrafish with traumatic brain injury was also rescued after implantation of the 3D-printed NSC-laden constructs. Therefore, this bioprinting technique appears promising in neural tissue engineering for the treatment of severe traumatic injuries or neurodegenerative diseases, such as brain stroke, Parkinson’s disease, and Alzheimer’s disease. Another work reported by this group involved the synthesis of a waterborne thermoresponsive amphiphilic polyurethane following the same methodology (Tsai et al., 2015). The soft segment was based on PCL diol (Mn 5 2000 Da) and poly(L-lactide-co-poly(ethylene oxide)) diol as a diblock copolymer PLLA-PEG (Mn 5 3387 Da) mixed in a 90:10 ratio. The rest of the reagents were the same as employed before. Polyurethane water dispersions (30% (w/v)) presented low viscosity under room temperature, reaching gel point near 37 C. Dispersions were mixed with human mesenchymal stem cells (hMSCs) to achieve a density of 2 3 106 cells/mL. Fibers were 3D printed at 37 C through a 26 G (260 μm inner diameter) nozzle and a nitrogen pressure of 241
275 kPa, and deposited layerby-layer as a gel. Fig. 8.3 displays the thermally-induced gelling of needleinjected PCL90LE10 fibers and the layer-by-layer stacking of the deposited fibers by manual injection and by 3D printing using a self-designed FDM platform. hMSCs viability and proliferation were confirmed in the printed construct. This hydrogel is also a promising 3D-printing ink for tissue printing. Although polyurethane aqueous dispersions could undergo sol-gel transition near body temperature allowing for cell encapsulation as cell-laden bioink for printing at 37 C, the long curing time and the narrow working window present

8.4 Additive Manufacturing of Composite Polyurethanes

FIGURE 8.3 Thermally-induced gelling of needle-injected PCL90LE10 fibers and the layer-by-layer stacking of the deposited fibers: (A) 40-layer stacking fibers fabricated by manual injection; (B) construct formed; (C) 3D printing by a self-designed FDM platform; (D) 3Dprinted fibers (two layers for cell visualization by optical microscope); (E) cells in the 3Dprinted fibers (30% solid content), and (F) cells in the 3D-printed fibers (25% solid content) during a period of 7 days. Reprinted from Tsai, Y.-C., Li, S., Hu, S-G., Chang, W.-C., Jeng, U-S., Hsu, S-h., 2015. Synthesis of thermoresponsive amphiphilic polyurethane gel as a new cell printing material near body temperature. ACS Appl. Mater. Interfaces 7, 27613
27623, with permission from ACS publications.

major drawbacks to this system. To improve the printability of the ink obtained by Hsieh et al., a soy protein isolate (SPI) was added to the dispersions (Lin et al., 2016). Soy protein presents inherent biocompatibility; hydrogen bonds and hydrophobic interactions, in addition to disulfide bonds present in soy protein could assist in the formation of 3D structures. The bioink was prepared from a SPU:SPI 1.3:1 ratio dispersion at 25 C, and L929 fibroblasts (mouse skin

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fibroblast cells) or murine NSCs were directly encapsulated at a density of 2 3 106 cells/mL. Cell-laden bioinks were printed at 37 C without preheating. A commercial 3D bioprinter (Regenovo, China), a syringe with a 337 μm innerdiameter nozzle, and 50
60 kPa pressure extrusion were employed. Layer-bylayer printed 3D hydrogel scaffolds were obtained with higher cell viability than that of SPU hydrogel. Based on rheological profiles, hybrid SPU/SPI hydrogel at a 1.3/1 ratio in the culture medium was selected as bioink for cell printing. Therefore, this hybrid hydrogel presents an improvement; being able to print for longer times and larger scaffolds layer-by-layer with structural integrity.

8.4.2.1.1 Liquid-frozen deposition manufacturing Hsu’s group also explored LFDM technology to obtain SPU scaffolds (Hung et al., 2014). Biodegradable SPU were synthesized using PCL diol (Mn 5 2000 Da) and poly(ethylene butylene adipate) diol (PEBA diol, Mn 5 2000 Da) as soft segments in a 2:3 molar ratio, following the methodology described by Hsieh and coworkers (2015). Scaffolds were printed from a feed containing SPU/PEO aqueous dispersions, using PEO as a viscosity enhancer. A 10% (w/w) PEO (Mn 5 900 kDa) solution was mixed with PU dispersions to obtain SPU/PEO dispersions with mass ratios ranging from 72/28 to 80/20, the total solid content being fixed to 20% (w/w). Extrusion was carried out employing a 300 μm heated nozzle at a constant pressure of 36 kPa on a low-temperature platform (
30 C). Highly elastic scaffolds with excellent chondrocyte-seeding efficiency, proliferation, and ECM secretion were obtained. SPU/PEO scaffolds displayed better properties than pure PU scaffolds (after PEO dissolution). This 3D-printing process presents great potential for fabricating scaffolds for cartilage tissue engineering applications, its main drawback being the impossibility of printing cell-laden scaffolds. This technique was also employed to incorporate bioactive agents into elastic SPU-based scaffolds to improve the functionality of the printed scaffolds (Hung et al., 2016). Printing inks were obtained from low-viscosity water dispersions of biodegradable nanoelastomer previously synthesized (Hung et al., 2014) (Fig. 8.4A). Scaffolds were 3D printed from a self-developed LFDM system using CAM/CAD techniques (Fig. 8.4B). Inks containing SPU, hyaluronan, and bioactive agent TGFβ3 or a small molecule drug Y27632 were employed to produce scaffolds for cartilage tissue engineering (Fig. 8.4C). Hyaluronan (HA) has been reported to promote cartilage repair. It is an anionic polysaccharide and the viscoelastic component of synovial fluid. The transforming growth factor β (TGFβ) family is the most frequently used bioactive factor for inducing hMSC chondrogenic differentiation, but it is expensive and can cause hypertrophy through sustained high content (Davidson et al., 2006). Molecular inhibitors have been proven to stimulate tissue regeneration or specific lineage differentiation (Zhang et al., 2015). Y27632 [1R,4r)-4-((R)-1-aminoethyl)N-(pyridin-4-yl)cyclohexanecarboxamide] is an inhibitor for Rho-associated coiled-coil containing protein kinase (ROCK) which increases the differentiation

8.4 Additive Manufacturing of Composite Polyurethanes

FIGURE 8.4 (A) Chemical structure and synthetic procedure for SPU nanoparticles as ink base for water-based 3D printing. (B) Schematics of the LFDM platform. (C) Appearance of compliant 3D-printed scaffolds. Reprinted from Hung, K.-C., Tseng, C.-S., Dai, L.-G., Hsu, S-h., 2016. Water-based polyurethane 3D printed scaffolds with controlled release function for customized cartilage tissue engineering. Biomaterials 83, 156
168, with permission from Elsevier.

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of chondroprogenitors (Chou et al., 2013), but its effect depends on cell density and morphology. The prepared scaffolds contained 5
25 ppm of Y27632 or 5
20 ng TGFβ3/mL. The scaffolds promoted the self-aggregation of hMSCs and the release of bioactive agents, inducing the chondrogenic differentiation of hMSCs. They resulted in effective regeneration of rabbit cartilage defects after knee implantation, which supports the potential of these novel 3D-printing composite scaffolds with controlled release bioactivity in cartilage regeneration.

8.4.2.1.2 Double-nozzle low-temperature deposition manufacturing Xu et al., explored different strategies using a DLDM technique to obtain 3D SPU-based scaffolds for tissue engineering. A 3D hybrid construct was developed for implantable liver manufacturing from a bioresorbable SPU and a naturally derived polymer gelatin (Xu et al., 2008a) to improve cell attachment with respect to previous development (Xu et al., 2008b). SPU was synthesized from PCL diol (Mn 5 2000 Da) and PEG diol (Mn 5 1000 Da) as soft segments in a 1:1 molar ratio with HDI and BDO as chain extenders (Yin et al., 2007). A home-made DLDM system was employed (Department of Mechanical Engineering, Tsinghua University, China). A 15% (w/w) SPU solution in dioxane, and a 20% (w/w) gelatin solution in water (containing 5%
10% (w/w) lysine) were deposited simultaneously through separate nozzles and solidified immediately on a lowtemperature platform (
28 C). Lysine and glutaraldehyde were added to improve stability and crosslink the constructs. 3D scaffolds with interconnected macropores were obtained, while micropores were formed after freeze-drying the constructs. PU was employed for mechanical support, while gelatin was used for accommodation of implant cells. DLDM was also employed to obtain double-layer SPU/collagen nerve conduits for peripheral nerve repair (Cui et al., 2009a). The outer layer of the nerve guide was made of SPU synthesized as described previously (Yin et al., 2007), with a porosity of about 75%. The inner layer was made of type I collagen to support cell attachment and proliferation. The SPU was dissolved in dioxane and the type I collagen was dissolved in acetic acid. First, the collagen solution was extruded through a nozzle, solidifying on the platform to generate an inner circle. Then, the other nozzle extruded the PU solution to produce an outer circle on the solidified collagen solution. When this procedure was finished the platform was lowered and the process was repeated to build a layer-by-layer 3D structure. Due to the deposition of the sample onto a cooled platform (
30 C), a TIPS took place, and the sample was then freeze dried to remove the solvents and crosslinked with glutaraldehyde. The tensile strength and elongation at break increased with increasing SPU solution concentration. This was due to a decrease in the total volume and size of solvent crystals formed due to the phase separation process, leading to a decrease in pore size and total porosity. SPU solutions of 12%
13 % (w/v) were optimal in mechanical behavior, thus, preventing deformation of the conduit architecture. The tensile strength and elongation at break of the samples increased slowly with increasing collagen concentration. The overall

8.4 Additive Manufacturing of Composite Polyurethanes

mechanical properties of the double-layer conduits were reflected mainly by the elastic properties of the SPU. The compressive strength of the conduits increased exponentially with increasing SPU and collagen concentration. Thus, SPU and collagen were combined to produce scaffold structures with good biocompatibility in the inner layer and the desired mechanical strength protruded by the outer layer. The technique ensures high manufacturing precision with control of wall thickness and a tight connection between the two layers. Wang et al., employed the same methodology and crosslinked collagen using EDC/NHS to avoid the use of toxic glutaraldehyde (Wang et al., 2009) before rat implantation. The inner oriented collagen nanofilament layer of the conduit presented micropores (20
100 μm) that permitted nutrient infiltration, while the outer SPU layer had micropores (15
20 μm) that prevented fibrous tissue invasion. Although no fibrous scar tissue invasion or other adverse tissue reactions were detected in either the double-layer SPU-collagen conduit or the single-layer SPU control, the first demonstrated better nerve repair. Cui et al., cultured RSC96 Schwann cells from rat sciatic nerves in layered SPU-collagen conduits and found a significant enhancement in their retention and viability compared to those made of pure SPU (Cui et al., 2009b). The inner collagen layer promoted Schwann cell adhesion, migration, and proliferation, while collagen filaments could serve as directional guides to enable the Schwann cells to function properly. These attractive results suggest that a double-layer SPU/collagen conduit is a promising scaffold for use as a guide conduit for promoting peripheral nerve regeneration. Wang’s group also prepared cell-seeded scaffolds through a DLDM technique, using gelatin, gelatin/hyaluronan, gelatin/chitosan, gelatin/alginate, gelatin/fibrinogen, and gelatin/alginate/fibrinogen hydrogels, at temperatures ranging from 0 C to 10 C (Wang et al., 2006, 2007; Xu et al., 2007, 2009; Yao et al., 2009; Li et al., 2009). Gelatin-based hydrogels solidified with high survival rates of encapsulated cells ( . 98%). Although the hydrogels provided ECM-like environments for the native cells, the mechanical properties of the resulting constructs were too poor to withstand blood and load stresses. The main limitation in integrating synthetic polymers into these constructs is that the latter do not have phasetransformation temperatures between 0 C and 10 C, and cells become damaged by ice crystals at lower temperatures. However, the use of cryoprotectants, such as dimethyl sulfoxide (DMSO), glycerol, and dextran-40, allows for the preservation of cells in suspensions or 3D structures at extremely low temperatures (,
70 C), and for their recovery with high survival rates. He and Wang fabricated a tubular SPU sandwich-like cell/hydrogel construct by DLDM employing DMSO or glycerol as cryoprotectants (He and Wang, 2011). Briefly, 20% (w/v) gelatin solution in phosphate-buffered saline (PBS), 5% (w/v) sodium alginate solution in PBS, and 5% (w/v) fibrinogen solution in fetal bovine serum (FBS) were mixed in a ratio of 2:1:1 (v/v/v) before use; and DMSO or glycerol were added at different concentrations (0
15% (v/v)). Rat adipose-derived stem cells (ADSCs) were incorporated into the cryoprotectantcontaining gelatin/alginate/fibrinogen solutions at a density of 1 3 105 cells/mL,

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and the solution was placed in a syringe at 4 C for 30 minutes. The other solution consisted of 15% (w/v) SPU (Yin et al., 2007) in tetraglycol, a cell-compatible organic solvent. The fabrication comprised of a layer-by-layer process at
20 C, where SPU inner and outer circle layers were formed, followed by cell-hydrogel deposition between them. After a week at
80 C, the constructs were then treated with calcium chloride and thrombin solutions at 37 C to crosslink the alginate molecules and to polymerize the fibrinogen, respectively. With the use of 5% (v/v) DMSO or 10% (v/v) glycerol, 73% and 62% cell viability was retained, respectively. Therefore, glycerol and DMSO directly incorporated into the gelatin/alginate/fibrin hydrogel, reduced the risk of damage produced by ice crystal formation during the cooling/warming stages. With the formation of a mechanically strong synthetic polymer outer coat, a cell-loaded construct can be connected to an in vivo vascular system or to an in vitro pulsatile culture system for complex organ regeneration, employing different cell types and polymers depending on organ requirements. Using this unique technology, it is now possible to manufacture very complex heterogeneous structures directly from CAD models with high resolution and sophistication. Huang et al., improved the technique for producing an elliptic hybrid hierarchical SPU
cell/hydrogel construct with a fluid one-way inlet and outlet architecture as shown in Fig. 8.5A and 8.5B (Huang et al., 2013; Wang et al., 2013). ADSCs were incorporated into the cryoprotectant-containing gelatin/alginate/ fibrinogen solutions at a density of 6 3 105 cells/mL, and the solution was placed in a syringe at 4 C for 20 minutes. A layer thickness of 0.13 mm ensured good

FIGURE 8.5 Hybrid hierarchical construct made by DLDM: (A) a hybrid construct with a SPU outer coat; the inset shows a SEM micrograph of the in vitro cultured sample with porous SPU and cell/hydrogel layers; (B) the middle part of (A) with branched/grid internal cell/ hydrogel channels; the inset image shows the propidium iodide staining result of the cell/ hydrogel section; (C) two complex constructs undergoing in vitro pulsatile culture. Reprinted and adapted from Huang, Y., He, K., Wang, X., 2013. Rapid prototyping of a hybrid hierarchical polyurethane-cell/hydrogel construct for regenerative medicine. Mat. Sci. Eng. C Mater. 33, 3220
3229, with permission from Elsevier.

8.4 Additive Manufacturing of Composite Polyurethanes

integration of dense micropores between the two material systems. Cell viability was mainly determined by the construct formation time. Higher inner-nozzle diameter led to higher cell viability, but also to higher layer thickness and thus lower formation quality. Moreover, cell viability decreased with increasing extrusion flux and duration of extrusion. A 20-minute formation time, 0.253 mm nozzle diameter, and 1.3 mm3/s extrusion flux were employed. Formation quality was also greatly influenced by the SPU solution concentration and gelatin/alginate/ fibrinogen ratio. Optimal processing parameters were 15% (w/v) SPU solution, and a gelatin/alginate/fibrinogen ratio of 2:1:1, extruded at 20 C and 4 C, respectively. Constructs with cell viability values greater than 80% were obtained with these conditions. An intrinsic grid network and branched channels ensured oxygen and nutrient supply to the accommodated cells across the whole construct, acting as a tissue regeneration template. The viability and proliferation of ADSCs in the elliptic constructs were demonstrated both in vitro and in vivo. A pulsatile bioreactor was adapted with a pulse amplitude range of 0%
7%, pulse frequency between 0 and 80 times/min, and circulatory flow condition of 0
0.2 MPa pressure (Fig. 8.5C). After a long culture period (14 days), the proliferation ability of ADSCs was better promoted by the pulsatile than the static culture. Construct implantation into the abdominal cavities of mice did not produce side reactions, such as fibrous encapsulation, inflammation, or delamination. The design and manufacturing technique reported by Huang et al., in this hybrid hierarchical construct formation process will revolutionize the fields of tissue engineering and regenerative medicine and play a central role in the field of complex organ manufacturing in the future (Huang et al., 2013).

8.4.2.1.3 Integrated organ printing DIW technology has also been applied in the form of 3D IOP to fabricate tissue constructs that potentially mimic structural and functional properties of native tissues. This technology is particularly useful for complex tissues that possess regional differences in cell types and mechanical properties. Merceron et al., employed an IOP system for the fabrication of a single integrated muscle
tendon unit (MTU) constructed from four different components (Merceron et al., 2015). A thermoplastic polyurethane (Tecoflex LM-95A, Lubrizol), and PCL were used as structural materials to mimic the mechanical properties of native tissue. C2C12 myoblasts and NIH3T3 fibroblasts (ATCC, Manassas, VA, USA) were mixed with a high viscosity hydrogel bioink (composed by hyaluronic acid, gelatin, and fibrinogen in a calcium-free high-glucose Dulbecco’s modified Eagle’s medium) at a density of 4 3 107 cells/mL for the fabrication of constructs. Four cartridges were used in a 3D bioprinter: two were loaded with the polymers (200 μm diameter nozzles and a heating system), and two were loaded with cell-hydrogel bioinks (300 μm diameter nozzles). PU was coprinted with C2C12 cell-based bioink for elasticity and to create a muscle side, while PCL was coprinted with NIH3T3 cell-based bioink for stiffness and to create a tendon side. An air pressure controller was employed for precise printing. After the composite construct was printed,

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the fibrin-based hydrogel was allowed to crosslink in a thrombin solution for 30 minutes. The final MTU was elastic on the muscle side, stiff on the tendon side, and intermediate in the interface region. Cell viability remained higher than 80% 7 days after printing, and the construct showed initial tissue development and differentiation. This study demonstrated the versatility of the IOP system to create integrated tissue constructs with region-specific biological and mechanical characteristics, such as those in the musculoskeletal system.

8.4.2.2 Fused deposition modeling Bioresorbable SPU commonly incorporate polyethers or polyesters as soft segments. In the case of polyester-based SPUs, moisture must be avoided before thermal processing due to the fact that these materials are susceptible to pyrohydrolysis. Although this is a problem for FDM processing, materials can be processed in a nitrogen atmosphere after drying. Moreover, polyether-based SPUs present higher hydrolytic resistance, but lower oxidative and thermal stability (Rychly´ et al., 2011). Studies on biomedical polyurethanes processed by FDM were carried out (Moore, 2005; Miller et al., 2017). Moore synthesized a PCL-based SPU containing 65% hard segments and extruded 1.7 mm diameter filaments to feed a FDM machine (Moore, 2005). A complex precise 3D-printed tissue-engineered scaffold was obtained by this technique. The creation of these scaffolds demonstrates their suitability for thermal processing. In vitro cell growth with fibroblasts on the printed scaffolds showed that the scaffolds were still biocompatible after extrusion and FDM printing. Cells secreted extracellular matrix, which is extremely important for structural integrity as the scaffold degrades. Chen et al., reported the fabrication of a tissue-engineered heart valve scaffold using a combination of FDM printing and electrospinning techniques (Chen et al., 2009). Scaffolds were prepared following a two-step procedure. First, a trileaflet scaffold was obtained through electrospinning of Tecoflex (Thermoplastic polyurethane); and second a heart valve ring was designed and constructed by FDM. The fabrication of a 3D-heart valve ring was constructed using Pro Engineer based on optimum hemodynamic analysis and converted to an STL file format. This method was proven as a promising fabrication process for fabricating a synthetic graft with biomechanical properties. One of the most important cons of FDM is the lack of flexibility of the obtained structures. Therefore, the use of FDM with hemodynamic analysis to design the heart valve ring and electrospun polyurethanes to fabricate the leaflets led to the combination of two advanced technologies for producing novel tissue-engineered heart valves. This combination of techniques and polyurethane materials was also used to fabricate coronary artery bypass grafts (Owida et al., 2011). The proposed manufacturing technique combines the advantages of both methods to obtain graft vessels with mechanical properties close to the human artery. Ovine endothelial cells were seeded and cultured. Endothelial cells adhered onto SPU surface, exhibiting a typical cobblestone morphology and demonstrating good cytocompatibility.

8.5 Remarks and Perspectives

A new line of composite filaments has been developed by inventor Kai Parthy (Poro-Lay, Cologne, Germany) for printing porous 3D objects. Among them, a polyurethane-based composite filament containing 30% (w/w) PVA, called GelLay, was employed to print models obtained using a Solidoodle table-top FDM printer (Solidoodle, Brooklyn, New York) (Castro et al., 2016). After dissolving PVA in warm water and ultrasonication, porous polyurethane scaffolds exhibited microfilamentous surface topography which may be suitable for applications requiring directionality of cell adhesion as in the case of neural tissue regeneration. Scaffolds were subjected to nucleation by immersion in SBF solution, and the microfilamentous structure was lost due to hydroxyapatite formation. The resulting nucleated scaffolds displayed excellent human fetal osteoblast (hFOB, CRL-11372) attachment, proliferation, and osteogenesis, in the first approach to test their potential for bone tissue regeneration (Castro et al., 2016). This work has great significance in 3D-printed implantable bioactive polymeric scaffolds where a lack of 3D printable cytocompatible material exists for use in tissue engineering. However, the manufacturer does not provide information on polyurethane composition, stating only that it is suitable for body parts. So, biostability and degradation products should be studied to assess its potential for tissue engineering.

8.4.3 PARTICLE BINDING The particle binding technique has been explored by Pfister et al., who produced a composite polyurethane scaffold from lysine ethyl ester diisocyanate (LDI) and a commercially available ZP11 powder, composed of a mixture of starch, short cellulose fibers, and dextrose as a binder (Pfister et al., 2004). First, an aqueous ink containing 2-propanol and glycerol was printed over lay-down patterns of ZP11 powder with a commercial Z402 3D printer (Z Corp.). The activated dextrose led to pattern formation, but the resulting objects were highly water soluble. To improve their mechanical properties and make them suitable for cell seeding, scaffolds were then infiltrated and reacted with a toluene solution containing LDI and DBTDL as a catalyst to form network structures and then cured at 60 C. Brittle and fragile porous cubes were obtained with a rough surface attributed to particle bonding. Scaffolds became rather flexible after swelling in water and Youngs modulus decreased significantly, which indicates a low crosslinking density restricted to starch interparticle spacing. Experiments with an osteoblast-like cell line proved that the scaffold is suitable for cell seeding.

8.5 REMARKS AND PERSPECTIVES Advances in biomaterials science, technology, imaging, computing, and other fields related to the biomedical area are beginning to transform life science. AM

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is a clear example of an integration of disciplinary approaches that will create new pathways and opportunities for scientific and technological advancements. AM technologies will enable the production of tissue-engineered constructs and biomedical devices, thus, bypassing poorly controlled manual cell-seeding procedures. Both biomaterials science and technological developments applied to AM are still in their infancy. Key issues in biomaterials research (matching degradation time to tissue growth and providing biomimetic mechanical response while achieving the rheological properties needed for processing), construct design (including vascularization), and system integration (multiple cells, materials, manufacturing, and combined manufacturing processes in sterile conditions), need to be urgently addressed. In the past decade, biodegradable SPU has become a useful and versatile biomaterial. Polyurethanes can be designed to have a specific biodegradation rate and can be 3D-printed at high temperatures from melts or from low-temperature solutions or dispersions to form complex structures. In particular, water-based SPU dispersions can be 3D-printed into highly compliant scaffolds. Biodegradable waterborne SPU have great potential for development into novel 3D-printing ink for customized tissue/organ constructs.

ACKNOWLEDGMENT This work is supported by the Argentinean National Agency of Scientific and Technological Promotion (Grant PICT2012-224), National Research Council CONICET (Grant PIP2013-089), and National University of Mar del Plata (Grant 15/G420). N.J. Lores thanks CONICET for the fellowship awarded.

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171. Bajaj, P., Schweller, R.M., Khademhosseini, A., West, J.L., Bashir, R., 2014. 3D Biofabrication Strategies for Tissue Engineering and Regenerative Medicine. Annu. Rev. Biomed. Eng. 16, 247
276. Bartolo, P., Kruth, J.-P., Silva, J., Levy, G., Malshe, A., Rajurkar, K., et al., 2012. Biomedical production of implants by additive electro-chemical and physical processes. CIRP Ann. Manuf. Techn. 61, 635
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9. Cui, T., Wang, X., Yan, Y., Zhang, R., 2009b. Rapid prototyping a double-layer polyurethane
collagen conduit and its Schwann cell compatibility. J. Bioact. Compat. Polym. 24, 5
17. Davidson, E.N.B., Vitters, E.L., van der Kraan, P.M., van den Berg, W.B., 2006. Expression of transforming growth factor-β (TGFβ) and the TGFβ signalling molecule SMAD-2P in spontaneous and instability-induced osteoarthritis: role in cartilage degradation, chondrogenesis and osteophyte formation. Ann. Rheum. Dis. 65, 1414
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6705. Ma, P.X., Elisseeff, J., 2006. Scaffolding in Tissue Engineering. CRC Press Taylor and Francis Group, Boca Raton, FL. Marcos-Ferna´ndez, A., Abraham, G.A., Valentı´n, J.L., San Roma´n, J., 2006. Synthesis and characterization of biodegradable non-toxic poly(ester- urethane-urea)s based on poly (ε-caprolactone) and amino acid derivatives. Polymer (Guildf) 47, 785
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CHAPTER

Composites based on bioderived polymers: potential role in tissue engineering: Vol VI: resorbable polymer fibers

9

Monika Yadav1, Kunwar Paritosh1, Nidhi Pareek2 and Vivekanand Vivekanand1 1

Centre for Energy and Environment, Malaviya National Institute of Technology, Jaipur, India 2 Department of Microbiology, Central University of Rajasthan, Ajmer, India

9.1 INTRODUCTION The prompt increase in the world’s population has led to the rapid consumption of organic chemicals and energy which is currently derived from fossil fuels. The constant dwindling of fossil fuels, the rise of petrochemical prices, and the increased deterioration of the environment have become motives for exploring new and renewable resources for the sustainable production of energy and organic chemicals (Chen and Patel, 2011). These developments have also prompted research into the production of synthetic polymers from bio-based monomers. A range of monomers and polymers are offered by naturally abundant biomass resources, such as polyesters (polylactides, polyglycolides, polyhydroxyalkanoates, polycaprolactone), furan derivatives, sugars, and lignin. The utilization of biopolymers as biomaterials has greatly influenced advancements in the biomedical field. The current market for implantation surgeries, cell culturing, tissue repair, and regeneration is around US$23 billion, and it is estimated to reach US$94.2 billion by 2025 (Manavitehrani et al., 2015). In order to meet the requirements of the biomedical community, the focus of research in the field of biomedical engineering has shifted toward biodegradable polymers. To be appropriate for biomedical applications and to get an appropriate host response, a biomaterial must fulfill a number of criterions from being biocompatible to being flexible in chemistry along with possessing good physical and mechanical properties. The fabrication of biodegradable and bioabsorbable biomaterials to aid in the recovery of damaged tissues is one of the greatest challenges in biomedical

Materials for Biomedical Engineering: Hydrogels and Polymer-based Scaffolds. DOI: https://doi.org/10.1016/B978-0-12-816901-8.00009-2 © 2019 Elsevier Inc. All rights reserved.

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engineering. Natural polymers, such as collagen, have been used for biomedical applications for centuries, however, the application of synthetic biopolymers is a relatively new area of research. Biopolymers offer a suitable alternative to traditional implanting materials, such as ceramics and metals. Synthetic biodegradable polymers for instance poly(lactic acid) (PLA), poly(glycolic acid) (PGA), poly (caprolactone) (PCL), and their copolymers are now frequently used in biomedical devices. These polymers can be degraded simply by the hydrolyzing of ester bonds without causing any inflammatory response in the host system. Furthermore, the degradation products of this process can be eliminated by the normal cellular activities of detoxification. After the discovery that lactic acid and glycolic acid are capable of being employed as degradable matrices for the delivery of bioactive substances, biopolymers have been investigated as a suitable bioresorbable polymer for fixation devices (Auras et al., 2004). In medical applications, synthetic degradable polymers have been used for about 30 years as sutures and bone fixing devices. These bioabsorbable fixation devices are extensively used by orthopedic surgeons as resorbable plates, screws, and dissolvable suture meshes, since they provide highly advantageous properties compared to titanium plates or other metallic implants. Being bioabsorbable, these devices do not erode bones when they are placed in the human body. Furthermore, their biodegradable properties forego the requirement of undergoing a second surgery for removing implants by allowing for the gradual recovery of tissue. After being resorbed, bioabsorbable plates do not cause interference with CT scans during subsequent medical evaluations, which adds to the desirability of these devices (Lasprilla et al., 2012). The application of biopolymers as scaffolds for cell transplantation has been investigated owing to their ability for stimulating cell regeneration and releasing drugs, such as antiinflammatories and antibiotics (Chen et al., 2006; Dai et al., 2010; Kulkarni et al., 2010). This chapter covers the synthesis, modifications, properties, and applications of the most commonly used biopolymer composites in biomedical fields, such as PLA, PLGA, collagen, silk fibroin, biocellulose, etc. (Fig. 9.1).

9.2 POLYESTERS Synthetic biodegradable polyesters are deliberated as the most commercially competitive polymers for biomedical applications since they can be fabricated from renewable resources in a cost-effective way. Their biocompatibility, biodegradability, as well as other physiochemical properties suitable for medical applications, make them attractive for use in the development of biomedical devices, such as sutures, plates, implants, bone fixing devices, stents, screws, and tissue repairs (Ratner et al., 2013; Sin et al., 2013; Diaz et al., 2014). Commercially, polyesters are also used as drug-releasing vehicles (Makadia and Siegel, 2011; Nazemi

9.2 Polyesters

FIGURE 9.1 Classification of biodegradable polymers.

et al., 2015). Furthermore, polyesters can be modified to tackle issues such as low cell adhesion, compatibility, hydrophobicity, and inflammatory responses in cellular and regenerative applications (Seyednejad et al., 2011). Polyesters are synthesized principally by polymerization, ring opening polymerization, and block copolymerization methods. Poly(propylene carbonate) (PPC) is commercially synthesized from the ring opening reaction between carbon dioxide and propylene oxide under the action of an active catalyst such as zinc glutarate (Zhong and Dehghani, 2010). The synthesis of PLA involves a multiple-step procedure that starts with the biosynthesis of lactic acid, which is further converted to its cyclic lactide form and then polymerized in the presence of a metal catalyst (Masutani and Kimura, 2015; Lasprilla et al., 2012). Polyβ-hydroxybutyric acid (PHB) is a carbon assimilation product which can be biosynthesized by employing diazotrophic bacteria of genera Acetobacter and Rhizobium. The polycondensation of two molecules of acetyl-CoA in the presence of 3-ketothiolase leads to the formation of acetoacetyl-CoA which can further be reduced to hydroxybutyric-CoA by acetoacetyl-CoA reductase. Finally, hydroxybutyric-CoA is polymerized to PHB in the presence of PHB synthase, liberating coenzyme-A in the process (Verlinden et al., 2007). Table 9.1 summarizes the properties of some commercially important polyesters and their applications in the medical field.

9.2.1 POLY(LACTIC ACID) PLA is a polymer that consists of lactic acid (2-hydroxy propionic acid) building blocks, a naturally occurring organic acid. Commercially, lactic acid is produced

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Table 9.1 Polyesters: Properties and Medical Applications (Farah et al., 2016) S. No.

Type of Polyester

Tensile Modulus (MPa)

In Vivo Degradation Rate

1

PLA

350
3500

98%
100% in 12 months

2

PGA

6000
7000

3

PLGA

2000

100% in 6
12 months 100% in 2
4 months

4

PHB

3500
4000

35% after 6 months

5

PPC

830

6

PCL

210
440

6% in 7 months 50% in 4 years

Applications Fracture fixation, meniscus repair, suture, anchors, vascular grafts, screws, adhesion barriers, bone graft substitutes, articular cartilage repair, bone augmentation, scaffolds Suture anchors, medical devices, meniscus repair, drug delivery Composition (85L:15G); screws, stents, plates, suture anchorsComposition (50L:50G); drug delivery, sutures, artificial cartilage repairComposition (90L:10G); artificial skin grafting, sutures, wound healing devices, tissue engineered vascular grafts Screws, sutures, scaffolds for tendon regeneration, surgical meshes, nerve repair, vascular grafts, bone tissue scaffolds Scaffolds Dental orthopedic implants, ear implants, suture coating, tissue repair, adhesion barrier, surgical meshes, stents, cardiac patches, tissue engineered scaffolds

by the bacterial fermentation of carbohydrates using various strains of genus Lactobacillus that produce lactic acid from hexose sugar. The type of carbohydrate that can be utilized depends upon the type of strain of Lactobacilli. PLA is one of the most promising biopolymers due to the fact that its monomer can be produced from nontoxic renewable feedstock. It is a thermoplastic comprising of high biocompatibility, biodegradability, and a wide profile of mechanical properties. Since lactic acid exists in the form of two enantiomers (L -lactic acid and D-lactic acid), polylactic acid usually refers to a family of polymers comprising of enantiomers: poly( L-lactic acid) (PLLA), poly( D-lactic acid) (PDLA), and poly(D, L-lactic acid) (PDLLA). After the reporting of the pseudoorthorhombic crystal structure of PLLA by Santis and Kovacs in 1968 (So¨dergard and Stolt, 2002), lactic acid based polymers became the very first fiber materials to be employed for resorbable sutures at a commercial scale. Since then a number of PLA based prosthetic devices has been synthesized.

9.2 Polyesters

9.2.1.1 Poly(lactic acid) fabrication A number of polymerization processes can be used for the formation of PLA from lactic acid, such as polycondensation, ring opening polymerization, azeotropic dehydration, and enzymatic polymerization (Garlotta, 2001) (Fig. 9.2). Polycondensation is the least expensive method of PLA synthesis and can be done by either direct polycondensation or melt polycondensation (Auras et al., 2004), which results in the production of oligomers with average molecular weight ranging from ten to thousands. HO 2 ðCHðCH3 ÞCOOÞn 2 H 1 HO 2 ðCHðCH3 ÞCOOÞm -HO 2 ðCHðCH3 ÞCOOÞn1m 2 H 1 H2 O

The process of direct polycondensation is carried out in three stages: the elimination of free water, oligomer polycondensation, followed by melt condensation of high molecular weight PLA. The major disadvantages associated with this method are: the generation of low molecular weight PLA and the manifestation of side reactions, for instance, transesterification that leads to the formation of ring structures (Auras et al., 2010). Attempts to achieve PLA of high molecular weight were made by sequential solid or melt polycondensation. In melt polycondensation, an additional fourth stage is employed apart from the three previously mentioned steps. This fourth stage, comprises of cooling down the melt polycondensed PLA to below its melting temperature (120 C
180 C depending on the presence of L- and D-lactide), followed by particle formation that later is subjected to the crystalline process.

FIGURE 9.2 PLA synthesis processes (Lasprilla et al., 2012).

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For achieving a high molecular weight product, the obtained low molecular weight PLA can also be subjected to a chain extension process in which the low molecular weight polymer is treated with chain extenders. Chain extenders are adequately competent to reconnect cleaved chains, thereby increasing the molecular weight of a polymer. In general, chain extender agents can possess bi or more functional groups, such as diisocyanate (Gu et al., 2008; Pesetskii et al., 2009), dianhydride (Bimestre and Saron, 2012), diamine (Li et al., 2006), epoxy, etc., to seam the two end groups of low molecular weight polymer chains (Khankrua et al., 2014). Chain extension is most commonly performed using diisocyanates. Gu et al. (2008), used 1,6-hexamethylene diisocyanate as a chain extender and obtained PLA with a molecular weight of 27,500 g/mol after 40 minutes of chain extension at 180 C. In a ring opening polymerization process, the ring of a lactic acid cyclic dimer is opened in the presence of a catalyst (Auras et al., 2010). This most commonly used method of PLA fabrication includes three steps: polycondensation, lactide synthesis, and ring opening polymerization. The presence of a catalyst results in PLA of controlled molecular weight. The sequence and ratio of D- and L-lactic acid units can be regulated by controlling the process parameters, such as residence time, type of catalyst, and temperature (Gupta et al., 2007). The process can be carried out in melt, bulk, and in solution. Azeotropic dehydration is a direct method for obtaining high molecular weight PLA without requiring any chain extender (Auras et al., 2010). In azeotropic dehydration, the principle stages are the same as in direct polycondensation. The difference lies in the abolition of the last stage melt polycondensation as the polycondensation is performed in a solution that facilitates the elimination of reaction water. This technique provides high molecular weight PLA, but with a considerable amount of catalytic impurities that further create toxicity and slow release properties in medical applications. These catalysts can be deactivated using phosphoric acid and pyrophosphoric acid. The catalyst can also be precipitated using strong acids.

9.2.1.2 Poly(lactic acid) processing Several processing techniques have been developed in order to convert PLA into molded parts, foams, fibers, films, etc. These processing technologies include drying and extrusion, injection molding, injection stretch blow molding, extrusion blown film, casting, thermoforming, foaming, fiber spinning, electrospinning, blending, compounding, and nanocompositing (Farah et al., 2016).

Drying and extrusion Before being subjected to melt processing, the polymer must be dried enough to prevent excessive hydrolysis, which can result in molecular weight loss thus causing changes in physical properties. Typically, the polymer is dried to 100 ppm moisture content. Processes having long residence times or high temperatures approaching 240 C require prior drying of resins to below 50 ppm in order to

9.2 Polyesters

minimize weight loss. Drying of PLA takes place in the temperature range of 80 C
100 C. The required drying time varies according to the drying temperature. Crystallized PLA resin pellets allow for drying at higher temperatures, while amorphous resin pellets must be dried at a temperature below Tg (B60 C) to prevent resin pellets from sticking to each other (Lim et al., 2008).

Injection molding Injection molding is the most widely used conversion mechanism for complex thermoplastic polymers that require high dimensional precision. Injection molding machines are equipped with an extruder for plasticizing the polymer melt. The design of injection molding machines for PLA are based on a reciprocating screw extruder that provides enough injection pressure to deliver polymer melt into mold cavities (Lim et al., 2008).

Stretch blow molding PLA bottle production is based on an injection stretch blow molding mechanism. The process produces PLA bottles with improved physical and barrier properties in comparison to the injection molding method. The process involves the formation of a preform using an injection molding machine. The preform is then transferred to a blow molding machine where it is stretched and blown to achieve the biaxial orientation of the polymer (Lim et al., 2008).

Cast film and sheet The fabrication process of PLA films and PLA sheets are basically indistinguishable. The main difference between them is in their stiffness and flexibility which pertains to differences in their thickness. Characteristically, films are # 0.076 mm in thickness, while sheets usually have a thickness of $ 0.25 mm. To fabricate PLA film, molten PLA is first extruded through a sheet die followed by quenching on polished chrome rollers that are arranged in such a way so as to be cooled down with circulating water (Lim et al., 2008).

Thermoforming In this process, a PLA sheet is heated to soften the polymer. The softened polymer is then forced into a mold, allowed to cool, removed from the mold, and then trimmed. This method is commonly used for the production of packing containers. The temperature required for thermoforming of PLA is lower than other conventional thermoformed plastics and ranges from 80 C to 110 C (Lim et al., 2008).

Foaming PLA foams have a special niche in biomedical applications owing to their biocompatibility and large surface area. Foaming is carried out by dissolving a blowing agent that forms bubbles in the PLA matrix. The nucleation of the bubbles formed in the PLA matrix is induced by increasing temperature or by

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CHAPTER 9 Composites based on bioderived polymers

decreasing the pressure. Fujiwara et al., (2006), studied PLA foaming by supercritical CO2 where the sample was heated to 50 C followed by the addition of CO2.

9.2.1.3 Poly(lactic acid) properties The properties of PLA depend on its constituent isomers, conditions during fabrication and processing, and its molecular weight. The stereochemistry of a polymer and its thermal history have a direct impact on PLA crystallinity, which is an indication of the quantity of crystalline region of a polymer with respect to amorphous region. Crystallinity further influences the physical and mechanical properties of the polymer, such as modulus, stiffness, hardness, tensile strength, melting point, etc. PLA in which PLLA content is higher than 90%, tend to be crystalline, while PLA with a lower PLLA content tend to be amorphous in nature. The glass transition temperature (Tg) and the melting temperature (Tm) of PLA decrease with decreasing PLLA percentage.

Physical proprties Physical characteristics, for instance heat capacity, density, and mechanical and rheological properties of PLA are dependent on its Tg. For amorphous PLA, its Tg is one of the most imperative parameters since dramatic changes in PLA behavior, pertaining to polymer chain mobility, take place at and above Tg. For semicrystalline PLA, both Tg and Tm are important parameters for predicting the physical properties of PLA (Auras et al., 2004; Bouapao et al., 2009; Yamane and Sasai, 2003). The density of crystalline and amorphous PLLA has been reported as 1.290 and 1.248 g/mL respectively. The density of solid polylactide was reported as 1.33 g/cm3 for meso-lactide, 1.36 g/cm3 for L-lactide, 1.25 g/cm3 for amorphous polylactide, and 1.36 g/cm3 for crystalline polylactide (Auras et al., 2004). PLA products are soluble in acetonitrile, dioxane, chloroform, 1,1,2-trichloroethane, methylene chloride, and dichloroacetic acid. Toluene, ethyl benzene, acetone, and tetrahydrofuran only partly dissolve polylactides. However, solubility in these solvents increases upon heating. Crystalline PLLA is insoluble in ethyl acetate, acetone, and tetrahydrofuran (Nampoothiri et al., 2010).

Thermal properties Semicrystalline PLA exhibits Tg and Tm like that of other thermoplastic polymers. Above Tg, which is around 58 C, PLA shows rubbery characteristics, while below Tg it behaves like glass, and below 45 C it behaves like a brittle polymer. The Tg of PLA depends both on the molecular weight and the optical purity of the polymer. PLA Tg increases with molecular weight since polymers with high molecular weight tend to be denser, having less free volume which restricts chain segment mobility. Furthermore, PLA having higher L-lactide contents are reported to have higher Tg values than the same polymer with the same amount of D-lactide (Dorgan et al., 2005).

9.2 Polyesters

Mechanical properties The mechanical properties and crystallization behavior of PLA depends on the molecular weight and stereochemistry of the polymer backbone (Garlotta, 2001). The tensile modulus of semicrystalline PLA is approximated to be around 3 GPa, while its tensile strength ranges from 50 to 70 MPa, its flexural modulus is estimated to be around 5 GPa, and flexural strength to be around 100 MPa (So¨dergard and Stolt, 2002). The tensile modulus of PLLA doubles when the molecular weight is raised from 50 to 100 kDa. Similarly, tensile strengths of 15.5, 80, and 150 MPa were reported for varying molecular weights of 50, 150, and 200 kDa respectively (Velde and Kiekens, 2002).

9.2.1.4 Poly(lactic acid) medical applications PLA possesses unique features, such as biocompatibility, biodegradability, and eco-friendliness, that enable its wide application in commodity plastics for packaging, surgical implant materials, and drug delivery systems. It has been investigated for use in porous scaffolds for the growth of neotissues (Gupta et al., 2007; Yamane and Sasai, 2003). Various devices have been prepared from PLA, such as biodegradable sutures, porous scaffolds for cellular applications, and drug releasing microparticles for drug delivery applications. Physical, chemical, plasma, and radiation induction methods can be used for surface modifications of PLA polymers. PLA polymer can be blended with other polymer and nonpolymeric components to achieve desirable properties (Cheng et al., 2009a,b; Saini et al., 2016). A number of patents have been filed in the past years by biomedical companies and academic institutions regarding biomedical applications of PLA blends. This section reviews the potential applications of PLA in the biomedical industry:

Wound healing and stents PLA offers various applications regarding wound healing, such as surgical suture, healing dental extraction wounds, and preventing postoperative adhesions. As a fiber, PLA is not preferred for sutures manufacturing, due to its slow degradation. Rate of degradation is dependent on the magnitude of applied stress. Another complication is the premature failure of PLA sutures in in vivo environments. However, PLLA fibers are favored for applications that require a long retention of strength, such as the reconstruction of tendons and ligaments as well as stents for urological and vascular surgery (Durselen et al., 2001).

Scaffolds for tissue engineering PLA matrix materials have garnered immense interest as support material in tissue engineering as they disappear from the transplantation site on their own with the passage of time, leaving behind a perfect patch of natural neotissue. For tissue engineering applications, the mechanical properties of PLA can be improved using a number of methods, such as blending, composites forming

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(Zhang et al., 2011), and copolymerization (Hamad et al., 2015). Blends of PLA and PEG were employed to fabricate tissue engineering scaffolds that show faster degradation rates in comparison to neat PLA scaffolds (Zhu et al., 2014) 3D porous scaffolds have been generated for cell culturing as required in cell-based gene therapy used for the regeneration of cardiovascular tissues, cartilage, and bone (Coutu et al., 2009; Papenburg et al., 2009). Stem cells planted on these scaffolds were found to be capable of recapitulating the developmental processes of bone formation after implantation in sites of bone defects (Caplan, 2009). The ability of lactic acid polymers to stimulate cell regeneration and the release of antibiotics and antiinflammatories has motivated the use of PLA scaffolds in cell transplantation (Kulkarni et al., 2010; Dai et al., 2010). The high strength of PLLA makes it suitable to be used in three-dimensional trays and cages. The degradation time of these PLA scaffolds depends on microstructural components, such as chemical composition, the presence of hydrolytically unstable bonds, porosity, molecular weight, and crystallinity. Other factors include the presence of impurities, additives, and/or catalysts; the geometry of the device; the site of transplantation, etc. Balancing all these factors during the tailoring of scaffolds is the key to biomedical devices.

Orthopedic implants and fixation devices Biodegradable polymers have been seen as replacements for metallic implants made up of steel or titanium used as bone fixation devices in form of plates, pins, screws, wires, etc. These implants have an advantage over metallic implants due to the fact that they do not require a second surgery for implant removal (Hamad et al., 2015). PLLA also finds application as injectable microspheres for fillings during facial reconstructive surgeries. These microspheres have also been employed as embolic materials for transcatheter arterial embolization which is used to manage arteriovenous malformations, tumors, and massive hemorrhages (Imola and Schramm, 2009). A patent has been filed regarding bone plates made by blending PLA and Ecoflex as well as a hot-melt adhesive polymer blend of the same material (McCarthy and Weinzweig, 2014).

Drug delivery PLA-based drug delivery systems are used for prolonged administration of various medical agents, such as local anesthetics, narcotic antagonists, contraceptives, vaccines, etc. In the case of PLA, hydrolytic ester cleavage results in the breaking of ester bonds which leads to the erosion of the device followed by drug release. PLA and its copolymers are used in the form of micro and nanoparticles produced using solvent evaporation techniques in the encapsulation of drugs, such as psychotic (Leroux et al., 1996), restenosis (Fishbein et al., 2000), hormones (Matsumoto et al., 1999), anticancer drugs (Wang et al., 2015) oridonin (Xing et al., 2007), dermatotherapy (Rancan et al., 2009), and proteins (BSA) (Gao et al., 2005). PLA blended with other natural and synthetic polymer materials have also been explored for matrix film- and nanofiber-based drug delivery

9.2 Polyesters

systems. The drug release profile of a PLA can be designed by tuning the blend composition which consequently alters not only the degradation rate but also the pattern of degradation (Saini et al., 2016). PLA/PEG-based porous matrix films were reported in wound healing applications for the delivery of gentamicin sulfate, which increased the rate of oxygen transmission and drug release in wounds due to their amorphous nature which was attributed to enhanced porosity and pore size (Chitrattha and Phaechamud, 2016). Similarly electrospun fibers prepared from PLA/poly(butylene adipate) (PBA) blends were investigated for the delivery of an antirheumatoid agent; Teriflunomide (TF) (Siafaka et al., 2016). PLA/PCL blends were investigated for the controlled release of amoxicillin, and these showed an intermediate response regarding degradation in comparison to rapidly degrading PCL and slowly degrading pure PLA materials (Valarezo et al., 2015). A patent has been filed for the formulation of polymer blended with an active compound for time controlled drug delivery (Kohn and Schachter, 2014) and intraocular delivery of drugs from polymeric microparticles blended with active agents useful for treating eye disorders (Lavik et al., 2013).

3D printing PLA 3D printing has been found feasible for improving the performance of implants and scaffolds for diagnostic, tissue engineering, and drug delivery systems (Chia and Wu, 2015). PLA printing is performed using direct or indirect 3D printing and fuse deposition technologies. PLA 3D printing can be used for the regeneration of complex tissues or for organ replacement, such as bone cartilage, vessels, livers, lymphoid organs, etc. Combining stem cells with PLA 3D scaffolds for regeneration is another anticipated goal in tissue engineering.

9.2.2 POLY(LACTIC-CO-GLYCOLIC ACID) (PLGA) COPOLYMERS PGA is a hydrophilic and highly crystalline polymer which is predominantly used for the manufacturing of sutures and carriers for drug delivery due to its rapidly degrading nature (Nair and Laurencin, 2007). It has been employed commercially as sutures since the 1970s (Ulery et al., 2011). However, it is not desirable for the fabrication of medical meshes owing to its brittle structure. Therefore, PLGA copolymer is preferred as compared to its constituent homopolymers for the manufacturing of bone substitute constructs (Gentile et al., 2014). PLGA is also preferred for sutures and drug delivery applications as it provides controlled degradation rates in comparison to PGA and PLA. Table 9.2 summarizes the comparison of various physical properties of PLA, PGA, and their copolymers.

9.2.2.1 Synthesis of PLGA PLGA is a linear copolymer which can be produced in different forms, depending on the ratio of its constituent monomers: lactide and glycolide. The process parameters during synthesis mechanisms strongly affect the physiochemical properties of the obtained PLGA product. The polycondensation of lactic acid and glycolic

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Table 9.2 Comparison of Properties of PLA, PGA, and Their Copolymers (Farah et al., 2016) S. No.

Properties

PLA

PGA

PLA/PGA (50:50)

PLA/PGA (75:25)

1 2

Melting point (Tm) ( C) Glass transition temperature (Tg) ( C) Tensile strength (MPa) Tensile modulus (GPa) Polymer density (g/cm3) Ultimate strain (%)

150
162 45
60

220
233 35
40


40
50


50
55

21
60 0.35
3.5 1.21
1.25 2.5
6

60
99.7 6
7 1.50
1.71 1.5
20

41.4
55.2 1
4.34 1.30
1.40 2
10

41.4
55.2 0.21
0.44 1.3 2.5
10

3 4 5 6

acid at temperatures above 120 C allows for the synthesis of PLGA having low molecular weight less than 10 kDa under water-removal conditions (Wang et al., 2006). PLGA with high molecular weight could be attained by employing a ringopening polymerization method that polymerizes lactide and glycolide in the presence of a metal catalyst, such as tin (II) 2-ethylhexanoate, tin (II) alkoxides, or aluminum isopropoxide, at high temperatures ranging from 130 C to 220 C (Kowalski et al., 2000). The enzymatic polymerization mechanism provides an alternative technique for obtaining PLGA uncontaminated with toxic metallic residues. The method involves enzymatic ring-opening (i.e., by employing lipase) under mild reaction conditions (temperature, pH, and pressure), which results in the production of PLGA with a low molecular weight (Duval et al., 2014).

9.2.2.2 Properties of PLGA PLGA can be dissolved with the use of a wide-range of solvents, such as chlorinated solvents, acetone, tetrahydrofuran, and ethyl acetate, unlike monomer constituents that show poor solubilities (Makadia and Siegel, 2011). It can be processed into a desirable shape and size and can be used to encapsulate biomolecules of any size. The physical properties of PLGA depends on different factors, including the initial molecular weight of the monomer units, the ratio of lactide and glycolide, the water exposure time, and the storage temperature (Houchin and Topp, 2009). The Tg of PLGA is reported to be above 37 C. Therefore, it shows glassy behavior by nature. Decreasing the lactide constituent and molecular weight of PLGA results in a decrease in its glass transition temperature (Park and Jonnalagadda, 2006). PLGA degrades due to hydrolysis of its ester linkages in aqueous environments. Water penetrates inside the amorphous region of the polymer and disrupts the hydrogen bonds and van der Waals forces, which results in a decrease in Tg. Furthermore, carboxylic end groups autocatalyze the hydrolysis process through massive disruption of backbone covalent bonds, which results in loss of integrity

9.3 Collagen

(Engineer et al., 2011). Degradation rate depends on different parameters, such as molecular weight, stereochemistry, and ratio of lactic acid (LA) and glycolic acid (GA). PLGA with a higher content of lactide is less hydrophilic, showing less water absorption thereby slowing down the degradation process, while a higher content of glycolide shows a faster degradation rate (Lu et al., 2000; Wu and Wang, 2001).

9.2.2.3 Medical Applications of PLGA PLGA is among the most widely used polyesters for fabricating sutures, drug delivery systems, and implants for bone tissue engineering. It is used in various forms, such as porous and fibrous scaffolds, hydrogels, microspheres, etc. For tissue engineering applications natural polymers, such as polysaccharides and proteins, are added to PLGA to improve cell interaction behavior. Dai et al. (2010) modified PLGA mesh with collagen type I to fabricate a biomaterial for use as supporting scaffolds for cartilage and bone regeneration applications. Nojehdehian et al. (2009) attempted surface modification of PLGA with poly-Llysine by employing a water/oil emulsion and solvent evaporation technique. The modification enhanced cell differentiation, however, mechanical properties were observed to be compromised. Huang et al. (2010) blended PLGA with PHBV for enhanced mechanical properties. The product showed higher compressive strength compared to pure PLGA scaffolds. For drug delivery PLA and PLGA are the most widely used polymers that serve in the treatment of various disorders, such as wound healing, cancer, diabetes, alcohol dependency, etc., in form of hydrogels, films, fibers, nanosystems, and particle based delivery systems (Saini et al., 2016).

9.3 COLLAGEN Collagen is a natural polymer which is found in skin, bone, ligaments, tendons, etc. It is the predominant protein found in bone, which is made up of long collagen fibrils. Repetitive units ([Gly-X-Y-]n, where X and Y are repeating proline and hydroxyproline residues) form cylindrical tropocollagen which helps in the formation of complex structural tissues such as intrafibrillar crystals (Olszta et al., 2007). These collagen fibrils provide sites for hydroxyapatite crystals to precipitate onto its surface to form mineralized structural units of bone. While calcium phosphates (CaP) are accountable for the compressive strength of bone, the resilient nature of bone is attributed to collagen which makes it tough and elastic (Barre`re et al., 2006). For biomedical applications, collagen is extracted from porcine or bovine skin, rat tail, or decalcified bovine or rabbit bone. Collagen is considered as an ideal material for numerous biomedical applications owing to its excellent properties, such as low antigenicity, cytocompatibility, and tissue regeneration potential (Jia et al., 2013). Since pure collagen has poor mechanical

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strength, it cannot be used as a bone substitute material in its native form. Therefore, research is focused on the production of composite scaffolds made up of collagen and bioactive ceramics.

9.3.1 COLLAGEN BIOACTIVE CERAMIC COMPOSITES These composites can be fabricated in two ways. The first method involves the addition of precipitated calcium phosphate powder into a solution of pure collagen followed by a crosslinking and lyophilization process which results in the development of a porous composite scaffold. Another method involves the precipitation of calcium and phosphate on a porous collagen material followed by a crosslinking and lyophilization process. Collagen bioactive ceramic composites can be constructed in the form of 3D scaffolds (Villa et al., 2015), hydrogels (Laydi et al., 2012), and dry powder form (Xu et al., 2015).

9.3.1.1 Collagen
HAP composites Yunoki et al. (2007) prepared a three-dimensional collagen-HAP (CHAP) nanocomposite scaffold through a coprecipitation method in which a phosphoric acid solution was mixed with collagen molecules and calcium hydroxide solution, which was then freeze-dried to create pores followed by a dehydrothermal crosslinking. The three-dimensional scaffolds prepared by this method showed high compressive moduli of 139 6 31 kPa with 94.7% porosity. New bone tissue formation in the pores was visible parrallely with the biodegradation of the implant, 4 weeks after implanting in rat femurs. Taira et al. (2009) developed CHAP scaffolds by dipping porous collagen sponge into calcium chloride and disodium hydrogen orthophosphate solutions five times. The scaffolds prepared by this method were found to be osteoconductive for 8 weeks when implanted in the tibia of a rabbit and started degrading after 12 weeks. Al-Munajjed et al. (2009) prepared CHAP composites by immersing porous collagen scaffolds in simulated bodily fluid for 4 days at 37 C. The process resulted in the deposition of the HAP layer onto the collagen strut. Cunniffe et al. (2010) prepared collagen HAP scaffolds using two different methods. In suspension methods, collagen
HAP scaffolds were formed through the lyophilization of collagen and HAP slurry. This scaffold showed 18 fold higher compressive moduli than collagen scaffolds. In the immersion technique, a porous collagen scaffold was immersed in HAP suspension and the resultant composite was found to have 12 fold higher compressive modulus in comparison to collagen scaffolds. Perez et al. (2013) developed micro- and nanostructured collagen HAP microcarriers for bone tissue engineering applications. They developed collagen
HAP by dispersing tricalcium phosphate (TCP) cement in collagen solution. Microcarriers were then fabricated with a water to oil emulsification technique. Apart from composite scaffolds, collagen
HAP composites were also used for coating titanium implants for enhancing osteointegration. Uezono et al., 2013

9.3 Collagen

coated Ti rods with CHAP through dipping in composite suspension followed by dehydrothermal crosslinking. Huang et al., 2014 coated Ti-Al-V cylinders with HAP using a plasma spraying method followed by dropping collagen solution onto it. Collagen-based composite scaffolds have also been used as carrier systems for cell or growth factors. In vitro cultures of induced pluripotent stem cell-derived mesenchymal stem cells (iPSCMSCs) have shown enhanced osteogenic differentiation when cultured on collagen/HAP/chitosan nanofibers (Xie et al., 2015). CHAP scaffolds soaked with total blood and platelet rich plasma (PRP) showed 2.7 and 3.3 fold higher bone formation respectively in comparison to CHAP scaffolds soaked in saline (Ohba et al., 2016). Similarly, CHAP scaffolds were also used as a carrier for bone morphogenetic protein-2 (BMP-2) which lead to a significant increase in bone volume (You et al., 2015). Similar positive effects on bone regeneration were reported by Tan et al., 2012; Hannink et al., 2013; and Taniyama et al., 2015.

9.3.1.2 Collagen TCP/BCP composites Composite scaffolds have been developed using calcium phosphates, such as β-tricalcium phosphate (β-TCP) and biphasic calcium phosphate (BCP) in conjunction with collagen. Collagen
β-TCP scaffolds in combination with bone marrow aspirate have been suggested as an alternative to autografts and has shown new Bone formation comparable to mineralized bone matrices (McDaniel et al., 2015). developed a new technique to produce collagen
CaP scaffolds through 3D printing of CaP scaffolds followed by its dipping in collagen solution. The resultant scaffolds showed good new bone growth when compared to allografts. Mate´-Sa´nchez de Val et al., 2014 compared collagen-HAP-TCP composites consisting of different composition ratios of HAP, TCP, and collagen and assessed the role of the ratio in determining bone formation capacity. The ideal ratio was found to be 40:30:30 corresponding to HAP wt%, TCP wt%, and collagen wt% respectively. The group further evaluated the effect of ratios on the resorption of scaffolds and found the fastest degradation rate with a ratio of 40:30:30, while a ratio of 60:20:20 showed the slowest resorption rate. Developed a scaffold comprising of a three-level hierarchy of CaP, collagen, and HAP. They combined a collagen network with porous calcium phosphate ceramic using a vacuum infusion method, followed by coating with HAP. After being implanted in the dorsal muscles of rabbits, this three-level hierarchical scaffold was found to have better mechanical strength and faster bone formation than normal CaP scaffolds. Tanaka et al. (2012) synthesized collagen and TCP composites for the delivery of FGF-2 in a rabbit tibial segmental defect. After 12 weeks, the defect was observed to be filled completely with new bone.

9.3.1.3 Collagen-bioglass based composites Bioactive glass, such as bioglass 45S5, 58S, S53P4, and Biosilicate, demonstrates excellent osteogenic (Xu et al., 2011) and angiogenic (Leu et al., 2008) potential.

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Therefore, researchers have employed these bioglass in conjunction with collagen to attain enhanced mineralization in comparison to collagen only scaffolds (Marelli et al., 2011; Marelli et al., 2010). These scaffolds have also shown increased metabolic activity during in vitro culture of MC3T3-E1 cells (Marelli et al., 2011). Long et al. (2014) developed a 3D macroporous composite scaffold of collagen and bioglass through a slurry dipping technique. The compressive strength of the scaffold was found to be 5.80 6 1.60 MPa, which is comparable to trabecular bone (4.00
12.00 MPa).Its water absorption was reduced from 889% to 52% because of the bioglass coating. Quinlan et al. (2015) developed cobalt ions incorporated into bioactive glass/collagen/glycosaminoglycan scaffolds that heightened the expression of vascular endothelial growth factor (VEGF) by means of stabilizing the hypoxia-inducible transcription factor. In vitro studies carried out using human umbilical vein and endothelial cells displayed superior tubule formation on cobalt-eluting collagen-bioactive glass scaffolds in comparison to cobalt free scaffolds. The scaffolds also showed higher osteogenetic activity than the cobalt-free scaffolds. Moreira et al. (2016) developed collagen-chitosan gels reinforced with bioactive glass nanoparticles for their application as injectable gels in place of prefabricated scaffolds. The gel showed thermosensitive behavior owing to β-glycerophosphate salts which were added to the gel. It remained fluid at room temperature but gelatinized at 37 C.

9.3.2 MEDICAL APPLICATIONS OF COLLAGEN Several collagen-based products have reached the market and Table 9.3 lists selected companies involved in bone tissue engineering. Clinical trials have been performed for Collograft, a product of Zimmer which consists of 60% HAP, 40% TCP, and purified bovine collagen. Collograft was mixed with the autogenous bone marrow of patients between the ages of 18 and 72, who had developed fractures in their long bones, such as the radius, ulna, humerus, tibia, or femur. It served as an effective bone graft substitute and its healing rate was comparable with that of autograft (Cornell et al., 1991). Another product available on the market from the same company is Puros Demineralized Bone Matrix (DBM) Putty which is prepared by the mineralized solvent dehydration of allografts. When compared to control grafts made of demineralized freeze-dried bone allografts (DFDBA) and Bio-Oss (an organic bovine bone mineral; a xenograft), it was found that grafts of Puros were resorbed and replaced with new bone significantly faster than DFDBA and Bio-Oss grafts (Noumbissi et al., 2005). Bio-Oss developed by Geistlich Biomaterials was clinically investigated in 42 patients suffering from severe horizontal bone atrophy. Autogenous block grafts with Bio-Oss collagen were successfully grafted with high efficacy providing complete healing within 5.8 months (Von Arx, Buser, 2006). Another product from the same manufacturers, namely Bio-Oss collagen, was

Table 9.3 Commercially Available Collagen Products and Their Applications (Kuttappan et al., 2016) S. No.

Product

Company

Constituent

Applications

1

Collagraft

Zimmer

2

GingivAid

3

Formagraft

Type I collagen, HAP/β-TCP

Traumatic osseous defects, fixation of acute long bone fractures, and bone void filler Dental implants, sinus lift. and alveolar ridge augmentation Substitute in bone grafting

4

Type I collagen, CaP

Bone void filler

5 6

CopiOs Sponge and Paste Ossigen OsseoFit

Maxigen Biotech Inc. Maxigen Biotech Inc. Zimmer

Bovine collagenType I and II, HAP, β-TCP Type I collagen, HAP/β-TCP

7

CONDUCTMatrix

Bone void filler Filler for bone voids or bone defects in the pelvis and extremities Bone void filler

8

Integra Mozaik

Collagen, bone mineral Type I bovine collagen, PLLA, β-TCP Type I collagen, carbonate apatite material Type I collagen, TCP

9 10

Vitoss Orthoss Collagen

Collagen, β-TCP, bioactive glass Porcine collagen, bovine HAP

11 12 13

MASTERGRAFT Putty Puros DBM Putty RegenOss

Bone void filler Bone void filler, reconstruction of tissues in orthopedic and spinal surgery, volume extender for composite bone grafting Bone void filler

14

Bio-Oss Collagen

Exactech DSM medical GLOBUS medical Integra LifeSciences Stryker Geistlich Surgery Medtronic

Type I bovine collagen, β-TCP

Zimmer JRI Orthopedics Geistlich Biomaterials

Demineralized bone matrix Type I collagen fibers, magnesium enriched HAP Porcine collagen, Geistlich Bio-Oss Particles

Bone void filler

Bone void filler Fixation devices for long bone fractures, spinal fusion Sinus floor Elevation, periodontal Regeneration, peri-implantitis, ridge Augmentation

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investigated for curing intrabony defects in single rooted tooth. The product was used alone and in combination with a collagen membrane and both curing procedures resulted in enhancing depth of probing pocket and bone fill (Nevins et al., 2003).

9.4 SILK FIBROIN Many insects, such as Bombyx mori, spiders, mites, and beetles, are known to have the ability to produce silk. The chemical composition and structure of silk fibroin (SF) vary among different species (Sutherland et al., 2010). SF from African wild silk moth, Gonometa postica, revealed a considerably high amount of serine, aspartic acid, and arginine (Mhuka et al., 2013), while Indian tropical tasar silk of Antheraea mylitta was found to possess arginine, glycin, and aspartic acid. In the past few years, there has been growing interest in the fabrication of chimeric silk, which is a composite fiber of B. mori silk and spider silk. Chimeric silk was reported to have superior mechanical properties compared to commercial grade silk (Teule´ et al., 2012). The earlier years of the application of silk fibroin as scaffolds saw reported allergic reactions followed by the exclusion of silk fibroins from biomedical applications (Soong and Kenyon, 1984; Wen et al., 1990; Hollander, 1994). Later it was revealed that the adverse effects of silk fibroins were caused by the presence of sericin (Aramwit et al., 2009). Therefore, the removal of sericin has become the first step in the fabrication of scaffolds. The glue-like sericin protein which covers fibroin is usually removed through the use of a degumming method that involves the thermochemical treatment of cocoons. Removing the watersoluble sericin protein covering from the cocoon leaves behind degummed fibers of silk that possess similar Young’s moduli. However, alterations in tensile strength ranging from 450 to 700 MPa was observed at an elongation of 12%
24% (Pe´rez-Rigueiro et al., 2001). These remarkable mechanical characteristics make SF desirable for the fabrication of load-bearing composites.

9.4.1 STRUCTURE OF SILK FIBROIN The structure of SF comprises of two types of subunits: a light chain and a heavy chain that are connected by means of a disulfide bond. The light chain consists of nonrepetitive and hydrophilic sequences while the heavy chain comprises of repeating hydrophobic blocks of Gly-Ala-Gly-Ala-Gly-Ser and repetitive blocks of Gly-Ala/Ser/Tyr dipeptides, which collectively form 12 crystalline domains (Zhou et al., 2001). High glycine content accounts for the tight packing of these domains into extremely stable β-sheet nanocrystals. The foremost molecular interactions in these crystalline β-sheets are hydrogen bonds that are responsible for enhancing the rigidity and tensile strength of silk. Silk possess the capability to

9.4 Silk Fibroin

self-assemble and self-heal by reforming its hydrogen bonds (Keten et al., 2010). SF can self-assemble itself into larger fibrous structures which are characterized as possessing enhanced mechanical properties (Sponner et al., 2007).

9.4.2 PROCESSING OF SILK FIBROIN SF can be processed into hydrogels, sponges, fibers, particles, microspheres, tubes, and electrospun fiber mats.

9.4.2.1 Hydrogelation Hydrogelation is induced in SF solution through high temperature, vortexing, sonication, freeze gelation, and electrogelation (Yucel et al., 2009; Wang et al., 2008a,b; Yucel et al., 2010). Structural changes during hydrogelation are caused by the disordered state of β-sheet conformation which further crosslinks and stabilizes the gel. The resultant hydrogels are three-dimensional polymer networks suitable for the delivery of cytokines and growth factors in tissue engineering applications. SF cryogels are obtained from frozen SF solutions that exhibit remarkable mechanical properties allowing them to resist compression without suffering from any crack development (Bhardwaj et al., 2011).

9.4.2.2 Electrospinning Electrospinning allows for the production of silk fibers with diameters ranging from micrometer to nanometer scale which mimics the properties of nanoscale fibrous ECM components. Fine mats of B. mori SF with a fiber diameter of less than 800 nm were designed by electrospinning with polyethylene oxide (Jin et al., 2002), while fibers of 80 nm diameter were produced by electrospinning with formic acid employed as a solvent (Min et al., 2004). Electrospun 3D scaffolds were reported to reveal increased adhesion and proliferation of preosteoblasts as well as the ALP activity of osteoblasts (Ki et al., 2008). In vivo studies in rats revealed increased bone regeneration in 3D electrospun scaffolds in comparison to porous 3D PLA scaffolds (Park et al., 2010).

9.4.2.3 Porogen leaching Porogen leaching is a method for fabricating porous SF sponges. The 3D porous structure of these sponge-like scaffolds allows for cell attachment and migration. The method involves pouring SF solution on porogens with defined pore sizes and allowing the solution to solidify. The porogen is later leached out to obtain SF sponges of controlled pore size and geometry (Correia et al., 2012). Process parameters can be varied by changing the type of solvent used, such as hexafluoro-2-propanol (HFP). HFP derived SF scaffolds have been reported to show slower degradation than scaffolds derived from aqueous solutions (Wang et al., 2008a,b).

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9.4.2.4 3D bioprinting Patterned SF nests (70
100 μm in diameter) modified with anionic and cationic side chains were employed as anchors for incubating and proliferating Escherichia coli cells by developing an inkjet process (Suntivich et al., 2014). 3D printing, in the absence of crosslinking agents or thickeners, was employed for recombinant spider silk protein, eADF4(C16), by pregelling the solution overnight at 37 C with 95% relative humidity (Schacht et al., 2015). The adhesion of different cell types, such as osteoblasts, fibroblasts, HeLa cells, myoblasts, or keratinocytes, which were seeded after the completion of the printing process, was investigated. Results revealed better adhesion of osteoblasts than of other cell types.

9.4.2.5 SF composites SF is frequently used in combination with other biomaterials, such as CaP, hydroxyapatite, and collagen, to fabricate composite scaffolds for tissue engineering applications. Hydroxyapatite can be incorporated into porous SF scaffolds by employing the direct deposition method (Jiang et al., 2013). The resulting scaffold leads to enhanced bone formation. Nanohydroxyapatite (NanoHA) has also been incorporated in SF hydrogels using ethanol as a gelating agent. The obtained SF/ NanoHA hydrogels were suggested as filling materials for bone defects (Ribeiro et al., 2015).

9.4.3 MEDICAL APPLICATIONS OF SILK FIBROIN 9.4.3.1 SF scaffolds for tissue engineering Porous SF scaffolds exhibit architecture comparable to trabecular and cortical bones, which enables them to facilitate osteogenesis in defect sites. Threedimensional electrospun scaffolds, possessing high porosity, controlled pore size, and ECM-like structure were reported as an alternative to PLA scaffolds with the same morphology (Park et al., 2010). Wu et al., 2011 incorporated mesoporous bioactive glass to induce rapid mineralization. After being implanted in mouse calvarial defects, the bioactive glass modified scaffolds facilitated mineralized bone formation in comparison to unmodified scaffolds.

9.4.3.2 Delivery of bioactive molecules SF has been reported to possess remarkable properties as a carrier for delivering bioactive molecules in numerous therapeutic applications. Various morphologies of SF, such as 3D scaffolds, hydrogels, planar films, microparticles, and electrospun fibers, have been investigated by researchers for the delivery of growth factors and enzymes (Haider et al., 2005; Shi et al., 2013; Farokhi et al., 2014). The delivery of BMP-2, which is known to play a decisive role in osteogenesis, has been investigated extensively (Wang et al., 2014). BMP-2 loaded SF particles showed superior bone formation in comparison to BMP-2 alone. In another study,

9.5 Biocellulose

BMP-2 was immobilized on planar SF films which resulted in amplified calcium deposition, increased ALP activity, and heightened expression of collagen I, osteopontin, and osteocalcin (Karageorgiou et al., 2004). BMP-2 growth factor was incorporated into an SF scaffold and new bone formation was observed to be accelerated (Karageorgiou et al., 2006).

9.4.3.3 Fixation devices Kim et al. (2005a,b) prepared nanofibrous membranes of electrospun SF to analyze their biocompatibility after implanting in rabbit calvarial defects. After 8 weeks of implantation, the SF membranes revealed complete defect coverage. Perrone et al. (2014) fabricated resorbable screws of SF for fixing rat femoral defects. After 8 weeks of implantation resorbable SF screws showed satisfactory bone remodeling comparable to commercially available PLGA fixation systems.

9.5 BIOCELLULOSE Cellulose, the Earth’s main biopolymer is assembled from glucose produced in plants during photosynthesis. It is a water-insoluble compound commonly found in the cell wall of plants (Keshk, 2014). Even though plants are the chief contributors of cellulose, numerous bacteria, such as Acetobacter, also retain the capability to produce cellulose (Esa et al., 2014). These nonphotosynthetic microorganisms acquire sugar, glucose, glycerol, or other simple organic substrates and convert them into pure cellulose. Unlike plant cellulose, bacterial cellulose is lacking of other contaminating polysaccharides and its isolation and purification processes are relatively easy to perform, and do not require high energy or severe chemical processes. Very few genera of bacteria are able to synthesize cellulose. Out of these species, the Gram-negative bacterium Gluconacetobacter xylinus (previously known as Acetobacter xylinum) secretes large quantities of cellulose. Other cellulose producing bacterial species include Achromobacter, Aerobacter, Agrobacterium, and Alcaligenes. It provides a wide range of applications as food matrix, dietary fibers, filter membranes, paper, etc. Due to its hydrophilic nature, biodegradability, nontoxicity, and biocompatibility, it has also emerged as a remarkable biopolymer for biomedical applications, such as artificial skin for temporary wound recovery, artificial blood vessels, scaffolds for tissue engineering applications, etc.

9.5.1 BIOCELLULOSE FIBRIL STRUCTURE The structure of biocellulose comprises of fibrils having β-1-4 glucopyranose chains that are held together by inter- and intramolecular hydrogen bonds (Ul-Islam et al., 2013). These units are arranged in a chair-like conformation in

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FIGURE 9.3 Chemical structure and intra- and intermolecular hydrogen bonds in crystalline cellulose.

which glucose residues rotated through an angle of 180 degrees around the molecular axis and hydroxyl group in an equatorial position. The hydrogen bond forming ability of the hydroxyl group plays a significant role in fibrillar and crystalline packing of cellulose. Microfibrils of bacterial cellulose are about 100 times smaller than plant cellulose (Gayathry and Gopalaswamy, 2014). Native cellulose consists of two types of crystal structures: Iα and Iβ (Keshk, 2014). Cellulose Iα is crystallized in large-sized microfibrils, whereas cellulose Iβ is formed in small-sized microfibrils. The structure of cellulose Iβ is thermodynamically more stable than that of cellulose Iα. The unit cell of cellulose Iα is triclinic, while in the case of cellulose Iβ the unit cell is monoclinic. The ratio of cellulose Iα and Iβ fluctuates significantly from species to species (Hirai et al., 1987) (Fig. 9.3).

9.5.2 PROPERTIES OF BIOCELLULOSE 9.5.2.1 Mechanical properties The Young’s modulus of crystalline cellulose is in the range of 100
200 GPa, which is comparable to the strength of steel. Its elastic modulus was determined by methods based on nanoscale indentation and was reported to be 139.5 6 3.5 GPa (Wu et al., 2013). Dri et al. (2013) employed an atomic structure model of cellulose with the help of quantum mechanics principles and predicted the Young’s modulus of crystalline cellulose to be around 206 GPa, which is similar to steel. A wide range of values were also reported for the longitudinal modulus of cellulose which were based on various theoretical and experimental approaches and the average modulus was reported to be 100 GPa (Cheng et al., 2009a,b). Attributed to these notable mechanical properties, cellulose has been considered as a potential material for load bearing applications. The incidences of high modulus imply proper stress transfer from matrices (host material) to cellulose (reinforced phase).

9.5 Biocellulose

9.5.2.2 Biocompatibility Biocellulose (BC) is broadly considered as a biocompatible material that only evokes moderate responses in vivo. Helenius et al. (2006) studied the biocompatibility of subcutaneous BC implants in rats with no observed foreign body reaction. Andrade et al. (2012) investigated the biocompatibility of BC membranes and peptide modified BC membrane implants in sheep and reported mild irritation in tissues with no change in inflammatory response. In another study, BC membranes did generate mild inflammation that decreased over time, however, no differences were observed regarding B-lymphocyte precursor populations and myeloid cells in bone marrow (Pe´rtile et al., 2012).

9.5.2.3 Hemocompatibility Hemocompatibility refers to the ability of a foreign material to exist in harmony with blood tissues, without causing any deteriorating effect. This property is specifically relevant for blood contacting biomaterials such as artificial blood vessels and pumps. Andrade et al. (2011) investigated the hemocompatibility of native bacterial cellulose and peptide modified bacterial cellulose using plasma recalcification time and whole blood clotting experiments, which exhibited favorable interactions with platelets establishing BC as a hemocompatible material.

9.5.2.4 Biodegradability Cellulose is a nonbiodegradable material for animals, as they do not possess the cellulose degrading enzyme called cellulase. However, the degree of degradation may vary depending on crystallinity, hydration, and swelling of cellulose. Miyamoto et al. (1989) studied cellulose degradation in canine specimens and reported that degradation depends on the crystalline structure of cellulose. It has been established that oxidized cellulose is more vulnerable to hydrolysis and, therefore, readily degradable. This finding has encouraged researchers to enhance the biodegradability of cellulose by employing oxidation reactions to make BC more vulnerable to hydrolysis (Li et al., 2009; Luo et al., 2013).

9.5.2.5 Nontoxicity There are no reports suggesting that BC is a toxic biomaterial as it does not procure any serious influence on cellular and genetic levels. However, inhalation of nanocellulose may cause pulmonary inflammation that results from nondegradable biomaterials.

9.5.3 BIOMEDICAL APPLICATIONS OF BIOCELLULOSE 9.5.3.1 Substitute biomaterials for medical applications Attributed to greater mechanical strength in wet state and generous permeability for liquids and gases, bacterial cellulose membranes are used as an artificial skin for temporary recovery of skin wounds (Fontana et al., 1990). Biofill and

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Gengiflex are two commercially available products of bacterial cellulose that are used in surgery and dental implants. Biofill has been used for efficacious cure of second and third degree burns as a temporary substitute to human skin. Bacterial cellulose can be used as a biomaterial to fabricate artificial vascular grafts that are used during coronary bypass graft surgery to treat cardiovascular disease. Klemm et al. (2001), was the first research group to scrutinize artificial vascular substitutes acquired with biomaterials from bacterial cellulose. They described a clinical product named BActerial SYnthesized Cellulose (BASYC) with high mechanical strength in wet state, low roughness of inner tube surface, and enormous water retention property. BASYC has been reported as successful for blood vessel replacement in rats and pigs (Schumann et al., 2009; Wippermann et al., 2009). BC/polyacrylamide gels have been prepared by combining PAA gels with bacterial cellulose in the presence of a cross-linker which provided a tensile fracture stress (40 6 10 MPa) similar to that of ligaments (38 6 10 MPa), which suggests these gels as potential substitutes during ligament replacements (Hagiwara et al., 2010). Similarly, nanocomposites of cellulose nanofibrils and collagen has revealed comparable mechanical properties to natural tendons and ligaments. These composites showed positive response during in vitro biocompatibility investigations providing adequate proliferation and differentiation of human endothelial cells (Mathew et al., 2012). Studies have elaborated a biocomposite hydrogel combined with carboxymethylated cellulose nanofibrils as a potential replacement for human nucleus pulposus present in intervertebral disks (Eyholzer et al., 2011). This gelatinous material is responsible for providing flexibility to the spine and is important for stress dissipation. Cellulose nanofibril reinforced biocomposite hydrogel can replace nucleus pulposus, providing similar mechanical properties and appropriate swelling ratio to restore the height of intervertebral disks (Borges et al., 2011).

9.5.3.2 Biocellulose-based scaffolds for bone tissue regeneration Biocellulose based tissue scaffolds have been investigated for the bone regeneration and healing required to compensate for bone loss caused by trauma, reconstructive surgery, neoplasia, congenital defects, etc. BC/hydroxyapatite membrane scaffolds were investigated for their bone regeneration capability and were reported to enhance the growth rate of osteoblast cells resulting in greater bone nodule formation. Saska et al. (2011) evaluated BC/HAP membrane scaffolds for the regeneration of bone in the tibia of rats and reported improved bone formation in defect sites.

9.5.3.3 Scaffolds for cell culture Cellulose based biomaterials offer a favorable environment to encourage efficient cell attachment and proliferation which is quite similar to natural tissue in terms of biocompatibility and mechanical properties. Cellulose based scaffolds for cell culture applications have been explored in the form of composites, hydrogels,

9.5 Biocellulose

electrospun fibers, membranes, and sponges. Electrospun fibers of maleic anhydride-grafted PLA reinforced with cellulose were investigated as potential scaffold materials for cell culture applications (Zhou et al., 2013). Human adult adipose derived mesenchymal stem cells (hASCs) seeded on these scaffolds showed enhanced cell proliferation. Nanocellulose based biomimetic nanocomposites were prepared by adding catalase and calcium peroxide (CaO2) into a matrix of cellulose nanofibrils. The incorporation of calcium peroxide and catalase induced the generation of hydrogen peroxide or oxygen, which can affect the survival of cells. While hydrogen peroxide negatively affected cell proliferation, the liberation of oxygen resulted in increased cell proliferation (Chang and Wang, 2013). Various biocellulose based scaffolds are used for applications in tissue engineering, which can be divided into three categories: pellicle/membrane scaffolds, surface modified pellicle scaffold, and matrix biocomposite scaffolds. Favi et al. (2013) investigated biocellulose based hydrogel scaffolds employed as supporting biomaterials for a culture of equine-derived bone marrow mesenchymal stem cells (EqMSCs). Chitosan, alginate, agarose, collagen, polypyrrole, gelatin, and poly(3hydroxybutyrate-co-4-hydroxybutyrate) have been investigated for utilization as matrices for BC/matrix scaffold systems. BC/polypyrrole membrane scaffolds were studied for seeding PC12 rat neuronal cells (Muller et al., 2013). Surface modifications, such as surface sulfation, phosphorylation, protein or peptide coatings, plasma treatment, etc., can be performed to enhance cell attachment. A BC scaffold coated with bone morphogenetic protein-2 (BMP-2) revealed a good biocompatibility, and the presence of BMP-2 has been reported to enhance the differentiation of mouse fibroblast-like C2C12 cells into osteoblasts. These BMP-2 modified scaffolds have also been reported to enhance bone formation during subcutaneous implantation (Shi et al., 2013).

9.5.3.4 Antimicrobial biomaterials The presence of bacterial communities in burn wounds, surgical procedures, and traumatic injuries cause wound infections that further delay the process of wound healing. This has led researchers to investigate wound dressing biomaterials exhibiting antimicrobial properties. The fabrication of cellulose based antimicrobial biomaterials can be achieved by incorporating antimicrobial materials into nanocellulose comprising of a porous network structure aided by physical and chemical approaches. Antimicrobial materials can be inorganic, such as silver particles and its derivative (Maneerung et al., 2008; Luan et al., 2012) compounds, or it can be an organic agent, such as lysozyme. Silver nanoparticles are the most intensively studied antimicrobial agents and can be incorporated into nanocellulose based biomaterials through chemical reduction by silver nitrate and impregnation method (Maneerung et al., 2008). Cellulose nanofibril materials possessing dendritic structured silver nanoparticles displayed better antimicrobial properties than spherical silver nanoparticles.

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9.5.3.5 Drug delivery applications Cellulose has an extended history of applications in the pharmaceutical industry as excipients to condense matrices loaded with drugs for the purpose of oral administration. There is still continuous research going on regarding the rate of tablet disintegration and prolonged drug release. Nanocellulose based drug carriers are categorized into three forms: microspheres, hydrogels, and membranes. Lin et al. (2011) developed a nanocellulose/sodium alginate microsphere-based drug delivery system. Muller et al. (2013) investigated bacterial cellulose as a hydrogel drug carrier for bovine serum albumin and found that drug release was a diffusion and swelling controlled process in regard to hydrogels. Huang et al. (2013a,b) investigated biocellulose membranes for utilization as a drug delivery system for the delivery of berberine hydrochloride and berberine sulfate and reported prolonged drug release. Even though attempts at using cellulose based materials for drug delivery systems has garnered positive outcomes, there are still many issues that require investigation in order to optimize the design process of these drug delivery systems. Specific attention needs to be given to the interaction between drug molecules and nanocellulose materials that can have both promoting and inhibitory effects on drug activity (Kolakovic et al., 2013).

9.5.3.6 Other biomedical applications The modification of nanocellulose based biomaterials by labeling with fluorescent materials enables their potential application in biosensors, optical bioimaging, and photodynamic therapy in biomedical fields. Labeling of the fluorescein-50isothiocyanate (FITC) molecule in cellulose nanofibrils based biomaterials was the first study reported regarding fluorescent labelling of cellulosic materials (Dong and Roman, 2007). Since then, various fluorescent compounds, such as pyrene dyes (Zhang et al., 2012), Rhodamine B isothiocyanate (Mahmoud et al., 2010), terpyridine and its derivatives (Hassan et al., 2012), 1-pyrenebu-tyric acid N-hydroxy succinimide ester (Yang and Pan, 2010), 5-(and-6)-carboxytetramethylrhod-amine succinimidyl ester, 5-(and-6)-carboxyfluorescein succinimidyl ester, 5-(and-6)carboxytetramethylrhod-amine succinimidyl ester (Nielsen et al., 2010), 5-(4, 6dichlorotriazinyl) aminofluorescein (Abitbol et al., 2013), PEI-chlorin p6 derivatives (Drogat et al., 2012), and 7-amino-4-methylcoumarin (Huang et al., 2013a,b) have been investigated. These fluorescent biocomposites have attracted enormous research interest from the scientific community of late, however, there is still a long way to go before achieving their practical application in the biomedical field.

9.6 CONCLUSIONS Being biodegradable and biocompatible, the fabrication and utilization of green polymers is a promising and sustainable technique to replace the conventional

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materials used for biomedical applications. Currently a wide range of biodegradable polymers exist that hold huge potential as biomaterials for biomedical applications. Biopolymers, such as PLA, PGA, PHB, collagen, biocellulose, etc., offer a variety of properties, such as biodegradability, biocompatibility, fabrication flexibility, and recyclability, which make them plausible alternates for conventional nonbiodegradable polymers. Additionally, current advancements in processing technologies and computer aided technology has enabled the formation of scaffolds with complex architecture that mimic the properties of their biological counterparts. Along with the research into polymers, the advancements in biological research have also led to a better understanding of the interactions between host and biomaterials on cellular, tissue, and systemic levels. However, there are still many limitations associated with bio-based polymers. For instance, the brittleness and poor thermal resistance of PLA have been a major bottleneck for its commercial application. Similarly, the mechanical strength of collagen also demands improvement before being suitable for orthopedic applications. These challenges regarding the improvement of physical and mechanical properties can be overcome with the blending of polymers with the desired properties, composites formation, polymer modifications, and by coating with other materials, etc. Apart from the challenges imposed by the inherent properties of bio-based polymers, there are also some technical challenges that block the commercial success of these polymers. One major technical challenge is maintaining the biodegradable properties of bio-based polymers during attempts to achieve higher mechanical and physical properties. Therefore, the field of biodegradable polymers demands active collaboration from material scientists, biologists, and chemists for carry on the progress in biomedical applications.

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CHAPTER

Composite scaffolds for bone and osteochondral defects

10

Vincenzo Guarino1, Silvia Scaglione2, Monica Sandri3, Simone Sprio3, Anna Tampieri3 and Luigi Ambrosio1 1

Institute of Polymers, Composites and Biomaterials, National Research Council of Italy, Naples, Italy 2CNR-National Research Council of Italy, IEIIT Institute, Genoa, Italy 3CNRNational Research Council of Italy, Institute of Science and Technology for Ceramic Materials (ISTEC), Faenza, Italy

10.1 INTRODUCTION The use of bioinspired scaffolds is mandatory to guide the regeneration of natural tissues. They have to be constituted by ideal materials that can control and promote specific cellular events, thus, having selected properties able to reflect the peculiar structure to be substituted (Guarino et al., 2007). Indeed, complex structures of tissues may be understood only by studying how natural processing phenomena concur to define the final shape and structure of the tissue, from macro- to nanoscale, and the nature of its various physical and chemical interactions (O’Brien, 2011). To date, all these approaches based on the “learning from nature” have been universally recognized as successful for the identification of basic guidelines to fabricate ex novo hard and mineralized tissues (i.e., bone, osteochondral). For instance, bone is a natural, anisotropic composite structure, with higher stiffness and tensile strength than soft tissues such as skin, cartilage, or blood vessels. For bone replacement, it is possible to combine a large variety of smart structural components—that is, cells, polymer chains, bioactive molecules, or magnetic responsive particles—in situ selectively to respond to the pertinent stimuli and, thus, triggering basic tissue formation mechanisms such as biomineralization (Raucci et al., 2012; Hai-Yan and Ning, 2014). In this context, a mimesis of the living tissue—the optimal replica of their mechanical, biological and functional properties—may be achieved by the support of open-pore biocompatible and biodegradable scaffolds able to provide a temporary substitute for the extracellular matrix (ECM) of natural tissue (Hutmacher, 2001). Firstly, these systems have to possess multifunctional properties to provide

Materials for Biomedical Engineering: Hydrogels and Polymer-based Scaffolds. DOI: https://doi.org/10.1016/B978-0-12-816901-8.00010-9 © 2019 Elsevier Inc. All rights reserved.

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a tailor-made substrate suitable for cell attachment, migration, proliferation, and differentiation (Sunderlacruz and Kaplan, 2009) Moreover, they have to present programmed biomechanical properties to guarantee the optimal transfer of the load-bearing function from the engineered material to the growing ECM (Guilak et al., 2014). With this aim, over the past two decades, degradable and partially degradable matrices have been variously designed by different manufacturing strategies to satisfy these requirements. More recently, it has been demonstrated that the most interesting way to satisfy the compelling multifunctional needs relies on the development of composite biomaterials for bone replacement. It is well-known that composite materials are generally composed of two phases, a continuous and a dispersed phase, separated by an interface with intermediate properties. The continuous phase is referred to the matrix and biodegradable and or bioresorbable polymers generally represent the material mostly used for its components. The dispersed phase, which is generally stiffer than the host matrix, can be discontinuous or continuous and it mainly works as a reinforcing agent to enhance the mechanical properties of the matrix (i.e., stiffness, strength) (Guarino et al., 2011). The interface between matrix and reinforcement plays a pivotal role in determining the mechanical performances and environmental stability of composite materials (Peter et al., 2010). For instance, it is well-known that physical and mechanical parameters (i.e., coefficient of thermal expansion, stiffness, strength, fatigue behavior) are strongly related to the interfacial features among different constituents of composite materials. From a mechanical point of view, the interface has to transfer a part of the stress to the reinforcement, thus, contributing to the composite stiffness and strength. Consequently, an important aspect in designing composite scaffolds is related to the stress transfer from matrix to reinforcement while trying to avoid discontinuities in the stress transfer and generation of stress concentration at the matrix/reinforcement interface. Accordingly, to obtain suitable interfaces, several techniques have been employed involving the reinforcement coating, modification of the matrix composition, and the chemical functionalization of the constituents. Among composite technologies for scaffold design, the most consolidated strategies consist in the enhancement of mechanical properties of the biocompatible polymer matrix by the use of reinforcement systems based on rigid bioactive particles (i.e., calcium phosphates) or short/long/continuous fibers which, in turn, might concur to the osteoconductivity and/or mechanical response of the polymer matrix. In all these cases, the main challenge is to impart these features by preserving a suitable porosity required to assure its success as a scaffold or for other orthopedic applications. Additionally, in another more biomimetic approach, the secondary phase consists in nanostructured apatite particles directly and homogeneously nucleated into the fibers of the main continuous bioresorbable polymeric matrix, suitable to improve the bioactivity and the mechanical behavior of the bioresorbable matrix.

10.2 Biodegradable Matrices

By pursuing the concept of “biomimetics,” indicating the ability of a synthetic material to closely reproduce the chemical composition, physical properties, and architecture of native bone tissues, this chapter highlights how the optimization of a lab-scale biomineralization processes allow the creation of 3D environments able to deliver signals stimulating cell chemotaxis and specific differentiation of autologous stem cells (Scaglione et al., 2012; Minardi et al., 2015; Sprio et al., 2016). This process of reproducing the cascade of phenomena acting in the formation of hybrid nanocomposites such as bone, and involving fully bioresorbable polymeric matrices can generate hybrid fibrous structures with excellent regenerative ability and improved safety because of the reduction of adverse side effects. Moreover, we highlight the recent development of novel superparamagnetic biohybrid devices and provide an overview of their potential future trends. The recent development of a bioactive superparamagnetic hydroxyapatite opens new possibilities to produce scaffolds with remote controlling, avoiding any cytotoxic effect. The application of local weak magnetic fields may provide a new tool to assist and direct cell behavior, thus, increasing the osteogenic and angiogenic capacity of the new bone scaffolds and this may open new perspectives in regenerative medicine. The main concept is that bone regeneration can be greatly aided by implantation of a biomimetic scaffold. In this case, the patient body behaves as a natural bioreactor guiding proper tissue regeneration without the need of complicated tissue engineering procedures or of the use of biological factors, thus, improving the safety of clinical approaches. In this chapter we will describe consolidated as well as emerging technologies to control the material and structural properties of composite scaffolds at micrometric and submicrometric scales for the repair and regeneration of bone and osteochondral tissues.

10.2 BIODEGRADABLE MATRICES Facing a complex biological and sensitive system such as the human body, biocompatibility is an imperative requirement for scaffold matrices in tissue engineering applications. It means that constituent materials must not elicit an unresolved inflammatory response or demonstrate immunogenicity or cytotoxicity. In addition, they have to present sufficient mechanical properties to, at least, not collapse during surgical procedures as well as allowing the patient’s normal activities, that is, biomechanical compatibility. In this context, a large set of polymeric materials have been successfully used as multifunctional matrices to build multifunctional scaffolds suitable to satisfy two key requirements: (1) material components with time and space-controlled biodegradability; and (2) percolative architectures with interconnecting pores of the right scale to promote tissue integration and vascularization. Additionally, they have to actively contribute to specific mechanical properties of the

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composite scaffolds, thus, offering an appropriate interface response to their implantation site, by exhibiting an appropriate surface chemistry and topology to support cell activities (i.e., attachment, differentiation, and proliferation). The main advantage of biodegradable polymers consists in their unique capability to be easily formed into a variety of shapes and sizes under less invasive processing conditions (Guarino et al., 2008a,b,c; Guarino and Ambrosio, 2010) to finely reproduce tissue defects with complex geometries and multilevel organization into composite scaffolds at micro and submicrometric level. Biodegradable matrices play a relevant role in the response of the scaffolds to specific external stimuli at the implant site. It is well-known that, once introduced into the body, bioactive polymers may actively or passively induce a biochemical response from the living tissue, thus, triggering strong biological adhesion/fixation mechanisms and reproducing the biological microenvironment—that is, changes in pH or cell-associated enzyme activity—around the scaffold during the integration process. Indeed, the biodegradation of polymer matrices plays a crucial role to regulate the interplay among biological signals exerted by cells by providing a dynamic modification of the scaffold morphology and a contextual release of small molecules that may be accepted by the local microenvironment. Hence, an adequate definition degradation rate is compulsory to address to achieve the desired physical properties of the scaffolds. For these reasons, biodegradable polymers have been used as a support matrix for the homing and delivery of cultured cells in 3D tissue reconstruction or as platforms for controlled drug delivery strategies (Langer and Tirrell, 2004). The increasing interest in degradable biomaterials mainly rises from their biological properties (Rezwan et al., 2006) and the opportunity to use them to fabricate multicomponent scaffolds with multiscale degradation kinetics for tissue repair and regeneration (Holy et al., 2003). In the current scenario of tissue engineering, composite scaffolds may be composed of biodegradable polymers either from natural or synthetic sources. Natural materials, including polysaccharides (starch, alginate, chitin/chitosan, and hyaluronic acid derivatives) or proteins (soya, collagen, fibrin gels, and silk), are generally used as matrices and only rarely as reinforcement agents—that is, in the form of biofibers such as lignocellulosic natural fibers—for the fabrication of scaffolds that are able to mimic the native composition of the ECM. This mainly depends on their capability to be easily decomposed into biologically recognized molecules without the production of harmful intermediates which can be metabolized and removed from the body via naturally pathways (i.e., metabolism or excretion) (Vert, 2005). Despite their excellent biological compatibility, they show several deficiencies in terms of mechanical responsiveness which drastically limits their use for hard tissues. For this reason, mineralized scaffolds for bone are often based on synthetic biodegradable polymers which can be synthesized under more finely controlled conditions and material impurities, thus exhibiting more predictable and reproducible mechanical and physical properties (i.e., tensile strength, elastic modulus, and degradation rate). However, biological risks such as

10.2 Biodegradable Matrices

toxicity, immunogenicity, and infections are more frequent in the presence of matrices from synthetic sources. However, the use of pure synthetic polymers with simple repeating units and structure may drastically improve their in vitro and in vivo biological performance (Table 10.1). Synthetic polymers mostly used for 3D composite scaffolds include saturated poly-α-hydroxyesters, such as polycaprolactone (PCL), poly(lactic acid) (PLA), and poly(glycolic acid) (PGA), as well as poly(lactic-coglycolide) (PLGA) copolymers which are characterized by hydrolytic degradation through deesterification (Mano et al., 2004). During the degradation process, polymer chains are broken into short fragments or monomeric components are removed by natural pathways (Seal et al., 2001). The body already contains highly regulated mechanisms for completely removing these monomeric components by converting them into metabolites. As a function of starting molecular weights of polymer chains and scaffolds manufacturing conditions, it is possible to regulate degradation rates in order to partially set physical and mechanical properties of the scaffold within a wide range. However, most undergo a bulk erosion process mainly due to the abrupt release of acidic degradation products which often favor strong inflammatory responses from the surrounding tissue and, consequently, premature failure of the implant (Martin et al., 1996). Likewise, poly-α-hydroxyesters and their copolymers (Kohn, 1996) are characterized by surface erosion mechanisms which also release acidic products with an adverse response at the surface. Although in vitro experiments have shown promising outcomes, in vivo tests have underlined some biological limitations, including acute inflammatory phenomena ascribable to slow degradation rates (Santavirta et al., 1990). In opposition to bulk degrading ones, they undergo a heterogeneous hydrolysis process predominantly confined to the polymer water interface. In this case, they are classified as “surface eroding polymers” and include poly(anhydrides), poly (ortho-esters) and polyphosphazene. The surface-eroding characteristics offer Table 10.1 Biodegradable Matrices: Summary of Main Physical Properties (Middleton and Tipton, 2000; Ramakrishna et al., 2004; Gunatillak and Adhikari, 2003) Biodegradation

Polymer

Melting Point ( C)

Time (Months)

Mechanism

Stiffness (MPa)

PCL PGA PLLA PLDGA PLGA Polyanhydrides Poly(ortho-esters) Polyphosphazene

58 230 175 Amorphous state Amorphous state 150 80 30 100 60 to 50

24 36 6 12 24 12 1 12 n.a. n.a. n.a.

Bulk Bulk Bulk Bulk Bulk Surface Surface Surface

100 500 300 900 40 2300 30 150 40 50 25 40 4 16 n.a.

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three key advantages over bulk degradation when used as scaffold materials: (1) retention of mechanical integrity over the degradative lifetime of the device due to the maintenance of mass to volume ratio; (2) minimal toxic effects (i.e., local acidity) due to lower solubility and concentration of degradation products; and (3) significantly enhanced bone ingrowth into the porous scaffolds due to the increment in pore size as the erosion proceeds (Shastri et al., 2002). This group of polymers can also be designed to be bulk degradable by introducing a high surface to bulk ratio in the scaffold, thus, making them potential drug-delivery vehicles (Liechty et al., 2010).

10.3 BIORESORBABLE MATRICES In the past decade, nature-inspired biomaterials have emerged as a new class of materials including a natural bioresorbable polymeric matrix in combination with an inorganic phase. Indeed, the need of using biomimetic materials and the limitations of the current fabrication methods are increasingly stimulating material scientists to explore this new class of compounds, able to form a polymer matrix that might be subjected to physiological, cell-mediated resorption in vivo until its complete degradation and metabolization, without the use of chemical dissolution pratices. In fact, the chemical leaching of polymeric scaffolds is one of the possible causes of failure in vivo as the dissolution process can be too fast with respect to the new bone formation process as well as the degradation products of many polymers resulting in harmful effects, jeopardizing the regenerative cascade so that fibrous scars may form rather than healthy, organized bone tissue. The inspiration for the design and development of nature-inspired biomaterials takes place from living organisms that are able to produce natural nanocomposites showing an amazing hierarchical arrangement of their organic and inorganic components from the nanoscale to the macroscopic scale. Structural arrangement is the the result of hybrid building blocks formed upon heterogeneous nucleation of inorganic nanophases (such as carbonates and apatites) onto natural polymers and driven by several control mechanisms (Mann, 2001). During bone tissue formation, type I collagen extruded by fibroblast cells acts as a template for the nucleation of the mineral phase through a hierarchical assembly of collagen molecules into fibrils and ever thicker fibers, whereas hydroxyapatite (HA) nanonuclei nucleate onto specific positively charged sites located in the collagen molecules. Natural bioresorbable polymers like collagen, gelatin, and hyaluronic acid, and their derivatives, offer the advantage of being very similar, or even identical, to macromolecular substances which the bio environment can recognize and deal with metabolically. The problems of toxicity and the stimulation of a chronic inflammatory reaction, which are frequently provoked by many synthetic polymers, may thereby be suppressed. Furthermore, the similarity to naturally

10.3 Bioresorbable Matrices

occurring substances introduces the capability of designing biomaterials which function biologically at the molecular, rather than the macroscopic, level. Another characteristic of natural polymers is their ability to be degraded by naturally occurring enzymes which guarantees that the implant in vivo will be metabolized by physiological mechanisms. This property may, initially, appear as a disadvantage since it reduces the durability of the implant. However, in terms of regeneration, the scaffold must provide a specific function for a temporary period of time, after which the implant is expected to degrade completely and to leave space at the new formed tissue. Furthermore, since it is possible to control the degradation rate of the implanted polymer by chemical cross-linking or other chemical modifications, they offer the opportunity to design implants with a controlled lifetime (Yannas, 1996). To resist the degradative attack by enzymes, collagenases (in the case of collagen) natural polymers are subjected to chemical modifications by cross-linking which is a well-known method to slow down the implant degradation rate when in vivo. A huge number of chemical reagents, added in suitable amounts, guarantees the preservation of the implant biocompatibility and have been investigated for the stabilization of collagen. For example, 1,4-butanediol diglycidyl ether (BDDGE), a symmetric di-epoxide molecule with well-known biocompatibility at low concentration, is able to react selectively with amine and/or carboxylic functions of collagen molecules depending on the pH of the reaction environment (Zeeman et al., 1999; Nicoletti et al., 2013) creating cross-linkage between adjacent collagen chains and has demonstrated its suitability in the improvement of collagen stability to collagenase. Likewise, other substances also derived from natural sources, like Ribose and Genipin (Vicens-Zygmunt et al., 2015; Cheng et al., 2012), demonstrate their biocompatibility and suitability as chemical crosslinkers of collagen and gelatin molecules (Shankar et al., 2017). Another simple and effective method for reducing the degradation rate of collagen by naturally occurring enzymes is a self-cross-linking procedure, dehydrative cross-linking (DHT), that is, based on thermal treatment under vacuum conditions and so removing moisture below about 1 wt.% and stabilizing collagen as well as gelatin by inducing formation of interchain peptide bonds. Moreover, experimental studies have investigated the relationship between cross-link density and type and another important aspect for positive behavior of implants, their mechanical properties, revealing that cross-links, allowing the formation of intermolecular and interfibrillar bonds, can create an interconnected fibrillar material and play an essential role in influencing collagen mechanical behavior (Depalle et al., 2015; Svensson et al., 2013). Besides these aspects, the porosity of the implant makes an indispensable contribution to its performance. Pores can be easily be incorporated in collagen gels and its hybrid derivatives by freeze-drying, consisting in freezing the suspension followed by sublimation, after which the resulting pore structure is a negative replica of the network of ice crystals.

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Consequently, the control of the conditions for ice nucleation and growth can lead to a large variety of pore structures. All these functionalities, governed by several control mechanisms, to guide the formation of bone-like hybrid nanocomposites, make collagen one of the elective natural polymer for the development of bioinspired hybrid materials.

10.4 APPLICATIONS IN TISSUE ENGINEERING 10.4.1 COMPOSITE SCAFFOLDS FOR BONE During the past decade, polymer-based composites have gained an increasingly important role in the development of a new generation of bioinspired scaffolds addressing bone regeneration. Composite scaffolds can be engineered more accurately than monolithic phase materials so that the effective tissue functionality (i.e., mechanical response) can be faithfully reproduced. As nature teaches, connective hard tissues like bone are constituted by the ECM (i.e., collagen, apatite, and water) and cells (osteocytes in bone and odontoblast in dentine) which control and adapt the structure performance based on mechano-sensitivity processes (Fung, 1993). Collagen fibrils of 100 nm diameter, reinforced by platelet apatite crystals at the nanostructural level (Rho et al., 1998) are distributed in an organized manner in mineralized tissues. Since the elastic moduli of collagen and HA show an appropriate gap from 15 to 110 GPa, the optimum combination of soft and ductile collagen with the stiff and brittle apatite crystals into a synthetic composite material may confer a range of properties found in natural tissue (De Santis et al., 2000). In this context, the employment of composite materials is promising because they can achieve an ideal balance of the strength and toughness of bone tissues (Ramakrishna et al., 2001) and so preserve porosity, biocompatibility, and biodegradability, which are mandatory factors to properly exhibit an in vitro or in vivo resorption rate able to match those of the tissues at the site of implantation (Hou et al., 2003). The appropriate selection of materials may be crucial to define the basic properties of the scaffold, which will to a great extent determine the ultimate properties of the newly formed bone (Salgado et al., 2004). Fiber and particulate reinforced polymeric composites represent an “engineering response” to develop mineralized tissue analogues. Similar to natural tissue, the interaction between the polymeric matrix and a reinforcing phase may be modulated to tailor the mechanical performance in terms of material strength and anisotropy. In particular, an adequate support of the mechanical stimuli by proper mechanical features (i.e., suitable stiffness, elasticity) is necessary to ensure a progressive transfer of stress to the newly forming bone tissue as the scaffold degrades. However, a designed pore architecture remains a basic prerequisite for assuring sufficient porosity and permeability to promote efficient cell colonization, complete vascular invasion, and satisfying all the main transport demands

10.4 Applications in Tissue Engineering

(i.e., nutrient and oxygen trafficking) of the remodeling tissue (Yang et al., 2001). Hence, it is really important to maximize surface-to-volume ratio in order to improve cell colonization and fluid transport, also preserving the mechanical response of the scaffold. This conflict often leads to a compromise in the optimal design solution (Karande et al., 2004). Currently, several technologies based on templating strategies allow for improving the biomechanical performance of the scaffolds through correctly balancing the porosity and mechanical properties. In order to guarantee the desired mechanical response, the highly porous network may be further sustained mechanically through the integration of ceramic particles (Guarino et al., 2008a,b,c) or continuous fibers (Zhao, et al., 2002) to prevent the collapse of the pore architecture during load application.

10.4.1.1 Calcium phosphate particle loaded porous/nonporous composites Over the past 20 years, calcium phosphates (CS) have been largely used as bone substitute materials due to their chemical similarity to the mineral phase of bone which assures the nontoxic, biocompatible, and bioactive behavior required to reach an intimate physicochemical bond between the implant and bone, that is, osteointegration. Hence, most synthetic polymers used for scaffold manufacturing (i.e., PLGA, PLLA and PCL) have been variously combined with CS (i.e., HA, a-TCP, b-TCP particles) to develop composite scaffolds for hard tissue replacements (Guarino et al., 2013) which are able to improve the bone-bonding properties and osteoconductivity of polymer-based scaffolds. Moreover, CS may influence the degradation kinetics of the polymer phase, contributing to form low molecular weight products which are easily eliminated in the physiological environment. However, the main advantage mainly concerns the high bioactive potential induced by the employment of ceramic materials which promote the formation of mineralization sites. Several studies have demonstrated the active role of CS fillers on the underlying in vitro degradation mechanisms by the simultaneous assessment of the influence of physicochemical properties of the porous scaffolds. The addition of HA particles was found to slightly modify the pore morphology with a slight reduction of the average pore size. More interestingly, other studies on scaffold mass losses showed that the presence of apatite phases embedded in the PCL matrix drastically increases the polymer crystallinity degree promoting the formation of more densely packed crystalline phases within the composite with a lower amount of amorphous regions that are potentially more susceptible to hydrolytic attacks due to better accessibility of the ester linkage (Guarino et al., 2009). In this case, the increase in crystallinity of the polymer matrix in HA-loaded scaffolds hinders the degradation of the composites, preferentially deflecting the fluids at the polymer ceramic interface which are more susceptible to hydrolytic attack. Indeed, crystal segments are chemically more stable than amorphous segments and generally reduce water permeation into the matrix in combination with the ionic strength, temperature, and pH of the medium.

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In this context, the addition of other compounds (e.g., bioactive glass) may further concur to stabilize local environmental conditions around the polymer in order to control its degradation. Meanwhile, the use of rigid bone-like particles embedded into a polymer matrix evidently improves the mechanical properties of the polymer matrix, promoting the use of composite scaffolds as a substrate for hard tissue replacement (Khan et al., 2004; Kikuchi et al., 1997). Their combination with biodegradable natural polymers may also improve the mechanical response. The brittle behavior, directly ascribable to the ceramic phase, can be softened by the coupling with polymeric phase, with an overall increase in the composite toughness. This provides a more valid basis for mimesis of the complex behavior of natural tissues to external mechanical stimuli (Wang et al., 2002). However, in the case of porous scaffolds, the contribution of mechanical response due to the ceramic phase may be partially hindered by the presence of macro- and microstructured pores, which represent a basic requirement to induce the regeneration mechanisms in tissue engineering applications. Some authors suggest the use of additive phases (i.e., short or long fibers) into the PCL matrix to stabilize the hardening effect of the ceramic phases without altering the pore architecture required for cellular ingrowth and matrix production (Ronca et al., 2014) Alternatively, the addition of bioactive apatite-like particles with peculiar shapes (i.e., needle-like crystals) may improve the interaction with the polymer matrix, further enhancing the mechanical response in compression by up to an order of magnitude (Guarino et al., 2016). This approach may overcome relevant limitations due to a lack of homogeneity distribution of hydroxyapatite particles in the polymeric matrix which dramatically compromises both the mechanical performance as well as the bioactive potential of the composite scaffolds (Guarino et al., 2008a,b,c). Moreover, this effect is also amplified by the polymer matrix degradation, more frequent losts of bioactive particles in the time, with the creation of voids within the polymeric structure (Guarino et al., 2010). All these factors may affect the mechanical response of the composite scaffold, influencing their integrity at longer times in the in vitro culture. Hence, it is emerging the idea that pore architectures with tailored porosity need to be complemented by bone-inducing agents to stimulate the native bone-formation activity (Liu et al., 2010). In particular, recent studies have emphasized the role of osteo-inductive phases in influencing surface and interface properties (i.e., hydrophobicity, surface charge, roughness) in order to more directly address specific cell activities toward osteogenesis (Jones, 2013). For instance, Veronesi et al. have been demonstrated that chemical modification of hydroxyapatite (HA) crystals with Mg21 and CO32 ion substitution into porous scaffolds with optimized morphological features (i.e., pore size, surface curvature, interconnectivity) induce improvements in the biological response of bone-like cells, better triggering the osteogenic response under in vitro and in vivo conditions (Veronesi et al., 2015). Indeed, Mg21-doped HA is able to retain more water at its surface than stoichiometric HA, with water molecules coordinated to cations into multilayers (Bertinetti et al., 2009) thus favoring cell anchorage and growth,

10.4 Applications in Tissue Engineering

arising from the ability of calcium phosphate to absorb several important extracellular matrix proteins. Guarino and Scaglione also demonstrated that ionic substitution stimulates cells to produce mineral extracellular matter through activation of early osteogenesis mechanisms (Guarino et al., 2014a,b). Meanwhile, a porous architecture with tailored porosity features (i.e., pore size, concavity) supports the in vivo growth of mature bone with hierarchical organization. In particular, anatural osteon-like structure of bone with lamellae centripetally assembled from the wall of the macropores toward the central bone marrow cavity was imparted just after the first 6 months of implantation.

10.4.1.2 Fiber-loaded composites An interesting strategy to improve the mechanical response of porous scaffolds derived from traditional composite materials consists in the inclusion of fibers, in a continuous or chopped form, with preordered spatial organization into the polymer matrix. In the past, partially resorbable composites has been variously obtained by combining a degradable polymeric matrix with high modulus slowly resorbable fibers (glass, carbon, amides) for the development of new fixation materials in orthopedic applications. However, the long-term effect of bioinert, biostable or slowly degradable fibers is not well-documented in living tissues. Recently, totally degradable reinforced composites are emerging as an alternative strategy to design composite scaffolds for bone tissue engineering mainly due to the drastic decay of long-term problems induced after their digestion by living tissues. In this context, it is crucial to properly select biodegradable polymer phases comprising composite scaffolds which should be able to degrade and resorb at a controlled rate at the same time as the tissue regenerates. Along this perspective, even few 15 years ago, Ambrosio et al. have suggested the use of composite tubular structures composed by a polyurethane matrix with continuous fibers of PLA and poly(glycolic acid) (PGA) helically wound embedded by a filament winding technique (Ambrosio, et al., 2001). By applying the basic theory of continuous fiber-reinforced composites to the scaffold design, they developed a composite material with optimal spatial fibers organization within the polymer matrix which is able to mimic the structural organization and performance of living bone. More recently, Guarino et al. proposed a porous system by the integration of resorbable PLA fibers using the filament winding technique (Fig. 10.1) coupled with the conventional salt leaching technique (Guarino et al., 2008a,b,c). In this case, degradation preferentially occurs at the fiber polymer interface resulting in a higher rate of degradation than for either material alone. Indeed, the characteristic degradation rate of composite scaffolds is commonly higher than monolithic ones, thus, not being ideal for clinical applications such as bone fracture fixation (i.e., required strength retention from a few weeks up to several months). In this case, the helicoidal organization of continuous fibers also assures a strong mechanical interlocking between fiber and matrix, thus, minimizing the breakdown occurring at the interface. Moreover, the faster degradation of the more hydrophilic component (i.e., PLA fibers) with respect to the matrix assures the formation of

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FIGURE 10.1 Porous composite scaffolds with a PCL matrix endowed with HA or MgHA crystals. With permission of Guarino, V., Scaglione, S., Sandri, M., Alvarez-Perez,M.A., Tampieri, A., R. Quarto, et al., 2014a. MgCHA particles dispersion in porous PCL scaffolds: in vitro mineralization and in vivo bone formation. J. Tissue. Eng. Regen. Med. 8 (4), 291 303; Guarino, V., Raucci, M.G., Ronca, A., Cirillo, V, Ambrosio, L., 2014b. Multifunctional scaffolds for bone regeneration. In: Mallick K. (Ed.). Bone Substitutes Biomaterials. Woodhead Publishing Series in Biomaterials, 78; Guarino, V., Veronesi, F., Marrese, M., Giavaresi, G., Ronca, A., Sandri, M., et al., 2016. Needle-like ion-doped hydroxyapatite crystals influence osteogenic properties of PCL composite scaffolds. Biomed. Mater. 11, 015018.

preferential channels into the porous structure which may more efficiently support fluid transport, cell colonization, and tissue invasion. Most interestingly, the biological behavior of these composite scaffolds depends on the presence of a multiscale porosity with tailored characteristics in terms of pore interconnections and pore size as a consequence of the peculiar degradation mechanisms of single phases. Hence, the well-organized pore network within the scaffold may potentially control cell colonization and fluid transport through its peculiar geometry (Guarino et al., 2012) with the combined effects of the reinforcing action of PLLA fibers and the slow degradation rate of the PCL matrix contributing to the improved mechanical performance under compression of the composite

10.4 Applications in Tissue Engineering

system. Changes in fiber and matrix features, such as diameter and fiber chemistry, during degradation could provide a guide for the bone remodeling process by achieving a composite structure with morphological and structural properties which then evolve during degradation toward a softer and more porous material in the long term. An improvement of this approach lies in the development of composite materials which comprise a biodegradable matrix incorporating bioactive rigid particles which combine the reinforcement action of bioactive phases (i.e., hydroxyapatite) with the tailored degradation kinetics of biodegradable and/or resorbable polymers (Guarino et al., 2014a,b). The inclusion of bioactive particles and fiber reinforcement may further improve the mechanical properties in the presence of high pore fraction, reaching intermediate levels of mechanical response, which is the result of a synergistic effect of continuous fibers and ceramic needle-like crystals amplifying the reinforcing role of the single elements in highly porous structures. Guarino et al. demonstrated that the precipitation of calcium-deficient HA from a-TCP assures the formation of needle-like crystals with a high surface/volume ratio which offer a significant contribution to the mechanical response of the scaffold due to the large surface area available to interact with the polymer matrix (Guarino and Ambrosio, 2008) (Fig. 10.2).

FIGURE 10.2 Composite scaffolds for long bone treatment fabricated by filament winding and salt leaching. With permission form Guarino, V., Causa, F., Salerno, A., Ambrosio, L., Netti P.A., 2008a. Design and manufacture of microporous polymeric materials with hierarchal complex structure for biomedical application. Mater. Sci. Technol., 24, 1111 1117; Guarino, V., Causa, F.,Taddei, P., Di Foggia, M., Ciapetti, G., Martini, D., et al., 2008b Polylactic acid fibre-reinforced polycaprolactone scaffolds for bone tissue engineering. Biomaterials 29, 3662 3670; Guarino, V., Causa, F., Netti, P.A., Ciapetti, G., Pagani, S., Martini, D., et al., 2008c. The role of hydroxyapatite as solid signal on performance of PCL porous scaffolds for bone tissue regeneration. J. Biomed. Mater. Res B: Appl. Biomater., 86, 548 557.

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A significant increase of elastic modulus, up to an order of magnitude higher, was obtained by integrating both PLA fibers and a-TCP particles with the PCL matrix. Despite the preferential distribution of these crystals along the inner pore surfaces, a dramatic synergistic effect between the two reinforcement systems may be recognized. This may be attributable to the highly ordered interfacial interactions which occur between the organized pretensioned continuous fibers and the growing needle-like crystals during salt leaching. These interactions result in a progressive increase in the stiffness of the final composite scaffolds up to the values necessary to mimic the mechanical behavior of trabecular bone. In the light of recent approaches in bone regeneration based on composite materials, the proposed strategy certainly offers the best compromise between structural and functional properties. Porous composite scaffolds made by incorporating bioactive particles (Lei et al., 2007) do not assure an adequate improvement in the mechanical performance for bone substitution. Meanwhile, composite materials formed by integrating high-modulus PLA fibers coated with calcium phosphate (CaP) into a PCL matrix (Kothapalli et al., 2008) show mechanical properties that are significantly higher than the values reported in the literature for PLA CaP composites, but lack structural porosity. In contrast, the proposed scaffolds show an intermediate level of mechanical response, which is the result of the synergistic effect of continuous fibers and ceramic needle-like crystals amplifying the reinforcing role of the single elements in highly porous structures.

10.4.1.3 Collagen-HA hybrid nanocomposite for bone Nature-inspired material science can be considered as the last frontier in biomaterials research: indeed, natural processes of assembling and mineralization are the basis for the building of innumerable biologic structures that exhibit insuperable functional and mechanical properties. While striving to design highperformance, multifunctional composites, material scientists are now realizing that materials with superior and adaptive properties already exist in nature and are constituted by apparently “poor” substances (Fratzl and Weinkamer, 2007; Jeronimidis, 2000). The majority of biological structures are characterized by a complex organization of hybrid (i.e., organic/inorganic) elements at different size scales, from nanometers to micro- and millimeter, presenting different structural features. The remarkable properties of living tissues, such as bone, cartilage, or tendon are the result of these complex interactions taking place across all levels of organization, with each level controlling the next one. Designing responsive, self-healing structures endowed with suitable mechanical behavior and life-time is one of the major objectives of materials research: particularly the identification of the material composition and morphological features able to stimulate cells and, by matrix remodeling, yield tissues with different functionality. In the continuing quest for improved performance, which may be specified by various criteria including higher bioactivity, less weight, more strength and lower cost, currently used biomaterials frequently reach the limit of

10.4 Applications in Tissue Engineering

their usefulness. Therefore, nature-inspired biocomposites represent the solution to the need to overcome the limitations of traditional materials (i.e., bulk or porous metals, polymers, ceramics); in fact, the combination of different materials at the nanoscale makes the development of implants with optimized performances and minimized drawbacks possible. It is well-known that natural bone is a biocomposite constituted of nanosized platelet-like crystals of hydroxyapatite (HA) grown in intimate contact with collagen fibers (Lowestam and Weiner, 1989). Collagen forms about 30% in weight of the total body protein and provides strength and structural stability to various tissues, from skin to bone. In mineralized tissues the close interaction between the apatite crystals and the proteinic component provides unusual structural properties and determines specific cell activity in vivo (Addadi and Weiner, 1996). The reproduction of the complex bone chemistry and morphology is a key issue for achieving effective regenerative ability; however, this task cannot be achieved through current manufacturing technologies. A biomimetic approach has been utilized to integrate calcium phosphates into different macromolecular matrices and to obtain oriented mineral deposition (Boskey, 1998). The physicochemical properties of hydroxyapatite/collagen composites are highly affected not only by the chemical interactions between hydroxyapatite crystals and the proteinic matrix but also by the structural organization of the matrix itself. Following the natural biomineralization process, bone-like hydroxyapatite nanocrystals (HA) have been nucleated on self-assembling collagen fibers and other kinds of peptide molecules like recombinant collagen type I derived protein (Ramı´rez-Rodrı´guez et al., 2016), exploiting the ability of the negatively charged carboxylate groups of protein to bind the calcium ions of HA (Rhee et al., 2000; Tampieri et al., 2003) and to synthetize hybrid materials reproducing the physicochemical features of natural bone tissue. In a lab-scaled process, type I collagen in the form of nanofibrils can be subjected to controlled assembly in an aqueous environment by pH variation and simultaneous mineralization with apatite nanophases where the content of foreign ions can be tailored to reach biocompetent compositions. In fact, the maintenance of a disordered crystal structure allows the entrapment of ions naturally present in the physiological environment (i.e., Mg21, CO32 , Sr21, Na1, K1, and SiO44 ) into the structure of the mineral phase (Fig. 10.3). The molecular habitus of type I collagen acts as a 3D substrate for the heterogeneous nucleation of the mineral phase but also as a constraint for the growth and long-range ordering of the mineral crystals. By this process CO32 ions can be introduced to preferably occupy the phosphate site of the HA lattice (B type position) (Boskey, 2006), thus, providing the mineral phase with enhanced activity for cell adhesion and resorbability. Among the foreign ions present in biologic apatite, Mg21 have the marked property of increasing the nucleation kinetics of HA on collagen fibers, but hamper crystal growth, thus generating nanosized HA nuclei that strongly enhance the bioavailability of the mineral phase (Bigi et al., 1992) (Fig. 10.4).

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FIGURE 10.3 Scheme of collagen assembling and mineralization (left) TEM image of biomineralized collagen fiber (right).

FIGURE 10.4 Biohybrid HA collagen composites. SEM images of a biomineralized collagen - as math or single fibre - completely covered by hydroxyapatite nanoparticles.

The bone-like features of HA collagen hybrid composites are reflected in the bioresorbability at physiological pH and high surface activity, referred to as the crystal size (i.e., ranging from 30 to 50 nm long, 15 30 nm wide, and 2 10 nm thick) (Lowestam and Weiner, 1989; Eppell et al., 2001) and to the specific orientation of the apatite nuclei in respect to the long axis of collagen. The preferential growth of apatite nuclei along the c axis, as induced by the presence of particular functional chemical groups on the surface of the organic template, affects the surface polarity of the final hybrid composites and, consequently, protein adhesion and cell attachment. Mechanical testing of hybrid HA collagen composites revealed a macroscopic pseudo-plastic behavior. The elastic modulus determined on the mineralized material well reproduces the value found for trabecular bone at correspondent values of porosity (Tampieri et al., 2008). Chatzipanagis et al., by using a novel micro-electromechanical device coupled to a confocal Raman microscope that enables in situ molecular investigations of

10.4 Applications in Tissue Engineering

micro-fibers under uniaxial tensile load, were able to perform mechanical studies on biomineralized HA/collagen micro-fibers by exerting forces in the mN range. Those investigations revealed that above 30 wt.% of mineralization of the proline-related Raman band shows a pronounced response to stress, which is not observed in nonmineralized collagen. This molecular response coincides with a strong increase in the Young’s modulus from 0.5 to 6 GPa for 0 and 70 wt.% mineralized collagen, respectively. Those results are consistent with the progressive interlocking of the collagen triple-helices by apatite nanocrystals as the degree of mineralization increases, demonstrating that the biomineralization process is a suitable approach for the mechanical reinforcement of a biopolymeric matrix like collagen (Chatzipanagis et al., 2016). The HA collagen composites can assume a fibrous structure as well as high and interconnected porosity, the amount and morphology of which can be tailored by customized freeze-drying processes. After Drying, The final scaffolds exhibit high cell recognition; therefore, they can be easily resorbed in vivo whereas new tissue forms. However, to limit the enzymatic degradation possibly preventing successful cell colonization and tissue regeneration, cross-linking methods can be applied by using physical or chemical approaches addressing specific links among functional groups of collagen, thus, enabling fiber bridging and tailored stability against resorption (Nicoletti et al., 2013). Moreover, the addition of a specific cross-linking agent could potentially guarantee proper chemical and mechanical stability without compromising the implant’s biocompatibility and bioresorbability, thus overcoming the limits of the bioresorbable matrix and improving their resistance to enzymatic digestion and mechanical behavior was ensured a more suitable degradation kinetic and efficacy of the implants. The in-lab reproduction of phenomena occurring in biological processes can be considered as a conceptually new approach for nanotechnology and may pave the way toward the development of new devices with outstanding properties. On the basis of the recognition of the different requirements to regenerate cartilaginous and bony parts, such processes can be directed to graded scaffolds reproducing different histological areas in the osteochondral tissue by simply varying the degree of mineralization and the alignment of collagen fibers (Tampieri et al., 2003, 2008). In this Section it has been discussed that biological mechanisms at the molecular scale occurring into the microenvironment can be influenced by the interaction with 3D scaffolds as a function of their peculiar chemical/physical properties. This implies that bioinspired syntheses are flexible processes that can be directed to fabricate specific devices on demand that would be suitable for the tailoring of the Young’s modulus in hybrid composites.

10.4.2 COMPOSITE SCAFFOLDS FOR OSTEOCHONDRAL DEFECTS Traumatic injuries to the joints and diseases such as osteoarthritis frequently involve structural damage to the articular tissues accompanied by pain, reduced mobility, and high socio-economic costs (Nesic et al., 2006; Temenoff and

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Mikos, 2000). Moreover, the treatment of osteo-chondral (OC) defects is problematic because the tissue damage extends to subchondral bone involving two different types of tissues that display different healing pathways (Mano and Reis, 2007; Martin et al., 2007; Kon et al., 2010). Current approaches to OC treatment are usually aimed at treating the medical condition rather than curing it, consequently, giving unpredictable results: removal and/or damage of this anatomical structure leads to degenerative changes of the articular cartilage and, ultimately, to osteoarthritis (Martin et al., 2007; Buckwalter and Mankin, 1997; O’Driscoll, 1998). Hence, the reconstruction of complex joints represents a major challenge in tissue engineering (TE) where several investigators have been working to develop scaffolds mimicking the osteochondral unit that includes subchondral bone, cartilage, and calcified cartilage (intermediate region) tissues. Due to the complex hierarchical structure and to the coexistence of several architectural organizations of the natural articular tissue (Fig. 10.5), a precise

FIGURE 10.5 Schematic representation of the osteochondral tissue and its components (top) and the view of osteochondral tissue from two histological sections (bottom). With permission from Di Luca, A., Van Blitterswijk, C., Moroni, L., 2015. The osteochondral interface as a gradient tissue: From development to the fabrication of gradient scaffolds for regenerative medicine. Birth Defects Res. C Embryo Today. 105 (1), 34 52.

10.4 Applications in Tissue Engineering

definition of biochemical and structural features across a spatial volume have to be carefully carried out for tuning of mechanical and functional properties of OC tissues. In particular, in the articular site bone and cartilage tissues work in a synchronized biofunctional manner keeping their specific properties and roles. A homogeneous biomaterial is not ideal for supporting the metabolic and morphogenic activities of chondrocytes and osteoblasts which need different microenvironments to grow. Moreover, the low OC tissue vascularization hinders tissue healing and union with the original tissue (Oliveira et al., 2006; Buckwalter, 1998; Redman et al., 2005); although integration in the vertical direction (where cartilage joins with subchondral bone tissue) can be quite successful, it is much more challenging in the lateral direction due to the low healing rate. In addition, the phenotype and characteristics of the newly formed tissue are significantly affected by the biological and mechanical cues to which they are exposed where the risk of formation of a fibrous, hypertrophic tissue is high (Hunziker, 2002). The promise of TE is rooted in the fact that engineered OC grafts can recapitulate the complex architecture of articular tissue, interacting with the host environment, and remodeling themselves while providing structural and mechanical functionality (Grayson et al., 2008). The cartilage and bone components can either be generated independently in vitro and combined together (i.e., multilayer porous scaffolds) or fabricated as a single composite graft (i.e., gradient porous/ nonporous composites).

10.4.2.1 Multilayer porous scaffolds The development of bioactive multilayered scaffolds for OC regeneration has been considered a desirable strategy by several groups since single-phase scaffolds have been lacking in recreating the intrinsically distinct structural and biochemical features of the functional microenvironment (Coburn et al., 2012; Malda et al., 2005). Multilayered OC scaffolds may be realized with several different techniques, including solvent casting and particulate leaching (Scaglione et al., 2010; Suh et al., 2002), gas foaming (Yoon and Park, 2001; Nam et al., 2000), phase separation (Nam and Park, 1999; Do Kim et al., 2004), and electro-spinning (Yoshimoto et al., 2003; Polini et al., 2011). Fabricating a functional OC bilayered scaffold requires simultaneous consideration of the appropriate features of the two compartments (i.e., osseous and cartilaginous) and of the continuous interface (Gao et al., 2001; Holland et al., 2005; Jiang et al., 2010; Kon et al., 2010; Mano and Reis, 2007; Oliveira et al., 2006; Schaefer et al., 2002; Sherwood et al., 2002; Tampieri et al., 2008; Giannoni et al., 2012; Nukavarapu and Dorcemus, 2013). The macro- and microstructure of the scaffold may significantly influence cell migration and viability, besides matrix deposition and blood vessel invasion, if required (Karageorgiou and Kaplan, 2005). A large fraction of pore volume represents a key requirement for effective nutrient and metabolite mass transport which are required for cell functions and for in vivo neo-tissue deposition. Moreover, the presence of pores

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accelerates the biodegradability of the graft, ultimately favoring a complete conversion of the scaffold into mature tissue (Bohner et al., 2011). Several studies reported different bilayered scaffolds realized with a wide range of pore diameters; some employed smaller pores in the chondral layer and larger pores in the osseous layer (Schek et al., 2004; Im et al., 2010; Qu et al., 2011; Liao et al., 2007), whereas others used the same sized pores throughout the whole scaffold (Shao et al., 2006; Chen et al., 2011; Jiang et al., 2007). Although pore size plays a significant role in biological delivery and tissue regeneration, it is unknown whether different pore sizes in the two layers may affect OC defect repair. However, to the best of our knowledge, no specific study has examined the effect of pore size on the in vivo efficacy of repairing osteochondral defects by using bilayered scaffolds, although some reviews variously describe the use of similar devices in different tissue iengineering applications (McMahon et al., 2008; Karageorgiou and Kaplan, 2005). Likewise, chemical composition is responsible for selective cellular differentiation (Lynn et al., 2010; Scaglione et al., 2010). The incorporation of osteoconductive materials (e.g., calcium-phosphate and Ca-P derivates) into bilayered polymeric scaffolds has been widely adopted to foster subchondral bone regeneration as well as overcoming the poor mechanical properties of biopolymers in the bony phase (Mohan et al., 2011; Giannoni et al., 2012). Some commercial bilayers scaffolds (e.g., Trufit and MaioRegen) have been clinically applied in acellular strategies (Kon et al., 2011; Bedi et al., 2010; Joshi et al., 2012). Kon et al. (Kon et al., 2011) showed that multilayer collagen/nanohydroxyapatite scaffolds promote bone and cartilage restoration. Recently, novel bilayer scaffolds composed of a silk layer and a silk-nanoCaP layer for OC regeneration have been realized by combining the outstanding osteoconductivity of ceramics with the excellent elasticity and biocompatibility of silk biopolymers (Yan et al., 2015). Chitosan-alginate composites have been also studied as useful biomaterials in cartilage and bone tissues and have shown promising results (Han et al., 2010). In 2007, Swieszkowski designed biphasic scaffolds composed of polycaprolactone (PCL) for the cartilage phase and PCL-TCP for the bony one. The geometric design process of the biphasic scaffold was performed by CAD computer systems. Both phases were separately fabricated, seeded with bone marrow-derived mesenchymal cells (BMSCs), and, finally, cultured in chondrogenic and osteogenic media for cartilage and bone regeneration, respectively. Then, the two phases were integrated into a single construct by using fibrin glue and implanted in critical size defects created in the medial condyle of a rabbit model. The quantification of tissue regeneration demonstrated the potential of PCL/PCL-TCP in promoting bone healing and the fast degradability of PCL in the cartilage phase (Swieszkowski et al., 2007). Chemically heterogeneous biphasic scaffolds have been also realized for goat femoral head articular cartilage repair (Ding et al., 2013); a polylactic acid-coated polyglycolic acid (PGA/PLA) material was realized for cartilage regeneration,

10.4 Applications in Tissue Engineering

while a poly-ε-caprolactone/hydroxyapatite (PCL/HA) was designed and realized for the regeneration of the femoral head. The PCL/HA scaffold was fabricated by FDM according to the 3D data achieved from the goat proximal femoral condyle and it was designed to contain regular 3D interconnecting micro-channels (200 400 μm in pore size) with a porosity of about 50%. Homogenous cartilage was regenerated on the surface layer of regenerated femoral heads and typical trabecular bony tissue was formed in the bony layer. Moreover, chondral and bony layers showed satisfactory integration and a typical osteochondral interface where cartilage tissue, immature calcified tissue, transitional trabecular bone, and hypertrophic chondrocytes were observed. These structures were present only at the interface area, showing the tissue-specific regeneration induced by different microenvironments (Ding et al., 2013). In 2002, Sherwood et al. developed a heterogeneous, osteochondral scaffold using a 3D printing process. The material composition, porosity, and macroarchitecture were varied throughout the scaffold structure in order to reproduce the hierarchical complexity of the OC tissue. In particular, the upper cartilage region was composed of d,l-PLGA/l-PLA, with macroscopic staggered channels to facilitate homogenous cell seeding. A similar pore conformation was highly adopted in cartilage substitutes with the aim to limit pore interconnection, finally restrictions for in vivo vascularization and nonisotropic mechanics (Giannoni et al., 2012; Scaglione et al., 2014). On the other hand, the lower bone portion was made by a l-PLGA/TCP composite and designed to maximize bone ingrowth while maintaining fairly strong mechanical support. The bone layer was 55% porous while the cartilage layer was 90% to reproduce the native mechanical properties of the two tissues. Due to the different porosity percentages, the tensile strength of the bone layer was similar to the fresh cancellous human bone, making these grafts suitable for in vivo applications, including full joint replacement (Sherwood et al., 2002). The transition region between these two sections was made by a material layer characterized by a porosity gradient in order to prevent delamination. The in vitro cell seeding showed that chondrocytes preferentially seeded into the cartilage portion of the device while cell attachment to the bone region was minimal. A bi-layered osteochondral scaffold, including a PVA/Gel/V layer for the cartilage and a n-HA/PA6 layer for the subchondral bone, was proposed by Li et al. (Li et al., 2015) to evaluate the potential of engineered osteochondral grafts in repairing articular osteochondral defects in rabbits. The two different layers of the scaffolds were seeded with allogenic bone marrow-derived stem cells (BMSCs) which were chondrogenically and osteogenically induced, respectively. The critically sized osteochondral defects were created in the knees of adult rabbits. Interestingly, growth factors may be also introduced into bilayered scaffolds to enhance articular tissue repair (Reyes R. et al., TERM, 2014; Guo et al., Biomat, 2009; Coluccino et al., 2016). Typically, recombinant types of growth factors are delivered in the culture medium to regulate cellular processes.

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One option to enhance the in vitro and in vivo efficacy of growth factors is to incorporate them into polymeric biomaterials in order to maintain their stability and to control their release kinetics. Growth factor delivery can also be accomplished in the form of microparticles, nanoparticles, or related material formats incorporated into the scaffold or via growth factor-secreting naturally or genetically engineered cells harbored within the scaffolds (Wang et al., 2009a,b). Specific biochemical signals (e.g., bone morphogenetic protein 2, insulin-like growth factor, transforming growth factor beta) may, thus, be introduced in the two layers to foster the cellular differentiation toward the chondrogenic and osteogenic lineage, respectively. Spatially controlled dual-growth factors or gene-release systems have been developed and the repair of both cartilage and subchondral bone layers were observed (Chen et al., 2011). However, despite positive preliminary results on the material-induced selective cellular differentiation and ECM deposition, multiphasic laminated/layered OC grafts still display some drawbacks as the possible structural delamination of the different layers in vivo due to inherent discontinuities across dissimilar materials. Recent studies focused on the formation of hierarchical structures allowing bonecartilage interfaces similar to the native OC. One of the most promising strategies is the development of heterogeneous scaffolds obtained by the combination of distinct, but integrated, layers corresponding to the cartilage and the bone regions (Panseri, Russo et al., 2012b; Cao et al., 2003; Tampieri et al., 2011). Such designs are based on the recognition of the different requirements to regenerate cartilage and bone parts of an osteochondral defect and preventing the risk of delamination of different components (Tampieri et al., 2011; Chen et al., 2011). For any designed scaffold system the integration between the cartilaginous and the osseous parts is critical; the transition between the osseous and the cartilaginous compartments occurs suddenly, being a “discrete” interface created by the joining of two distinct materials. Moreover, the introduction of glues or interfacing materials between porous scaffold compartments presents the risk of limiting nutrient diffusion across the interface, hindering the effectiveness of the implant (Harley et al., 2010). The persistent incidence of delamination of the two layers significantly reduces the overall in vivo efficacy of these engineered grafts for treating osteochondral injuries. Therefore, a new generation of osteochondral scaffolds has attempted to overcome these limitations by predesigning the interface between the two distinct layers (Grayson et al., 2008; Harley et al., 2010; Jiang et al., 2010). Basically, this conceptual approach consists of ensuring a continuous and stable interface between the two layers forming the osteochondral graft. Following this approach, Giannoni et al. presented a prototype of bilayer monolithic osteochondral scaffold (Fig. 10.6) with the aim to obtain a chemicophysical continuum between the two different compartments of the osteochondral scaffold (Giannoni et al., 2012).

10.4 Applications in Tissue Engineering

FIGURE 10.6 In vivo tissue formation within an osteochondral graft. MSCs deposited newly formed bone (black arrows) in the bony layer around the HA granules, while articular chondrocytes differentiated and deposited cartilage-like tissue onto the chondral layer of the graft (empty arrows) which was positive under toluidine blue staining. H&E (A, C) and toluidine blue (B, D) staining. (C, D) Enlarged views of the bone and cartilage matrixes newly deposited, where typical cells of mature tissues are visible. Bars 5 100 mm (A, B); 50 mm (C, D) (Giannoni et al., 2012).

Prototypes were prepared by combining solvent casting and particulate leaching methods to reach a high percentage of total porosity within the entire osteochondral graft. Composition and microstructure of the scaffold was properly assessed: the bony layer was designed to offer a bioactive mineralized surface to the cells so that they could differentiate toward the osteogenic lineage and deposit bone matrix within the available volume of the graft. Since the formation of bony tissues demands effective in vivo blood vessel colonization within the implant, a fully interconnected pore structure has been designed to favor the vascularization process in the bony layer. In the chondral layer they aimed to design a porous structure able to maximize the available volume for cartilage matrix deposition, besides reducing the interconnections for limited vascularization. Improvement in the mechanical properties or stability of multilayer scaffolds and optimization and incorporation of bioactive factors in scaffolds still represent big challenges (Chen et al., 2011; Moutos et al., 2007; Kon et al., 2011). Other problems are related to achieving a good interface between the different layers (Yang et al., 2009).

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10.4.2.2 Gradient porous/nonporous composites Articular cartilage and subchondral bone constitute a complex tissue structure that involves a progressive gradient of material and physiological properties (Mohan et al., 2011). The native osteochondral interface consists of a calcified cartilage layer (Yang et al., 2009) that is flanked by an undulating tidemark and an even more irregular cement line adjacent to the subchondral bone (Wang et al., 2009a,b). Calcified cartilage becomes established at the edges of a “permanent” epiphyseal bone layer (i.e., proximal reserve zone and articular cartilage hypertrophic zone) and the tidemark serves as a barrier to vascular invasion and the calcification of hyaline cartilage (Hoemann et al., 2012). This structure is semi-permeable and permits the passage of small molecules (,500 Da) from the subchondral bone to the articular cartilage layer (Arkill and Winlove, 2008). Moreover, this special barrier is essential for maintaining the integrity of repaired cartilage over time and preventing osseous upgrowth into full-thickness defects (Hunziker et al., 2001). Thus, a number of studies have developed strategies to facilitate osteochondral regeneration and tissue integration at the osteochondral interface (Ding et al., 2013; Khanarian et al., 2012; Dormer et al., 2011). Based on the native interface structure, the ideal cartilage-to-bone interface scaffold would support chondrocyte viability and promote a calcified cartilage matrix formation with appropriate physical properties. Thus, designing and fabricating such a scaffold involving a gradient interfacial structure is a prerequisite for the success of osteochondral tissue engineering (Fig. 10.7). The use of gradient scaffolds with already differentiated or progenitor cells has been recently proposed. Some of these approaches have also been translated in clinical trials without satisfactory results. Further efforts in the development of constructs, will be needed in the near future to guarantee a functional regeneration of the OC interface by gradients more closely resembling its native environment (Di Luca et al., 2015). Some graded scaffold prototypes have been realized by adopting different manufacturing procedures: electrospinning (He et al., 2014), gas foaming combined with particulate leaching (Salerno et al., 2012), centrifugation-based techniques (Oh et al., 2007), thermally induced phase separation (Mannella et al., 2015), and rapid prototyping (RP) technology (Melchels et al., 2010). Among them, gas foaming and particulate leaching do not allow to accurately predetermine the final internal structure of the porous grafts, while RP techniques offer a spatial precision at the micro-scale. On the other hand, electrospinning and centrifugation procedures allow producing fibrous scaffolds that recapitulate the protein network of native ECM. While electrospinning is mainly focused on the realization of nanofibrous films, thermally induced phase separation and centrifugation methods offer significant advantages for the fabrication of 3D scaffolds, which may be functionally graded for porosity and composition (Mannella et al., 2015; Oh et al., 2007). From a chemical point of view, the logical criterion for the selection of natural polymer candidates passes through the components of the ECM. For the

FIGURE 10.7 On the left side, schematic representation of the osteochondral tissue. In the central side, directions of its gradients. The stiffness gradient continues in the subchondral bone, whereas the nutrient gradient stops in the radial zone. From Mansfield, J.C., Bell, J.S., Winlove, C.P., 2015. The micromechanics of the superficial zone of articular cartilage, Osteoarthritis and Cartilage, 23 (10), 1806 1816. On the right side, SEM images representing the development of the morphological gradient for PCL and the chemico-morphological one for COL/HA scaffolds. Line bars: 1mm. From Marrella, A., Aiello, M., Quarto, R., Scaglione, S., 2016. Chemical and morphological gradient scaffolds to mimic hierarchically complex tissues: from theoretical modeling to their fabrication. Biotechnol. Bioeng. 113 (10), 2286 2297.

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bone compartment, collagen type I combined with calcium phosphate derivates was used to produce scaffolds or to cover the underlying polymer by coating the pore surface to mask the synthetic polymer and present an ECM-like environment for cell attachment, growth, and differentiation (Ekin et al., 2016; Marrella et al., 2016; Tampieri et al., 2008). For the regeneration of the cartilage portion, collagen type II glycosaminoglycans, or hyaluronic acid have been used (Maturavongsadit et al., 2016). For both bone and hyaline cartilage, the most common synthetic polymers used are poly(lactic acid), poly(glycolic acid), and polycaprolactone (Cao et al., 2003; Ekin et al., 2016; Giannoni et al., 2012). In 2016, Marrella et al. (2016) described chemical and morphological gradient scaffolds based on collagen and ceramics (i.e., HA powders), combining centrifugation and freeze-drying techniques with the aim to recapitulate the hierarchically complex OC tissue. In 2009, Liu et al. realized OC scaffolds composed of three parts, each with different features and materials to mimic bone, cartilage and the interface between them, respectively. The manufacturing approach consisted of two highly porous layers with larger pores for bone and cartilage repair and compact separate layer with low porosity and smaller pores for the interface. PLGA-TCP and PLGANaCl slurries were prepared and extruded in a room at low temperature to fabricate the osteochondral scaffold as defined in the model. The gradient scaffold was cross-linked and combined with collagen sponges. Even if PLGA is a widely used polymer for cartilage and bone tissue engineering (Sherwood et al., 2002; Park, 1995), its surface chemistry does not promote cell adhesion (Lee et al., 2002) so collagen sponges were introduced to promote cell adhesion. On the other hand, TCP was added into the bony layer to improve the biocompatibility, mechanical properties, and the osteoinductive capacity in vivo (Yuan et al., 2001; Qu et al., 2002). These OC scaffolds combined with progenitor cells were implanted in the knee joints of five New Zealand white rabbits. At 6 weeks postoperation, the samples contained cartilage-like tissues and some bone-like tissues in the bone region, proving that this system is promising to fabricate scaffolds with heterogeneous materials and gradient hierarchical porous structures useful to mimic osteochondral tissue (Liu et al., 2009). The combination of human mesenchymal stem cells (hMSCs) with additive manufacturing technologies was also adopted to fabricate scaffolds with a gradient in pore shape (Di Luca et al., 2016). The design of scaffolds with structural gradients, such as pore shape, may stimulate the differentiation of hMSCs and lead to improved tissue regeneration strategies. Interestingly, integrated scaffolds with embedded gradients of growth factors at the interface have been proposed for triggering simultaneous tissue formation and have a synergistic effect on tissue regeneration (Dormer et al., 2011). Using gradients with multiple bioactive factors, multiple tissue regeneration can be addressed via a single-cell source, where, for example, stem cells can be differentiated along different lineages within the same constructs.

10.4 Applications in Tissue Engineering

From the perspective of OC tissue engineering, an in vitro study reported that only co-cultures with chondrocytes (as opposed to fibroblasts or osteoblasts) was successful at promoting osteogenic differentiation of mesenchymal stem cells in a selective manner (Gerstenfeld et al., 2002) indicating the importance of simultaneous triggering of osteo- and chondro-induction for osteochondral tissue regeneration.

10.4.2.3 Magnetic bioinspired hybrid nanocomposites for osteochondral tissue The use of biomimetic scaffolds can be an effective approach for bone tissue regeneration; however, the individual patients’ metabolism plays an important role in the regulation of the kinetics and extent of new bone formation. Indeed, metabolic diseases and degenerative conditions induced by aging can seriously penalize new bone formation and fracture healing. In consideration of the everincreasing aging of the world population, the occurrence of degenerative diseases is expected to steadily rise in the next few decades, thus, new therapeutic approaches are required to boost tissue regeneration in patients with reduced endogenous regenerative potential. Tissue engineering approaches and the use of drug delivery systems able to deliver growth factors are two main approaches for enhancing tissue regeneration. Particularly, much effort is being dedicated to the development of scaffolds with the ability of controlled biochemical stimulation and temporo-spatially controlled molecular delivery. In this respect, recent advances in material science suggest that the use of weak magnetic fields is appealing as remote signaling for noninvasive controlling of biomedical devices in vivo (Tampieri et al., 2011). The use of magnetic materials in nanomedicine is, thus, raising steadily growing interest due to the numerous possible applications including cancer therapy by hyperthermia, magnetic resonance imaging, and other diagnostic approaches based on the guiding of such particles to specific targeted areas in vivo and their use as nanoprobes (Pankhurst et al., 2003; Meng et al., 2013; Panseri et al., 2013). This represents a promising tool for new personalized therapies. A serious drawback in the use of magnetic materials in nanomedicine is their long-term cytotoxicity (Lewinski et al., 2008; Singh et al., 2010). Intense effort has, therefore, been dedicated to engineering SPIONs (superparamagnetic iron oxide nanoparticles) with surface treatments to achieve enhanced biocompatibility and affinity with cells. A significant advance can be the development of magnetic materials with intrinsic biocompatibility and bioresorbability. In this respect, it has been shown that the doping of the apatite lattice with Fe21/Fe31 ions in specific calcium sites yields a new phase with intrinsic paramagnetic behavior (FeHA) (Tampieri et al., 2012). FeHA is characterized by excellent in vitro biocompatibility, as confirmed by in vitro response of FeHA nanoparticles which do not reduce cell viability and at the same time enhance cell proliferation compared to undoped HA particles. Moreover, a pilot animal study of bone repair (a rabbit critical bone defect model) demonstrated the in vivo biocompatibility and biodegradability of FeHA

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(Panseri, Cunha et al., 2012a). The achievement of biocompatible nanobiomaterials with magnetic properties opens new perspectives in regenerative medicine. Particularly, the development of bone scaffolds with the ability of remote magnetic activation is now an emerging concept in regenerative medicine since it has been demonstrated that weak magnetic or pulsed electromagnetic fields are effective in promoting bone fracture healing and bone ingrowth in various animal models (Clavijo-Jordan et al., 2012; Assiotis et al., 2012). In this respect FeHA can also be synthesized by suitable modification of the biomineralization process (see above Par. 4.1.3) to induce heterogeneous nucleation of FeHA nanophase on type I collagen (Tampieri et al., 2014) generating biomimetic hybrid scaffolds with superparamagnetic ability. The flexibility of the synthesis process allowed for tailoring the extent of mineralization of the superparamagnetic hybrid composites; consequently, 3D hybrid constructs with features mimicking different mineralized and nonmineralized tissues could be developed and merged into morphologically, compositionally, and magnetically graded implants able to mimic multifunctional anatomical regions (Fig. 10.8). Another key requirement to be fulfilled to achieve fast and effective osteointegration is the early and stable fixation of the scaffold; the implant fixation represents a major clinical problem due to the difficulties in obtaining a stable interface between host bone and the scaffold. Magnetic forces can also aid in improving the fixation of scaffolds in the implant site; a superparamagnetic implant can be stabilized by external magnetic bandages; and this can help

FIGURE 10.8 (A) Graded scaffold for osteochondral regeneration based on hybrid magnetic composites. (B) SEM details of the three integrated layers, subchondral bone (top), mineralized cartilage (middle), and articular cartilage (bottom). (C) Implant in an osteochondral defect. (D) Magnetic external fixator.

10.5 Conclusions

to minimize the use of external fixating media and possibly reduce hospitalization time and the patient’s pain. The possibility to fix a magnetic scaffold was preliminary investigated. For this goals, biomimetic and bioresorbable scaffolds were associated with magnetic nanoparticles. The magnetic scaffolds were placed within a trabecular size defect in a rabbit model in contact with cylindrical NdFeB permanent magnets producing a static magnetic fields of 1.2T (Panseri et al., 2013; Russo et al., 2016). The results (Fig. 10.9) demonstrate that magnetic forces influence the orientation of scaffold architecture in vivo and induce the formation of a well-ordered tissue architecture that may shorten the tissue remodeling time and accelerate the formation of mature bone. In conclusion, the use of magnetic scaffold in conjunction with a permanent magnet represent a suitable route for faster tissue regeneration. The presence of a mineral phase with bone-like features and the ability to be activated by remote magnetic signal make this new biomaterial very promising to boost regenerative processes in extended bone and osteochondral regions, even in patients with reduced endogenous regenerative potential.

10.5 CONCLUSIONS In this chapter, we described more succesfully strategies currently used to design composite scaffolds for bone and osteochondral defect regeneration. All of proposed approaches confirm that an optimum combination of properties exerted by biorecognized matrices (biodegradable or bioresorbable) and by bioactive reinforcement systems with different shape factors (i.e., granules, needles, and short or long fibers) concur to realize biologically adaptive structures with an efficient interface with living tissues, occurring on the nano-, micro- and macroscales. A future challenge lies in the reappraisal of the traditional concepts of composite materials in the light of recent advances in nanotechnology. Successful implementation of nanoinspired strategies are enabling the creation and manipulation of local microenvironments, as will productive investigation of the relevant biological mechanisms, including tissue morphogenesis, differentiation, and maintenance. For instance, the recent development of iron-substituted hydroxyapatite nanophases with excellent biocompatibility and intrinsic magnetic properties or electroactive polymers (EAP) with intrinsically conductive properties are demonstrating their ability to be activated by remote electro-magnetic signaling, thus representing a new switching tool for the development of a multifunctional platform generating smart biodevices for various applications in regenerative medicine and theragnostic applications. This new material to overcome the limitations of toxic metallic nanoparticles currently used in nanomedicine is promising for the future establishment of new, more effective, and personalized approaches for bone regeneration. Moreover, the possibility of supporting bone regeneration by external magnetic stimulation will be a key issue to find new therapeutic solutions

325

FIGURE 10.9 Schematic summary of the in vivo experiment to investigate the effect of the combination of magnetic hybrid bone-like material with permanent magnet in femoral condyle. Optical images of retrieved bone sections at 4 (B) and 12 (A, C) weeks of follow-up (scale bar 100 μm). From Russo, A., Bianchi, M., Sartori, M., Parilli, A., Panseri, S., Ortolani, A., et al., 2016. Magnetic forces and magnetized biomaterials provide dynamic flux information during bone regeneration. J. Mater. Sci. Mater. Med. 27, 51. Available from: https://doi.org/10.1007/s10856-015-5659-0.

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Panseri, S., Cunha, C., D’Alessandro, T., Sandri, M., Giavaresi, G., Marcacci, M., et al., 2012a. Intrinsically superparamagnetic Fe-hydroxyapatite nanoparticles positively influence osteoblast-like cell behavior. J. Nanobiotechnol. 10, 32 42. Panseri, S., Russo, A., Cunha, C., Bondi, A., Di Martino, A., Patella, S., et al., 2012b. Osteochondral tissue engineering approaches for articular cartilage and subchondral bone regeneration. Knee. Surg. Sports Traumatol. Arthrosc. 20, 1182 1191. Panseri, S., Russo, A., Sartori, M., Giavaresi, G., Sandri, M., Fini, M., et al., 2013. Modifying bone scaffold architecture in vivo with permanent magnets to facilitate fixation of magnetic scaffolds. Bone 56, 432 439. Park, T.G., 1995. Degradation of poly (lactic-co-glycolic acid) microspheres: effect of copolymer composition. Biomaterials 16 (15), 1123 1130. Peter, M., Binulal, N.S., Nair, S.V., Selvamurugan, N., Tamura, H., Jayakumar, R., 2010. Novel biodegradable chitosan gelatin/nano-bioactive glass ceramic composite scaffolds for alveolar bone tissue engineering. Chem. Eng. J. 158 (2), 353 361. Polini, A., Pisignano, D., Parodi, M., Quarto, R., Scaglione, S., 2011. Osteoinduction of human mesenchymal stem cells by bioactive composite scaffolds without supplemental osteogenic growth factors. PLoS One 6 (10), e26211. Qu, D., Li, J., Li, Y., Khadka, A., Zuo, Y., Wang, H., et al., 2011. Ectopic osteochondral formation of biomimetic porous PVA-n-HA/ PA6 bilayered scaffold and BMSCs construct in rabbit. J. Biomed. Mater. Res. B. Appl. Biomater. 96, 9 15. Qu, S., Leng, Y., Guo, X., Cheng, J., Chen, W., Yang, Z., et al., 2002. Histological and ultrastructural analysis of heterotopic osteogenesis in porous calcium phosphate ceramics. J. Mater. Sci. Lett. 21 (2), 153 155. Ramakrishna, S., Majer, J., Wintermantel, E., Leong, K.W., 2001. Biomedical applications of polymer-composite materials: a review. Compos. Sci. Technol. 61, 1189 1224. Ramakrishna S, Huang ZM, Kumar GV, Batchelor AW, Mayer J. An Introduction to Biocomposites. London: Imperial College Press; 2004 (. 36). Ramı´rez-Rodrı´guez, G.B., Delgado-Lo´pez, J.M., Iafisco, M., Montesi, M., Sandri, M., Sprio, S., et al., 2016. Biomimetic mineralization of recombinant collagen type I derived protein to obtain hybrid matrices for bone regeneration. J. Struct. Biol. 196, 138 146. Raucci, M.G., Guarino, V., Ambrosi, L., 2012. Biomimetic strategies for bone repair and regeneration. J. Funct. Biomateri.s (Open Access) 3, 688 705. Redman, S.N., Oldfield, S.F., Archer, C.W., 2005. Current strategies for articular cartilage repair. Eur. Cell. Mater. 9 (23 32), 23 32. Reyes, R., Delgado, A., Sa´nchez, E., Ferna´ndez, A., Herna´ndez, A., Evora, C., 2014. Repair of an osteochondral defect by sustained delivery of BMP 2 or TGFβ1 from a bilayered alginate PLGA scaffold. J. Tissue Eng. Regen. Med. 8 (7), 521 533. Rezwan, K., Chen, Q.Z., Blaker, J.J., Boccaccini, A.R., 2006. B iodegradable and bioactive porous polymer/inorganic composite scaffolds for bone tissue engineering. Biomaterials 27, 3413 3431. Rhee, S.H., Do Lee, J., Tanaka, J., 2000. Nucleation of hydroxyapatite crystal though chemical interaction with collagen. J. Am. Ceram. Soc. 83, 2890 2892. Rho, J.Y., Kuhn-Spearing, L., Zioupos, P., 1998. Mechanical properties and the hierarchical structure of bone. Med. Eng. Phys. 2, 92 102. Ronca, A., Guarino, V., Raucci, M.G., Salamanna, F., Martini, L., Fini, M., et al., 2014. Large defect tailored composite scaffolds for in vivo bone regeneration. J. Biomater. Applicat. 29 (5), 715 727.

References

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Sunderlacruz, S., Kaplan, D.L., 2009. Stem cell- and scaffold-based tissue engineering approaches to osteochondral regenerative medicine. Semin. Cell. Dev. Biol. 20 (6), 646 655. Svensson, R.B., Mulder, H., Kovanen, V., Magnusson, S.P., 2013. Fracture mechanics of collagen fibrils: influence of natural cross-links. Biophys. J. 104, 2476 2484. Swieszkowski, W., Tuan, B.H., Kurzydlowski, K.J., Hutmacher, D.W., 2007. Repair and regeneration of osteochondral defects in the articular joints. Biomol. Eng. 24 (5), 489 495. Tampieri, A., Celotti, G., Landi, E., Sandri, M., Roveri, N., Falini, G., 2003. Biologically inspired synthesis of bone-like composite: self-assembled collagenfibers/hydroxyapatite nanocrystals. J. Biomed. Mater. Res. A. 67, 618 625. Tampieri, A., Sandri, M., Landi, E., Pressato, D., Francioli, S., Quarto, R., et al., 2008. Design of graded biomimetic osteochondral composite scaffolds. Biomaterials 29 (26), 3539 3546. Tampieri, A., Sprio, S., Sandri, M., Valentini, F., 2011. Mimicking natural biomineralization processes: a new tool for osteochondral scaffold development. Trends Biotechnol. 29, 526 535. Tampieri, A., D’Alessandro, T., Sandri, M., Sprio, S., Landi, E., Bertinetti, L., et al., 2012. Intrinsic magnetism and hyperthermia in bioactive Fe-doped hydroxyapatite. Acta Biomater. 8 (2), 843 851. Tampieri, A., Iafisco, M., Sandri, M., Panseri, S., Cunha, C., Sprio, S., et al., 2014. Magnetic bio-inspired hybrid nanostructured collagen-hydroxyapatite scaffolds supporting cell proliferation and tuning regenerative process. ACS Appl. Mater. Interf. 6 (18), 15697 15707. Temenoff, J.S., Mikos, A.G., 2000. Review: tissue engineering for regeneration of articular cartilage. Biomaterials 21 (5), 431 440. Veronesi, F., Giavaresi, G., Guarino, V., Raucci, M.G., Sandri, M., Tampieri, A., et al., 2015. Bioactivity and bone healing properties of biomimetic porous composite scaffold: in vitro and in vivo studies. J. Biomed. Mater. Res. A. 103 (9), 2932 2941. Vert, M., 2005. Aliphatic polyesters: great degradable polymers that cannot do everything. Biomacromolecules 6, 538 546. Vicens-Zygmunt, V., Estany, S., Colom, A., Montes-Worboys, A., Machahua, C., Sanabria, A.J., et al., 2015. Fibroblast viability and phenotypic changes within glycated stiffened three-dimensional collagen matrices. Respir. Res. 16 (82), 1 15. Wang, X., Li, Y., Wei, J., De Groot, K., 2002. Development of biomimetic nano-hydroxyapatite/poly(hexamethylene adipamide) composites. Biomaterials 23, 4787 4791. Wang, F., Ying, Z., Duan, X., Tan, H., Yang, B., Guo, L., et al., 2009a. Histomorphometric analysis of adult articular calcified cartilage zone. J. Struct. Biol. 168, 359 365. Wang, X., Wenk, E., Zhang, X., Meinel, L., Vunjak-Novakovic, G., Kaplan, D., 2009b. Growth factor gradients via microsphere delivery in biopolymer scaffolds for osteochondral tissue engineering. J. Control Release 134, 81 90. Yan, L.P., Oliveira, J.M., Oliveira, A.L., Reis, R.L., 2015. In vitro evaluation of the biological performance of macro/micro-porous silk fibroin and silk-nano calcium phosphate scaffolds. J. Biomed. Mater. Res. B. Appl. Biomater. 103 (4), 888 898. Yang, S., Leong, K.F., Du, Z., Chua, C.K., 2001. The design of scaffolds for use in tissue engineering. Part I. Traditional factors. Tissue Eng. 7, 679 689. Yang, P.J., Temenoff, J.S., 2009. Engineering Orthopedic Tissue Interfaces. Tissue Eng. Part B: Rev. 15 (2), 127 141.

Further Reading

Yannas, I.V., 1996. Biomaterial Science, An Introduction toMaterials in Medicine, In: Buddy D., Ratner B.D., Hoffman A.S., Schoen F.J., Lemons J.E. (Eds.), par 2.7 Natural Materials. Yoon, J.J., Park, T.G., 2001. Degradation behaviors of biodegradable macroporous scaffolds prepared by gas foaming of effervescent salts. J. Biomed. Mater. Res. 55 (3), 401 408. Yoshimoto, H., Shin, Y.M., Terai, H., Vacanti, J.P., 2003. A biodegradable nanofiber scaffold by electrospinning and its potential for bone tissue engineering. Biomaterials. 24 (12), 2077 2082. Yuan, H., De Bruijn, J.D., Zhang, X., Van Blitterswijk, C.A., De Groot, K., 2001. Use of an osteoinductive biomaterial as a bone morphogenetic protein carrier. J. Mater. Sci. Mater. Med. 12 (9), 761 766. Zeeman, R., Dijkstra, P.J., Van Wachem, P.B., Van Luyn, M.J.A., Hendriks, M., Cahalan, P.T., et al., 1999. Crosslinking and modification of dermal sheep collagen using 1,4butanediol diglycidyl ether. J. Biomed. Mater. Res. 46, 424 433. Zhao, F., Yin, Y., Lu, W.W., Leong, J.C., Zhang, W., Zhang, J., et al., 2002. Preparation and histological evaluation of biomimetic three-dimensional hydroxyapatite/chitosangelatin network composite scaffolds. Biomaterials. 23, 3227 3234.

FURTHER READING Kon, E., Delcogliano, M., Filardo, G., Altadonna, G., Marcacci, M., 2009. Novel nanocomposite multi-layered biomaterial for the treatment of multifocal degenerative cartilage lesions. Knee Surg. Sports Traumatol. Arthrosc. 17 (11), 1312 1315.

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Plasma treated and untreated thermoplastic biopolymers/biocomposites in tissue engineering and biodegradable implants

11

Binay Bhushan1 and Rakesh Kumar2 1

Department of Physics, Birla Institute of Technology, Mesra, Patna Campus, India 2 Department of Biotechnology, Central University of South Bihar, Gaya, India

11.1 INTRODUCTION Poly(lactic acid) or polylactide (PLA) and polyhydroxyalkanoates (PHAs) fall in the category of thermoplastic biopolymers obtained from renewable resources (Farah et al., 2016; Gigli et al., 2016). Both these polymers have proven potential either to replace conventional petrochemical-based polymers for industrial applications or as leading biomaterials for numerous applications in medicine (Lopes et al., 2012; Drumright et al., 2000). PLA and PHAs have been employed to develop devices in the form of sutures, cardiovascular patches and orthopedic pins, tissue engineered nerves, tendon and articular cartilages, repair patches, slings, adhesion barriers, stents, wound dressings, and carriers for controlled drug release (Hazer et al., 2012; Valappil et al., 2006; Jaffredo and Guillaume, 2014). PLA has better thermal processability compared to PHAs. PLA can be easily processed on standard equipment to yield molded parts, films, or fibers (Handbook of Polymer Applications in Medicine and Medical Devices, 2014; Garlotta, 2002). PHAs can also be processed by injection molding, film extrusion, blow molding, thermoforming, fiber spinning, and film forming (Auras et al., 2004). PLA hydrolyzes to its constituent α-hydroxy acid when implanted in living organisms, including the human body. It is then incorporated into the tricarboxylic acid cycle and excreted. Moreover, PLA degradation products are nontoxic (at a low composition) making it a natural choice for biomedical applications (Rasal et al., 2010). Molecular weight affects the properties of PLA. Low molecular weight PLA is desirable for biomedical applications because high molecular weight PLA has a complete resorption time of 2 8 years. Hence, the prolonged existence of PLA in vivo in some organs may lead to inflammation and infection.

Materials for Biomedical Engineering: Hydrogels and Polymer-based Scaffolds. DOI: https://doi.org/10.1016/B978-0-12-816901-8.00011-0 © 2019 Elsevier Inc. All rights reserved.

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The Food and Drug Administration (FDA), USA, has also approved PLA for direct contact with biological fluids (Gupta et al., 2007). PLA or PHAs, as biodegradable materials, have drawn enormous interest because of the fact that the support disappears from the transplantation site with the passage of time, leaving behind a perfect patch of natural tissue (Gupta et al., 2007). The surface properties of PLA and PHA materials play a critical role in determining their biomedical applications. Different surface modification strategies, such as physical, chemical, plasma, and radiation induced methods, have been employed to create desirable surface properties. Chemical modification of polyesters is considered more favorable than physical modification because it can be scaled up in a more reproducible manner. Pristine and surface modified PLA and PHAs have extensive applications in the biomedical field, including suture, bone fixation material, drug delivery microsphere, and tissue engineering (Hazer et al., 2012; Valappil et al., 2006; Jaffredo and Guillaume, 2014; Manavitehrani et al., 2016). PLA or PHAs in modified and unmodified forms may be fabricated into fiber form so as to be used as sutures. There are several areas where PLA or PHAs materials can be used. PLA fibers are the preferred material for applications that require the prolonged retention of strength, such as ligament and tendon reconstruction and stents for vascular and urological surgery (Durselen et al., 2001). Three-dimensional porous scaffolds of PLA can be used in cell based gene therapy for cardiovascular diseases; muscle tissue, bone, and cartilage regeneration; and other treatments of cardiovascular, neurological, and orthopedic conditions (Coutu et al., 2009; Kellomaki et al., 2000; Papenburg et al., 2009). The future of PLA or PHAs lies in their combination with other biomaterials or modifications in order to increase/maintain their strengths and mitigate their weaknesses. This may be a potent strategy for the engineering of complex tissues.

11.2 STRUCTURE OF PLA AND PHAS The chemical formula of PLA is (C3H4O2)n and it has one COOH group at one end and an OH group at the other end (Kumar et al., 2014). The functional groups situated at both ends are responsible for the reactive/interactive behavior of poly(lactic acid). PHAs are also long molecules that are made up of small monomers. In the case of PHAs, the monomers are 3-hydroxyalkanoates. The monomers have a hydroxyl group at the third carbon making the structure: 3-hydroxy alkanoates. The hydroxyl group of one monomer is attached to the carboxyl group of another by an ester bond. Hence, PLA or PHAs based biopolymers are categorized as polyester. There are several reports in which PLA is characterized structurally by Fourier transform infrared (FTIR). Hydroxyl ( OH of alcoholic and carboxylic)

11.3 Synthesis of PLA and PHAs

and aCQO bands in PLA were assigned at 3447 and 1752 cm21, respectively. The band at 1455 cm21 in FTIR represented C H bending vibrations in PLA (Kumar et al., 2014). In the FTIR spectra of the extracted PHA, the observed 3420.20 cm21 absorption band is indicative of the presence of the hydroxyl group in a polymer chain (Salazar et al., 2014). The absorption band at 2955.76 cm21 was assigned to the asymmetric methyl group of PHA. The bands at 2925.98 and 2855.99 cm21 were assigned to asymmetric CH2 of the lateral monomeric chains and symmetrical CH3, respectively. The absorption band at 1741.44 cm21 had been reported to be a carbonyl ester (R CQO) stretching vibration band. Since the synthesis of PHA was carried out by bacteria, the vibration at 1469.77 cm21 had been assigned to bacterial protein amide II (aNQCQO) in the cell. Absorption at 1378.83 and 1259.89 cm21 were assigned to terminal CH3 groups and asymmetric C O C stretching vibration, respectively. The series of absorption bands at 1166.87 619.39 cm21 had been assigned to C O and C C stretching vibration in the amorphous phase.

11.3 SYNTHESIS OF PLA AND PHAS The chemistry of PLA involves the processing and polymerization of lactic acid monomers. Since lactic acid is a chiral molecule, PLA has three stereoisomers and they are: poly(L-lactide) (PLLA), poly(D-lactide) (PDLA), and poly(DL-lactide) (PDLLA) (Lopes et al., 2012). Lactic acid can be made through the fermentation of sugars obtained from renewable resources, such as sugarcane or corn starch. PLA was synthesized in 1932 by Carothers (DuPont, USA). In the past, only low molecular weight PLA was obtained by heating lactic acid. Ringopening polymerization has been used to produce high molecular weight PLA (Jamshidian et al., 2010). Isotactic and optically active PLLA and PDLA are crystalline, whereas relatively atactic and optically inactive PDLLA is amorphous (Farah et al., 2016). There are several companies all over the world, such as Cargill Dow Polymer LLC, Shimadzu Corp, Mitsui Chemicals, Musashino Co. (Kumar and Kumar, 2012), that have been producing PLA. PLA can be prepared from lactic acid through different polymerization processes, such as polycondensation, ring-opening polymerization, azeotropic dehydration, and enzymatic polymerization. Currently, direct polymerization and ringopening polymerization are the most used production techniques. The presence of both hydroxyl and carboxyl groups in lactic acid enables it to be converted directly into polyester via a polycondensation reaction. It has been reported that direct condensation polymerization has fewer manufacturing steps and lower cost as compared to ring-opening polymerization, in addition to being easier to manipulate and commercialize. The low molecular weight of the resultant polymer in

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direct condensation polymerization is due to equilibrium among the free acid, the oligomers, and the water produced during the reaction or after certain special treatment. A controlled molecular weight of PLA can be obtained by catalytic ring-opening polymerization of the lactide intermediate (Kim et al., 2009). By controlling the residence time and temperatures in combination with catalyst type and concentration, it is possible to control the ratio and sequence of D- and L-lactic acid units in the final polymer (Farah et al., 2016). PHAs can be synthesized either by chemical means or by biological approaches. Biosynthesis of PHAs leads to high molecular weight compared to that achieved by chemical means. There are several microorganisms responsible for the synthesis of PHAs. These include Pseudomonas putida CA-3 (NCIMB 41162), Cupriavidus necator, and many more. Starch, glucose, fructose, fatty acids, and many other nutrients are responsible for the growth of PHAs. Synthesized PHAs may be of medium chain length or short chain length depending on the number of carbons. In one of the reported literatures, P. putida CA-3 cultures were grown in a medium (pH 7.1) at 30 C with shaking at 200 rpm to produce PHA (Ward and O’Connor, 2005). It was grown in batch culture, under nitrogen limiting conditions. Sodium ammonium phosphate (SAP) was supplied at a concentration of 1.0 g/L as the nitrogen source in the growth medium. After 48 hours, cells were harvested and the PHA obtained was quantified. The production of PHA is possible in limiting nitrogen conditions and, hence, the concentration of SAP, which had been used as the nitrogen source, was determined. The results stated that the depletion of nitrogen in the growth medium coincided with PHA accumulation. Fig. 11.1 shows the percentage cell dry weight of PHA and the PHA yield accumulated by P. putida CA-3 cells for aromatic substrates. It was observed that increasing the initial concentration of the substrate increased the production of PHAs. The highest level of PHA (% cell dry wt.), accumulated by P. putida CA3 cells grown on aromatic substrates, was 59%. Similarly, Fig. 11.2 shows the production of PHA; percent cell dry weight of PHA; and the PHA yield accumulated by P. putida CA-3 cells on aliphatic substrates. In one of the studies, two low-cost carbon substrates (i.e., carob pulp and fique juice) were tested for lab-scale production of PHA with Bacillus megaterium (Ward and O’Connor, 2005). The authors suggested that PHA production using carob pulp may be as high as with sugarcane molasses, and it could also serve as a substrate for the synthesis of the most commercialized type of polyhydroxybutyrate and poly(hydroxy-butyrate-co-valerate). The biomass growth medium for the production of PHA contained 10.0 g/L yeast extract, 15.0 g/L nutrient broth, and 5.0 g/L ammonium sulfate (Salazar et al., 2014). The stock inoculum was aseptically introduced into this medium at a concentration of 3% (v/v) and incubated at 30 C, with 250 rpm for 24 hours in a shaker incubator. After that the cell biomass was harvested at 4 C in a centrifuge (9000 3 g) for 10 minutes. The harvested biomass was then used to seed the minimal medium.

11.3 Synthesis of PLA and PHAs

70

Phenylvaleric acid ( ) Phenylhexanoic acid ( ) Phenylheptanoic acid ( ) Phenyldecanoic acid ( × ) Phenyloc tanoic acid ( )

PHA (% cell dry wt.)

60 50 40 30 20 10 0 0

5

15 10 Substrate concentration (mM)

20

FIGURE 11.1 Percentage cell dry weight of PHA and PHA yield accumulated by Pseudomonas putida CA-3 cells for aromatic substrates. Reprinted from Ward, P.G., O’Connor, K.E., 2005. Bacterial synthesis of polyhydroxyalkanoates containing aromatic and aliphatic monomers by Pseudomonas putida CA-3. Int. J. Biol. Macromol. 35, 127 133.

The minimal medium for PHA production contained 3.5 g/L NaNH4HPO4
4H2O; 5.7 g/L K2HPO4; 3.7 g/L KH2PO4; and 10 mM g/L carboxylic acid having different carbon contents of C8:0, C12:0, C16:0, and C18:1, respectively. The pure carboxylic acid was used as the sole carbon and energy source during the fermentation process (Gumel et al., 2014). Sterile 1.0% (v/v) Mg2SO4
7H2O solution and 0.1% (v/v) trace elements (MT) solution containing (g/L): CaCl2
2H2O 1.47; CoCl2
6H2O 2.38; CuCl2
2H2O 0.17; FeSO4
7H2O 2.78; MnCl2
4H2O 1.98; and ZnSO4
7H2O 0.29, dissolved in 1 M HCl were aseptically and separately added prior to inoculum seeding. 3% (v/v) harvested biomass suspended in a phosphate buffer was seeded into this medium aseptically and then incubated at 30 C, 250 rpm for a total period of 48 hours except otherwise stated. PHA accumulated biomass was harvested by centrifugation at 9000 3 g for 10 minutes (Salazar, et al., 2014). The supernatant was discarded while 10% (v/v) n-hexane in distilled water was used to wash the pellets three times to remove residual fatty acids. Later, the biomass was dried in a vacuum in the presence of phosphorus pentaoxide. About 10% (w/v) dried biomass was suspended in chloroform and refluxed at 80 C for 4 hours, after which a Buchner flask equipped with PTFE filter paper was used to filter the reflux mixture under vacuum. The filtrate was concentrated in a rotary evaporator at 4 C under reduced pressure till it reached about 5% of the original volume. The polymer was then precipitated in

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70 60 PHA (% cell dry wt.)

344

50 40 30 20 10 0

0

5

10 15 Substrate concentration (mM)

20

25

Nonanoic acid ( ) and decanoic acid ( ) butyric acid ( ), hexanoic acid (×), heptanoic acid ( ), octanoic acid ( ), valeric acid ( )

FIGURE 11.2 Percentage cell dry weight of PHA and PHA yield accumulated by Pseudomonas putida CA-3 cells for aliphatic substrates. Reprinted from Ward, P.G., O’Connor, K.E., 2005. Bacterial synthesis of polyhydroxyalkanoates containing aromatic and aliphatic monomers by Pseudomonas putida CA-3. Int. J. Biol. Macromol. 35, 127 133.

cold methanol while being gently stirred. The PHA solution was allowed to settle under gravity for 24 hours after which the supernatant was decanted leaving behind a layer of PHA film. This film was air-dried before the purification steps were repeated by chloroform dissolution and methanol extraction.

11.4 PROPERTIES OF PLA AND PHAS Polylactide is a clear, colorless thermoplastic having an appearance similar, in many respects, to polystyrene, while polyhydroxybutyrate (an important polymer in the category of PHAs) is opaque and creamy in color. The density of amorphous and crystalline PLLA has been reported as 1.248 and 1.290 g/mL, respectively (Auras et al., 2004), while that of PHB is around 1.248 g/mL. In general, PLA and PHA products are soluble in organic solvents, such as dioxane, acetonitrile, chloroform, methylene chloride, 1,1,2-trichloroethane, and dichloroacetic acid. PLA and PHAs get partly dissolved in ethyl benzene, toluene, acetone, and tetrahydrofuran in cold conditions, though they are readily soluble in these solvents when heated to boiling temperatures.

11.4 Properties of PLA and PHAs

11.4.1 MECHANICAL PROPERTIES The mechanical properties of PLA or PHAs are crucial to the success of their application in tissue engineering. To be used successfully in tissue engineering, it is critical that these biomaterial scaffolds temporarily withstand and conduct the loads and stresses that the new tissue will ultimately bear (Dhandayuthapani et al., 2011). There are several mechanical properties that are important: 1. Elastic modulus—measures strain in response to a given tensile or compressive stress along the force. 2. Flexural modulus—measures the relationship between a bending stress and the resulting strain in response to a given tensile or compressive stress perpendicular under load. 3. Tensile strength—maximum stress that the material can withstand before it breaks. 4. Maximum strain—ductility of a material or total strain exhibited prior to fracture. Fig. 11.3 shows the mechanical properties of neat PLA. PLA alone has extremely low percentage elongation at break (2.8%) but has a tensile strength of 39.2 MPa. Hence, it is categorized as a brittle material. The introduction of additives, in many cases, resulted in significant decreases in tensile strength with a lower percentage elongation at break, that is, 0.7% 0.8%. The decrease in tensile strength of PLA in the presence of additives showed that the additive acted as a filler but not as reinforcement. The flexural modulus of PLA as reported in the literature was found to be 50 60 MPa (Kumar et al., 2013; Plackett et al., 2003). The trend in the mechanical properties of PLA in the presence of a plasticizer demonstrated that the plasticizer enhanced the segmental mobility of PLA chains which in turn increased the ability of amorphous PLA to follow plastic deformation, thus, decreasing the yield stress and increasing the elongation at break. PHAs showed tensile strength of around 30 40 MPa with around 10% 15% elongation at break. Among all PHAs, polyhydroxybutyrate is widely studied in addition to polyhydroxy succinate and polyhydroxy valerate. Elongation at break is higher for PHAs in comparison to PLA (Gigli et al., 2016). There are several literatures that have explored the incorporation of nanoparticles (up to 10 wt.%) in PLA or PHAs in order to enhance their mechanical properties (Armentano et al., 2013). This is due to the fact that nanoparticles possess high interfacial areas as compared to their volume. Nanoclay has been used widely to reinforce PLA. Increase in tensile modulus (37%) and elongation at break (48%) was observed after the addition of 5 wt.% of organoclay (Armentano et al., 2013). However, Rhim et al. (2009) observed a decrease in the tensile strength and the elongation at break after adding 5 wt.% of unmodified clay and organoclays (C20A and C30B) to the PLA matrix.

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40

Tensile strength (MPa)

35 30 25 20 15 10 5 0 –1

0

1

2 Strain (%)

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FIGURE 11.3 Mechanical properties of neat PLA.

11.4.2 THERMAL PROPERTIES The glass transition temperature (Tg) of amorphous PLA is one of the most important parameters. It is well known that dramatic changes in any polymer chain mobility take place at and above Tg. PLA is not completely amorphous, hence, for semicrystalline PLA, both Tg and melting temperature (Tm) are important. The melt enthalpy estimated for a pure PLA of 100% crystallinity is 93 J/g and in some papers it has been reported as high as 148 J/g. The differential scanning

11.4 Properties of PLA and PHAs

Heat flow (mW, au)

calorimetry (DSC) results of PLA are shown in Fig. 11.4. A glass transition temperature of around 50 C can be clearly observed in Fig. 11.4. Cold crystallization and melting curves (147.6 C) were also observed in PLA. The thermal properties of PHA based polymeric systems were significantly affected by the synthetic strategy. The glass transition temperature of PHAs is extremely low and it is below room temperature. In general, it varies from 2 C to 8 C, depending on the type of PHAs. The degree of crystallinity affects the melting temperature of PHAs. Generally, the melting temperature ranges from 160 C to 175 C. The shorter the length of the repeat units in PHAs, the lower the melting point is. This is attributed to the formation of less and less perfect crystals due to the increased difficulty of PHAs chain folding (Gigli et al., 2012, 2013; Soccio et al., 2012; Gualandi et al., 2012). Thermogravimetric analyses (TGA) are used to assess the thermal stability and degradation profiles of PLA/PHAs materials. Themogravimetric curve of the PLA matrix under inert atmosphere exhibits a typical single weight-loss step, with a maximum decomposition rate at about 370 C (Fig. 11.5). The addition of nanoparticles increases the thermal degradation temperature of PLA and the extent of this increase usually depends on the exfoliation degree of the organoclays. In order to understand the thermal degradation of PHAs based polymers, the study of mass loss data obtained from dynamic TGA is important. Single step degradation is generally observed after 300 C in PHAs. There will be the influence of additives on PHAs degradation. If that is there, then the degradation mechanisms of both the polymer and the additive have to be analyzed.

147.6

81.8

25

FIGURE 11.4 DSC curves of neat PLA.

50

75

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125 150 175 Temperature (°C)

200

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FIGURE 11.5 TGA curves of neat PLA.

In one of the literatures, the melting temperature of the medium chain length poly-3-hydroxyalkanoates polymer was observed to range from 43.3 C to 66.5 C and the glass transition temperature from 1.0 C to 1.0 C (Salazar et al., 2014). Low melting and glass transition temperatures were observed and it was attributed to the random composition of 3-hydroxyalkanoates. A relatively low enthalpy of fusion was observed as compared to 145.3 J/g for PHA, which could be due to the amorphous nature of the polymer. A relationship between the fatty acid chain length of carboxylic acid used for the production of PHB and the polymer thermal stability was clearly observed. When the initial substrate was taken as lauric acid, a polymer with a degradation temperature of 264.6 C was observed. On changing the fatty acid to palmitic acid or oleic acid, the polymer degradation temperature increased to 302.2 C. This increase in thermal stability with increased fatty acid chain length was suggested to be due to the increase in the longer chain monomer fraction. This indirectly favored side-chain crystallization resulting in a polymer with relatively high degradation temperature.

11.4.3 TRANSPARENCY PLA is a transparent matrix, whereas PHAs are opaque in nature. PLA retains its transparency with no reduction in the amount of light transmitted when a low volume percentage of well dispersed nanoreinforcements are added. Hence, it can be said that the transparency of nanoparticle reinforced PLA is also an indication of nanofiller dispersion. Sanchez-Garcia and Lagaron (2010) reported a brownish color for nanoclay reinforced PLA and a gradual increase of yellowness index

11.4 Properties of PLA and PHAs

with increasing nanoclay contents. A reduction of 32% of visible light transmitted (at 650 nm) was also observed when 5 wt.% of nanoclays were added.

11.4.4 BIOCOMPATIBILITY PLA has been demonstrated to be biocompatible and to degrade into nontoxic components. PLA hydrolysis occurs in vivo which means that there is no requirement for enzymes to catalyze it. As reported earlier, the duration of degradation of PLA can last from 1 to 2 years (Papageorgiou et al., 2010). The degradation of PLA produces lactic acid. In the metabolic activity of human beings, lactic acid is the starting material for the Krebs cycle (Armentano et al., 2013). PHAs have excellent cell compatibility with a variety of primary cells and cell lines, such as cartilage cells, bone-marrow-derived mesenchymal stem cells, fibroblasts, and osteoblasts. The behaviors of cells are different for different PHAs. Electrospun PHA membranes are better substrates for the culturing of human mesenchymal stem cells compared to compression-molded PHA membranes. The reason is that higher cell viability is maintained in electrospun PHAs (Yu et al., 2010). Also, different cells had different responses to PHA materials. For fibroblasts, the smoothest copolymer of PHB with 20% hydroxy hexanoate was preferred, while for osteoblasts, a PHB copolymer with 12% hydroxy hexanoate surface with appropriate roughness was the most favorable. This roughness may also promote differentiation in osteoblasts. Human embryonic stem cells, spontaneously differentiated human embryonic stem cells, and mesenchymal stem cells, were cocultured with a copolymer of PHB-hydroxy hexanoate copolymer and collagen (Lomas et al., 2013). Following exposure to the appropriate induction medium, retention of osteogenic, chondrogenic, and adipogenic differentiation by the expression of genes were observed. This suggested that PHAs have the potential for use as biocompatible scaffolds in future tissue engineering.

11.4.5 PROCESSABILITY The main conversion approaches of PLA or PHAs are based on melt processing. In this process, the processing temperature is kept above the melting point of the polymers concerned. Commercial grades of PLA or PHAs can typically be processed using a conventional single screw or twin-screw extruder. Drying of PLA or PHAs pellets before processing is of major importance so as to prevent hydrolysis during processing. The processing of PLA or PHAs in different shapes and sizes makes PLA or PHAs suitable biomaterials. The shapes and sizes of the organs vary markedly; hence, it is important for biopolymers to assume the desired shape and sizes. However, there are drawbacks in the form of their high brittleness, which limits their use in flexible biomaterials. Many research efforts have been directed toward overcoming these limitations, such as the addition of

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natural additives and plasticizers, copolymerization, and blending with biopolymers (Lomas et al., 2013). In one research paper, the authors explored the possibility of adding soy protein isolate (SPI) as a filler in PLA (Kumar, 2013). They demonstrated that higher amounts of SPI worsen the processability of blends. The mechanical properties of the blends as reported were almost the same. However, the thermal properties of the blends having PLA and SPI were better compared to unblended PLA. Triacetin as a plasticizer effectively eliminated the brittle nature of PLA which is its main drawback. But there was an appreciable decrease in the tensile stress with the addition of a plasticizer.

11.5 APPLICATION OF PLA AND PHAS IN TISSUE ENGINEERING In human beings, organ failure and tissue loss are extremely problematic and the current approach to treat this problem is based on transplantation. In tissue engineering patients with organ defects and malfunctions are treated with their own cells, grown on a polymer support. After that, tissues are regenerated from natural cells without the help of a donor (Farah et al., 2016). However, tissue engineering is an interdisciplinary field which requires the knowledge and principles of life science and engineering so that biological substitutes can be developed in order to maintain, restore, or improve tissue function. Bone and cartilage are tissues that have highly complex and hierarchically arranged structures. Bone can be regarded as a composite material consisting of 60 wt.% inorganic calcium phosphate embedded in 30 wt.% organic phase, mainly collagen and, finally, 10 wt.% of water. On the contrary, cartilage consists of a hydrated proteoglycan gel (Salerno and Pascual, 2015). Biopolymers, such as PLA, are widely used in tissue engineering for bone and cartilage repair. PHAs are biopolymers belonging to the aliphatic polyester family and their use as scaffolds in bone and cartilage regeneration has been reported (Salerno and Pascual, 2015). Porous PHA matrices are fabricated by particulate leaching and phase separation (You et al., 2011), as well as composites with hydroxyapatite particles and functionalized with peptides for enhanced biocompatibility and stem cell differentiation. Biopolymers from PLA or PHAs can provide unique features in terms of biochemical and biophysical properties for tissue engineering. These materials can be used to mimic the chemical and structural functionality of native tissue, ultimately enhancing tissue regeneration. The possibility of synthesizing novel biomedical devices from PLA and PHAs are very promising. The literatures previously discussed demonstrate that PLA and PHAs based biopolymers provide suitable therapeutic options for tissue applications in bone, cartilage, blood vessel, and nerve regeneration.

11.6 Biodegradability of PLA and PHAs

Poly(3-hydroxybutyrate-co-3-hydroxyhexanoate) copolymer consisting of random copolymers of 3-hydroxybutyrate and 3-hydroxyhexanoate plays a significant role in the field of biomedical materials. It is potentially useful for a wide range of biomaterial applications. The development of this copolymer in tissue engineering applications has been explored in the form of bone, cartilage, tendon, nerve, and vessel repair (Yang et al., 2014). Nanofibrous scaffolds based on PLA for tissue engineering have been shown to be a versatile tool for tissue engineering. It can be used as a three-dimensional topographical surface for cell attachment, a depot for drug delivery, and a substrate for biofunctionalization (Patricio et al., 2013; Santoro et al., 2016). Electrospinning is a widely used technique for the production of nanofibrous scaffolds. These processes have enabled the creation of designer scaffolds for addressing the needs of specific tissues, such as having fiber alignment and incorporating bioactive molecules. Electrospinning of fiber is dependent on the applied voltage as well as the spinning distance and both of these parameters affect the thickness of the electrospun fibers.

11.6 BIODEGRADABILITY OF PLA AND PHAS PLA or PHAs primarily degrade by hydrolysis, after several months of exposure to moisture. The degradation of these biopolymers occurs in two stages (Farah et al., 2016). First, random nonenzymatic chain scissions of the ester groups lead to a reduction in the molecular weight of the biopolymers. In the second stage, the biopolymers having reduced molecular weight are naturally metabolized by microorganisms to yield carbon dioxide and water. A biopolymer degradation rate is mainly determined by its reactivity with water, if the degrading medium is water. If the degrading medium is not water, then thermal activation, hydrolysis, biological activity (i.e., enzymes), oxidation, photolysis, or radiolysis affect the degradation rate of the biopolymer. Several other factors accelerate biopolymer degradation: 1. 2. 3. 4. 5. 6.

An increase in hydrophilic monomers. An increase in hydrophilic, acidic end groups. An increase in reactive hydrolytic group in the backbone. Molecular weight Less crystallinity Smaller device size when the biopolymers are molded into the organs shape.

The location of the device in human beings can play an important role in the degradation rate of biopolymer implants. Large devices, fabricated from biopolymers, if implanted in areas with poor vascularization may degrade and overwhelm the body’s ability to flush away degradants.

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PLA degradation was studied in animal and human bodies for medical applications, like implants, surgical sutures, and drug delivery materials (Hamad et al., 2015). It was observed that in these environments, PLA is initially degraded by hydrolysis and the soluble oligomers formed are metabolized by cells. The mechanism is similar to what has been discussed earlier. If structurally observed, PLA has a relatively long half-life of hydrolysis due to steric effects. The alkyl group hinders the attack by water. It was reported in the literature that PLLA stents or fibers exposed to in vivo conditions do not begin to degrade until 12 months approximately (Bergstrom, 2013). PLA is extremely stable and retains its weightaverage molecular weight and physical properties for years. This is typified by its growing use in clothing and durable applications. The high weight average molecular weight of PLA is also naturally resistant to supporting bacterial and fungal growth, which allows it to be safely used for applications such as food packaging and sanitation. For implantable medical devices, once implanted in the body, the biodegradable device should maintain mechanical properties until it is no longer needed and then it should be degraded, absorbed, and excreted by the body, leaving no trace. Semicrystalline PLA biodegradation occurs in two phases. In the first phase, water penetrates the bulk of the device. In this step long polymer chains are converted into shorter, ultimately water-soluble fragments. The reduction in molecular weight of PLA is soon followed by a reduction in physical properties as water begins to fragment the device. Mechanical properties, such as yield strain, yield stress, and elongation to failure for PLA based medical devices start to decrease. These changes will have consequences for devices that are load bearing throughout their degradation process (Bergstrom and Hayman, 2016). A study of the degradation of PHB and its copolymer; poly(hydroxybutyrateco-hydroxyvalerate) (PHBV), has been carried out in pilot-scale composting conditions (Weng et al., 2010). In the pilot-scale composting conditions, parameters such as pH value, temperature, and the amount of oxygen and CO2 produced, were determined periodically. The amount of CO2 evolved gives an indication of the degree of biodegradation of PHBV. The degree of disintegration of PHBV film could reach 100% and 81% under pilot-scale and laboratory-scale composting conditions, respectively. After 12 weeks of biodegradation there were no residual PHVB films. This indicates that within 2 3 months there will be complete degradation of PHA based polymers. The size of the samples became smaller with the increase of composting time and after 12 weeks it was difficult to find the remnants of the polymer. Before degradation the surface of PHBV films was smooth, but after 20 days of degradation the surface was eroded and many cavities were found. Biodegradability patterns of two PHAs: a polymer of 3-hydroxybutyric acid (3-PHB) and a copolymer of 3-hydroxybutyric and 3-hydroxyvaleric acids (3PHB/3-PHV) containing 11 mol% of hydroxyvalerate, were studied in a tropical marine environment (Volova et al., 2010). The average temperature of the water was 28.75 C 6 1.65 C, with a minimum of 27.1 C and a maximum of 30.4 C.

11.6 Biodegradability of PLA and PHAs

Water pH values were close to neutral. The average salinity of the water in the study period was 34%. Dissolved oxygen concentration varied from 5.4 to 8.3 mg/mL. There were no significant differences between the degradation rates of 3-PHB and 3-PHB/3-PHV specimens. Overall in this research paper, it was found that biodegradation is influenced by the shape of the polymer item and the preparation technique used rather than by the chemical composition of the polymer. If the shape is quite thick then more time will be needed to degrade the polymers. After biodegradation, the specimens lose their mass and the molecular weight of the polymers decreased. Researchers have isolated PHA-degrading microorganisms and they were identified as Enterobacter sp. (four strains), Bacillus sp., and Gracilibacillus sp. Poly (3-hydroxybutyrate-co-3-hydroxyvalerate) (PHBV) degrading microorganisms from soil were isolated and the degradation was analyzed using scanning electron microscopy (SEM) and FTIR spectroscopy (Shah et al., 2010). The samples were buried in soil for 120 days at 37 C. After the designated time, there was an increase in biomass on the surface of PHBV film pieces and in the medium. This result indicated that microbes utilize PHBV as a source of carbon and energy. At the end of the incubation process, SEM results showed many changes in surface morphology, such as erosion and extensive roughening of the surface with pit formation, as compared to the untreated samples. The FTIR spectra of PHBV film showed a decrease in the peak from 1725 cm21 (untreated plastic film) to 1721 cm21. In the FTIR study, the disappearance of a peak at 2745 cm21 in the treated sample showed the breakdown of ester bonds. It can be concluded that soil contains microorganisms that can potentially be used for the biodegradation of plastic wastes. It was concluded in this paper that PHBV can be degraded by the action of a variety of microorganisms (bacteria/fungi) isolated from soil. The degradation could be performed either by growing the bacteria in the presence of the polymer or by incubating the polymer with the depolymerase enzymes. The biodegradation behavior of poly(hydroxybutyrate); PHBV with different contents of HV (40 mol% HV; 20 mol% HV; 3 mol% HV), and P(3HB,4HB) with 10 mol% 4HB were investigated under controlled composting conditions according to ISO 14855-1 (Weng et al., 2011). This study was carried out in order to understand the influence of chemical structure on the biodegradability of PHB and PHBV. The procedure mentioned in this research article is to determine the ultimate biodegradability of the material under conditions simulating an intensive aerobic composting process. The test apparatus used in this research paper includes 2.5 L composting vessel, an air-supply system, and the apparatus for determination of carbon dioxide with a continuous infrared analyzer. In the wellestablished process, 100 g of test material was mixed with 600 g of inoculum and then introduced into a static composting vessel where it was intensively composted under the optimum oxygen level using an air-supply system. The concentration of oxygen, temperature, and the test period were 6%, 58 C 6 2 C, and 6 months. The compressed air was supplied at a constant low pressure at a rate of

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150 L/min. The carbon dioxide in the air was first removed by passing the air through a solution of sodium hydroxide, and at the same time the air was humidified. It was found that PHAs with different chemical structures could be biodegraded under controlled composting conditions. The results state that the order of biodegradability was P(3HB,4HB)BPHBV-40 . PHBV-20 . PHBV-3 . PHB, which was well correlated with the HV and 4HB content. Erosion from the surface to the interior of PHAs indicated the biodegradation of PHAs. The chemical structure of PHAs after degradation was not changed. However, their thermal stability and molecular weight decreased during the composting.

11.7 PLASMA TREATMENT OF PLA AND PHAS PLA and PHAs are hydrophobic thermoplastic polyesters. During practical applications, PLA and PHAs preserve their mechanical strength and other material performances. These polymers finally degrade to low molecular weight compounds such as H2O, CO2, and other nontoxic byproducts (Ikada and Tsuji, 2000). This degradation property indicates that a device made of such biodegradable polymers can be implanted in a human body without necessitating a second surgery to remove the device. Biomaterials from PLA and PHAs have been well characterized and fabricated to match the biochemical properties of soft tissue. But there is generally a lack of mechanical compatibility between PLA and PHA based polymer implants and living tissues, which is attributed to their high hydrophobicity and low surface energy. It is generally difficult for cells to attach, spread, and proliferate on these biodegradable polymers. Therefore, the surface of these polymers needs to be modified in order to enhance their hydrophilicity and surface energy. Chemical modification of these thermoplastic polyesters results in the introduction of specific functional groups on their surfaces which increases their cytocompatibility. Two possible wet-chemical routes can be followed for surface modifications: surface aminolysis and surface hydrolysis. Surface aminolysis leads to the formation of free amino groups on the surface of the polyester. On the other hand, surface hydrolysis with NaOH results in the formation of carboxylic acid and hydroxyl groups. The presence of NH2, OH, and COOH groups in these wet-chemical processes results in enhanced hydrophilicity and improved cell-material interactions (Zhu et al., 2002, 2004). However, these wetchemical processes have a lot of disadvantages too. Surface modification through these processes are quite rough which can lead to unwanted side effects, such as faster degradation rate, irregular surface etching, and a reduction of mechanical performance (Desmet et al., 2009). The degree of modification induced by these processes may not be reproducible because of their strong dependence on molecular weight, crystallinity, or tacticity as discussed in the previous sections. From an environmental point of view, these techniques use substantial amounts of water or other liquids which generate hazardous chemical waste. Apart from wet-chemical

11.7 Plasma Treatment of PLA and PHAs

routes, many other techniques of surface treatment have been reported for improving polymer properties, each having their own advantages and disadvantages (Slepicka et al., 2013; Kasalkova et al., 2013; Goddard and Hotchkiss, 2007; Rasal et al., 2010). Plasma treatment of polymer surfaces is one of the newer technologies that offer innovative solutions to adhesion and wetting problems in pharmaceutical industries. It provides an economical solution for the cleaning and activation of component surfaces before further processing. The major advantages of the plasma treatment process are its simplicity, reliability, and affordability. Plasma in contact with a material surface can change its surface properties, such as wetting, metal adhesion, dye ability, chemical inertness, lubricity, and biocompatibility (Slepicka et al., 2012a,b). By changing certain parameters, like temperature, pressure, power, gas, and substrate type, either surface functionalization, codeposition, ablation, or crosslinking takes place. The upper layer of polymer surface can be modified by plasma treatment without using solvent or generating chemical waste and with less degradation and roughening of the material than with many wet-chemical treatments. (Goddard and Hotchkiss, 2007). Improving adhesion characteristics, increasing hydrophilicity, introducing special functional groups, or modifying surface morphology are the basic purposes of the plasma treatment process (Rimpelova et al., 2013; Kolska et al., 2014; Kasalkova et al., 2014).

11.7.1 PLASMA AND PLASMA SURFACE INTERACTIONS Plasma refers to an ionized gas or an electrically neutral medium of positive and negative particles. It is considered to be the fourth state of matter after solid, liquid, and gas. It is in the form of gaseous or fluid-like mixtures of free electrons, ions, and radicals and also contains neutral particles like atoms or molecules. The typical luminosity of plasma results from the relaxation of these particles from their excited states to ground state. Plasma is subdivided into two categories (Geyter and Morent, 2012); equilibrium (or nonthermal/low temperature/cold) and nonequilibrium (or thermal/high-temperature/hot). Thermal equilibrium refers to when the constituent particles of plasma are at the same temperature. High temperatures used to create these types of plasma are destructive for heat-sensitive polymers and, therefore, most of the applications for surface modification of polymers make use of nonthermal or cold plasma. Since nonthermal plasma contains a mixture of reactive species, four different kinds of interactions take place between plasma and a surface, these are briefly discussed (Geyter and Morent, 2012): 1. Plasma treatment: This process is basically used to enhance the surface energy of a polymer. When a polymer surface is exposed to cold plasma generated in O2, N2, or NH3, polar hydrophilic groups are introduced onto the polymer surface, thus, enhancing its hydrophilicity. On the other hand, plasma

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generated in He or Ar leads to the creation of free radicals that can be used for crosslinking or grafting of oxygen-containing groups. Also in this case, the surface is exposed to oxygen or air after the treatment (Desmet et al., 2009). However, surface modifications induced by this process are not permanent because the treated surface tends to recover to their untreated state during storage and they will also undergo postplasma oxidation reactions (Geyter et al., 2008). 2. Plasma postirradiation grafting: This is a two-step process in which the first step is the plasma treatment as discussed above, followed by the second step in which the activated polymer surface is brought into direct contact with a monomer. The monomer can be in the gas phase or the substrate can be immersed into a monomer solution (Vasilets et al., 1997). Since the monomer is not subjected to the reactive plasma environment, the grafter polymers will have similar characteristics to polymers synthesized by conventional polymerization processes (Desmet et al., 2009). In contrast to plasma treatment, this process results in a permanent effect. 3. Plasma syn-irradiation: In this process, a monomer is first adsorbed to a material, and then the substrate is exposed to plasma. Therefore, contrary to the previous process, the monomer is directly subjected to the plasma. This type of plasma treatment process generates radicals in the adsorbed monomer layer and the surface of the substrate leads to crosslinked polymer in the form of top-layer surface modification (Desmet et al., 2009). 4. Plasma polymerization: This process is used to prepare thin films of polymers known as plasma polymers. During this process, gaseous or liquid monomers are inserted via a carrier gas into the discharge zone where they are converted into reactive fragments (Morent et al., 2009). These reactive fragments recombine to deposit a polymer film on the substrate exposed to the plasma. However, polymers formed by this process differ in their chemical structure and composition as compared to polymers formed via conventional polymerization processes (Desmet et al., 2009).

11.7.2 CHARACTERIZATION TECHNIQUES FOR PLASMA TREATED POLYMER SURFACES Plasma treated polymer surfaces are generally characterized by four measurement techniques: 1. Contact angle and surface free energy: Surface polarity (hydrophilicity or wettability) is determined by the measurement of water contact angle by goniometry using static water drop method as well as dynamic method. The water contact angle data is helpful in calculation of surface free energy. Contact angle is a quantitative measurement of the wetting of a solid by a liquid. It is defined geometrically as the angle formed by a liquid at the three phase boundary where a liquid, gas, and solid intersect. Static contact angles

11.7 Plasma Treatment of PLA and PHAs

are measured when a liquid droplet is standing on the surface and the three phase boundary is not moving; dynamic contact angles are measured during the movement of the three phase boundary. In the well-developed process, a drop of water is put on the polymer surface of which one wants to determine the contact angle. A camera is fitted to observe the shape of the drop of water with respect to time. The shape of the water droplet indirectly refers to the contact angle of the polymeric surface. The experiment to observe the contact angle is stopped when the change in the contact angle with time becomes negligible. 2. Surface Morphology: The surface morphology (roughness or smoothness) of a surface is examined using SEM or atomic force microscopy (AFM). SEM uses electrons rather than light to form an image. A beam of electrons is produced at the top of a microscope by the heating of a metallic filament. The electron beam follows a vertical path through the column of the microscope. It makes its way through electromagnetic lenses which focus and direct the beam down toward the sample. Once it hits the sample, other electrons (backscattered or secondary) are ejected from the sample. Detectors collect the secondary or backscattered electrons, and convert them to a signal that is sent to a viewing screen similar to an ordinary television, thereby producing an image. SEM has a large depth of field, which allows a wide area of the sample to be in focus at a time. It also produces images of high resolution, which means that closely spaced features can be examined at a high magnification. For the determination of morphology through SEM, samples are coated with gold. But nowadays, samples without gold sputtering are also being used. About 10 15 kV accelerating voltage is maintained for gold sputtered samples while about 5 kV is maintained for uncoated samples. On the other hand, AFM works in the same way as touch when a person uses their fingers to feel and probe the environment when they cannot see it. Using a finger to visualize an object, the brain is able to deduce its topography while touching it. The resolution that is obtained with this technique is determined by the radius of the fingertip. AFM measures ultrasmall forces (less than 1 nN) present between the AFM tip and a sample surface. These small forces are measured by measuring the motion of an extremely flexible cantilever beam having an ultra-small mass. AFM is capable of investigating the surfaces of both conductors and insulators on an atomic scale if suitable techniques for the measurement of cantilever motion are used. AFM can be used either in a static or dynamic mode. In the static mode a sharp tip at the end of the cantilever is brought into contact with a sample surface. In the dynamic mode, the tip is brought in close proximity (within a few nm) to, not in contact with, the sample. 3. Surface Chemistry: The presence of functional groups on a plasma treated surface is determined using X-ray photoelectron spectroscopy (XPS). XPS is a surface analytical technique that is based on the photoelectric effect. Each atom in the surface has core electrons with a characteristic binding energy that

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is conceptually, not strictly, equal to the ionization energy of that electron. When an X-ray beam is directed toward the sample surface, the energy of the X-ray photon is adsorbed completely by the core electrons of an atom. Photoelectron spectroscopy utilizes photoionization and analysis of the kinetic energy distribution of the emitted photoelectrons to study the composition and electronic state of the surface of a sample. Generally, carbon, hydrogen, oxygen, and nitrogen atoms are studied using the XPS technique. These four atoms generally constitute the functional groups in compounds. 4. Gravimetry: This technique involves the measurement of the thickness of the ablated surface layer after plasma treatment. The thickness of the ablated layer is calculated from the changes in the weight of samples before and after treatment. In order to enhance the sensitivity of the measurement, the samples are exposed to plasma from both sides.

11.7.3 PLASMA TREATMENT OF PLA Sarapirom et al. (2014) studied the effect of the surface modification of PLA through low-pressure ammonia plasma on adsorption of human serum albumin (HSA). They found that plasma treatment resulted in the increase of surface roughness. The increase in roughness had a slight dependence on the plasma generation power. The AFM images of untreated PLA and NH3-plasma treated PLA with various radio frequency (RF) powers before and after HSA adsorption are shown in Fig. 11.6. The AFM images do not show globular structure, indicating

FIGURE 11.6 AFM images of untreated PLA and NH3-plasma treated PLA with various RF powers before and after protein adsorption. Reprinted from Sarapirom, S., Yu, L.D., Boonyawan, D., Chaiwong, C., 2014. Effect of surface modification of poly(lactic acid) by low-pressure ammonia plasma on adsorption of human serum albumin. Appl. Surf. Sci. 310, 42 50.

11.7 Plasma Treatment of PLA and PHAs

that HSA were uniformly distributed over the surface of the samples without aggregation. The HSA-covered plasma-treated samples showed higher root-meansquare roughness as compared to that of the HSA-covered untreated samples. The smoother morphology of the HSA-covered untreated samples was attributed to higher coverage of HSA, whereas the coarser morphology was caused by a lower HSA coverage which led to some material surface areas being uncovered, resulting from protein aggregation into islands (Gonzalez et al., 2010). The surface roughness change qualitatively indicated that HSA adsorption was reduced on the plasma treated PLA surface. The decreased HSA adsorption is expected to increase cell attachment because of the reduction in the cell-adhesion barrier. The authors also measured the change in the contact angle of water after plasma treatment and the result is shown in Table 11.1. The table shows that the water contact angle decreased with increasing RF powers. For the untreated sample, the angle was 69 degrees, which decreased substantially to 15.5 degrees after 10 minutes of plasma treatment with RF of 100 W. Wang et al. (2016) studied cold atmospheric plasma (CAP) surface nanomodified 3D printed PLA scaffolds. The objective of the study was to use CAP as a quick and inexpensive way to modify the nanoscale roughness and chemical composition of a 3D printed scaffold surface. The resulting scaffold should mimic the micron structure of natural tissues as well as the nanoscale extracellular matrix properties of the tissues where these are intended to be replaced. The authors showed that the water contact angle of a normal 3D printed PLA scaffold dropped significantly after CAP treatment from 70 to 24 degrees. The SEM images of the untreated PLA scaffolds showed a smooth texture with no nanosized features observed on the surface. However, the CAP modified PLA scaffolds exhibited a rough surface, with an increasing nanopatterned area with respect to the exposure time of CAP treatment. The CAP-induced nanoscale morphology might be a result of charged particle-surface collision and heat accumulation due to long Table 11.1 Contact Angle of Untreated, Ar-Pretreated and NH3 Plasma-Treated PLA With Different RF Powers Under 10 min Treatment Time RF Power (W)

Water Contact Angle (Degree) 6 SD

0 (Untreated) Ar pretreated 50 75 100

69.0 6 1.9 52.2 6 1.7 20.0 6 1.0 15.0 6 1.0 15.5 6 0.7

Reprinted from Sarapirom, S., Yu, L.D., Boonyawan, D., Chaiwong, C., 2014. Effect of surface modification of poly(lactic acid) by low-pressure ammonia plasma on adsorption of human serum albumin. Appl. Surf. Sci. 310, 42 50

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term CAP treatment. It was concluded in this research paper that the nanopatterning of the PLA surface after CAP exposure may be favorable to cell growth and proliferation (San-Miguel and Behrens, 2012). Nanoscale surface roughness was further vindicated by AFM images, which showed that the surface roughness (Rq) of the untreated scaffolds drastically increased (up to 250%) after 1, 3, and 5 minutes of CAP treatment from 1.20 to 10.50, 22.90, and 27.60 nm respectively. The CAP treatment of PLA not only changed nanoscale roughness but also chemistry, as the XPS analysis showed that the ratio of oxygen to carbon increased significantly after CAP treatment. The results clearly confirmed that both hydrophilicity and nanoscale roughness changes to scaffolds after CAP treatment might play an important role in enhancing bone cell and mesenchymal stem cell attachment and functions. Slepicka et al. (2013) determined the surface properties of various polymers after Ar plasma treatment in order to conduct cytocompatibility tests. In the case of PLA, they found that the water contact angle decreased to about 32 and 28 degrees after Ar plasma treatment (150 seconds) of RF powers of 5 and 10 W respectively. The lower plasma power led to a lesser decrease of contact angle for all PLA biopolymers. This phenomenon is probably caused by the lower kinetic energy, charge, and density of impacted argon ions toward the polymer surface. The aging behavior of all PLA substrates treated by Ar plasma treatment was also studied. For PLA, the contact angle increased rapidly with increasing aging time and after 100 hours of aging, it achieved saturation at about 83 and 98 degrees for 5 and 10 W discharge power, respectively, as compared to 71 degrees for pristine PLA. Chaiwong et al. (2010) investigated the influence of SF6 plasma generated by an inductively coupled discharge on the hydrophobicity of PLA. They found that SF6 plasma treatment enhanced the hydrophobicity of PLA as observed in the increase of the water contact angle. The PLA surface was modified by the incorporation of fluorinated functional groups (CF, CF2, CF3) that enhanced the hydrophobicity character of the polymer surface. The reason behind this is that these fluorinated groups are apolar and, hence, reduce the surface energy of PLA resulting in an increased contact angle. The fluorinated polymer surface exhibited good chemical stability and improved barrier properties. These characteristics might be useful for packaging industries, however, for tissue engineering, SF6 plasma treatment cannot be used because of the increased hydrophobicity of the polymer surface. Surface modification of PLA through air atmospheric plasma treatment has been reported (Jorda et al., 2014). The adhesion properties of the treated surface were investigated by optimizing the process parameters in terms of the nozzlesubstrate distance and the sample advance rate. The authors used four different liquids (water, glycerol, diiodomethane, and formamide) for contact angle measurements and subsequent surface energy calculations. A circular nozzle with a rotation speed of 1900 rpm was used for surface modification while atmospheric plasma was applied at different nozzle-substrate distances between 10 and 20 mm

11.7 Plasma Treatment of PLA and PHAs

and at different sample advance rates ranging from 100 to 1000 mm/s. The best results were achieved for conditions of low advance rates (100 300 mm/s) and low nozzle-substrate distances of around 10 mm. Hirotsu et al. (2002) treated melt extruded PLA sheets with O2, He, and N2 plasmas to improve wettability. According to their observations, the plasma treatment did not affect PLA biodegradation in soil. An anhydrous NH3 plasma treatment was used by Yang et al. (2002) to improve the hydrophilicity and cell (human skin fibroblast) affinity of complex shapes, like porous PLA scaffolds, prepared using a particular technique. They found that reactive amine groups were created by the NH3 plasma on PLA scaffolds that anchored collagen through polar and hydrogen bonding interactions.

11.7.4 PLASMA TREATMENT OF PHAS As has been previously discussed, PHB and PHV are well known polymers in the category of PHAs. Most of the works related to plasma treatment have been done for PHB and PHV. Slepicka et al. (2014) studied the cytocompatibility of PHB modified by Ar plasma discharge as a function of plasma discharge power and the time of plasma exposure. The cytocompatibility study on pristine and modified PHB was carried out on mouse embryonic fibroblasts (NIT 3T3). They found that plasma treatment leads to an increase in PHB surface polarity. Exposure to an Ar plasma discharge power of 3 W for 40 seconds leads to a decrease of the contact angle to about 30 degrees as compared to the average contact angle of 66 degrees for pristine PHB. However, as the exposure time was increased beyond 40 seconds, the contact angle began to slowly increase. A similar trend was observed at a higher discharge power of 5 W. This increase of contact angle after the threshold exposure time might be because of the surface ablation and changes in the surface roughness (Siegel et al., 2008). The effect of plasma treatment on PHB for mouse embryonic fibroblasts (NIH 3T3) is shown in Fig. 11.7. The figure shows that the proliferation and growth of NIH 3T3 is satisfactory even on pristine PHB. However, plasma treatment leads to further improvement of PHB surface biocompatibility, manifesting in cell proliferation and growth. Fig. 11.8 shows images of cell proliferation and growth on pristine and plasma-modified samples. The images show that plasma activation enhances cell proliferation as well as cell homogeneity. A comparative study of the modification of PHB films by various plasmas, such as Ar, O2, H2O/O2, H2O, and H2O2, has been reported by Mas et al. (1997). The comparison between these plasma treatments indicated a better surface wettability, correlated to C OH bonds, obtained with H2O and H2O2 plasma than that obtained with Ar and O2 plasma which incorporated oxygen atoms in CQO bonds more easily. The lowest contact angles were obtained with H2O2 and H2O/ O2 plasma. The increase in plasma power had a negligible effect on the contact angles with H2O/O2 plasma, while it gave lower contact angles with Ar and O2 plasma.

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FIGURE 11.7 Number of adhered and proliferated NIH 3T3 cells 24, 48, and 72 h after seeding on pristine PHB, PHB, and plasma-modified PHB (8 W for 240 s). The values for tissue polystyrene (PS) are also shown for comparison. Reprinted from Slepicka, P., Styblova, S., Kasalkova, N.S., Rimplelova, S., Svorcik, V., 2014. Cytocompatibility of polyhydroxybutyrate modified by plasma discharge. Polym. Eng. Sci. 54, 1231 1238.

Two different kinds of PHB surfaces, PHB foil and PHB nonwoven fabric, modified by Ar plasma treatment have been studied by Slepicka et al. (2015). They found that the surface energy of both PHB foil and nonwoven fabric increased significantly after plasma treatment. In fact, PHB nonwoven fabric exhibited an almost immeasurable contact angle after plasma treatment. For PHB foil, the contact angle decreased sharply to 37 degrees as a result of 8 W plasma treatment with an exposure time of 240 seconds. They found that a higher plasma power exhibited nonlinear changes in the contact angle dependence curve. On the other hand, when the authors treated the PHB fabric with the same exposure time as was used for the PHB foils with the same plasma power, the surface became completely hydrophilic. The authors did not observe any influence on contact angle even after 2 weeks of aging the plasma treated PHB fabric. For the study of cytocompatibility, the PHB foils were seeded with NIH 3T3 cells. Plasma treatment had a positive influence on cell adhesion and proliferation in addition to enhancing cytocompatibility. They also found that the plasma treated PHB foils

11.7 Plasma Treatment of PLA and PHAs

FIGURE 11.8 Images of adhered and proliferated NIH 3T3 cells taken after the first, second, and third day from seeding on pristine PHB (first column), on plasma-modified PHB with 240 s of 8 W (second column), and on tissue polystyrene (third column). The white line represents 50 μm. Reprinted from Slepicka, P., Styblova, S., Kasalkova, N.S., Rimplelova, S., Svorcik, V., 2014. Cytocompatibility of polyhydroxybutyrate modified by plasma discharge. Polym. Eng. Sci. 54, 1231 1238.

preserved their cytocompatibility even after being annealed at 100 C. Therefore, PHB foil may find use as a novel biodegradable scaffold for cell growth. Further, the antibacterial properties of PHB fabric, using a Gram-negative bacterial strain of Escherichia coli, was also studied. For these antibacterial tests, two different types of surface modification were selected: plasma treated PHB fabric (240 seconds, 8 W) and another set of the same samples sputtered with silver (approximately 3 nm). The samples were in direct contact with the bacteria statically incubated in physiological solution for 3 hours followed by seeding on solid agar plates. The number of surviving bacterial colonies was evaluated for each sample and compared to a control (bacteria incubated without tested samples). The results of the antibacterial testing showed that the silver sputtering on PHB fabric led to the construction of an optimal antibacterial surface.

11.7.5 DISADVANTAGES OF PLASMA TREATMENT It is clear from the discussion in Section 11.7.4, that the plasma surface modification of polymers, which began in the 1960s, has been extensively and successfully

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used to improve the hydrophilicity and cell affinity of polyester surfaces. However, one cannot rule out the main disadvantage of this technique; that is, the effectiveness of surface modification is lost, to some extent, because of surface rearrangement (Ximing et al., 1992). The reason for this rearrangement of surface modifying species is to minimize interfacial energy by thermally activated macromolecular motions (Yang et al., 2002; Ximing et al., 1992; Occhiello et al., 1991; Safinia et al., 2005). Slepicka et al. (2013) observed an increase in contact angle due to the reorientation of polar groups on the surface layer of polymers. Attempts have been made to restrict the mobility of surface modifying species. It has been observed that modifying effects could be maintained if samples are preserved at low temperature (0 C 4 C) (Yang et al., 2002). However, this stabilization approach might not be practical for biomedical applications since this temperature range is much lower than physiological as well as room temperature. Another disadvantage of plasma treatment is that the treatment process itself can also affect the degradation of PLA. It has been observed that PLA degradation increased with an increase of plasma power and treatment time (Wan et al., 2006). The plasma-modification depth increased with treatment time, while plasma power influenced the depth only slightly. Therefore, it is obvious that some undesirable effects of surface modification make the plasma treatment process unsuitable for certain biomedical applications.

11.8 CONCLUSIONS PLA and PHAs polymers have proven potential as leading biomaterials for numerous applications in medical devices in the form of sutures; cardiovascular patches and orthopedic pins; tissue engineered nerves, tendon, and articular cartilages; repair patches; slings; adhesion barriers; stents; wound dressings; and as carriers for controlled drug release. PLA can be synthesized chemically while PHAs can be obtained through fermentation processes as an intracellular fermentation product. Low molecular weight biopolymers are desirable for biomedical applications because high molecular weight biopolymers have a complete resorption time of 2 8 years. As such, the prolonged existence of biopolymers in vivo in certain organs may lead to inflammation and infection. For biopolymers to be used successfully in tissue engineering, it is essential that these biomaterial scaffolds temporarily withstand and conduct the loads and stresses that the new tissue will ultimately bear. However, there are drawbacks in the form of their high brittleness which limits their use in flexible biomaterials. Many research efforts, such as the addition of natural additives and plasticizers, copolymerization, and blending with biopolymers, have been directed toward overcoming the high brittleness of PLA or PHAs based biopolymers. Plasma treatment of biopolymer surfaces is a technology that offers innovative solutions to adhesion and wetting problems in pharmaceutical industries. Polymer surfaces can be exposed to plasma generated

References

in O2, N2, He, Ar, or NH3. Plasma treatment results in an increase in surface roughness. Chemical modification of these biopolymers results in the introduction of specific functional groups onto their surfaces which increases their cytocompatibility. The SEM and AFM images as well as contact angle studies discussed in this chapter show that plasma activated biopolymers have enhanced cell proliferation and cell homogeneity. The main disadvantage of plasma treatment techniques is that the effectiveness of the surface modification is lost, to some extent, because of surface rearrangement. Another disadvantage is that the treatment process itself can also affect the degradation of biopolymers. Overall, it can be concluded that plasma treated PLA or PHAs can be used effectively in tissue engineering and other areas of biomedical engineering.

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The design of two different structural scaffolds using β-tricalcium phosphate (β-TCP) and collagen for bone tissue engineering

12

Takaaki Arahira1 and Mitsugu Todo2 1

Section of Bioengineering, Department of Dental Engineering, Fukuoka Dental College, Fukuoka, Japan 2Research Institute for Applied Mechanics, Kyushu University, Fukuoka, Japan

12.1 INTRODUCTION If a load-bearing bone, such as the tibia or femur, suffers a malignant tumor, such as an osteosarcoma, in orthopedic surgical treatment, a part of bone tissue containing the sarcoma is removed to prevent further metastasis and then reconstructed using bone grafts, such as autografts and allografts (Poffyn et al., 2011; Capanna et al., 2011). In the grafting techniques shown in Table. 12.1; autografts, allografts, and artificial bone substitutes are mainly used to reconstruct such damaged bones (Poffyn et al., 2011; Capanna et al., 2011). Although autograft and allograft are ideal treatments to obtain faster regeneration, there are some problems, such as for autograft, the extraction of healthy bone and the pain of the extracted parts and for the allograft, infections and rejections (Tiwari, 2012). Therefore, the need for artificial bone substitutes has been growing rapidly and different kinds of porous bioceramics have been developed as bone substitutes (Yamasaki et al., 2009); however, the rate of bone regeneration is not comparable with that of autografts and allografts and furthermore, the brittle characteristic with low fracture energy of bioceramics may cause the sudden collapse of implanted substitutes. To overcome such limitations of artificial bone substitutes, tissue engineering techniques have been developed to achieve faster regeneration of bone tissue with ideal mechanical properties compatible with bone tissues (Roohani-Esfahani et al., 2012; Collins et al., 2009; Bretcanu et al., 2009). In a typical tissue engineered bone regeneration approach, a regenerated bone graft is prepared by culturing and differentiating mesenchymal stem cells (MSCs) in a porous scaffold for a certain period of time and then they are implanted into the damaged region (Lu et al., 2011). The constituents and structure of the scaffold may strongly Materials for Biomedical Engineering: Hydrogels and Polymer-based Scaffolds. DOI: https://doi.org/10.1016/B978-0-12-816901-8.00012-2 © 2019 Elsevier Inc. All rights reserved.

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Table 12.1 Overview of the Treatment of Bone Regeneration Grafting Technique

Advantage

Problem

Autograft

• Faster regeneration • No rejection

Allografts

• Noninvasive

Artificial bone substitute

• Noninvasive • No rejection • Enough grafting • Faster regeneration • No rejection

• • • • • • •

Regenerated bone (mesenchymal stem cell loaded artificial bone substitute)

Extraction of healthy bone Pain of extracted parts Limited amount of extraction Infection Rejection Low mechanical property Insufficient balance between bone regeneration and degradation of material

Few studies have been done.

influence the proliferation and differentiation behavior of stem cells by functioning as an artificial extracellular matrix (ECM) (Arpornmaeklong et al., 2008). An ideal scaffold is biodegradable with an adjustable degradation rate that matches the speed of tissue regeneration, and has the proper mechanical properties to achieve mechanical compatibility with the surrounding tissues as well as a suitably porous structure to provide enough space for cell proliferation and ECM formation (Correlo et al., 2009; He et al., 2010). As shown in Table. 12.2, natural biopolymers (Ge´rard and Doillon, 2010; Girotto et al., 2003; Heinemann et al., 2008), synthetic biodegradable polymers (Lohan et al., 2011; Park and Todo, 2011; Porter et al., 2009), bioactive ceramics (Udoh et al., 2010; Zhao et al., 2011; Zhang et al., 2011), and their composite materials (Arahira and Todo, 2014; Arahira and Todo, 2016; Arpornmaeklong et al., 2008; Akkouch et al., 2011; Hiraoka et al., 2003; Intan et al., 2016; Salerno et al., 2011; Wiria et al., 2008) have been used as raw materials for scaffolds in tissue engineering. One of the most promising materials for scaffolds is type I collagen because it is the primary organic constituent of bone tissue. Collagen matrix scaffolds guarantee excellent biological compatibility with appropriate porosity and interconnected porous structures for cell proliferation and differentiation (Arahira and Todo, 2014; Hiraoka et al., 2003; Keogh et al., 2010; Liang et al., 2010). On the other hand, β-tricalcium phosphate (β-TCP) is a bioactive ceramic that has also been widely used in bone tissue engineering because of its good osteoconductivity, cellular adhesion, and accelerated differentiation (Lohfeld et al., 2012; Kang al., 2012). Furthermore, β-TCP has faster degradation than crystalline hydroxyapatite (Cao and Kuboyama, 2010). Organic/inorganic composite scaffolds, such as β-TCP/collagen with MSCs, have been considered as promising candidates in bone tissue engineering. Matsuno et al., developed composite scaffolds with collagen and β-TCP particles and examined the optimal ratio

12.1 Introduction

Table. 12.2 Summary of Biomaterials for Tissue Engineering Material

Application

Reference

Chitosan

Vascularized tissue Bone tissue

Hyaluronic acid

Cartilage tissue

Poly(L-lactide) (PLLA)

Bone tissue

Poly(ε-caprolactone) (PCL)

Bone tissue

Poly(glycolic acid) (PGA)

Cartilage tissue

Bioceramic

Hydroxyapatite (HA)

Composite materials

β-tricalcium phosphate (β-TCP) α-tricalcium phosphate (α-TCP) Chitosan/collagen

Bone substitute Bone tissue

Gérard and Doillon (2010) Heinemann et al. (2008) Girotto et al. (2003) Park and Todo (2011) Porter et al. (2009) Lohan et al. (2011) Zhao et al. (2011)

Natural polymer

Synthetic biodegradable polymer

Collagen

Bone substitute

PCL-HA

Bone tissue

Collagen/β-TCP

Bone tissue Bone tissue Bone tissue

Poly(L-lactide-coε-caprolactone) (PLCL) Collagen/HA/poly(lactideco-ε-caprolactone) Poly(vinyl alcohol)/HA β-TCP/PLLA

Cardiovascular tissue Bone tissue Tissue engineering Bone tissue

Zhang et al. (2011) Udoh et al. (2010) Arpornmaeklong et al. (2008) Salerno et al. (2011) Hiraoka et al. (2003) Arahira and Todo (2014) Arahira and Todo (2016) Intan et al. (2016) Akkouch et al. (2011) Wiria et al. (2008) Arahira et al. (2015)

of β-TCP granules to collagen solution (Matsuno et al., 2006). They also assessed the compressive mechanical properties of the composite scaffold in dry and wet conditions and found that such properties influence bone regeneration after implantation in vivo. They concluded that collagen/β-TCP scaffolds have adequate mechanical strength and compatibility for treating bone defects. In their study, the compressive mechanical properties of scaffolds without cells were

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evaluated. However, the mechanisms for the proliferation, differentiation of osteoblast, and ECM formation of MSCs in such composite scaffolds have not been well understood yet. MSCs have been widely investigated in tissue engineering techniques because MSCs have multilinage potential, such as osteoblasts, chondrocytes, and adipocytes (Caplan, 1991). It has been reported that the regeneration of osteochondral tissues may be effectively achieved by combining biodegradable scaffolds and MSCs (Pabbruwe et al., 2010; Wang et al., 2010). Organic/inorganic composite scaffolds, such as collagen/β-TCP with MSCs, have been considered as promising candidates in bone tissue engineering. In a typical bone reconstruction process, MSCs retrieved from a patient are cultured in a composite scaffold until bone-like structure is regenerated as a combination of MSCs, ECM, and the scaffold, and then the regenerated bone substitute is implanted into the damaged region. It is, therefore, ideal that the bone graft has mechanical compatibility with the surrounding bone tissues such that mechanical stimulus is transferred into the bone graft to complete bone regeneration. However, detailed mechanisms of proliferation, differentiation, and ECM formation of MSCs in a composite scaffold have not been well understood yet. In this chapter, two different structural scaffolds were fabricated by using β-TCP and collagen. One scaffold is collagen-based with β-TCP particles fabricated using freeze-drying methods. A collagen/β-TCP scaffold is known as a “particle distributed scaffold” (Arahira and Todo, 2014; Todo and Arahira, 2013; Arahira et al., 2011). The other scaffold is β-TCP-based with a porous collagen structure fabricated using the polyurethane template method and a subsequent freeze-drying method. A β-TCP/collagen scaffold is known as a “two phase structural scaffold” (Arahira and Todo, 2016). Rat bone marrow mesenchymal stem cells (rMSC) were cultured in two different structural composite scaffolds for up to 28 days in order to evaluate the time-dependent behavior of the ECM formation and the mechanical performance of the scaffold
cell system.

12.2 COLLAGEN-BASED POROUS SCAFFOLD 12.2.1 FABRICATION AND CHARACTERIZATION OF PARTICLE DISTRIBUTED SCAFFOLD 12.2.1.1 Fabrication of particle distributed scaffold A type-I collagen solution (Nippon Meat Packers Inc., Japan) was used to fabricate porous collagen and collagen/β-TCP composite scaffolds using a freezedrying method (Lu et al., 2010; Arahira and Todo, 2014). Briefly, the type-I collagen was derived from porcine skin, the concentration of collagen was 1 wt%, and the solvent was hydrochloric acid with pH between 2.5 and 3.5. β-TCP powder (β-TCP-100, Taihei Chemical Industrial Co., Japan) was used to fabricate the collagen/β-TCP composite scaffold. The collagen solution and the β-TCP

12.2 Collagen-Based Porous Scaffold

powder were mixed following a fixed weight ratio of 90:10 using a magnetic stirrer. The mixed solution was poured into silicon rubber molds of 10 mm diameter and 5 mm depth, and then frozen at
80 C for 2 hours in a deep freezer. These frozen samples were then freeze-dried at
50 C for 24 hours using a lyophilizer equipped with a vacuum pump (EYELA FDU-1200, Tokyo Rikakikai Co., Japan). The freeze-dried scaffolds were then crosslinked under an atmosphere of glutaraldehyde (Wako Pure Chemical Industries., Japan) vapor at 37 C for 4 hours. After the crosslinking process, the scaffolds were treated with 0.1 M glycine aqueous solution to block unreacted aldehyde, after which they were washed with deionized water and lyophilized (Lu et al., 2010; Arahira and Todo, 2014).

12.2.1.2 Characterization of particle distributed scaffold The surfaces of the specimens were observed using a field emission scanning electron microscope (FE-SEM) (S-4100, Hitachi, Ltd., Japan) in order to assess the porous microstructures of the scaffolds. Scaffolds with proliferated cells were also washed with a phosphate buffer saline (PBS, pH 5 7.4) solution to remove the media, and then dehydrated with ethanol, immersed in a t-butyl alcohol solution, and finally freeze-dried using a freeze-drying machine (ES-2030, Hitachi, Ltd., Japan). Scaffolds were mounted on aluminum stages and sputter-coated with Pt-Pd using an anion sputter coater (E-1030, Hitachi, Ltd., Japan). The porosity values of the scaffolds were evaluated by Archimedes’ Principle (Oh et al., 2007). Briefly, each scaffold was thoroughly immersed in a glass bottle filled with ethanol. The scaffolds with ethanol soaked into the pores were removed from the bottle. The porosity was then determined using: Porosityð%Þ 5

ðW2 2 W3 2 WS Þ=ρe 3 100 ðW1 2 W3 Þ=ρe

(12.1)

where W1 and W2 are the weights of the glass bottle filled with ethanol before and after immersion of the scaffold, respectively. W3 is the weight of the glass bottle with ethanol after the removal of the scaffold. WS is the weight of the dried scaffold and ρe is the density of ethanol. The swelling behavior of each scaffold was evaluated using the swelling ratio quantitatively. The scaffolds were immersed in PBS solution for 1, 3, 24, and 48 hours at 37 C in a humidified atmosphere of 5% CO2. At each sampling point, those scaffolds were taken out of the PBS solution, and then excess PBS was removed using filter paper. The swelling ratio is then evaluated using the equation (Sang et al., 2011): Swelling ratio 5

Wwet 2 Wdry Wdry

(12.2)

where Wdry is the initial weight of the dried scaffold and Wwet is the weight of the scaffold with soaked PBS. Compressive moduli of the scaffolds were periodically measured using a universal testing machine (EZ-Test, Shimadzu Co., Japan) at a loading rate of 1 mm/

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min under dry and wet conditions. Load-displacement relations were recorded using a personal computer and stress
strain relations were then evaluated from the load-displacement relations. The stress (σ) and strain (ε) were evaluated using: σ5

4F d2

(12.3)

ε5

ΔL L

(12.4)

where F is the load under the compressive test; d is the diameter of the scaffold; L is the height of the scaffold; and ΔL is displacement after loading at each time point.

12.2.2 IN VITRO CELL EXPERIMENT 12.2.2.1 Cell culture rMSCs (DS Pharma Biomedical Co., Japan) were precultured in a tissue culture flask with cell growth medium consisting of alpha-minimal essential medium (α-MEM) (Wako Pure Chemical Industries., Japan) supplemented with 10% fetal bovine serum and 1% penicillin-streptomycin (MP Biomedicals LLC., Japan) at 37 C in a humidified atmosphere of 5% CO2. In this study, rMSCs were passage 3
5. The medium was changed twice per week. After rMSCs reached 80%
90% confluence, they were removed from the flask using a cell desquamation solution (Accutase Merck Millipore Co., Japan) and collected. 1 3 105 cells suspended in 10 μL of α-MEM medium were seeded in each of the scaffolds and then they were incubated for 1 hour to make the rMSCs adhere to the surfaces of the scaffolds. After the preincubation process, these scaffolds were transferred to 12-well plates containing 2 mL of differentiation medium per well. The differentiation medium was composed of the cell growth medium previously described and a supplement of osteoblast differentiation (KE-200, DS Pharma Biomedical Co., Japan), including MEM, β-glycerophosphate, L-ascorbic acid, and dexamethasone. The plates were then incubated at 37 C in a humidified atmosphere of 5% CO2. The medium was also changed twice per week.

12.2.2.2 Compression test Compression tests of the scaffolds with cells were conducted periodically by using a universal testing machine (EZ-Test, SHIMADZU Co., Japan) at a loading rate of 1 mm/min after the cell culture periods for 7, 14, 21, and 28 days. Loaddisplacement relations were recorded using a personal computer and stress
strain relations were then evaluated from the load-displacement relations. The stress (σ) and strain (ε) were evaluated using Eqs. (12.3) and (12.4). For each of the specimens, the average compressive modulus was obtained as the initial linear slope of the stress
strain relation. Three specimens were tested for each condition.

12.2 Collagen-Based Porous Scaffold

The variations in the increase of compressive modulus between scaffolds with and without rMSCs during culture were also evaluated.

12.2.2.3 Microstructural characterization The surfaces of the scaffolds were observed using a FE-SEM (S-4100, Hitachi, Ltd., Japan) in order to characterize porous microstructures and proliferation behavior, including the formation of ECM by rMSCs in the scaffolds. Briefly, samples for FE-SEM observation were prepared using this procedure (Arahira and Todo, 2014): Scaffolds with proliferated cells were washed by PBS to remove the media and dehydrated with ethanol, then immersed in a t-butyl alcohol solution, and finally freeze-dried using a freeze-drying machine (ES-2030, Hitachi, Ltd., Japan). The freeze-dried scaffolds were mounted on aluminum stages and sputter-coated with Pt-Pd using an anion sputter coater (E-1030, Hitachi, Ltd.).

12.2.2.4 Evaluation of cell number and alkaline phosphatase activity Cell number and alkaline phosphatase (ALP) activity were evaluated using a spectrophotometric micro plate reader (2030 ARVO X2, Perkin Elmer Co., Japan). The cell number was estimated using a cell counting kit (Dojindo Laboratories, Japan). Scaffolds with rMSCs were removed from the 12-well plate after the culture periods of up to 28 days, washed several times, and then sliced into small pieces. These pieces, soaked in PBS, were placed into centrifuging tubes with reagent solution and reacted for 2 hours at 37 C in a humidified atmosphere of 5% CO2. The light absorptions of the reaction solutions were then measured by the plate reader at a wavelength of 450 nm. ALP was measured to evaluate the osteogenic differentiation of rMSCs quantitatively. Specimens for the assay of ALP were prepared by means of Fujita’s protocol (Fujita et al., 2005). Briefly, scaffolds with rMSCs were washed by PBS and then frozen at
30 C. After repeating the freezing-thawing process three times, the specimens were assayed using an ALP kit (The Labassay ALP kit, Wako Pure Chemical Ind., Japan). After each specimen was placed in the centrifuging tube, the buffered substrate (p-nitrophenylphosphate disodium 6.7 mmol/L, pH 9.8) was added and reacted for 15 minutes at 37 C in a humidified atmosphere of 5% CO2. The reaction of the solution was stopped with aqueous sodium hydroxide (0.2 mol/L) and the production of p-nitrophenol was measured by the plate reader at a wavelength of 405 nm.

12.2.2.5 Gene expression analysis Gene expression of osteogenesis was analyzed using real-time PCR reactions with primers of β-actin, collagen type I, and osteocalcin. Collagen and collagen/β-TCP scaffolds with rMSCs were washed with PBS and then frozen in liquid nitrogen. Each of the frozen scaffolds were crushed into powder in a mortar using a muddler. The RNA of the cells was isolated from the powder from each scaffold using NucleoSpin RNAII (MACHEREY-NAGEL GmbH & Co.). For each scaffold, three samples were prepared for the RT-PCR measurement. RT-PCR

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CHAPTER 12 The design of two different structural scaffolds

amplification was performed using the reagent (SYBR premix EX Taq Perfect real-time kit, Takara Bio Inc.) with a RT-PCR system (the thermal cycler Dice RealTime System TP960, Takara Bio Inc.). β-actin was used as the house keeping control and results were quantified for collagen I and osteocalcin using the ΔΔCt relative quantification method (Bonanomi et al., 2003).

12.2.2.6 Statistics All the data are presented as mean 6 standard deviation and were derived from three independent samples. Data analysis was performed using analysis of variance (ANOVA). Any difference was considered statistically significant when the P value was ,.05.

12.3 EXPERIMENTAL RESULTS 12.3.1 CHARACTERIZATION OF PARTICLE DISTRIBUTED SCAFFOLD Optical and SEM micrographs of collagen and collagen/β-TCP scaffolds are shown in Fig. 12.1A. The mean diameters of the collagen and collagen/β-TCP scaffolds are 7 and 8 mm, respectively. Both scaffolds show similar porous structure, however, the walls of the composite scaffold are straighter, in terms of alignment, than the collagen scaffold, indicating improved stiffness by distribution of β-TCP particles. The representative pore size is distributed in the range of 50
200 μm in the collagen scaffold and 50
150 μm in the composite scaffold. The surface morphology of the collagen scaffold (Fig. 12.1A) is characterized by entangled collagen fibrils, while the composite scaffold (Fig. 12.1A) shows rough surface with bared β-TCP particles. The porosity of the collagen scaffold evaluated from Eq. (12.1) was about 99.6%, while that of the composite scaffold was 97.0%, indicating that the addition of β-TCP particles slightly reduces the porosity of collagen scaffolds. Swelling ratio is shown in Fig. 12.1B. The swelling ratio of the collagen/ β-TCP scaffold reaches its maximum of about 4.1 at 3 hours, and then keeps constant. On the other hand, the swelling ratio of the collagen scaffold exhibits rapid increases up to 15.7, and then slowly increases up to about 18.0, and then becomes constant. It was, thus, clearly indicated that the collagen scaffold shows much higher swelling ratio, indicating that the deformability of the collagen scaffold tends to increase the amount of absorbed PBS. It was also noted that the addition of β-TCP particles results in the structural stability of the composite scaffold. The compressive modulus is presented in Fig. 12.1C. The compressive modulus, under the dry condition for the collagen/β-TCP scaffold, increases compared to that of the collagen scaffold. In addition, under the wet condition, that of the collagen/β-TCP scaffold is higher than that of the collagen scaffold.

12.3 Experimental Results

FIGURE 12.1 Material characteristic evaluation of collagen and collagen/β-TCP scaffolds. (A) Optical and SEM micrographs of collagen and collagen/β-TCP scaffolds. (B) Swelling ratio. The result is presented as the mean 6 standard deviation for n 5 5. # indicate P ,.01 vs collagen/β-TCP scaffold at each day. (C) Compressive mechanical properties of collagen and collagen/β-TCP scaffolds.  , # indicate P ,.05 vs collagen scaffold (dry) and P ,.05 vs collagen/β-TCP scaffold (wet), respectively (Arahira and Todo, 2014).

12.3.2 RESULTS OF IN VITRO CELL EXPERIMENT The variations of compressive modulus during cell culture are shown in Fig. 12.2A and B. First of all, it was noted that the modulus of the composite scaffold was much higher than that of the collagen scaffold, clearly indicating that the distribution of β-TCP particles dramatically enhanced the elastic modulus under compression. The moduli of the collagen scaffolds with and without cells tend to decrease gradually until 14 days. The modulus of the collagen scaffold with cells slightly increased with increasing culture period after 14 days, while that of the collagen scaffold without cells kept almost constant. On the contrary, the moduli of the collagen/β-TCP scaffolds with and without cells obviously increased during the culture period of 7 days. The modulus of the composite scaffold with cells decreased from 7 to 21 days, and then increased until 28 days. It was indicated that the modulus at 28 days was recovered to be almost the same as that at 7 days. The modulus of the composite scaffold without cells decreased from 7 to 14 days, and slightly increased until 28 days.

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CHAPTER 12 The design of two different structural scaffolds

(A)

(B) 25 Collagen (with cells) Collagen

2.5

Compressive modulus (kPa)

Compressive modulus (kPa)

3

2 1.5 1 0.5

Collagen/β-TCP (with cells) Collagen/β-TCP

20

15

10

5

0

0 0

5

(C)

10 15 20 Days in culture

25

30

10 15 20 Days in culture

(D)

25

30

1 Collagen Collagen/β-TCP

Collagen Collagen/β-TCP

50

#

0.8 p-Nitrophenol (mmol/l)

Cell number ( x 104)

5

0

60

40 30

#

20

0.6

#

#

0.4 #

0.2 10 0

0 0

5

(E)

10 15 20 Days in culture

25

30

1.4

0

5

(F)

1.6

10 15 20 Days in culture

25

30

3

Collagen Collagen/β-TCP

2.5

Collagen Collagen/β-TCP

Gene expression

1.2 Gene expression

380

1 0.8 0.6 0.4

2 1.5 1 0.5

0.2 0 7

14

21

Days in culture

28

0 7

14

21

28

Days in culture

FIGURE 12.2 (Continued)

12.3 Experimental Results

L

Variations of cell proliferation are shown in Fig. 12.2C. The cell number in the collagen scaffold increased until 14 days and then kept constant until 28 days. On the contrary, the cell number in the collagen/β-TCP scaffold kept increasing until 28 days. The variations of ALP activity are shown in Fig. 12.2D. The ALP activity in the collagen scaffold increased until 14 days and was 2.5 times larger than the initial value at the first day of cell culture. The ALP activity decreased from 14 to 28 days. In the case of the collagen/β-TCP scaffold, the ALP activity kept increasing until 28 days with stepwise behavior. The ALP activity became about twice that of the initial value at 21 and 28 days. It was also found that the maximum value of ALP activity in the collagen/β-TCP scaffold was about twice that of the collagen scaffold at 14 days. Results of the RT-PCR related to osteogenic genes expression, such as type I collagen and osteocalcin, are shown in Fig. 12.2E and F, respectively. The expression of type I collagen in the composite scaffold was higher than in the collagen scaffold at all days examined. In the case of the collagen scaffold, type I collagen tended to decrease after 14 days, while in that of the collagen/β-TCP scaffold, it kept increasing until 21 days and became almost constant up until 28 days. Osteocalcin showed higher expression in the collagen scaffold than in the composite scaffold from 7 to 14 days. These results clearly show that the generation of osteogenic genes was much more activated in the collagen/β-TCP scaffold than in the collagen scaffold. FE-SEM micrographs of the surface region of the collagen and the collagen/ β-TCP scaffolds with cells are shown in Fig. 12.3. In the collagen scaffold, round-shaped cells were observed at 1 hour (Fig. 12.3A and a). Collagen matrices and network structures constructed by the collagen fibrils existed on the surface at 7 and 14 days (Fig. 12.3B, b, C, and c). Small spherical structures were observed with the collagen networks at 21 and 28 days (Fig. 12.3D, d, E, and e). These spherical structures were considered to be matrix vesicles produced by osteoblasts. However, the porous structure of collagen scaffold was unstable in the medium due to swelling. On the contrary, in the collagen/β-TCP scaffold, flattened cells with filopodia were seen at 1 hour (Fig. 12.3A and a). After 7 days, collagen matrices were created actively and thin network structures were formed

Results of cell experiments. (A, B) Variation of compressive modulus of collagen and collagen/β-TCP scaffolds during cell culture. (C) Variation of cell number during cell culture. The result is presented as the mean 6 standard deviation for n 5 3. #,  indicate P ,.05 vs collagen scaffold at 1 day and P ,.05 vs collagen scaffold at 1 day, respectively. (D) Variation of ALP activity during cell culture. The result is presented as the mean 6 standard deviation for n 5 3.  , # indicate P , 0.05 vs collagen/β-TCP scaffold at 1 day and P , 0.05 vs collagen scaffold at each day, respectively. (E, F) Gene expression of type I collagen and osteocalcin. The results are presented as the mean 6 standard deviation for n 5 3.  indicate P ,.05 vs collagen scaffold at 7 days (Arahira and Todo, 2014).

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CHAPTER 12 The design of two different structural scaffolds

FIGURE 12.3 Field emission scanning electron microscope micrographs of collagen and collagen/β-TCP scaffolds with rat bone marrow mesenchymal stem cells (Arahira and Todo, 2014).

(e.g., Fig. 12.3C and c). Matrix vesicles were also formed on the collagen networks and turned into calcified matrices (Fig. 12.3d and e), which are much larger than matrix vesicles. The surfaces of the collagen/β-TCP scaffold were fully covered by proliferated cells and ECM more than those of the collagen scaffolds. It was also observed that the tissue-like structures consisting of collagen fibrils and calcified matrices became thick gradually as culture period increased.

12.4 MECHANISM OF VARIATIONAL MECHANICAL BEHAVIOR BETWEEN SCAFFOLD STRUCTURE AND CELL RESPONSE In this section, a porous collagen/β-TCP scaffold, known as the particle distributed scaffold, was fabricated using the freeze-drying method and compared with a collagen scaffold in order to clarify the effectiveness of β-TCP particles on structural, mechanical, and biological properties. It was shown that the collagen/β-TCP

12.4 Mechanism of Variational Mechanical Behavior

scaffold possessed continuous porous structure with a range of pore sizes from 50 to 150 μm, which is known to be suitable for tissue ingrowth and cell attachment (Ge´rard and Doillon, 2010). A scaffold is usually used to culture cells and to create a tissue-like structure in vitro, and then it is expected to be implanted into the damage part of the target tissue. It is, therefore, important for the scaffold to maintain its original structure during cell culture in vitro (Chen et al., 2000). The experimental results obtained clearly indicated that the addition of β-TCP particles into the collagen matrix effectively improved structural stability as shown in Fig. 12.1A, indicating that the composite scaffold is suitable for cell culture. The cell number in the collagen/β-TCP scaffold increased until 28 days, on the other hand, that of the collagen scaffold increased until 14 days, and then almost kept constant to 28 days. In the case of the collagen scaffold, the proliferation of rMSCs became slow after reaching confluent at 14 days because cell proliferation inside the scaffold was thought to be suppressed because of the shrinkage and ductile deformation of the porous structure caused by swelling. On the contrary, in that of the collagen/β-TCP scaffold, rMSCs were well proliferated into the porous structure of the composite scaffold because of the improved structural stability achieved by the addition of β-TCP powder. In the collagen scaffold, pore size increased with the increase in swelling time. On the contrary, in the composite scaffold, the pore size kept constant (Arahira and Todo, 2014). The large pore size also improved cell proliferation and attachment (Murphy et al. 2010). It was also noted that structural stability is another important factor for effective proliferation. Therefore, the collagen/β-TCP scaffold showed optimal pore size and good structural stability due to β-TCP particles. ALP is a well-known enzyme produced by osteoblasts and is recognized as one of the most important osteoblastic makers. Wang et al. (2003) reported that ALP increased rapidly and peaked at about 2 weeks of culture. In the case of the collagen scaffold, ALP showed its peak at 14 days, and then decreased until 28 days. On the contrary, in that of the collagen/β-TCP scaffold, ALP dramatically increased from 14 to 21 days and became twice that of the initial value at 21 days of culture. β-TCP is thought to provide attractive environments for osteoblastic cells, such as adhereable surface morphology, nucleation sites of apatite crystals, and culture medium with high concentrations of calcium and phosphate ions released from non-sintered β-TCP particles (Wang et al., 2003; Zhang and Zhang, 2001). Those environmental factors are considered to activate the differentiation of rMSCs into osteoblasts. Cell growth behavior and formation of ECM were characterized by FE-SEM. In both the scaffolds, collagen fibrils, matrix vesicles, and mineralized nodules were recognized as components of ECMs. Network structure connecting cells with association of collagen fibrils were observed at 7 days. Many spherical structures appeared at 14 days and they were thought to be matrix vesicles and mineralized nodules produced by osteoblasts differentiated from rMSCs. A layered structure of ECM consisting of collagen matrices was formed at 21 days. Spherical structures with diameters between 40 and 200 nm were thought to be

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CHAPTER 12 The design of two different structural scaffolds

matrix vesicles, which are considered to play an important role in the calcification process of bone formation (Ozawa et al., 2008). The RT-PCR results showed that all osteoblastic markers in the collagen/ β-TCP scaffold were much higher than those in the collagen scaffold. Osteocalcin is known to appear in the late stage of osteoblastic differentiation (Wang et al., 2003). The expression of osteocalcin in both the scaffolds increased until 21 days. The osteocalcin level in the collagen scaffold decreased after 21 days, however, that in the collagen/β-TCP scaffold kept constant until 28 days, indicating that the collagen/β-TCP scaffold can maintain the activation of osteoblastic differentiation. It was noted that this difference is due to the osteoblastic differentiation ability of β-TCP. In this chapter, the focus is on the effects of the proliferation and differentiation of rMSCs on the compressive mechanical property of collagen/β-TCP scaffolds in comparison to collagen scaffolds. Variations in the increase of compressive modulus between scaffolds with and without rMSCs are shown in Fig. 12.4. The variational mechanical behaviors of the collagen and collagen/ β-TCP scaffolds were characterized by dividing them into two and three stages, respectively. In stage A of the collagen scaffold, the degradation of the scaffold strongly affected the modulus change more than the strengthening effect caused by cell proliferation and migration. In stage B, the compressive modulus increased due to the formation of collagen networks by osteoblasts differentiated from rMSCs. The compressive modulus was almost recovered to the initial value at 28 days of culture. In stage A of the collagen/β-TCP scaffold, the compressive modulus dramatically increased due to cell proliferation and the formation of an ECM network structure composed of a collagen matrix. It was noted that the cell (B)

3 Degradation of scaffold

2.5 2

Formation of collagen network by osteoblasts

Cell proliferation and migration

1.5 1

Stage a

Stage b

0.5 0

–0.5 0

5

10

15

20

Days in culture

25

15

Cell proliferation Differentiation into osteoblast Degradation of scaffold

Increase of modulus (kPa)

(A)

Increase of modulus (kPa)

384

30

10

Stage A

Stage B

Stage C

5

0 Differentiation Calcification

Cell proliferation Formation of ECM network structure

–5

0

5

10

15

20

25

30

Days in culture

FIGURE 12.4 Variation of increase of modulus during cell culture. (A) Collagen scaffold. (B) Collagen/ β-TCP scaffold (Arahira and Todo, 2014).

12.5 β-TCP-Based Porous Scaffold

proliferation and the formation of the ECM network structure were supported by the structural stability improved by the addition of β-TCP powders. In stage B, it was thought that rMSCs continued to proliferate and differentiate into osteoblasts, however, the degradation of the scaffold progressed gradually due to the decomposition of β-TCP powders. In stage C, the compressive modulus increased again due to a hardening effect mainly achieved by calcification.

12.5 β-TCP-BASED POROUS SCAFFOLD 12.5.1 FABRICATION AND CHARACTERIZATION OF TWO PHASE STRUCTURAL SCAFFOLD 12.5.1.1 Fabrication of two phase structural scaffold The β-TCP scaffolds were fabricated using the template method using polyurethane (PU) sponges (HR-20, Bridgestone, Japan) (Arahira and Todo, 2016; Nikaido et al., 2013; Udoh et al., 2010). The pore diameters of the PU sponge ranged from 200 to 500 μm and the fiber width was 80 μm. A 5 wt% poly vinyl alcohol (PVA) solution was used as the binder. β-TCP powder and PVA solution (10 g/10 mL) were mixed using a rotator. The PU sponge (10 3 10 3 10 mm3) was immersed in this slurry, and then excess slurry was removed carefully. The PU sponge coated with slurry was dried at room temperature for 24 hours. After drying, the β-TCP-coated PU sponge was heated from room temperature to 400 C for 6 hours and 40 minutes, and then from 400 C to 1100 C for 5 hours. After sintering, the β-TCP scaffolds were allowed to slowly cool down inside the furnace to room temperature. A type I collagen solution was used to fabricate β-TCP/collagen scaffolds using the same method as described in Section 12.2.1.1 (Lu et al., 2010). Briefly, the type I collagen solution was poured into the β-TCP scaffold and then frozen at 280 C for 2 hours in a deep freezer. These samples were then freeze-dried at 250 C for 24 hours using a lyophilizer equipped with a vacuum pump (EYELA FDU-1200, Tokyo Rikakikai Co., Japan). The freezedried scaffolds were then crosslinked under glutaraldehyde atmosphere (Wako Pure Chemical Industries, Japan) vapor at 37 C for 4 hours, and then were treated with a 0.1 M glycine water solution to block the unreacted aldehyde. After the blocking, the scaffolds were rinsed with deionized water and lyophilized (Lu et al., 2010).

12.5.1.2 Characterization of two phase structural scaffold •

•

The surfaces of the scaffolds were examined using an FE-SEM (S-4100, Hitachi, Ltd., Japan) to characterize the porous microstructures of the scaffolds. The scaffolds were mounted on aluminum stages and sputter-coated with Pt-Pd using an anion sputter coater (E-1030, Hitachi, Ltd., Japan). Porosity values of the scaffolds were evaluated using Archimedes’ Principle as described in Section 12.2.1.2 (Arahira and Todo, 2014; Oh et al., 2007).

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•

•

•

Briefly, each scaffold was thoroughly immersed into a glass bottle filled with ethanol. The scaffolds with ethanol soaked into the pores were removed from the bottle. Their porosity was then determined using Eq. (12.1). The swelling behavior of each scaffold was quantitatively evaluated by the same method as described in Section 12.2.1.2. Briefly, the scaffolds were immersed in PBS solution for 1, 3, 24, and 48 hours at 37 C in a humidified atmosphere of 5% CO2. At each sampling point, the scaffolds were removed from the PBS solution, and the excess PBS was removed using filter paper. The swelling ratio was then evaluated using Eq. (12.2) (Arahira and Todo, 2014; Sang et al., 2011). The compositions of the β-TCP-based scaffolds were evaluated using powder X-ray diffraction (XRD). The β-TCP and β-TCP/collagen scaffolds were ground into fine powder before the XRD analysis. The XRD patterns of each scaffold were recorded using a vertically mounted diffractometer system (RINT 2500 V, Rigaku, Tokyo, Japan) with Ni-filtered CuKα radiation generated at 40 kV and 100 mA. Samples were scanned from 20 to 80 2θ in continuous mode. The compressive modulus and strength of the scaffolds were periodically measured using a universal testing machine (EZ-Test, Shimadzu Co., Japan) at a loading rate of 1 mm/min under dry and wet conditions. Load-displacement relations were recorded using a personal computer and stress
strain relations were then calculated from the load-displacement relations. The stress (σ) and strain (ε) were then calculated using the formula: σ5

F a2

(12.5)

ε5

ΔL L

(12.6)

where F is the load under the compressive test; a is the side of the scaffold; L is the height of the scaffold; and ΔL is displacement after loading at each time point. For each specimen, the average compressive modulus was obtained as the initial linear slope of the stress
strain relation. The compressive strength was also obtained as the maximum load value. Three specimens were tested for each condition.

12.6 IN VITRO CELL EXPERIMENT 12.6.1 CELL CULTURE rMSCs (DS Pharma Biomedical Co., Japan) were precultured using the same procedure as described in Section 12.2.2.1. Briefly, they were cultured in a tissue culture flask with a medium comprising of α-MEM (Wako Pure Chemical

12.6 In Vitro Cell Experiment

Industries., Japan) supplemented with 10% fetal bovine serum and 1% penicillinstreptomycin (MP Biomedicals LLC., Japan) at 37 C in a humidified atmosphere of 5% CO2. After the rMSCs reached 80%
90% confluence, they were removed from the flask and collected. In total, 1 3 105 cells, suspended in 10 μL of α-MEM medium, were seeded in each of the scaffolds, and then incubated for 1 hour to cause the rMSCs to adhere to the surfaces of the scaffolds. After this process, these scaffolds were transferred to 12-well plates containing 2 mL of differentiation medium per well. The differentiation medium used was the same compositions as described in Section 12.2.2.1. The plates were then incubated at 37 C in a humidified atmosphere of 5% CO2. The medium was changed twice per week.

12.6.2 EVALUATION OF MECHANICAL CHARACTERISTICS The compressive modulus and strength of the scaffolds with proliferated cells were evaluated using the same procedure as described in Section 12.5.1.2. The specimen sampling times were at 7, 14, 21, and 28 days. The stress
strain relations were evaluated based on the load-displacement relations. For each of the specimens, the average compressive modulus was obtained as the initial linear slope of the stress
strain relation. Compressive strength was obtained as the maximum load value. Absorption energy was also calculated by the area between the stress
strain curve and transverse. Three specimens were tested for each condition.

12.6.3 MICROSTRUCTURAL CHARACTERIZATION The surfaces of the scaffolds were examined using the same procedure as described in Section 12.2.2.3. Briefly, the samples for FE-SEM observation were prepared using this procedure: Scaffolds with proliferated cells were washed with PBS to remove the medium and then dehydrated with ethanol, immersed in tbutyl alcohol, and finally freeze-dried. The freeze-dried scaffolds were mounted on aluminum stages and sputter-coated with Pt-Pd.

12.6.4 EVALUATION OF CELL NUMBER AND ALKALINE PHOSPHATASE ACTIVITY Cell number and ALP activity were also evaluated using the same procedure as described in Section 12.2.2.4. Cell number and ALP activity were measured using a cell counting kit (Dojindo Laboratories, Japan) and an ALP kit (Labassay ALP kit, Wako Pure Chemical Ind., Japan), respectively. In the case of cell number, light absorptions of the reaction solutions were then measured by the plate reader at a wavelength of 450 nm. In the case of ALP activity, the production of p-nitrophenol was measured by the plate reader at a wavelength of 405 nm.

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CHAPTER 12 The design of two different structural scaffolds

12.6.5 GENE EXPRESSION ANALYSIS The gene expression of osteogenesis was analyzed using the same procedure as described in Section 12.2.2.5. The RNA of the cells was isolated from each scaffold using NucleoSpin RNAII (MACHEREY-NAGEL GmbH & Co.). For each scaffold, three samples were prepared for RT-PCR measurement. RT-PCR amplification was performed using a SYBR premix EX Taq Perfect real-time kit (Takara Bio Inc.) with a RT-PCR system (the thermal cycler Dice RealTime System TP960, Takara Bio Inc.). β-Actin was used as the control, and the results were quantified for collagen type I and osteocalcin using the ΔΔCt relative quantification method (Bonanomi et al., 2003).

12.6.6 ALIZARIN RED S STAINING The area of osteoblastic mineralization was examined using Alizarin red S staining. The specimens were fixed with 95% ethanol and stained with 1% Alizarin red S (Wako Pure Chemicals Industries). After staining, the samples were examined under an optical microscope.

12.6.7 STATISTICS All the data are presented as mean 6 standard deviation and were derived from three independent samples. Data analysis was performed using Fisher’s least significant difference method. Any difference was considered statistically significant when the P value was ,.05 and ,.01.

12.7 EXPERIMENTAL RESULTS 12.7.1 CHARACTERIZATION OF TWO PHASE STRUCTURAL SCAFFOLD FE-SEM micrographs of the β-TCP and β-TCP/collagen scaffolds are presented in Fig. 12.5A. Both scaffolds have similar porous structures, and the β-TCP/collagen scaffold has another porous structure by collagen fibril. The average pore size is 200 μm in the β-TCP part but is distributed in the range of 50
200 μm in the collagen phase. The surface morphology of the β-TCP scaffold is connected by β-TCP particles, and a gap is observed between the β-TCP particles in certain areas. The β-TCP/collagen scaffold contains a β-TCP frame filled with collagen fiber. The porosity of the β-TCP scaffold calculated from Eq. (12.1) was approximately 97.2%, whereas that of the β-TCP/collagen scaffold was 98.1%. This indicates that the insertion of the collagen phase slightly increases the porosity of the β-TCP scaffold.

12.7 Experimental Results

FIGURE 12.5 Material characteristic evaluation of β-TCP and β-TCP/collagen scaffolds. (A) Optical and SEM micrographs of β-TCP (left side) and β-TCP/collagen scaffolds (right side). (B) X-ray powder diffraction patterns of β-TCP and β-TCP/collagen scaffolds. (C) Variation of swelling ratio during soaking experiment. The results are presented as mean 6 standard deviation for n 5 5. (D, E) Compressive mechanical properties of β-TCP and β-TCP/ collagen scaffolds.  , #, and $ indicate P ,.05 vs β-TCP scaffold (dry), P ,.05 vs β-TCP scaffold (wet), and P ,.05 vs β-TCP scaffold (dry), respectively (Arahira and Todo, 2016).

Fig. 12.5B presents the XRD patterns of the β-TCP and β-TCP/collagen scaffolds. XRD patterns for α-TCP and β-TCP powders are shown as a reference. The XRD pattern of the scaffold exhibited several broad peaks between 10 and 60 degrees assigned to the diffraction peaks for β-TCP. It was indicated that by heating the scaffold at 1100 C for 5 hours the β-TCP phase is obtained. Swelling ratio is shown in Fig. 12.5C. The swelling ratio of the β-TCP/collagen scaffold reaches its maximum of approximately 4 at 24 hours and then remains constant. The swelling ratio of the β-TCP scaffold also exhibits an increase to approximately 3 and then slowly increases to approximately 3.5. The β-TCP/collagen scaffold, thus, exhibits a much higher swelling ratio, indicating that the deformability of the collagen phase tends to increase the amount of absorbed PBS. It was also noted that the insertion of the collagen phase leads to the structural stability of the β-TCP/collagen scaffold. Compressive modulus and compressive strength are presented in Fig. 12.5D and E. The compressive modulus under the dry condition for the β-TCP/collagen scaffold showed dramatic increases compared to that of the β-TCP scaffold. However, under the wet condition, the value is almost the same. The β-TCP/collagen scaffold under the wet condition also maintained a high strength compared with that of the β-TCP scaffold.

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12.7.2 RESULTS OF IN VITRO CELL EXPERIMENT The variation of compressive modulus and compressive strength in vitro are shown in Fig. 12.6A, a and B, and b, respectively. The compressive modulus of the β-TCP scaffold with and without rMSCs remained almost constant until 28 days; however, that of the β-TCP/collagen scaffold with rMSCs decreased from 0 to 14 days and then increased for up to 28 days even though the compressive modulus of the β-TCP/collagen scaffold without rMSCs remained almost constant until 28 days. In particular, the modulus of the β-TCP/collagen scaffold with rMSCs was higher than that of the β-TCP/collagen scaffold without rMSCs from 14 to 28 days. The compressive strength of the β-TCP scaffold without rMSCs decreased for 7 days and then remained almost constant; however, that of the β-TCP scaffold with rMSCs decreased for 7 days and then gradually increased. For the β-TCP/collagen scaffold without rMSCs, the compressive strength remained almost constant until 28 days; however, that of the β-TCP/collagen scaffold with rMSCs decreased for 7 days and then gradually increased. The variational behavior of absorption energy is shown in Fig. 12.6C and c. In both the β-TCP and β-TCP/collagen scaffolds without rMSCs, absorption energy remained almost constant until 28 days. The β-TCP scaffold with rMSCs decreased at 14 days and then recovered its initial value. The β-TCP/collagen scaffold with rMSCs gradually decreased for 14 days and then dramatically increased until 28 days. The variations of cell viability during cell culture are presented in Fig. 12.7A. The cell number in the β-TCP scaffold increased until 21 days and then remained constant until 28 days. In contrast, the cell number in the β-TCP/collagen scaffold continued to increase for up to 28 days. The variations of ALP activity are shown in Fig. 12.7B. The ALP activity in both scaffolds increased gradually until 14 days and then increased sharply until 28 days. There is no difference in ALP activity between the β-TCP and β-TCP/collagen scaffolds. The RT-PCR measurements related to osteogenic gene expressions, such as type I collagen and osteocalcin are presented in Fig. 12.7C and D. The expression of type I collagen in the β-TCP/collagen scaffold was higher than that in the β-TCP scaffold after all the days examined except for at 7 days. In the β-TCP scaffold, the expression of osteocalcin tended to increase gradually until 28 days; however, that in the β-TCP/collagen scaffold remained almost constant until 21 days and remained almost constant up to 28 days after it had slightly increased up to 14 days. FE-SEM micrographs of the surface region of the β-TCP and β-TCP/collagen scaffolds with rMSCs are presented in Fig. 12.8. White arrows in SEM images indicate the cell and ECM. In the β-TCP scaffold, collagen fibrils and the network structure constructed by rMSCs were observed on the surface at 7 days. Collagen membrane structures were also observed with collagen fibrils at 21 and 28 days. In contrast, in the β-TCP/collagen scaffold, after 7 days, collagen fibrils were created actively, and thin network structures were formed on both the β-TCP frame

Compressive modulus (kPa)

Compressive strength (kPa)

300 250 200 150 100 50 0

400

5

10 15 20 Days in culture

25

350

250 200 150 100 50 0

5

10 15 20 Days in culture

5 4 3 2 1 0

30

300

0

6

5

(b)

β-TCP/collagen (with cells) β-TCP/collagen

25

30

0.06

β-TCP (with cells) β-TCP

7

0

30

Compressive strength (kPa)

Compressive modulus (kPa)

350

0

(a)

β-TCP (with cells) β-TCP

(C)

8

Absorption energy (N•mm)

(B)

400

10 15 20 Days in culture

25

25

15 10 5 0

5

10 15 20 Days in culture

0.03

25

30

#

0.02 0.01 0

(c)

20

0

0.04

30

β-TCP/collagen (with cells) β-TCP/collagen

β-TCP (with cells) β-TCP

0.05

Absorption energy (N•mm)

(A)

0

0.8

5

10 15 20 Days in culture

25

30

β-TCP/collagen (with cells) β-TCP/collagen

0.7 0.6 0.5 0.4 0.3 0.2 0.1 0

0

5

10 15 20 Days in culture

25

30

FIGURE 12.6 Variation of compressive mechanical properties of β-TCP and β-TCP/collagen scaffolds. (A, a) Compressive modulus during cell culture. (B, b) Compressive strength during cell culture. (C, c) Absorption energy during cell culture. The results are presented as mean 6 standard deviation for n 5 3. #,  indicate P ,.05 vs β-TCP scaffold at 14 days culture and P ,.05 vs β-TCP/collagen scaffold at 21 days culture, respectively (Arahira and Todo, 2016).

CHAPTER 12 The design of two different structural scaffolds

(A)

(B)

80

β-TCP β-TCP/collagen

60 50 40 30 20

(C)

β-TCP β-TCP/collagen

1 0.8 0.6 0.4 0.2

10 0

1.4 1.2

p-Nitrophenol (mmol/L)

Cell number (× 104)

70

0

5

10 15 20 Days in culture

25

0

30 (D)

1

0.8

0.6 0.4

10 15 20 Days in culture

25

30

β-TCP/collagen

0.6 0.4 0.2

0.2 0

5

β-TCP

β-TCP/collagen

Gene expression

0.8

0

1

β-TCP

Gene expression

392

7

14 21 Days in culture

28

0

7

14 21 Days in culture

28

FIGURE 12.7 Cell experimental evaluation of β-TCP and β-TCP/collagen scaffolds. (A) Variation of cell number during cell culture. The results are presented as mean 6 standard deviation for n 5 3. $ and $$ indicate P ,.01 vs β-TCP scaffold after 1 and 7 days culture, respectively. # and ## indicate P ,.01 vs β-TCP/collagen scaffold after 1 and 7 days culture and P ,.01 vs β-TCP/collagen scaffold after 1, 7, and 14 days culture. (B) Variation of ALP activity during cell culture. The results are presented as the mean 6 standard deviation for n 5 3. $ and $$ indicate P ,.05 vs β-TCP scaffold after 1, 7, and 14 days culture and P ,.01 vs β-TCP scaffold after 1, 7, and 14 days culture, respectively. # indicates P ,.01 vs β-TCP/collagen scaffold after 1, 7, and 14 days culture. (C, D) Variation of gene expression during cell culture. C and D show type-I collagen and osteocalcin, respectively. The results are presented as the mean 6 standard deviation for n 5 3. $ and  indicate P ,.05 vs β-TCP/collagen scaffold after 7, 14, and 21 days culture and P ,.01 vs β-TCP scaffold after 28 days culture, respectively (Arahira and Todo, 2016).

and collagen phase. Matrix vesicles were also formed on the collagen network structures and turned into calcified matrices, which are much larger than matrix vesicles. The surfaces were covered by proliferated cells and ECMs more so than

FIGURE 12.8 Field emission scanning electron microscope micrographs of β-TCP scaffold and β-TCP/collagen scaffold with rat bone marrow mesenchymal stem cells. White arrows indicate cells and extracellular matrix (Arahira and Todo, 2016).

394

CHAPTER 12 The design of two different structural scaffolds

those of the β-TCP scaffolds. It was also observed that the tissue-like structures comprising of collagen networks and calcified matrices gradually thickened as the culture period increased. In the results of the Alizarin red staining, the staining area of the β-TCP/collagen scaffold is larger than that of the β-TCP scaffold because of the collagen porous structure of the β-TCP scaffold (Arahira and Todo, 2016).

12.8 MECHANISM OF VARIATIONAL MECHANICAL BEHAVIOR BETWEEN SCAFFOLD STRUCTURE AND CELL RESPONSE In this section, a β-TCP/collagen scaffold was fabricated using the template and freeze-drying methods and compared with a β-TCP scaffold to clarify the effectiveness of the collagen porous structure on structural, mechanical, and biological properties. In the template method, fully interconnected porous structure similar to cancerous bone was obtained. It should be noted that this structure is necessary for cell growth and nutrient supply. In addition, a porous collagen structure was introduced in the β-TCP scaffold to enlarge the area of cell proliferation. It was demonstrated that the porous β-TCP/collagen composite scaffold possessed a continuous porous structure with a pore size from 50 to 200 μm, which is known to be suitable for tissue regeneration and cell proliferation and attachment (Ge´rard and Doillon, 2010). A scaffold is usually used to culture cells and create a tissuelike structure under an in vitro condition that is then expected to be implanted into the damaged part of the target tissue. It is, therefore, essential for the scaffold to maintain its original structure (Chen et al., 2000). In the results of XRD analysis, the XRD pattern of the scaffold exhibited several broad peaks between 10 and 60 degrees assigned to the diffraction peaks for β-TCP. It should be noted that sintering the scaffold at 1150 C was shown the peak of β-TCP phase as a single phase although this sintering temperature was very close to α-β phase transition temperature (1180 C) (Nikaido et al., 2013). Fig. 12.9 shows a summary of the variational behavior of the β-TCP-based porous scaffold-rMSC system. Firstly, the compressive mechanical property is thought to change after immersion in the medium. According to Fig. 12.5D and E, in the β-TCP scaffold, the compressive strength decreased with immersion in the medium even though the compressive modulus was almost constant. It was indicated that immersion in the medium affects mechanical strength due to structural change; however, no effect on elastic deformation was observed. The mass change was calculated by: Mass change ð%Þ 5

Wd 2 W 0d 3 100 Wd

(12.7)

12.8 Mechanism of Variational Mechanical Behavior

FIGURE 12.9 Summary of cell experimental results of β-TCP scaffold and β-TCP/collagen scaffold.

where Wd and W0 d are the weights of the scaffold before and after immersion of ethanol, respectively. The mass change of the β-TCP scaffold was approximately 8.8%, whereas that of the β-TCP/collagen scaffold was 6.4%. This indicated that disconnected β-TCP particles in the β-TCP scaffold were more than those in the β-TCP/collagen scaffold. As shown in Fig. 12.5A, disconnected β-TCP particles that were released into the medium were immediately observed after the β-TCP scaffold was immersed in the medium, which indicates that the mechanical strength decreased with local defects triggering fracture. Considering the slight change of strength after 7 days in culture in Fig. 12.6A, it can be presumed that the local release of β-TCP particles finished after approximately 7 days, and then the β-TCP scaffold obtained structural stability. However, the compressive modulus of the β-TCP/collagen scaffold was almost constant even though the compressive strength tended to decrease as seen in Fig. 12.6A. The decrease of compressive strength caused the release of β-TCP particles similar to that in the β-TCP scaffold.

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According to Fig. 12.7A, the cell number of the β-TCP/collagen scaffold increased for up to 28 days even though that of the β-TCP scaffold remained constant from 21 to 28 days. It was noted that the area for cell proliferation in the β-TCP/collagen scaffold is larger than that in the β-TCP scaffold because a collagen porous structure was introduced as shown in Fig. 12.5A. In the RT-PCR results, the expression of osteocalcin and type I collagen of the β-TCP/collagen scaffold is greater than that of the β-TCP scaffold, as observed in Fig. 12.7C and D. In the results of Alizarin red staining, the mineralized deposition by osteoblasts of the β-TCP scaffold was limited to the area of β-TCP frame, on the other hand, that of the β-TCP/collagen scaffold was localized at the β-TCP frame and collagen phase. It was also noted that the range of cell activity was spread, and cell proliferation and differentiation into osteoblasts were improved by introducing collagen, which is superior for cell affinity. In this chapter, the focus was on the effects of the proliferation, differentiation, and ECM formation of rMSCs on the compressive mechanical property of the β-TCP/collagen scaffold compared with the β-TCP scaffold. For the β-TCP scaffold, the compressive modulus tended to decrease for up to 14 days and then remained constant until 28 days. This result indicated that the proliferation and production of collagen by cells caused a decrease of the compressive modulus average due to the formation of membrane structure with low modulus on the surface of the β-TCP frame and pore structure. Fig. 12.8 shows the membrane structure produced by rMSCs on the surface of the β-TCP scaffold at 7 days. However, compressive strength decreased by approximately 2 kPa for up to 7 days and then increased by approximately 1 kPa. The differentiation of rMSCs into osteoblasts and ECM formation improved the mechanical property, which corresponded to the increase of gene expression of osteocalcin and type I collagen (Fig. 12.7C and D), ALP activity (Fig. 12.7B), and calcification (Fig. 12.9) after 14 days. In contrast, for the β-TCP/collagen scaffold, the modulus decreased for up to 14 days compared with the scaffold without rMSCs, which indicated that structural brittleness progressed due to cell proliferation and the generation of collagen. However, the modulus recovered between 14 and 21 days, which suggested that structural stiffness was improved by the promotion of calcification by osteoblasts. The compressive strength of the β-TCP/ collagen scaffold with rMSCs tended to decrease for up to 7 days because of the low strength of the β-TCP frame similar to that of the β-TCP/collagen scaffold without rMSCs. However, compressive strength from 7 to 28 days was higher than that at 0 days resulting from the formation of the ECM on the β-TCP frame and porous collagen phase. This study clearly demonstrates that the porous collagen phase is effective in the proliferation of rMSCs due to the increase of surface area. Moreover, β-TCP improved the differentiation of rMSC into osteoblasts and subsequent ECM formation, resulting in the strengthening of structure and the improvement of mechanical properties.

12.10 Present Study

12.9 SUMMARY The effect of rMSCs culture on the variational behavior of the compressive modulus of the composite scaffold composed of β-TCP and collagen was examined in this study. The conclusions were summarized as: 1. The distribution of β-TCP powders in the collagen matrix effectively improved the structural stability of the composite scaffold, while the collagen scaffold exhibited structural instability with the collapse of porous structure when immersed in the culture medium. 2. The collagen/β-TCP scaffold showed higher ALP activity and more active generation of osteoblastic markers than the collagen scaffold, indicating that β-TCP can activate the differentiation of rMSCs into osteoblasts and ECM formation, such as type I collagen and the following mineralization. 3. The variational behavior of the compressive modulus of the collagen/β-TCP scaffold in vitro was affected by the degradation of the scaffold, the proliferation of cells, and the ECM formation. In the first stage, the modulus of the collagen/β-TCP scaffold tended to increase due to cell proliferation and the following formation of network structure. In the second stage, the modulus tended to decrease because the material degradation such as ductile deformation of collagen and decomposition of β-TCP were more effective on the mechanical property than the ECM formation. In the third stage, active calcification by formation and growth of mineralized nodules resulted in the recovery of modulus. 4. The collagen porous structure in the β-TCP scaffold effectively improved the structural stability of the composite scaffold, whereas the β-TCP scaffold exhibited structural instability with the collapse of the porous structure when immersed in the culture medium. 5. The β-TCP/collagen composite scaffold exhibited more active generation of osteoblastic markers than the β-TCP scaffold. 6. The variational behaviors of the compressive modulus, strength, and absorption energy of the β-TCP/collagen composite scaffold were affected by cell proliferation and ECM formation.

12.10 PRESENT STUDY In the present study, bone
cartilage tissue construction was fabricated using the freeze-drying and template methods. In the case of bone layer, the porous β-TCP structure was first constructed using the template method, and then the porous collagen structure was introduced into the continuous pores of the β-TCP structure. In the cartilage layer, type II collagen sponge was fabricated using the freeze-drying method. First, rMSCs were cultured in each scaffold and differentiated into osteoblasts or chondrocytes for 14 days. After that, these regenerated

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CHAPTER 12 The design of two different structural scaffolds

tissues were combined and cultured for 28 days. Compressive mechanical properties and bone or cartilage-related markers were evaluated every 7 days for 28 days. It was found from the experimental results that the compressive modulus of both scaffolds tended to increase with rMSC culture until 14 days and then decrease to 28 days. It was concluded that two layered scaffolds can provide good structural stability as regenerated bone
cartilage layered tissue.

12.11 FUTURE WORK These scaffold-MSC systems will be applied to other regenerated tissue constructs, such as vascular tissue, ligament tissue, and periodontal tissue, to assess the relationship between mechanical property and cell growth behavior. In tissue engineering, the variational behavior of mechanical property is an extremely important factor to evaluate the mechanical stability of regenerated tissue at the stage of regeneration process. The variational mechanical property of the layered tissue construction will also be evaluated in the future.

ACKNOWLEDGMENT This work was supported by JSPS Fellows (231708).

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1355. Ge´rard, C., Doillon, C.J., 2010. Facilitating tissue infiltration and angiogenesis in a tubular collagen scaffold. J. Biomed. Mater. Res. A 93, 615
624. Girotto, D., Urbani, S., Brun, P., Renier, D., Barbucci, R., Abatangelo, G., 2003. Tissuespecific gene expression in chondrocytes grown on three-dimensional hyaluronic acid scaffolds. Biomaterials 24, 3265
3275. He, X., Lu, H., Kawazoe, N., Tateishi, T., Chen, G., 2010. A novel cylinder-type poly(Llactic acid)-collagen hybrid sponge for cartilage tissue engineering. Tissue Eng. C 16, 329
338. Heinemann, C., Heinemann, S., Bernhardt, A., Worch, H., Hanke, T., 2008. Novel textile chitosan scaffolds promote spreading, proliferation, and differentiation of osteoblasts. Biomacromolecules 9, 2913
2920. Hiraoka, Y., Kimura, Y., Ueda, H., Tabata, Y., 2003. Fabrication and biocompatibility of collagen sponge reinforced with poly(glycolic acid) fiber. Tissue Eng. 9, 1101
1112. Intan, P.A., Arahira, T., Todo, M., 2016. Characterization of tensile mechanical behavior of MSCs/PLCL hybrid layered sheet. J Funct. Biomater. 7, 14. Kang, Y., Kim, S., Bishop, J., Khademhosseini, A., Yang, Y., 2012. The osteogenic differentiation of human bone marrow MSCs on HUVEC-derived ECM and β-TCP scaffold. Biomaterials 33, 6998
7007. Keogh, M.B., O’Brien, F.J., Daly, J.S., 2010. Substrate stiffness and contractile behaviour modulate the functional maturation of osteoblasts on a collagen-GAG scaffold. Acta Biomater. 6, 4305
4313. Liang, W.-H., Kienitz, B.L., Penick, K.J., Welter, J.F., Zawodzinski, T.A., Baskaran, H., 2010. Concentrated collagen-chondroitin sulfate scaffolds for tissue engineering applications. J. Biomed. Mater. Res. A 94, 1050
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Lohan, A., Marzahn, U., El Sayed, K., Haisch, A., Kohl, B., Mu¨ller, R.D., et al., 2011. In vitro and in vivo neo-cartilage formation by heterotopic chondrocytes seeded on PGA scaffolds. Histochem. Cell. Biol. 136, 57
69. Lohfeld, S., Cahill, S., Barron, V., McHugh, P., Du¨rselen, L., Kreja, L., et al., 2012. Fabrication, mechanical and in vivo performance of polycaprolactone/tricalcium phosphate composite scaffolds. Acta Biomater. 8, 3446
3456. Lu, H., Ko, Y.-G., Kawazoe, N., Chen, G., 2010. Cartilage tissue engineering using funnellike collagen sponges prepared with embossing ice particulate templates. Biomaterials 31, 5825
5835. Lu, H., Hoshiba, T., Kawazoe, N., Koda, I., Song, M., Chen, G., 2011. Cultured cellderived extracellular matrix scaffolds for tissue engineering. Biomaterials 32, 9658
9666. Matsuno, T., Nakamura, T., Kuremoto, K.-I., Notazawa, S., Nakahara, T., Hashimoto, Y., et al., 2006. Development of β-tricalcium phosphate/collagen sponge composite for bone regeneration. Dent. Mater. J. 25, 138
144. Murphy, C.M., Haugh, M.G., O’Brien, F.J., 2010. The effect of mean pore size on cell attachment, proliferation and migration in collagen-glycosaminoglycan scaffolds for bone tissue engineering. Biomaterials 31, 461
466. Nikaido, T., Tsuru, K., Kawach, G., Munar, M., Matsuya, S., Nakamura, S., et al., 2013. Fabrication of βTCP foam using αTCP foam as a precursor by heat treatment. Key Eng. Mater. 529
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18. Oh, S.H., Park, I.K., Kim, J.M., Lee, J.H., 2007. In vitro and in vivo characteristics of PCL scaffolds with pore size gradient fabricated by a centrifugation method. Biomaterials 28, 1664
1671. Ozawa, H., Hoshi, K., Amizuka, N., 2008. Current concepts of bone biomineralization. J. Oral Biosci. 50, 1
14. Pabbruwe, M.B., Kafienah, W., Tarlton, J.F., Mistry, S., Fox, D.J., Hollander, A.P., 2010. Repair of meniscal cartilage white zone tears using a stem cell/collagen-scaffold implant. Biomaterials 31, 2583
2591. Park, J.-E., Todo, M., 2011. Development and characterization of reinforced poly(l-lactide) scaffolds for bone tissue engineering. J. Mater. Sci. Mater. Med. 22, 1171
1182. Poffyn, B., Sys, G., Mulliez, A., Van Maele, G., Van Hoorebeke, L., Forsyth, R., et al., 2011. Extracorporeally irradiated autografts for the treatment of bone tumours: tips and tricks. Int. Orthop. 35, 889
895. Porter, J.R., Henson, A., Popat, K.C., 2009. Biodegradable poly(ε-caprolactone) nanowires for bone tissue engineering applications. Biomaterials 30, 780
788. Roohani-Esfahani, S.I., Lu, Z.F., Li, J.J., Ellis-Behnke, R., Kaplan, D.L., Zreiqat, H., 2012. Effect of self-assembled nanofibrous silk/polycaprolactone layer on the osteoconductivity and mechanical properties of biphasic calcium phosphate scaffolds. Acta Biomater. 8, 302
312. Salerno, A., Zeppetelli, S., Di Maio, E., Iannace, S., Netti, P.A., 2011. Processing/structure/property relationship of multi-scaled PCL and PCL-HA composite scaffolds prepared via gas foaming and NaCl reverse templating. Biotech. Bioeng. 108, 963
976. Sang, L., Luo, D., Xu, S., Wang, X., Li, X., 2011. Fabrication and evaluation of biomimetic scaffolds by using collagen
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9. Todo, M., Arahira, T., 2013. In vitro bone formation by mesenchymal stem cells with 3D collagen/β-TCP composite scaffold. In: Proc. 35th IEEE Eng. Med. Biol. Soc. (EMBC), pp. 409
412. Udoh, K.-I., Munar, M.L., Maruta, M., Matsuya, S., Ishikawa, K., 2010. Effects of sintering temperature on physical and compositional properties of α-tricalcium phosphate foam. Dent. Mater. J. 29, 154
159. Wang, Y., Uemura, T., Dong, J., Kojima, H., Tanaka, J., Tateishi, T., 2003. Application of perfusion culture system improves in vitro and in vivo osteogenesis of bone marrowderived osteoblastic cells in porous ceramic materials. Tissue Eng. 9, 1205
1214. Wang, L., Fan, H., Zhang, Z.-Y., Lou, A.-J., Pei, G.-X., Jiang, S., et al., 2010. Osteogenesis and angiogenesis of tissue-engineered bone constructed by prevascularized β-tricalcium phosphate scaffold and mesenchymal stem cells. Biomaterials 31, 9452
9461. Wiria, F.E., Chua, C.K., Leong, K.F., Quah, Z.Y., Chandrasekaran, M., Lee, M.W., 2008. Improved biocomposite development of poly(vinyl alcohol) and hydroxyapatite for tissue engineering scaffold fabrication using selective laser sintering. J. Mater. Sci. Mater. Med. 19, 989
996. Yamasaki, N., Hirao, M., Nanno, K., Sugiyasu, K., Tamai, N., Hashimoto, N., et al., 2009. A comparative assessment of synthetic ceramic bone substitutes with different composition and microstructure in rabbit femoral condyle model. J. Biomed. Mater. Res. B. Appl. Biomater. 91, 788
798. Zhang, Y., Zhang, M., 2001. Synthesis and characterization of macroporous chitosan/calcium phosphate composite scaffolds for tissue engineering. J. Biomed. Mater. Res. 55, 304
312. Zhang, X., Guo, Y., Li, D.X., Wang, R., Fan, H.S., Xiao, Y.M., et al., 2011. The effect of loading icariin on biocompatibility and bioactivity of porous β-TCP ceramic. J. Mater. Sci. Mater. Med. 22, 371
379. Zhao, J., Duan, K., Zhang, J.W., Guo, L.Y., Weng, J., 2011. Preparation of highly interconnected porous hydroxyapatite scaffolds by chitin gel-casting. Mater. Sci. Eng. C 31, 697
701.

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13

Composite materials based on hydroxyapatite embedded in biopolymer matrices: ways of synthesis and application

A. Yanovska and S. Bolshanina Sumy State University, Ministry of Education and Science of Ukraine, Sumy, Ukraine

13.1 TYPES OF BIOPOLYMER MATRICES (COLLAGEN, GELATIN, CHITOSAN, ALGINATE, AND THEIR COMBINATIONS) In the world, millions of people suffer from bone injuries, disorders, and musculoskeletal problems. They are usually in need of partial or total replacement of the bone tissue. In Europe, 20% 30% of adults are affected by musculoskeletal pain ˚ kesson, 2007; Ferreira et al., 2012). In the United States, around 6.3 (Woolf and A million fractures are registered each year (Ferreira et al., 2012), so bone is the most commonly replaced organ in the body. Bone is composed of hydroxyapatite (HA) (69% 80%), collagen (Col) (17 20 wt.%), and other substances (water, proteins, etc.) (Karageorgiou and Kaplan, 2005). Composite materials based on biopolymers and calcium phosphates are widely used for bone replacement (Haroun and Migonney, 2010) Natural polymers are a good substitute for synthetic polymers. Their composition mimics extracellular matrix (ECM), they are bioabsorbable, biocompatible, biodegradable, and able to adsorb bioactive molecules. Biopolymer composites for medical applications should have similarities to the complex architecture of the human body and polymer composites. Furthermore, biopolymer-based drug and bioactive agent release systems should be developed as multifunctional release systems for maintaining the functionality of biological molecules (Park et al., 2016). Proteins and polysaccharides are widely applied in regenerative medicine and tissue engineering (Mano et al., 2007). Polysaccharides perform different physiological functions and may have various applications in regenerative medicine and tissue engineering (Table 13.1) (Park et al., 2016). Various applications of polysaccharides are presented in Table 13.1. Materials for Biomedical Engineering: Hydrogels and Polymer-based Scaffolds. DOI: https://doi.org/10.1016/B978-0-12-816901-8.00013-4 © 2019 Elsevier Inc. All rights reserved.

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Table 13.1 Main Characteristics of Polysaccharides for Biomedical Application Biopolymer Chitosan (CS)

Alginate (Alg)

Hyaluronic Acid (HyA)

Monomer unit Chemical formula

2-amido-2-deoxyβ-D-glucose residues linked through β-1,4 linkages

1,4-glycosidically linked β-Dmannuronic acid (Mblocks) and α-Lguluronic acid (Gblocks) monomers

Composed of Dglucuronic acid and D-Nacetyl-glucosamine linked together by alternating β-1,4 and β-1,3 glycosidic bonds

Main properties

Cationic biopolymer with bacteriostatic properties, soluble in aqueous acidic media, biodegradable, biocompatible

Affects cell proliferation, differentiation, tissue repair, soluble in water, hydrogel forming ability.

Application

Drug delivery, gene therapy, tissue engineering, and wound healing. blood coagulation, cytokine induction Scaffolds, wound dressings, beads, films, coatings, membranes, sponges, microgels, nanogels, nanofibers, microparticles, nanoparticles

Anionic, hydrophilic polymer, biocompatible, nontoxic, low cost, able to form hydrogels in the presence of multivalent cations Cell transplantation in hydrogels, drug delivery

Photocrosslinked hydrogels, hyaluronic acid scaffolds, hyaluronan based sponges

82 166

Fibrils, microspheres, granules, scaffolds, films, sponges, hydrogels, cell encapsulation, injectables cell delivery, drug and antibacterial components delivery 31 37

5 8

11 17

37 43

Forms of biomaterial

Tensile strength, MPa Elongation at break, %

Functional component of tissues, applied for cartilage support by injection

102 120

13.1 Types of Biopolymer Matrices

Table 13.2 Protein Based Biopolymers for Biomedical Applications: Main Characteristics Biopolymer Collagen (Col)

Gelatin (Gel)

Monomer unit Chemical formula

α-chains are formed by repetitions of the tripeptide: (Gly-X-Y-), and are linked to each other, building the characteristic triple helix of type I, II, and III collagen.

Main properties

Fibrillar collagens are insoluble in their native structure but soluble in aqueous solution in denaturated form (procollagens). Combinations of collagen with other materials increases their mechanical performance, abundance, biocompatibility, high porosity, facility for combination with other materials, easy processing, hydrophilicity, low antigenicity, absorbability in the body, etc.

Gelatin is a polyampholyte with cationic (13% lysine and arginine), anionic (B12% glutamic and aspartic acid), and hydrophobic groups (B11% of the chain comprising of leucine, isoleucine, methionine, and valine) present in a 1:1:1 ratio. The other part of the chain contains glycine, proline, and hydroxyproline. Obtained by the partial hydrolysis of collagen, biodegradable, biocompatible, soluble in warm water (t . 35 C). Thermoreversible gelling agent in pharmaceutical and medical applications. Gelatin is obtained through a controlled hydrolysis of collagen, which is a major component of skin, bones, and connective tissue. It exhibits excellent qualities, such as biocompatibility, biodegradability, and low antigenicity. It is used for incorporating nanoparticles or microspheres, materials for drug delivery, scaffolds, and films.

Application

Forms of biomaterial

Tensile strength, MPa Elongation at break, %

Networks with highly organized 3D architectures suitable for use in scaffolds.Injectable hydrogels, membranes, and films, sponges, scaffolds, and micro/nanospheres. 202 224

0.66

37 39

62.5

Proteins are large biological molecules composed of one or more long chains of amino acid residues. They are extensively used in drug delivery and tissue engineering scaffolds due to their elasticity, cell protection abilities, and scaffold formation (Ninan et al., 2015; Gupta and Nayak, 2015) The most widely used proteins are collagen and gelatin (Table 13.2).

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Collagen is the most abundant fibrous protein in mammals, constituting, by weight, more than one-third of body protein tissue (Patino et al., 2002). There are around 28 types of collagen (Gelse et al., 2003) but the most prevalent type found in the ECM of tendon and bone tissues is type I collagen (Gelse et al., 2003). Hydroxyapatite is deposited into the holes of type I collagen in the process of bone biomineralization (Patino et al., 2002). Therefore, due to compositional and some structural analogy to natural bones, the composites of collagen and HA are of special interest among all kinds of bone substitutes (Ferreira et al., 2012; Glowacki and Mizuno, 2008.). The primary structure of collagen is an amino acid triplet; (-Gly-X-Y-)n where glycine (Gly) is every third residue with strict repeating along the amino acid chain. A high proportion of proline (X) and hydroxyproline (Y) residues (B20%) in tripeptide sequences exist. Hydroxyproline is not commonly found in other proteins, while in collagen it constitutes more than 50% of the total amino acid content (Olszta et al., 2007). The secondary structure is an α-helix, where (-Gly-X-Y-)n tripeptides are linked to each other, with the formation of a triple helix of type I, II, and III collagen (Fratzl, 2003). The nonhelical domains are at the end of the a-chains, where the C-terminus initiates triple-helix formation and the N-terminus is involved in the regulation of primary fibril diameters. The short nonhelical telopeptides of collagen are linked by covalent bonds with other collagen molecules and other molecules present in the ECM (Ferreira et al., 2012). The tertiary structure of collagen molecules is a triple helix that is usually formed as a heterotrimer of two identical a 1 (I)- and a 1 (II)-chains and one a 2 (I)-chain with about 1000 amino acids, and is approximately 300 nm in length (L) and 1.5 nm in diameter, especially in collagen type I (Kadler et al., 1996). In the quaternary structure of collagen fibrils molecules are able to self-assemble into a supramolecular form via a quarter-stagger package pattern of five triple-helical collagen molecules highly oriented with D-periodic banding spaces (Kadler et al., 1996; Bozec et al., 2007). Fibrillar collagen is insoluble in its native structure but in the denaturated form of procollagens it can be solubilized in aqueous solutions (Kadler, 2004). Due to the denaturation of collagen by thermal treatment (37 C for bovine collagen) it is converted into a randomly coiled gelatin (Ratner, 2004). Collagen has found a wide variety of applications in the field of medicine, including: sutures, hemostatic agents, tissue replacement and regeneration (bone, cartilage, skin, blood vessels, trachea, esophagus, etc.), cosmetic surgery (lips, skin), dental composites, skin regeneration templates, membrane oxygenators, contraceptives (barrier method), biodegradable matrices, protective wrapping of nerves, implants, corneal bandage, contact lens, drug delivery, etc. (Ratner, 2004; Meena et al., 1999; Pannone, 2007). It can be easily modified into different physical forms, such as powders, particles, fibers, gels, solutions, films, membranes, sponges, blends with other polymers, and ceramic based composites with ceramics (Ferreira et al., 2012; Glowacki and Mizuno, 2008; Kadler, 2004; Meena et al., 1999). Natural bones are a complex assembly of parallel type I collagen nanofibrils and HA crystals precipitated on their surface (Currey, 2012). Thus, composite

13.1 Types of Biopolymer Matrices

materials for bone replacement should be included in the process of bone formation promoted by osteoblasts (mineralization) to create an environment for the crystallization of calcium phosphates (Ferreira et al., 2012). Two types of cells are involved in the bone formation process: osteoblasts (bone-forming) and osteoclasts (bone-resorbing). During the process of ossification, osteoblasts secrete type I collagen with noncollagenous proteins, such as osteocalcin, bone sialoprotein, and osteopontin. Osteoblast-secreted ECM may initially be amorphous and noncrystalline, but gradually transforms into more crystalline forms (Boskey, 2003). One of the main challenges to bone tissue engineering is to develop scaffolds with optimal mechanical properties, biodegradability, and appropriate architecture for cell colonization and organization, which can ensure the integration of a scaffold with host tissue (Ferreira et al., 2012). Major advances in bone tissue engineering with scaffolds are achieved through growth factors and cells. The biomechanical system of bone is complex so that these requirements of an ideal scaffold are diverse: A synthetic bone scaffold should provide porous architecture and mechanical support in order to promote bone cell migration and differentiation into the scaffold, thereby, encouraging osteoinduction and enhancing osteointegration with the host tissue. In addition it should be sterilizable without the loss of bioactivity, the release of bioactive molecules, or the use of drugs, and should degrade in a controlled manner without inflammation and the production of toxic products of degradation (Porter et al., 2009; Mistry and Mikos, 2005). The general requirements of a bone tissue engineered scaffold are (Venkatesan et al., 2015; Dorati et al., 2011): • •

• • • • • • •

Three-dimensional structure, biocompatibility, and sufficient surface area. Porosity (pore size .100 μm) with interconnection between pores to allow cell adhesion, migration, and proliferation of bone cells in the desired direction. Biodegradability and nontoxicity. Mechanical strength comparable with cortical bone. Holders for growth factors, drugs, antibacterial components, and bioactive molecules. Equal balance of osteoblastic and osteoclastic differentiation. Cell attachment and cell-matrix interaction. Structural anisotropy and influence on the orientation of cells and ECM deposition. Low cost and ability to sterilize without loss of bioactivity.

Ideal wound dressings should be nontoxic, nonallergenic, biocompatible, sterile, possess good adhesion strength to the skin and low adherence to wound surfaces, provide mechanical protection due to high mechanical strength, provide easy and accessible wound coverage area, provide a barrier against infection due to optimal shape and surface bacteria or fungi, control of fluid loss due to the capability of

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absorbing excess exudate, and promote wound healing. They should be costeffective, acceptable to the patient, and allow for reducing pain during treatment (Dorati et al., 2011; Abdelrahman and Newton, 2011; Caramella et al., 2016). Natural materials, such as collagen, chitosan, fibrinogen, alginates, and hyaluronic acid, are prospective biomimetic scaffolding materials due to their bioactivity and capability to interact with cells, as well as their biocompatibility and biodegradability (Abdelrahman and Newton, 2011). However, natural polymers have several disadvantages, such as poor mechanical properties, low reproducibility depending on the natural source, water solubility or high hydrophilicity (swelling), poor processability, possible denaturation during processing, immunogenicity and, in some cases, potential risk of transmitting animal-originated pathogens (Dorati et al., 2011; Abdelrahman and Newton, 2011; Caramella et al., 2016). Polymers and bioceramics have been combined to fabricate biomimetic scaffolds for bone tissue engineering, as native bone is a combination of naturally occurring polymers and biological apatite. Moreover, polymers and ceramics that have the ability to degrade in vivo are ideal candidates for composite scaffolds because they can gradually degrade while new tissue is formed. Certain inorganic/ceramic materials, such as hydroxyapatite or calcium phosphate, which have good osteoconductivity and have been studied for mineralized tissue engineering, show drawbacks, such as poor processability into highly porous structures and brittleness. In contrast, polymers offer great design flexibility because of their composition and structure and, therefore, they have been extensively studied for various tissue engineering applications, including bone tissue engineering (Ferreira et al., 2012; Abdelrahman and Newton, 2011). Porous collagen scaffolds with ceramic particles have been produced for bone tissue engineering purposes (Ferreira et al., 2012; Zhang et al., 2003). Threedimensional collagen scaffold materials have been designed to mimic one or more bone-forming components. Porous scaffolds are rather widely applied for the regeneration of bone tissue since the migration of osteogenic cells contributing to bone tissue growth is possible in them. Pore size is an important issue for composite materials formation: if the scaffold pores are too small, pore occlusion by the cells may occur, preventing cellular penetration, neovascularization of the inner areas of the scaffold, and extracellular matrix production. For bone tissue engineering purposes, the pore size should be in the 200 900 μm range (Ferreira et al., 2012). Porosity is an essential parameter that distinguishes the properties of obtained biomaterials. Solid conjunctive tissues are porous, for example, the porosity of cancellous bone is between 50% and 90% (Kailasanathan et al., 2012). Another group of collagen based materials are hydrogels, which have been used in a wide variety of tissue engineering applications (Ferreira et al., 2012). Hydrogels represent an important class of biomaterials in biotechnology and medicine because most of them exhibit excellent biocompatibility with minimal inflammatory responses and tissue damage, and thus, many studies on bone tissue engineering applications have been undertaken (Ferreira et al., 2012; Saito et al., 2004).

13.1 Types of Biopolymer Matrices

Hydrogels are hydrophilic polymer networks that may absorb from 10% or 20%, up to thousands of times their dry weight in water, this allows for cells to adhere, proliferate, and differentiate onto hydrogels (Drury and Mooney, 2003). Minimally invasive treatments have been developed using injectable systems for bone tissue engineering. Several injectable gels have been used to carry cells in order to engineer bone. A collagen hydrogel has excellent swelling ability in water, suitable physical properties (e.g., mechanical properties, gelling ability), susceptibility to enzymatic degradation, and suitable biological properties, therefore, it is an excellent candidate for cell encapsulation (Ferreira et al., 2012; Nicodemus and Bryant, 2008.). Chitosan collagen composite hydrogel materials have potential applications in regenerative medicine, particularly in applications where injectable cell carriers are required. Hydrogels with various weight ratios of chitosan (CS) and collagen were fabricated by initiating gelation using β-glycerophosphate and a weak base under various temperature conditions (Huang et al., 2011). The presence of collagen in chitosan collagen materials was associated with increased cell proliferation, increased gel compaction, and a resulting stiffer matrix. Such materials can be used as in situ gel-forming materials for tissue regeneration or for cell encapsulation (Ferreira et al., 2012; Huang et al., 2011). Injectable hydrogels with biomimetic properties, containing chitosan, collagen, and nanohydroxyapatite were successfully obtained for bone regeneration. The CS-HA-Col solution rapidly formed a stable gel at body temperature. The scaffold obtained was a noninvasive scaffold with microstructure, composition, and surface properties similar to those of physiological bone (Huang et al., 2011). Injectable HA-Col-Alg hydrogels were used as carriers of bone morphogenic protein-2 after gelation in 30 minutes by ionic crosslinking of Ca21 with alginate in CaCl2 solution (Sotome et al., 2004). After 5 weeks of implantation they showed bone formation throughout the implant without obvious deformation of the material, whereas bone formation was observed only in a section of collagen sponge (Ferreira et al., 2012; Sotome et al., 2004). Resorbable collagen membranes are used in bone regeneration procedures due to their biocompatibility and bioresorbability. Collagen barrier membranes for periodontal defect regeneration have been widely used in oral surgery to avoid the need for a second surgery (Gentile et al., 2011). Despite favorable regenerative results there are drawbacks of using collagen membranes for bone regeneration, these include: (1) the implantation of animal-derived collagen includes a potential risk of disease transmission from animals to humans; (2) the loss of space-maintaining ability in aqueous solutions; (3) poor mechanical strength; and (4) the degradation rate of collagen membranes did not match that of normal tissue-healing processes (Ferreira et al., 2012; Do¨ri et al., 2007). Composite membranes based on apatite crystals and collagen preserved the biological functions of damaged tissues in a more efficient and biomimetic way due to their bioactivity, osteoconduction, osteoinduction, and biocompatibility (Ferreira et al., 2012). Collagen microspheres are commonly formed by phase separation, emulsifying methods, crosslinking native collagen molecules, and are applied as carriers for

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drug/protein delivery. Particle size is controlled by the molecular weight of the collagen employed, so, collagen denaturation to gelatin is used for the production of smaller spheres with a diameter of about 0.1 μm with higher stability, and ability for sterilization (Ferreira et al., 2012; Sehgal and Srinivasan, 2009). In bone tissue engineering, collagen microparticles have been generally introduced into scaffolds based on synthetic polymers and/or ceramics (e.g., HA) to enhance osteoblast cell growth within bone-filling materials (Ferreira et al., 2012; Wu et al., 2004). Gelatin is prepared through chemical and thermal denaturation of collagen and lacks the fibrillar structure required for the orderly deposition of Hydroxyapatite (HA) crystals. However, increasing the nucleation and growth of inorganic phases in a gelatin gel system is possible in the presence of lipids, proteoglycans, sodium polyacrylate, and polyaspartate (Teng et al., 2006). In order to enhance the interaction between the inorganic apatite and the organic polymer gelatin, some researchers have introduced hydrophilic polar groups, such as phosphate, carboxyl, and hydroxyl groups, onto hydrophobic substrates (Haroun and Migonney, 2010). Being derived from collagen, gelatin has several structural similarities with it and its physical and chemical characteristics, such as molecular weight distribution and amino acid sequence depending on the collagen source. Gelatin becomes soluble in warm water (T .35 C) due to the helix-to-coil conformation transition as a result of reduced molecular interactions (Zaupa et al., 2011). The high solubility of gelatin requires stabilization through chemical crosslinking with alginate dialdehyde; a nontoxic crosslinking agent which is obtained through periodate oxidation of alginate (Boanini et al., 2010). Alginate dialdehyde crosslinking occurs through the reaction of aldehyde groups with free amino groups of lysine or hydroxylysine amino acid residues of gelatin polypeptide chains, and this was previously successfully applied to gelatin hydrogels and films (Boanini et al., 2010; Boanini and Bigi, 2011). Gelatin has high biocompatibility and biodegradability due to its structural similarity to collagen. It is also naturally occurring, permeable to nutrients and other substances, and easy for researchers to manipulate or with which to work. In the past few years it has been widely used for tissue engineering (Dong et al., 2017). Gelatin is obtained through a controlled hydrolysis of collagen, which is a major component of skin, bones, and connective tissue. It exhibits excellent qualities, such as biocompatibility, biodegradability, low antigenicity (Gashti et al., 2016), and is more economically effective than collagen. Today, much interesting research is focused on the composites of calcium phosphate in conjugation with gelatin, in addition to the fact that gelatin chains can strongly interact with each other in water through hydrogen bonds (Babaei et al., 2013). Chitosan-based composites are nontoxic, biocompatible, biodegradable, chemically inert and low cost. They also have good film-forming property. Second organic polymer—Chitosan acting as reinforced phase in HA-based composite (Babaei et al., 2013). Gelatin is successfully used for incorporating nanoparticles or microspheres. Its easily modifiability also makes it a good material for drug delivery

13.1 Types of Biopolymer Matrices

(Wang et al., 2016; Alemdar, 2016). A wide variety of gelatin-based composites have been developed for tissue engineering, such as gelatin-siloxane hybrids, β-TCP/chitosan/gelatin, chitosan/gelatin/alginate/hydroxyapatite (Sharma et al., 2016), gelatin/chitosan/nanobioglass 3D porous scaffolds (Maji et al., 2016), and hybrid macroporous gelatin/bioactive-glass/nanosilver scaffolds (Dong et al., 2017; Yazdimamaghani et al., 2014). Gelatin is a polyampholyte (Elzoghby et al., 2012) with cationic (13% lysine and arginine), anionic (B12% glutamic and aspartic acid), and hydrophobic groups (B11% of the chain comprising of leucine, isoleucine, methionine, and valine) present in a 1:1:1 ratio. The other parts of the chain contain glycine, proline, and hydroxyproline (Yanovska et al., 2016; Elzoghby, 2013). Due to its relatively high content of functional groups, gelatin is a promising candidate for template based synthesis (Tlatlik et al., 2006). Thus, it is not surprising that this biopolymer is widely employed for a variety of applications, including for biomimetic mineralization aimed at developing scaffolds for tissue engineering. Im et al. (2012) set out to create a hydrogel-based scaffold that not only included the use of nanomaterials but also HA and chitosan. The use of HA in coordination with chitosan increased the osteoconductive properties and overall strength of the obtained scaffold. Scaffolds based on calcium phosphate loaded with gentamycin sulfate were combined with liposomes as an additional drug carrier. It was discovered that the calcium phosphate scaffold, after attaching to the infection site, releases the liposomes in vacuole form, which then move on to releasing the gentamycin. This release mechanism had greater antibiofilm activity than the control group (pure gentamycin), but that is to be expected considering the direct interaction (Zhu, 2010). Composite materials based on HA and Gel with silver and zirconium oxide additions were obtained (Yanovska et al., 2016). The study investigates the use of low concentrations of gelatin under low powered ultrasonic irradiation in the coprecipitation synthesis. Polysaccharide nanocomposites have become an important material in tissue engineering over the past decade. They offer an alternative to synthetic polymers in the preparation of soft nanomaterials for bone tissue replacement. They have also been used in composites with metal nanoparticles and carbon-based nanomaterials. Polysaccharides are comprised of monosaccharide units joined with glycosidic linkages and due to their unique features differ from the other families of biopolymers. Polysaccharides are widely used in the field of building up of biomaterials due to the possibility of changing their properties in relation to the final application. This versatility is not only related to chemical or process modification of a single polymer, but also to the large range of polymers available. Polymers are widely used for drug delivery applications, such as targeted drug delivery systems (Dvir et al., 2011; Tripodo et al., 2013), for the delivery of protein drugs (LoPresti et al., 2011), or for the intracellular delivery of drugs, as well as in the biomedical field in orthopedic regenerative medicine for example (Dorati et al., 2014).

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Most polysaccharides are recovered natural products that are produced as structural biopolymers or for energy storage by plants, microbes, and animals. They could be monofunctional (having only one functional group like an OH group) or polyfunctional (having OH, COOH, and/or NH2 groups). Carbohydrates also have a high level of chirality with linear or branched structures, they could be water soluble or insoluble, environmentally safe, and nonimmunogenic (Dumitriu, (Ed.), 2004). Such diversity of properties and structures of carbohydrates allow for nanomaterials and nanocomposites to be obtained with biological and molecular advantages over nucleic acids and proteins for applications in materials science. They are more stable than proteins or nucleic acids and are usually not irreversibly denatured on heating (Zheng et al., 2015). Alginate is a natural polysaccharide extracted from algae that also contains a large variety of nutriments, such as vitamins, salts, iodine, and sterols (Rinaudo, 2014). It is isolated from brown seaweed through extraction in dilute alkaline condition, which makes alginic acid soluble in water. Usually, the extracted alginic acid is converted into sodium salt which is the form most used and marketed. Alginate is nontoxic and biocompatible, even if, due to extraction from natural sources, a variety of impurities, such as heavy metals, proteins, and endotoxins, could be present representing a major problem for pharmaceutical applications. It is a linear polymer consisting of (1-4) α-L-guluronic (G) and (1-4) β-Dmannuronic acid blocks (M), and heteropolymeric sequences of M and G (MG blocks). The source from which alginate is extracted will determine different sequences in its constitutive blocks. The chemical structure of alginate is composed of blocks of D-mannuronic acid and L-guluronic acid but their disposition along the polymer chain is not predictable, in fact, portions of homogeneous blocks are composed of either acid residue alone are separated by blocks made of random or alternating units of mannuronic and guluronic acids (Rinaudo, 2014). Chemical modification of alginates is performed in order to enhance existing properties (improvement of ionic gel strength by additional covalent crosslinking, increase hydrophobicity of the backbone, improve biodegradation, or achieve greater HA nucleation and growth). Modified alginates introduce completely new properties otherwise not existing in unmodified alginates (anticoagulant properties, formation of chemical or biochemical anchors to interact with cell surfaces, change of temperature dependent characteristics such as lower critical solution temperature). The physical properties in aqueous media and biological activity of alginates strongly depend on the M/G ratio and distribution of M and G blocks along the chain (Aarstad et al., 2012). The water solubility of alginates depends on three parameters: pH of the solvent, ionic strength of the medium, and the presence of divalent ions in the solvent. Alginates are soluble above a certain critical value of pH when the carboxylic acid groups are deprotonated. Polymer conformation, chain extension, viscosity, and therefore, solubility are strongly influenced by changes in ionic strength (Pawar and Edgar, 2012).

13.1 Types of Biopolymer Matrices

Alginate chelates with divalent cations to form hydrogels. So, an aqueous solvent should be free of crosslinking ions to enable dissolution. In the presence of divalent cations gel formation is driven by the interactions between G-blocks with the formation of tightly held junctions (Donati et al., 2005). In addition to G-blocks, MG blocks also participate, forming weak junctions (Donati et al., 2005). Thus, alginates with high G contents yield stronger gels. The affinity of alginates toward divalent ions decreases in the order: Pb . Cu . Cd . Ba . Sr . Ca . Co, Ni, Zn . Mn (Pawar and Edgar, 2012). Ionotropic gels and acid gels from alginate will behave differently in terms of swelling properties and physical and chemical features. The degradation of alginate ionotropic gels could arise from the breaking of the divalent crosslinking (Rinaudo, 2014) (Fig. 13.1). Calcium crosslinking of alginates can be performed by two methods. The first is a “diffusion” method, wherein crosslinking ions diffuse into the alginate solution from an outside reservoir. The second is the “internal setting” method, where the ion source is located within the alginate solution and a controlled trigger (typically pH or solubility of the ion source) sets off the release of crosslinking ions into the solution (Pawar and Edgar, 2012). Diffusion set gel is made by dropping a Na-alginate solution into a CaCl2 bath. Internal set gels typically use insoluble calcium salts, such as CaCO3, as a calcium source. Sodium alginate gelation usually happens due to ion-exchange. This process is reversible in the presence of EDTA (sodium citrate, sodium oxalate). The gelation rate is a critical factor in controlling gel uniformity and strength when using divalent cations, and slower gelation produces more uniform structures and greater mechanical integrity (Lee and Mooney, 2012; Kuo and Ma, 2001). The polyanionic nature of alginate and its chemical structure make alginate suitable for forming highly hydrated hydrogels. The water entrapped in alginate gels is physically bonded with polymer network, but it is still able to migrate

FIGURE 13.1 Schematic crosslinking of alginate in the presence of Ca21 ions complexed with Lguluronic blocks. Reproduced with permission from Babaei, Z., Jahanshahi, M., Rabiee, S.M., 2013. The fabrication of nanocomposites via calcium phosphate formation on gelatin chitosan network and the gelatin influence on the properties of biphasic composites. Mater. Sci. Eng. C 33, 370 375.

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outside making the obtained gels as a first-line material for cells encapsulation (Augst et al., 2006). The gelation rate and mechanical properties of gels depend on gelation temperature (Lee and Mooney, 2012). At lower temperatures, the reactivity of ionic crosslinkers (e.g., Ca21) is reduced, and crosslinking becomes slower. The resulting crosslinked network structure has greater order, leading to enhanced mechanical properties (Drury et al., 2004). The mechanical properties of ionically crosslinked alginate gels can vary significantly depending on the chemical structure of alginate. For example, gels prepared from alginate with a high content of G residues exhibit higher stiffness than those with a low amount of G residues (Perucca et al., 2014). The limited long-term stability under physiological conditions is the critical drawback of ionically crosslinked alginate gels because such gels can be dissolved in the surrounding media due to the release of divalent ions and exchange reactions with monovalent cations. In addition, the release of calcium ions from the gel may promote hemostasis, while the gel serves as a matrix for aggregation of erythrocytes (Lee and Mooney, 2012). The solubility of alginates strongly depends on the state of its carboxylic acid groups. Alginic acid with protonated form carboxylic groups is not fully soluble in any examined solvent system, including water. Na-alginate is well dissolved in water. Sodium alginate (SA) in the form of dry powder can be stored without degradation in a cool, dry place and away from sunlight for several months. Its shelf life can be extended to several years by storing it in a freezer. The rate of degradation increases rapidly above pH 10.0 and below pH 5.0. Above pH 10 degradation arises mostly from the β-elimination mechanism, while below 5.0 degradation is mostly due to acid catalyzed hydrolysis (Pawar and Edgar, 2012). Alginates are susceptible to chain degradation not only in the presence of acids or bases, but also at neutral pH values in the presence of reducing compounds. In addition, sterilization techniques such as heat treatment, autoclaving, ethylene oxide treatment, and γ-irradiation cause alginate degradation (Pawar and Edgar, 2012; Drury et al., 2004). Alginates are hydrophilic polysaccharides. In addition to hydrophilicity resulting from the two hydroxyl groups, the carboxylate ion enhances water solubility at pH 5 and above. The motivation for hydrophobic modification of alginates is, therefore, to transform the polysaccharide from its predominantly hydrophilic nature to a molecule with amphiphilic or hydrophobic characteristics. The most straightforward way to achieve this transformation is by covalent attachment of hydrophobic moieties, such as long alkyl chains or aromatic groups to the polymer backbone (Pawar and Edgar, 2012). Alginate has found several applications in tissue repair and regeneration, in particular, as a bulking agent, for drug delivery, as a carrier for cell therapies, and as a model of extracellular matrix (Perucca et al., 2014). Alginates are widely used due to their ability to form beads, films, fibers, hydrogels, and composite materials in the presence of bivalent metal ions that interact with zones rich in G-blocks in a greater amount than with zones rich in M-blocks and MG-blocks (Rinaudo, 2014).

13.1 Types of Biopolymer Matrices

Alginates could be applied as a food additive in jams for stabilization, emulsification, gelation, and for increasing the viscosity of food (Lucconi et al., 2014) Edible films based on alginate combined with nanoparticles are used for antimicrobial packaging of fruits and vegetables (Sabra and Deckwer, 2005). The most significant application of alginates is biomaterial production (Norajit and Ryu, 2011; Lee and Mooney, 2012). They are used as stabilizers in the pharmaceutical industry, for controlled drug release, scaffold production, cell encapsulation (Rinaudo, 2014; Kuo and Ma, 2001), wound dressings, and hemostatic materials (Rinaudo, 2014; Lee and Mooney, 2012). Alginate based fibers are nontoxic, nonallergic, bacteriostatic, and biocompatible with high absorptive properties (Rinaudo, 2014; Draget and Taylor, 2011). Nanocomposite beads (SA/HA) are used as drug-controlled release matrices (Zhang et al., 2010a,b). A series of novel pH-sensitive nanocomposite beads based on sodium alginate/hydroxyapatite were prepared during the sol gel transition process of SA by the in situ generation of HA microparticles in the beads. Diclofenac sodium was used as the model drug. The synergistic effect of the biopolymer and inorganic material as well as the strong interfacial interactions between them via electrostatic interaction and hydrogen bonding could improve the mechanical properties, swelling behavior, controlled release behavior, and drug loading efficiency of the biopolymer matrices. HA is an ideal material for the preparation of drug scaffolds because of its excellent properties, such as the ability to adsorb a variety of chemical species and biocompatibility. The release of drugs from HA is fast, owing to the weak interaction between the drugs and the HA particles (Mizushima et al., 2006). The combination of biopolymer and HA seems to be a feasible way to prolong the release of drugs to make biopolymer/HA composites applicable for long-term controlled release carriers. The microparticles of HA act as inorganic crosslinkers in nanocomposites and decrease the movability of SA polymer chains, thereby changing their surface morphology and swelling ratio. The reaction time, temperature, and concentration of Ca21 ions, influenced the entrapment efficiency and release rate of drugs. The entrapment efficiency of diclofenac sodium was improved in sodium alginate/ hydroxyapatite nanocomposite beads (Gonza´lez-Rodrı´guez et al., 2002). HA and HA-Alg materials can be used as excellent natural adsorbents to remove Zn (II) from wastewaters with good efficiency and low cost. The amount of Zn21 adsorption was found to increase with increases in initial metal ion concentration and contact time but found to decrease with increases in the amount of adsorbent and temperature. The maximum adsorption capacity of hydroxyapatite was found to be 62.5 mg/g with an initial Zn21 concentration range of 30 90 ppm, whereas for HA-Alg was found to be 56.49 mg/g with the same Zn21 ion concentration range (Sen and Khoo, 2013). Being an anionic biopolymer sodium alginate is successfully combined with cationic chitosan. Chitosan is a deacetylated derivative of chitin used in material science, food, agriculture, environmental protection, medicine, and biotechnology (Aranaz et al., 2009; Honarkar et al., 2009; Dash et al., 2011). It is widely applied

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in tissue engineering, wound healing, drug delivery, and gene delivery (Muzzarelli, 2009) in various forms: gels, beads, membranes, nanofibers, nanofibrils, scaffolds, nanoparticles, microparticles and sponges (Honarkar et al., 2009; Dash et al., 2011; Muzzarelli, 2009; Cai et al., 2010; Dong et al., 2010; Radhakumary et al., 2011; Dev et al., 2010a,b; Ehrlich et al., 2006; Jayakumar et al., 2010; Anitha et al., 2011). It is a linear polysaccharide comprised of β-1,4 linked N-acetyl-2-amino-2-deoxy-D-glucose residues (Aranaz et al., 2009). Chitosan matrices have several advantages: flexibility, porosity, compatibility, processability biocompatibility, and biodegradability, but they are mechanically weak and unstable. To enhance mechanical strength polymeric blends of CS with Alg, poly(lactic-co-glycolic acid), poly(lactic acid), silk fibroin, carboxymethyl cellulose, carbon nanotubes, and gelatin are used as well as blends with bioactive and osteoconductive ceramics based on HA and other calcium phosphates with antibacterial activity (Srinivasan et al., 2012; Rezwan et al., 2006; Venkatesan et al., 2011; Jiang et al., 2008). The majority of naturally occurring polysaccharides, such as pectin, dextrin, agar, agarose, carrageenan, and cellulose are acidic in nature, while chitosan is a highly basic polysaccharide. Chitosan exhibits special properties, such as solubility in various media, viscosity, mucoadhesivity, polyelectrolyte behavior, ability to form films, metal chelation, and optical and structural characteristics. It also has the potential to bind with microbial and mammalian cells. Chitosan is considered a good candidate for tissue engineering, due to its regenerative effect on connective gum tissue. It plays an important role in osteoblast formation and shows fungistatic, hemostatic, analgesic, antitumor, antibacterial, antifungal, and anticholesteremic activities. It can be easily biodegraded into nontoxic residues and its rate of degradation is related to the molecular mass of the polymer along with deacetylation degree (Kim et al., 2011; Li et al., 2005; Shukla et al., 2013). Crosslinked scaffolds made of chitosan and various polymers (e.g., pectin, alginate) could be loaded with drugs. Alginate-crosslinked chitosan scaffolds and CS-pectin scaffolds are used as pentoxifylline delivery carriers (Lin and Yeh, 2010a,b). Ampicillin-loaded alginate microspheres coated with chitosan, chitosan scaffolds impregnated with dexametasone, CS/HA scaffolds loaded with tetracycline hydrochloride are also widely used as temporary drug carriers (Anitha et al., 2014). Nifedipine incorporated Col-CS membranes were developed for transdermal drug delivery applications using alginate gel as the drug reservoir. Modification of CS membranes with collagen makes the release of drugs more controlled. The permeability of membranes was found to depend on the concentration of Col and CS in membranes. Double layered scaffolds contain: a first mucoadhesive layer made of CS with or without an anionic crosslinking polymer (SA, gellan gum) and a backing layer made of ethyl cellulose. For membrane production CS could be combined with pectin, HA and HA coatings, Alg, and Col for drug delivery purposes (delivery of ketoprofen, ampicillin, tetracycline hydrochloride, fenbufen, amikacin, vitamin B2, lidocaine-HCl, berberine, nifedipine, vancomycin hydrochloride, and sodium salicylate) (Niranjan et al., 2013).

13.2 Calcium Phosphates as an Essential Part of Composite Materials

Doping and the addition of antibacterial metal ions, such as nanophase Cu21, Zn , and Ag1, to the CS polymer significantly improved its antimicrobial properties (Niranjan et al., 2013; Saravanan et al., 2013; Thian et al., 2013; Logith Kumar et al., 2016). Chitosan/hyaluronic acid scaffolds, upon the addition of calcium phosphate cement, exhibited a significant increase in ALP activity with no significant change in the rate of osteoblastic cell proliferation (Hesaraki and Nezafati, 2014). The presence of fucoidan in a CS/alginate scaffold stimulated the proliferation of MG-63 osteoblast-like cells with significant enhancement in ALP activity and apatite deposition over the scaffold (Venkatesan et al., 2015). The addition of nHAp to a CS/gelatin matrix not only increased the mechanical property of the scaffolds but also stimulated the proliferation and differentiation of induced pluripotent stem cells of gingival fibroblasts to osteocytes (Ji et al., 2015). Fucoidan in β-TCP-CS scaffolds increased the compressibility, apatite deposition, and osteocalcin release, which favor osteogenic differentiation of human mesenchymal stromal cells in vitro (Thian et al., 2013; Puvaneswary et al., 2015). Hyaluronic acid (HyA) is a major component of the ECM of the skin, cartilage, and vitreous humor. Almost 50% of the HyA in the human body is present in the skin, especially in the intracellular space, where its concentration reaches up to 2.5 g/L. HyA can affect cell proliferation, differentiation, and tissue repair (Collins and Birkinshaw, 2013; Chong et al., 2005). HyA is a nonsulfated, linear, natural polysaccharide composed of D-glucuronic acid and D-N-acetylglucosamine linked together by alternating β-1,4 and β-1,3 glycosidic bonds. It has high molecular mass and interesting viscoelastic properties that are influenced by its polymeric and polyelectrolyte characteristics. Only high molecular-weight HyA possesses mucoadherence and antiinflammatory properties (Collins and Birkinshaw, 2013). The high solubility of HyA impedes the development of polymers for tissue engineering. Therefore, crosslinking and esterification are performed in order to prevent water ingress and to reduce the solubility of HyA. Hyaluronic acid could be indirectly crosslinked by attaching thiols, methacrylates, or tyramines or it could be directly crosslinked with formaldehyde or divinyl sulfone. HyA is an essential functional component of almost all tissues in vertebrates (Park et al., 2016). 21

13.2 CALCIUM PHOSPHATES AS AN ESSENTIAL PART OF COMPOSITE MATERIALS Calcium phosphate (CaP) ceramics have been widely used as bone substitutes because of the similarity of their chemical composition to the mineral phase of bone (Samavedi et al., 2013; Dziadek et al., 2017), their biocompatibility, and ability to bond with bone tissue under certain conditions. However, because of their brittleness, their clinical applications have been limited to the non- or

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low-load bearing parts of the skeleton. Many types of calcium phosphates have been considered as biomaterials for bone reconstruction in dental, orthopedic, and maxillofacial applications due to different behavior in living organisms, including bioactivity, biodegradability, and biological response. Bioactivity, degradation behavior, osteoconductivity, and osteoinductivity of CaP ceramics generally depend on their Ca/P ratio, crystallinity, and phase composition (Samavedi et al., 2013; Dorozhkin, 2012). Calcium phosphate has the general formula Ca102x(HPO4)x(PO4)62x(OH)22x, with 0 # x , 2 (Lim et al., 2010). Three calcium phosphate cements, brushite (BR), tricalcium phosphate (TCP), and hydroxyapatite, were recently developed as bone substitution materials in orthopedic surgery. Hydroxyapatite (Ca10(PO4)6(OH)2) has been widely applied as a biomaterial for the substitution of human bone tissues due to its excellent biocompatibility and bioactivity (Babaei et al., 2013). Synthetic HA (Ca10(PO4)6(OH)2) shows good stability in the body, while tricalcium phosphates (α-TCP, β-TCP, Ca3(PO4)2) are more soluble, whereas biphasic calcium phosphate (BCP: a mixture of HA and β-TCP) exhibits intermediate properties depending on the weight ratio of stable/degradable phases. Thus, the dissolution rate decreases in the order: α-TCP . β-TCP . BCP . HA (Samavedi et al., 2013). Natural bone hydroxyapatite is nonstoichiometric and contains, beside the main components (Ca(PO4)32, (OH)2), some other groups and trace elements (e.g., CO32 , F2, Mg21, Na1, K1, Sr21, Zn21). So the new trends in calcium phosphates preparation give possibility to obtain substituted CaP ceramics. They do not only have a different solubility and bioactivity than their parent materials, but also cause modified biological response due to the release of biologically active ions during dissolution (Dziadek et al., 2017). CaP ceramics exhibit different biological effects in vivo, and while most of them are osteoconductive, only certain types are osteoinductive. Samavedi et al., proposed that the range of osteoinductive potentials of CaPs decrease in the order: BCP . TCP . HA (Samavedi et al., 2013). Currently HA, TCP, and BCP ceramics are quite common types of materials used for various biomedical applications and they are readily available on the market (Liuyun et al., 2013). In the past few years, many efforts have been directed at developing calcium phosphate-containing composite materials, especially polymer-matrix composites. Brushite (dicalcium phosphate dihydrate: DCPD, (CaHPO4))2H2O) is a phase of calcium phosphate that has a higher solubility than HA at physiological pH and an ideal in vivo resorption rate that can match the rate of new bone formation. Besides having bioresorpability characteristics, Brushite-based orthopedic cements are biocompatible, osteoconductive, and bioresorbable, and have mechanical properties similar to those of cancellous bone. Compared to HA, tricalcium phosphate (TCP: Ca3(PO4)2) is generally considered as a resorbable bioceramic. β-TCP is the preferred bioceramic among all TCPs because of its mechanical strength, good tissue compatibility, and ability to bond directly to tissue for the regeneration of bone without any intermediate connective tissue.

13.2 Calcium Phosphates as an Essential Part of Composite Materials

In addition, fast bone regeneration and proper bioresorption rate are other additional attributes of β-TCP (Babaei et al., 2013). Stoichiometric HA (Ca/P molar ratio of 1.67) is considered to be osteoconductive but not osteoinductive. Furthermore, because HA is the most stable among the CaP ceramics, its surfaces provide highly effective nucleating sites for the precipitation of apatite crystals in contact with culture medium and bodily fluids (Samavedi et al., 2013). The solubility, bioactivity, and biological response of HA can be modified by anionic and cationic substitution. Mainly for these reasons, HA is widely used for preparing polymer-ceramic composite materials usually with the aim of imparting bioactivity and osteoconductivity and also to improve mechanical properties (Thein-Han and Misra, 2009). One of the most promising bone grafting materials is a collagen/nanosized HA composite due to its ability to mimic the composition and structure of natural bone. New biomineralization and self-assembly technics have been used to obtain collagen/HA composites with oriented and hierarchical structure between organic and inorganic phases (Ficai et al., 2010; Zhang et al., 2010a,b). A highly oriented collagen/HA composite with morphology similar to compact bones was obtained using the self-assembling method (Ficai et al., 2010). A method of precrystallization allowed for a composite material to be obtained, in which HA nanocrystals are located between collagen fibers with the creation of a suitable structure for bone defect regeneration (Zhang et al., 2010a,b). Several studies were conducted on the use of another natural polymer, namely CS, as a component of HA-containing composites. The presence of nanosized HA in a CS matrix increased the compressive strength and Young’s modulus of the composites. This can be related to hydrogen-bonding interactions between NH2 and OH groups of HA and chelation between NH2 and Ca21 when a coprecipitation method of composite preparation was used. The incorporation of HA affected the degradation rate of the composites, induced bioactivity, and favored osteoblast cell attachment and proliferation (Thein-Han and Misra, 2009). HA fillers have also been shown to modify numerous properties of synthetic polymer matrices, especially poly(α-hydroxy esters). The comparison of hydroxyapatite with different morphologies, namely nanoparticles (nHA) and whiskers (wHA), as poly(lactic-co-glycolic acid (PLGA)) matrix modifiers, showed that composites with nanosized HA had higher bending strength in comparison to wHA-containing materials. This can be attributed to a more homogeneous distribution of nHA in the polymer matrix and also to the enhanced crystallization of the PLGA matrix. It was also shown that the morphology of HA fillers influenced the in vitro degradation behavior of the composites (Liuyun et al., 2014). In turn, the use of various contents (0 30 wt.%) and sizes (0 50 μm, 5 μm, and N200 nm) of HA particles led to changes in the thermal properties and/or crystallinity, as well as the mechanical strength of the poly(L-lactic acid) (PLLA) and PLGA-based composites (Damadzadeh et al., 2010). Tricalcium phosphate (β-TCP, TCP) is a well-known CaP-based bioceramic. β-tricalcium phosphate has been widely investigated and successfully used in

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clinical applications as a biomaterial for bone repair applications due to its remarkable biocompatibility, in vivo resorbability, bioactivity, and good osteoconductivity (Dorozhkin, 2010; Zheng et al., 2011). Moreover, results have indicated the osteoinductivity of TCP (Dziadek et al., 2017). It has been successfully combined with other ceramic materials: Al2O3, HA, ZrO2, to produce small orthopedic or dental implants (Jang et al., 2011; Kulkova et al., 2014; Mina et al., 2015) in the form of composite layers on the implants (Kulkova et al., 2014; Mina et al., 2015; Jang et al., 2011). The most prospective form of TCP is nanoparticles which modify a biodegradable polymer matrix (Kulkova et al., 2014; Jang et al., 2011; Hu et al., 2010). Significant osteogenic benefits and the formation of new bone on the implant surface was observed due to the incorporation of ceramic particles into polymer-based composite materials (Cao and Kuboyama, 2010; Homaeigohar et al., 2006). In addition, during the degradation process of composite implants/scaffolds based on biodegradable aliphatic polyesters, particles of TCP which modified the polymer matrix, improved not only osteoconductivity but also neutralized the acidic pH of the environment typical for degradation of a polymer matrix (Kunze et al., 2003; Bhumiratana et al., 2011; Lu and Zreiqat, 2010). These phenomena decrease the risk of local complications (inflammatory reactions) and increase degradation rate compared to pure polymer implants (Kim et al., 2012). Moreover, preparation techniques under room temperature may allow for the incorporation of growth factors (BMP-2) and thermally unstable antibiotics. Octacalcium phosphate (OCP) is considered to be a mineral precursor to carbonate-containing calcium-deficient HA, which is a prototype for apatite crystals in bones and teeth. The structure of OCP is closely associated with HA, because OCP is composed of apatite layers stacked alternately with hydrated layers and can be converted into Ca-deficient HA in neutral aqueous conditions (Handa et al., 2012). Biphasic calcium phosphate ceramics are a family of two-phase materials that combine the low solubility and osteoconductive character of HA with the osteoinductivity of a more soluble phase, namely β-TCP (Arafat et al., 2011). The main advantage is the possibility of controlling the bioactivity, resorption rate, and mechanical properties of CaP ceramics by manipulating the HA/β-TCP ratio (Dziadek et al., 2017). BCP properties, especially biological effects, can be also modified by ion substitution (Dziadek et al., 2017). A study on polycaprolactone/ TCP beads loaded with vancomycin (1 4 wt.%) showed that drug release had a diffusional character. It was observed over a period of 4 11 weeks depending on the composite matrix homogeneity and porosity (Dziadek et al., 2017). A commercially available collagen/BCP composite (Osteont II collagen, Genoss. Co. Ltd.) was studied using a rabbit calvarial defect model (Dorozhkin, 2010). The composite materials showed slow resorption, good osteoconductive properties, and maintained the space needed for bone tissue regeneration in contrast to a collagen sponge (Lee et al., 2015a). It was also shown that the collagen/BCP composite is a promising candidate for a carrier providing there is a constant release

13.3 Formation of Composite Materials

profile of recombinant human BMP-2 (Lee et al., 2015b). A gelatin/pectin/BCP nanocomposite scaffold was fabricated for growth factor (BMP-2 and VEGF) delivery. The studies observed that the scaffold had constant release properties. Furthermore, in vitro and in vivo tests conducted with osteoblast-like MC3T3-E1 and rat models, respectively, showed that the composites enhanced cell proliferation and new bone formation (Amirian et al., 2015).

13.3 FORMATION OF COMPOSITE MATERIALS The most popular method of nanocomposite formation is the biomimetic process. Gelatin based composites were prepared through the precipitation of hydroxyapatite nanocrystals in an aqueous solution of gelatin at pH 8 and 38 C. The development of HA nanocrystals in an aqueous solution of gelatin was influenced by the Gel/HA ratio. A higher gelatin concentration induced tiny crystallite formation (4 9 nm), while a lower concentration contributed to the development of bigger crystallites (30 70 nm). The coprecipitated HA-Gel nanocomposites showed chemical bond formation between HA nanocrystals and Gel macromolecules and had a self-organized structure along gelatin fibrils (Chang et al., 2003). Highly prospective are composites of calcium phosphate in conjugation with gelatin due to the interaction between gelatin chains by hydrogen bonds (Babaei et al., 2013). In combination with CS and glutaraldehyde reinforced composite material is obtained (Figs. 13.2 and 13.6). Composites of CS-Gel/calcium phosphate were synthesized using an in situ precipitation method. The amount of Ca(NO3)2
4H2O, (NH4)2HPO4, CS, and Gel were calculated to make composites with different ratios of organic/inorganic phase. In the first step, a chitosan solution was prepared by dissolving chitosan in acetic acid solution with stirring at room temperature. Then (NH4)2HPO4 was added to the CS solution under agitation until the salt was entirely dissolved. Subsequently, the Ca(NO3)2
4H2O solution was added drop by drop into the CS solution. Gel was dissolved in distilled water at 40 C for 1 hours. Upon complete dissolution, the gelatin solution was added dropwise into the chitosan solution of Ca21 and PO432 with stirring at ambient temperature. Next, 150 μL glutaraldehyde (25 wt.%) as a crosslinker was added to the previously mixed solution. The synthesized hydrogel was stored under ambient conditions for 24 hours to achieve complete crosslinking. It was then kept in an ammonia solution for 48 hours at 25 C. Under this alkaline condition, CaP crystals precipitated within the hydrogel matrix. Then the composite was gradually rinsed with distilled water. Finally, the hydrogel-like composites were set into petri dishes and lyophilized with a freeze dryer and the CS-Gel/CaP was obtained (Babaei et al., 2013). A series of hydroxyapatite/gelatin/alginate nanocomposites with different amounts of alginate were synthesized through the method of coprecipitation. With the increase in amount of alginate, a crosslinked alginate/gelatin polymer network

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FIGURE 13.2 Schematic representation of the formation mechanism of homogeneous inorganic/organic composites by in situ precipitation (Babaei et al., 2013).

is formed. On increasing alginate content, the dimensions of the crystals increased and their morphology changed from needle-like to long fiber-like, and at high alginate content, the crystals tended to aggregate in separate clusters (Ehrlich et al., 2006). Alginate scaffolds were prepared by adding crosslinking agents (calcium based substances) freeze-dry lyophilization, and electrospinning. (Venkatesan et al., 2015) Sodium alginate has the distinctive ability to form hydrogels and is also prepared via ionotropic crosslinking in the presence of divalent cations, such as calcium ions (Venkatesan et al., 2015). Hydrogels based on calcium-crosslinked alginate have been widely investigated for various drug delivery purposes (El-Sherbiny, 2010; El-Sherbiny et al., 2010, 2011). Besides the external addition of calcium to crosslink the alginate hydrogel, it has been reported that sodium alginate can form an injectable system with Ca21 slowly and partially released from calcium-containing materials, such as CaCO3, CaSO4, and hydroxyapatite in the presence of D-gluconic acid D-lactone (GDL) (Lee and Mooney, 2012; Lee et al., 2011).

13.3 Formation of Composite Materials

Calcium alginate beads greater than 1.0 mm in diameter can be prepared using a syringe with a needle or a pipette (Gombotz and Wee, 2012). Sodium alginate solution that contains the solubilized protein is transferred drop-wise into a gently agitated divalent crosslinking solution. The diameter of the beads formed is dependent on the size of the needle used and the viscosity of the alginate solution. A larger diameter needle and higher viscosity solutions will produce larger diameter beads. The viscosity of sodium alginate can also influence the shape of the microbeads produced. The beads become more spherical as the concentration of sodium alginate solution is increased (Badwan et al., 1985). However, in general, sodium alginate solutions of greater than 5% (w/v) are difficult to prepare. The beads that are formed are allowed to be fully cured in the crosslinking solution for a short period of time, usually minutes, before they are rinsed with distilled water (Gombotz and Wee, 2012). Another method of microbead preparation involves protein encapsulation using an oil-in-water emulsification technique (Monshipouri and Price, 1995). This encapsulation method may work better for stable peptides and proteins or synthetic low molecular weight drugs since it involves the use of harsher chemical reagents, such as ethyl ether, to remove the oil at the end of the process. The size of microbeads formed by emulsification is highly dependent on the stirring speed and the rate of the addition of the crosslinking solution. Complex coacervation of oppositely charged polyelectrolytes has been commonly used as a method for preparing microbeads. Under specific conditions of polyion concentration, pH, and ionic strength, the polyelectrolyte mixture can separate into two distinct phases; a dense coacervate phase which contains the microbeads and a dilute equilibrium phase (Gombotz and Wee, 2012). Various types of proteins could be encapsulated in alginate microbeads: albumin, bovine serum albumin (BSA), endothelial cell growth factor, epidermal growth factor, fibrinogen, gamma globulin, insulin, leukemia inhibitory factor (LIF), myoglobin, nerve growth factor, interleukin-2 (IL-2) (Gombotz and Wee, 2012). Complex coacervation could be observed between alginic acid, gelatin, chitosan, and albumin. In the alginate-chitosan system, the complex is formed by spraying a sodium alginate solution into the chitosan solution. The resultant alginate-chitosan microbeads are mechanically strong and stable over a wide pH range. With the alginate-albumin system, coacervation is found to be limited compared to other polypeptide-polysaccharide systems due to the high viscosity of the albumin-alginic acid complex and a propensity to precipitate. The optimum conditions for maximum coacervate yield are pH 3.9, an ionic strength of 1 mM, and a 0.15% (w/v) total polyion concentration (Gombotz and Wee, 2012). Alginate matrices have proven to be useful for the slow release of several potential therapeutic proteins and several studies have demonstrated the in vitro and in vivo efficacy of these systems. The HA/calcium alginate microsphere composite was prepared through a crosslinking method with a particle size in the range of 125 425 μm (Liu et al., 2007). The HA/alginate composite was formed as a result of the in situ nucleation

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of HA on the alginate polymeric chain, which can be used as a scaffold for bone and tissue growth. HA crystallites are needle-shaped with a typical length of 20 50 nm and a width of 5 nm (Wang et al., 2009a,b). Porous HA/alginate composites with good mechanical properties are developed by integrating HA powder with alginate crosslinked with calcium. The mechanical properties can be increased with an increase in the concentration of alginate (Rajkumar et al., 2011). The preparation of the HA/alginate composite from antibiotic-loaded HA spherical granules and alginate ensures that the bone replacement material remains in the location of the implantation and also avoids any postoperative infection (Venkatesan et al., 2015). Alginate hydrogels can be prepared through the crosslinking with two moieties of polymer strands (Venkatesan et al., 2015). Injectable self-gelling alginate formulations were obtained by mixing alginate microspheres (as calcium reservoirs) with soluble alginate solutions for potential use in immunotherapy (Hori et al., 2009). The redistribution of Ca21 ions from microspheres into the alginate solution led to crosslinking and stable hydrogel formation, the mechanical properties of which correlated with the concentration of calcium-reservoir microspheres added to the solution. Soluble factors (cytokine interleukin-2) were incorporated into alginate matrices by simply mixing them prior to gelation. Due to self-gelling ability alginate formed soft macroporous gels allowing ready access to microspheres and carrying therapeutic factors embedded in the matrix. They are useful for stimulating immune cells, as well as for other soft tissue regeneration applications (Hori et al., 2009). Porous scaffolds of alginate combined with calcium phosphate were prepared by mixing calcium phosphate powder with sodium alginate following with injection into a calcium containing bath in order to obtain fibrous structures. (Venkatesan et al., 2015; Lee et al., 2011). The rapid reaction between calcium ions and alginate allows for a 3D porous scaffold with different porosity level to be obtained. The addition of alginate shortens setting time, increases hardening reaction, stromal cell attachment, proliferation, and improves the different kinds of shapes, such as microspheres or fibers (Venkatesan et al., 2015; Lee and Mooney, 2012; Lee et al., 2011). In the work Yanovska et al. (2016), composite materials based on carbonate HA and gelatin were prepared in situ by coprecipitation from aqueous solutions with the addition of ZrO2 and silver ions. The first solution was obtained through the addition of 0.84 g NaHCO3 to 0.1 M solution of Ca(NO3)2
4H2O with stirring until fully dissolved. Gelatin was added during the synthesis of carbonate apatite: 5% gelatin solution (80 C) was added to 0.06 M solution of Na2HPO4
4H2O with stirring and maintaining constant temperature (second solution). The obtained solution was added drop-wise to the first one with constant stirring and 10 M solution of NaOH until pH 5 8 was obtained. The obtained product was uniform fine composite that was divided into two parts. In the first part 5 mL of AgNO3 solution with 10 g/L concentration

13.3 Formation of Composite Materials

FIGURE 13.3 The synthesis scheme of composite materials HA-Gel (B), and HA-Gel-Ag (C) obtained by coprecipitation from aqueous solutions. Reproduced with permission from Yanovska, A., Kuznetsov, V., Stanislavov, A., Husak, E., Pogorielov, M., Starikov, V., et al., 2016. Synthesis and characterization of hydroxyapatite-gelatine composite materials for orthopaedic application. Mater. Chem. Phys. 183(1), 93 100.

(sample c) was added, the second part remained unchanged. The obtained composites were rinsed with distilled water, aged for 48 hours at 22 C and freeze dried. The synthesis scheme is presented in Fig. 13.3. In the Fig. 13.4 morphologyHA-gelatin composite materials is present: general view of obtained composite materials (1a 1e) shows that most of them have porous (1 a, c, d) or laminated structure (1e). Micrographs at magnification 3 4 (2a 2e) and 3 10 (3a 3e ) were made by optical microscopy. SEM micrographs (4a 4e) were made at magnification 3 50. Porous structure is preferred for ingrowths of newly formed bone tissue. The calculated porosity of the obtained materials is about 75% 80%, which corresponds to the porosity of a spongy bone (Yanovska et al., 2016). The study concerns comparisons of HA and gelatin based composite materials obtained in situ from aqueous solutions. Some materials, like HA/Gel/ZrO2, HA/ Gel/Ag, HA/Gel, have biodegradation and osteoconductive features allowing for the consideration of mixing ultrasound and coprecipitation methods as appropriate for apatite formation. The formation of the composites through coprecipitation of HA and gelatin gives porous, biocompatible, and biodegradable HA/Gel/Ag and HA/Gel materials (Yanovska et al., 2016). Composite scaffolds CS-nHA with different fractions of nHA (0, 0.5, 1, and 2 wt. % of nanosized HA) were synthesized by Thein-Han and Misra (2009). The nHA powder was ultrasonically dispersed in deionized water for 2 hours. Then, the dispersed particles were added dropwise to the chitosan solution, with the solution being

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FIGURE 13.4 Morphology of obtained composites, general view: (1A) HA-Gel-ZrO2, (1B) HA-Gel, (1C) HA-Gel-Ag, (1D) HA-Gel-ZrO2-Ag, (1E) HA-Gel, and optical micrographs of listed composite materials at magnification 3 4 (2A E) and at magnification 3 10 (3A E). SEM micrographs (4A E). Reproduced with permission from Yanovska, A., Kuznetsov, V., Stanislavov, A., Husak, E., Pogorielov, M., Starikov, V., et al., 2016. Synthesis and characterization of hydroxyapatite-gelatine composite materials for orthopaedic application. Mater. Chem. Phys. 183(1), 93 100

agitated and mixed using a magnetic stirrer for 1.5 2 hours. The obtained pure chitosan solution or CS-nHA dispersion was transferred to polystyrene petri dishes (2 cm2), frozen at 20 C for 24 hours, and lyophilized in a freeze dryer. Composite materials showed interconnected pore structures with uniform and nonagglomerated distribution of nHA particles within the chitosan matrices, including the pore wall. It is believed that the electrically charged nature of the chitosan network is helpful in the uniform distribution of HA. The initial cell attachment of CS and CS-nHA scaffolds after 4 hours was found to be 75% 6 5% and 85% 6 4%,

13.4 Biomedical Applications of Obtained Composite Materials

respectively. Therefore, a CS-nHA composite is a potential scaffold material for bone regeneration (Thein-Han and Misra, 2009). Porous HA-CS-Alg composite scaffolds were prepared through in situ coprecipitation and freeze-drying for bone tissue engineering. At low HA concentrations the HA-CS-Alg composite scaffolds were highly porous and interconnected with a pore size of around 50 220 μm. The porosity of the scaffolds decreased with increasing HA content from B85% to B75% and density increased from 0.1 to 0.24 g/cm3. First a chitosan aqueous solution was prepared by dissolving chitosan powder in an acetic acid solution, then alginate powder was dissolved in a NaOH solution. Orthophosphoric acid and calcium hydroxide were added to the chitosan and alginate solutions respectively to form HA-CS-Alg composite scaffolds. The ratios of CS to H3PO4 and Alg to Ca(OH)2 were adjusted so that the final HACS-Alg weight ratios were 10/90 and 70/30 respectively. The resulting suspension was mixed under constant stirring in a blender for 1 hour. Acetic acid was gradually added dropwise to the suspension until pH 7.4 was obtained. The slurry was placed into cell culture plates and stored in a freezer at 15 C until frozen. The samples were lyophilized and dried, crosslinked with 1% CaCl2 solution, and then immersed in distilled water for 24 hours, washed, and finally freeze-dried at 80 C for 24 hours to obtain HA-CS-Alg composite scaffolds (Jin et al., 2012). The composite scaffolds were highly porous and interconnected with a pore size of around 50 220 μm at low concentrations of HA. As the HA content increased, the porosity of the scaffolds decreased from 84.98% to 74.54%. The CS-Alg polyelectrolyte complex was highly porous and interconnected with a pore size of around 50 220 μm. At low HA contents, the pore structure of the composite scaffolds (Fig. 13.5B and C) was similar to the CS-Alg polyelectrolyte complex (Fig. 13.5A). Pore size decreased with increasing HA content above 30 wt.% and pore structure locally collapsed and agglomerated (Fig. 13.5D and E). The obtained scaffolds had no cytotoxic effects on MG-63 cells, and they had good biocompatibility. An implantation experiment in mouse skulls revealed that the composite scaffold provides a strong positive effect on bone formation in vivo in mice. Furthermore, the HA-CS-Alg composite scaffold has been shown to be more effective for new bone generation than the CS-Alg scaffold (Jin et al., 2012).

13.4 BIOMEDICAL APPLICATIONS OF OBTAINED COMPOSITE MATERIALS The development of synthetic materials with strong osteogenic characteristics remains a challenging subject in the field of biomaterials science. Some reports have shown that synthetic calcium phosphate ceramics alone have osteogenic characteristics when used in large animal models, such as sheep (Yuan et al., 2010). Nevertheless, the use of synthetic materials alone, without cells or growth

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FIGURE 13.5 Photograph of HA-CS-Alg composite scaffolds and SEM morphology of composite scaffolds with different HA contents: (B) 0, (C) 10, (D) 30, (E) 50, and (F) 70 wt.%. Reproduced with permission from Jin, H.-H., Kim, D.-H., Kim, T.-W., Shin, K.-K., Jung, J.S., Park, H.-C., et al., 2012. In vivo evaluation of porous hydroxyapatite/chitosan alginate composite scaffolds for bone tissue engineering. Int. J. Biol. Macromol. 51, 1079 1085.

factors, is still favorable, since these types of combinations increase the handling complexity and costs for the users. The composite OCP/Gel disks obtained by vacuum dehydrothermal treatment were implanted in Wistar rat calvarial critical-sized defects (CSD) for 4, 8, and 16 weeks. Then they were subjected to radiologic, histologic, histomorphometric, and histochemical assessment. After 1 day of incubation, the attachment of mouse bone marrow stromal ST-2 cells on the disks of the OCP/Gel composites was also examined. OCP/Gel composites containing 24, 31, and 40 wt.% of OCP with pore sizes of 10 500 μm were obtained. Plate-like crystals were observed to be closely associated with Gel matrices. The OCP (40 wt.%)/Gel composite repaired 71% of the bone defect in conjunction with material degradation by osteoclastic cells. Thus, an OCP/Gel composite can repair rat calvarial crania sized defects highly efficiently and has favorable biodegradation characteristics (Handa et al., 2012). The chemistry and relatively mild crosslinking conditions of alginate have enabled this naturally occurring biopolymer to be used for the encapsulation of a wide variety of biologically active agents, including proteins, cells, and DNA oligonucleotides. Through the selection of the appropriate alginate type, gelation conditions, added excipients, and coating agents, matrices of various morphologies, pore size, water content, and dehydration rates can be fabricated. This high degree of flexibility can result in the delivery of active agents over time periods ranging from minutes to months (Gombotz and Wee, 2012). Alginate hydrogels encapsulated with adipose derived mesenchymal stem cells were used for the delivery of biologically active molecules (BMP). They were

13.4 Biomedical Applications of Obtained Composite Materials

obtained by the mixing of cells with potassium alginate followed by sedimentation in CaCl2 solution (Venkatesan et al., 2015, 160). Alg-HA composite materials could be used in bone tissue engineering as a scaffold material to deliver cells, BMP-2, and growth factors (Venkatesan et al., 2015). Cytocompatibility, osteoinduction, and bone mineralization (ALP and osteocalcin assay) of HA-Ag(10%) composites were illustrated (Thrivikraman et al., 2014). In the work by Yanovska et al. (2016), composite materials containing 0.01% 0.1% silver were synthesized. Such and amount of silver is nontoxic but has antibacterial activity and decreases inflammation at the initial stage of implantation. Zirconium oxide falls under the category of bioinert materials (Thrivikraman et al., 2014). Zirconium materials are widely used as implant materials. Zirconium compounds are generally considered to be of low toxicity. Their cytotoxicity, carcinogenicity, mutagenic, or chromosomal alterations in fibroblasts or blood cells have not been observed (Vagkopoulou et al., 2009). Composite material HA-Gel was implanted as a powder, while samples HAGel-ZrO2, HA-Gel-Ag, HA-Gel-ZrO2-Ag, and HA-Gel as porous sponges. After 21 days of composite materials HA-Gel-ZrO2 and HA-Gel-Ag implantation, the defect was filled with newly-formed bone tissue with irregular staining as compared to the host bone (Yanovska et al., 2016). There was a network of vessels visible inside the new bone tissue and red bone marrow cells that filled the intertrabecular spaces. The newly-formed tissue was rich in osteoblasts, which indicated an active remodeling process (Plut et al., 2015). Connective-tissue capsule formed around sample HA-Gel 21 days after implantation. It was made by random oriented collagen fibrils and a thickness of 28.65 6 14.55 μm was reported. There were some vessels, fibroblasts, lymphocytes, and neutrophils visualized in the connective-tissue capsule and some macrophages near the area contacting with the implanted material. There was no tissue ingrowths and cell migration into the implant pores, which indicates the absence of osteoconductive properties needed for successful implant ingrowths in bone tissue (Jahan and Tabrizian, 2015). Partial biodegradation of material HA-Gel-ZrO2-Ag resulted in the formation of a connective-tissue capsule after 21 days of implantation. The pores of the material were filled with connective tissue that had a high density of leukocytes and macrophages. A newly-formed bone tissue was visualized around the remnants of the composite material. The vascular network, which complicates regeneration processes, was not observed. Since angiogenesis is connected with the bone healing process its suppression leads to decelerated healing (Geris et al., 2008). Composite material HA-Gel (sponge) was completely substituted by newlyformed bone tissue up to 21 days after implantation. New bone tissue has irregular staining that may indicate incomplete mineralization. A lot of vessels, being a source of osteogenic cells were observed in the new tissue. Compared to the host bone, the new one has a large amount of osteoblasts taking part in remodeling of newly-formed bone tissue.

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After implantation, samples HA-Gel (powder) and HA-Gel-ZrO2-Ag did not reveal osteoconductive properties, and a connective-tissue capsule is formed around the implant. Slow biodegradation is observed for HA-Gel samples implanted as powder or sponge. Given that the porosity of the samples is sufficient, degradation of the properties may be explained by either insufficient or increased density. This feature may slow down or even prevent vascular ingrowth, therefore, osteogenic elements penetrate the defect zone slowly. As the composites HA-Gel-ZrO2, HA-Gel-Ag, HA-Gel (sponge) were introduced into the bone defect, there was no any sign of inflammation in the implantation zone and in the neighboring tissues by the 21st day. The biocompatibility of the composites with bone tissue was observed. In addition, full biodegradation of the introduced composites takes place. The formation of a tubular-bone tissue at such an early stage of implantation indicates the apparent reparative and hemostatic features of the composites (Yanovska et al., 2016). In Fig. 13.6 the morphology was different on pure CS and nanocomposite scaffolds is presented. In general, preosteoblasts attached, spread, and grew well on both types of scaffolds and their morphology changed with time, irrespective of the type of scaffold. The morphology day 1 was spherical on pure CS and flat and star-like

FIGURE 13.6 SEM illustrating morphology of preosteoblasts seeded on high molecular weight chitosan and CS-nHA scaffolds (saggital section). Preosteoblasts on CS surface after (A) day 1, (B) day 3, (C) day 7, and (D) day 21 of cell culture; and on CS-nHA surface after (E) day 1, (F) day 3, (G) day 7, and (H) day 21 of cell culture. Reproduced with permission from Thein-Han, W.W., Misra, R.D.K., 2009. Biomimetic chitosan nanohydroxyapatite composite scaffolds for bone tissue engineering. Acta Biomater. 5, 1182 1197.

References

on the nanocomposite scaffold. After day 21, the morphology of the pure CS and nanocomposite scaffolds had the resemblance of a sheet. The surface of the cells on the CS scaffolds appeared to be less rough as compared to those present on the CS-nHA nanocomposite scaffolds. In pure CS, cells were present as clusters or agglomerates, while in the nanocomposite scaffold the cells coalesced to form a large and flat lump (Fig. 13.6F). After three days of culture, the cell morphology of the CS-nHA nanocomposite scaffold differed from that of the chitosan scaffold from the viewpoint of cell contact, extracellular matrix formation, and cell penetration (Fig. 13.6B and F). Preosteoblasts on the CS surface after day 3 and day 21 showed cell proliferation and cell cell contacts with extracellular matrix formation. After 7 days of culture, the cell population was higher on the nanocomposite scaffolds than on the CS scaffolds. Fibrous collagenous extracellular matrix was observed on both types of scaffolds’ surfaces (Fig. 13.6). The highly proliferated preosteoblasts with their extracellular matrices had small globules of mineral deposits on them together with cellular attachments. After 21 days, the majority of the CS-nHA composite scaffold areas were covered by layers of osteogenic cells. CS-nHA composite scaffolds were compared to pure CS scaffolds for bone tissue engineering. Chitosan scaffolds with nHA content 0.5, 1, and 2 wt.% were successfully fabricated through freezing and lyophilization. The nanocomposite scaffolds showed a highly porous structure with a pore size similar to the scaffolds with varying nHA content. The nanocomposite scaffolds exhibited greater compression modulus, slower degradation rate, and reduced water uptake, but their water retention ability was similar to that of pure CS scaffolds. Favorable biological responses of preosteoblast (MC 3T3-E1) on nanocomposite scaffolds included improved cell adhesion, higher proliferation, and well spreading morphology in relation to the pure CS scaffolds. This behavior was observed in spite of the similarity between the porous structure of pure chitosan and the nanocomposite scaffold, suggesting the determining role of nHA in facilitating the greater cellularity. This chapter provides a comparative assessment of the structure property process relationship of 3D CS-nHA and pure CS scaffolds in conjunction with their respective biological responses and advances as well as insight on aspects that concern bone tissue engineering (Thein-Han and Misra, 2009). Porous HA-CS-Alg composite scaffolds were prepared through in situ coprecipitation and freeze-drying for bone tissue engineering (Jin et al., 2012). The obtained scaffolds have no cytotoxic effects on MG-63 cells, and they have good biocompatibility.

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FURTHER READING Benoit, D.S.W., Nuttelman, C.R., Collins, S.D., Anseth, K.S., 2006. Synthesis and characterization of a fluvastatin-releasing hydrogel delivery system to modulate hMSC differentiation and function for bone regeneration. Biomaterials. 27, 6102 6110. Morch, Y.A., Donati, I., Strand, B.L., Skja˚k-Bræk, G., 2006. Effect of Ca21, Ba21, and Sr21 on alginate microbeads. Biomacromolecules 7, 1471 1480. Wang, L., Stegemann, J.P., 2010. Thermogelling chitosan and collagen composite hydrogels initiated with β-glycerophosphate for bone tissue engineering. Biomaterials. 31, 3976 3985. Zilberman, Y., Turgeman, G., Pelled, G., Xu, N., Moutsatsos, I.K., Hortelano, G., et al., 2002. Polymer-encapsulated engineered adult mesenchymal stem cells secrete exogenously regulated rhBMP-2, and induce osteogenic and angiogenic tissue formation. Polym. Adv. Tchnol. 13, 863 870.

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14

Study of microstructural, structural, mechanical, and vibrational properties of defatted trabecular bovine bones: natural sponges

Sandra M. London˜o-Restrepo1, Cristian F. Ramirez-Gutierrez1, Herminso Villarraga-Go´mez2 and Mario E. Rodriguez-Garcı´a3 1

Posgrado en Ciencia e Ingenierı´a de Materiales, Centro de Fı´sica Aplicada y Tecnologı´a Avanzada, Universidad Nacional Auto´noma de Me´xico, Quere´taro, Me´xico 2Nikon Metrology, Inc., Brighton, MI, United States 3Departamento de Nanotecnologı´a, Centro de Fı´sica Aplicada y Tecnologı´a Avanzada, Universidad Nacional Auto´noma de Me´xico, Quere´taro, Me´xico

14.1 INTRODUCTION Autografts, allografts, and xenografts have been used for tissue repair. Although the use of autografts implies faster bone integration than in the case of allografts and xenografts, the surgical risk is higher. Allografts have a lot of advantages, such as a minor risk of infection, low costs, biosecurity, among others, but the demand of tissue donors for allografts exceeds the number of donors. On the other hand, xenografts have become a safer alternative to bone regeneration due to their similitude with human bone and because they have a better osteogenic response. Tissue engineering has designed synthetic scaffolds to promote bone regeneration for clinical applications, but its attempts have not yet been able to reproduce all the physicochemical properties of natural bone. As such, xenografts are still considered the best option for medical applications, and the work presented here provides a complete physicochemical characterization of cube sponges extracted from bovine trabecular bone. According to O’Brien (2011), there are five requirements that a scaffold must meet in order to be used in tissue engineering: biocompatibility, biodegradability, mechanical properties, architecture, and manufacturing technology. Biocompatibility refers to biological acceptability; it means that the tissue has cellular activity with the scaffold without adverse effects. Here, osteoconductivity and osteoinductivity play a vital role. It is well known that hydroxyapatite is biologically active and compatible, but when it is obtained from mammalian bones, the organic phase must be wholly removed to guarantee an adequate Materials for Biomedical Engineering: Hydrogels and Polymer-based Scaffolds. DOI: https://doi.org/10.1016/B978-0-12-816901-8.00014-6 © 2019 Elsevier Inc. All rights reserved.

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response by tissue. Biodegradability is related to the process through which a biological system degrades a scaffold (e.g., cellular, enzymatic, bacterial, or viral). The mechanical properties of a biomaterial must match that of the host bone properties (Bose et al., 2012), and the anatomical site where it will be implanted. Designing and obtaining materials with the appropriate mechanical properties represents a big challenge for tissue engineering. Regarding scaffold architecture, it has been reported that it should be an interconnected network with high porosity and the mean pore size, but more detailed requirements about the architecture and structural properties of scaffolds are needed. This chapter will provide such details. Another criterion that should be included in the design of scaffolds is the presence of nanometric HA crystals which provide a large surface area to react with host tissue. Finally, economic viability needs to be taken into consideration for the design of scaffolds; this is related to cost-effectiveness and the ability to scale-up. This is not an issue for natural scaffolds from trabecular bone due to the fact that it is a biowaste which can be used for medical applications if suitable treatments are applied to guarantee the harmlessness of the material. Tissue engineering has employed and developed a substantial amount of materials to fabricate scaffolds, for example, bioceramics (calcium phosphate, tricalcium phosphate, hydroxyapatite, bioglass, and demineralized bone particles), synthetic polymers (poly(α-hydroxy ester)s, polyanhydride, poly(vinyl alcohol), and biodegradable polyurethane, among others), and natural polymers (collagen, alginate, chitosan, and agarose, among others) (Khang, 2017). If the properties of a natural bone are not taken into account the scaffold will barely be successful. Mammalian bones are a composite system formed by an inorganic phase and an organic matrix. Two structures with different architectures are exhibited in the inorganic phase: trabecular and cortical bone. Trabecular bone is a percolated macrolattice of trabeculae formed by nonstoichiometric hydroxyapatite crystals called biohydroxyapatite (BIO-HA), which are a kind of carbonated calcium phosphate that contains many substitutional ions (London˜o-Restrepo et al., 2018). Meanwhile, the cortical bone is formed by a mesoporous matrix. Trabecular and cortical bones have preferential micro and structural orientations (Ramı´rez et al., 2007). In the case of surgery applications of cancellous bone, two different primary routes have been established. The first one uses cancellous bone from human and animal sources, and the second is related to the development of synthetic scaffolds with various levels of porosity, which attempt to have a similar architecture to natural 3D sponges. However, so far it is not possible to manipulate the preferred orientation and microarchitecture of biohydroxyapatite crystals to reproduce the mechanical and functional properties of bone. As it was previously mentioned, cancellous bones have been replaced by synthetic scaffolds. Several methods have been developed to fabricate hydroxyapatite scaffolds, such as water-soluble polymers, gel casting, polymer sponges, salt leaching, particle sintering, and 3D printing methods, among others (Ramay and Zhang, 2003; Chung et al., 2011; Wu et al., 2014; Stratton et al., 2016; Rabionet et al., 2018). Tissue engineering has established some requirements for scaffold

14.1 Introduction

design, such as biocompatibility, bioresorbability, mechanical properties, and porosity. Mechanical properties found in these materials are similar to those in natural cancellous bone. However, in studies of the mechanical properties of spongy bone and synthetic scaffolds, their complex 3D architecture has not been taken into consideration. Collagen/hydroxyapatite composite scaffolds were developed through methods such as microwave assisted cotitration and freezedrying techniques (Wang and Liu, 2014; Qiu et al., 2015; Kozlowska, 2018; He et al., 2018), and the final chemical composition of the scaffolds did not include other chemical elements, such as Mg, Al, Ba, Cu, Fe, K, Mn, Zn, and Na, which are present in bone (Giraldo-Betancur et al., 2013; London˜o-Restrepo et al., 2018). In all these studies, information about the three-dimensional microarchitecture or preferential crystalline orientation of the hydroxyapatite crystals of these scaffolds was not reported on. Lindahl (1976), studied the mechanical properties of trabecular bone obtained from human vertebrae and tibias. A deterioration of the mechanical properties of the bone as a function of age for men and women was found for the uniaxial force, which is parallel to the bone. However, a detailed study of the mechanical properties of these bones for other directions has not been reported. Raman spectroscopy is a powerful tool used to determine bone quality; with this technique, mineral/organic ratio, CO322/PO432 ratio, crystallinity, and collagen quality parameters can be determined (Morris and Mandair, 2011). Biochemical changes in trabecular bone, such as mineralization and carbonate accumulation in different regions of the femoral heads, have been studied using Raman microspectroscopy to determine the composition and structure of this kind of bone (Buchwald et al., 2012). After the mineralization process of bone, structural changes happened as a result of ionic exchange due to the chemical environment (Penel et al., 2005). It has been reported that some properties of bone, like ν1PO432/amide I ratio, depend on the orientation of the bone (Kozielski et al., 2011). However, information about the mechanical and structural properties of these bone sponges has not been related to results provided by Raman spectra. The chemical composition, density, and orientation of trabeculae, govern the mechanical properties of sponge. As such, mechanical properties change as a function of the position and direction of the sponge within the bone. Giesen et al. (2011) studied the mechanical properties for two loading directions: axial and transverse loadings. However, the transversal loading of a bone has two components, and only one of them was studied. According to their study, differences in mechanical properties as well as bone anisotropy can be considered mechanical adaptations due to loading during the use of the condyle. The mechanical properties of the spongy bone were dependent on the tested directions as will be shown in Section 13.3. Internal bone structure has more influence than chemical composition on the mechanical properties of sponge. Considering the aforementioned studies, the aim of this chapter is to provide a complete report of the physicochemical properties of natural sponge obtained from bovine trabecular bone. Some of the techniques that will be discussed in this

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chapter are: differential scanning calorimetric (DSC); inductively coupled plasma (ICP), Raman spectroscopy, X-ray diffraction, scanning electronic microscopy (SEM), transition electronic microscopy (TEM), and X-ray tomography.

14.2 BONE COMPOSITION Bone tissue is composed of mineral and organic phases and water. About 10% of bone tissue is comprised of water. The mineral phase (70%) is mainly composed of carbonated hydroxyapatite (80%), calcium carbonate (15%), other minority phases (5% dicalcium phosphate, calcium dibasic phosphate, and tricalcium phosphate, among others), and some substitutional ions, such as Mg, K, and N, among others (London˜o-Restrepo et al., 2018). Collagen type I (90%), lipids and noncollagen proteins are the main components of the organic phase (20%). Both phases provide strength and resilience to the bone. Hence, the skeleton can absorb impact without breaking. The minerals are not directly bound to the collagen, it is rather through the noncollagen proteins that active sites for mineralization and cell union are provided. The skeleton is the structural support that protects the body’s organs, and it is a subjection point for muscles. Large bones are associated with high movements, such as jumping and running. Large bones such as the femur are divided into three zones, namely: diaphysis, which is the central one; epiphysis, which is the external part of the bone; and finally, the metaphysis, which is located between the diaphysis and epiphysis. The most compacted zone of bone is called cortical bone, and spongy bone is called trabecular or cancellous bone; which is the focus this work. Hierarchical levels in bone are designed to perform many functions. In the first place, carbonated hydroxyapatite nanocrystals are aligned with collagen fibers which are organized in concentric parallel layers named lamellae that are located around the blood vessels forming osteons. The osteons are densely packed forming compact or cortical bone. Trabecular bone is built by a cancellous network with low bone density (Olszta et al., 2007). Trabecular bone has higher porosity and bigger pore size than cortical bone; the pore size for trabecular bone is in the order of microns to millimeters.

14.2.1 CORTICAL BONE Human cortical bones contain three major anatomical cavities: Haversian/ Volkmann canals, osteocyte lacunae, and canaliculi (see Fig. 14.1) (Wang and Ni, 2003). Haversian and Volkmann canals are microscopic tubes in cortical bone where blood vessels and nerves travel through the bone along with many other substances, such as sugar, amino acids, fatty acids, inorganic compounds, hormones, neurotransmitters, and coenzymes. Porosity due to the Haversian and

14.2 Bone Composition

FIGURE 14.1 SEM image of anatomical bone cavities in bovine cortical bone (femur).

Volkmann canals is known as vascular porosity (VP) and it is the largest porosity in cortical bone (Cardoso et al., 2013). These canals are major contributors to total porosity. The Haversian and Volkmann canals communicate with bone cells through connections called canaliculi. Osteocytes are the most abundant kind of cells within bones, which are located in the osteocyte lacunae, and from them leave canaliculi in all directions. These canaliculi allow for the communication between the lacunae and the central conducts (Haversian canals). It is important to recall that this branched system provides many pathways for the diffusion of

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oxygen, blood nutrients, and other aforementioned elements into the extracellular fluid to the osteocytes, and it also allows for the expelling of cellular waste through the blood vessels. At a hierarchical level, cortical bone has three levels of porosity. The first one is due to the Haversian and Volkmann canals or vascular porosity as previously mentioned. The second porosity is due to the osteocyte lacunae and canaliculi known as “lacunar-canalicular” porosity (LCP). Finally, the third porosity is known as collagen-apatite porosity (CAP) and is the smallest porosity in cortical bone (Cardoso et al., 2013). Fig. 14.1 shows the anatomical bone cavities previously mentioned. Bone surface is dense, but has microholes of about 30 μm in diameter; these holes are Haversian canals which are the central ducts of bone (see Fig. 14.1A). A longitudinal view of a bone is shown, where it is possible to identify Haversian canals, lacunae, and canaliculi. It is also possible to see a rough surface composed of collagen nanofibers and hydroxyapatite nanocrystals (see Fig. 14.1B). Inside the Haversian canals, it is possible to see collagen fibers and a lot of canaliculi connecting the Haversian canal with the lacunae (see Fig. 14.1C). As such, it is possible that the hierarchical structure of bone influences its mechanical and structural properties.

14.2.2 TRABECULAR BONE In large bones, like the femur, the trabecular bone is found in the epiphysis and the metaphysis. The femur head distributes mechanical load to the whole bone, in particular to the cortical region. The mechanical properties of bone depend on the macroarchitecture and crystalline orientation of the trabeculae. It is well known that spongy bone has higher metabolic activity than cortical bone, which is one of the reasons to study it in detail. Mechanical resistance is an important parameter to consider for tissue engineering and for clinical applications because this parameter is directly related to the risk of fracture. However, the 3D structure of natural sponge has not been considered yet.

14.2.3 BONE POROSITY Wang and Ni (2003) studied human cortical bone porosity and pore size distribution using a low field pulsed NMR approach with human cadaver femurs from donors ranging from 16 to 89 years old. They reported that pore sizes in bone may be divided into three groups with porosity ranges of 1.3% 3%, 4.1% 25.5%, and 0.7% 9.8%, respectively. They associated the first group with lacunae and the second one with Haversian canals, but they did not associate the third group with anything. According to Rodrı´guez et al. (2011), porosity increases with age, which is about 8% for young people and 24% 28% for elderly individuals. Haversian canals tend to increase with age while lacunae porosity decreases. Pore size diameters for Haversian canals vary from 27.1 to 77.5 μm and for lacunae from 3.07 to 4.45 μm. Bone has another porosity

14.2 Bone Composition

Table 14.1 Porosities Reported for Bones Porous Size (μm)

Porosity (%)

Source

Reference

55 70

Trabecular bone

Heinl et al. (2008) Ural et al. (2007)

50 200 34 31 1

3.07 4.45 (lacunae) 27.1 77.5 (Haversian canal)

1.8 6.8 About 70 1.3 3

Wachter et al. (2001) Bae et al. (2012) Ooi et al. (2007) Wang and Ni (2003)

4.1 25.5

4.8 (average) 1.5 50 100 nm 1 nm 1.5 0.7 17.2

Haversian canal diameter; human bone Fifth percentile of Haversian canal diameters Haversian canal diameter; human cortical bone from the lateral diaphysis Human cortical bone Bovine cancellous bone Mid-diaphysis of human femurs: 16 89 years old.

Cortical bone from rat tibias Lacunar porosity of rat tibias Vascular porosity Lacunar-canalicular porosity Collagen-apatite porosity Rat cortical femur; lacunar porosity

2.76

Mouse femoral diaphysis; canalicular porosity Rat tibiae

54.6 (average)

Human lumbar spines; average age: 63 years old

PalacioMancheno et al. (2014) Cardoso et al. (2013) Tommasini et al. (2012) Schneider et al. (2011) Britz et al. (2010) Rodríguez et al. (2011)

associated with the trabeculae zone that will be shown in the morphological characterization section. Table 14.1 shows porosities and pore sizes reported in the literature for cortical and trabecular bone; these values can be used as a reference in the case of scaffold engineering.

14.2.4 HYDROXYAPATITE Hydroxyapatite (HA) is a mineral that belongs to the apatites family which is a group composed of phosphates, arsenates, and vanadates with hexagonal and pseudohexagonal monoclinic structures. “Apatite” is the name of chlorapatite, fluorapatite, and hydroxyapatite, and they exhibit the general

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formula: A5(BO4)3(OH, F, Cl), where A represents metallic cations, such as calcium (Ca), barium (Ba), sodium (Na), and lead (Pb), among others. While B represents phosphorous (P), vanadium (V), or arsenic (As) (Go´mez Ortega et al., 2004). Stoichiometric hydroxyapatite has the chemical formula: Ca10(PO4)6(OH)2 or Ca5(PO4)3OH, because in fact, the base of the structure is Z 5 2 with a Ca/P rate of 1.667; while hydroxyapatite from mammalian bones, called “biohydroxyapatite” (BIO-HA), is nonstoichiometric (Giraldo-Betancur et al., 2013) and its Ca/P ratio varies from 1.8 to 3. At a crystallographic level, HA has a compact hexagonal packing with oxygen atoms at tetrahedral and octahedral holes with a spatial group P63/m and lattice parameters ˚ , c 5 6.88 A ˚ , α 5 β 5 90 degrees, and γ 5 120 degrees of a 5 b 5 9.37 A (Rujitanapanich et al., 2014). The chemical formula for stoichiometric hydroxyapatite is Ca10(PO4)6(OH)2, but to differentiate the structural component of the HA structure, it is important to rewrite this chemical formula as Ca(I)4Ca(II)6[PO(I)O(II)O(III)2]6(O(IV)H)2. Regarding the structure of hydroxyapatite, there are two calcium types as a function of the chemical environment. The notation Ca(I) and Ca(II) indicates Ca atoms type 1 and type 2, respectively. This division in two calcium atoms is due to each one being surrounded by a different number of oxygen atoms in the apatite structure (Fig. 14.2). Ca(I) has a 9-coordination number, where all sites are occupied by oxygen atoms, which results in a tricapped trigonal prism structure. Ca(II) has a 7-coordination number and represents a pentagonal bipyramid (Campa Molina et al., 2007). On the other hand, P atoms are surrounded by four oxygens, and they form a tetrahedron structure that is the characteristic structure of the phosphate group (PO432).

FIGURE 14.2 Crystalline structure of hydroxyapatite.

14.2 Bone Composition

14.2.5 BIOHYDROXYAPATITE Biohydroxyapatite (BIO-HA) is a nonstoichiometric apatite due to its numerous ionic substitutions that are obtained from natural sources. Carbonate groups and halogens can replace the phosphate and hydroxyl groups, respectively. The ionic character of hydroxyapatite turns it into a hard, refractory ceramic with a fusion point above 1500 C; this also allows for partial or complete ionic substitution in the lattice with other ions of similar atomic ratio. BIO-HA has many medical applications due to its physicochemical properties being similar to human hydroxyapatite, but its applications are limited to direct replacement because it has low mechanical strength. Its main uses in this field are as a filling material, for coating metal prosthesis, and maxillofacial surgery, among others, where mechanical strength is not required. This ceramic makes bonds with surrounding bone tissue to promote material integration and the growth of new bone tissue. Nanometric hydroxyapatite (10 100 nm) has better functional properties than HA with micrometric sizes ( . 1 μm). Its superficial reactivity and ultra-fine structure, for example, help in tissue-graft interaction when this material is implanted. In addition, nanometric HA promotes more adherence, mineral deposition, osseointegration, osteoblast differentiation, and proliferation in comparison to micrometric HA (Nasiri-Tabrizi et al., 2014). Pore sizes of at least 50 μm in diameter are required for the elaboration of replacement parts to guarantee fibrovascular tissue penetration, and for pores to be interconnected (Thuault et al., 2014). Bone powder is composed of HA nanometric polycrystals infused within organic matrices (fat and protein, see Fig. 14.3A and B). Lipids look like an amorphous coating in the material surface, while protein can be seen as oriented nanofibrils. When bone is raw, the organic matrix does not allow for the viewing of HA nanocrystals. After a cleaning process, like the hydrothermal treatment that involves pressure and temperature reported by London˜o-Restrepo et al. (2016), it is possible to see the HA crystals in the surface (see Fig. 14.3C and D), but the morphology of the HA crystals in images taken at 10,000 3 is not well defined yet. In this way, high-resolution microscopy is needed. When bone powder, after a hydrothermal treatment, is calcinated at 600 C for 50 hours, fat and protein are entirely removed, and in this manner, it is possible to see micrometric HA crystals with semispherical morphologies (see Fig. 14.3E and F). In natural bones, nanopolycrystalline HA looks like small clusters that have preferential orientation. It is well known that for temperatures above 750 C, HA polycrystals can be transformed into single crystals through a coalescence phenomenon (London˜oRestrepo et al., 2018). HA single crystals have hexagonal and pseudohexagonal monoclinic rod-like shapes with micrometric sizes. These rods are obtained in order to understand the structural properties of the material, but it is well known that for medical applications, it is much better to use nanopolycrystalline HA due to its composition, large reaction surface, and semispherical morphology. This last characteristic is important because rod tips tend to damage bone cells.

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FIGURE 14.3 Bovine bone powders: (A) 5000 3 , (B) 10,000 3 . Bovine bone powders with hydrothermal process: (C) 5000 3 , (D) 10,000 3 . Hydroxyapatite from bovine bone: (E) 10,000 3 , (F) 25,000 3 .

The physicochemical properties and extraction efficiency of BIO-HA depend on the extraction technique used, heating rate, annealing temperature, sintering time, cooling rate, and bone nature (Ramirez-Gutierrez et al., 2016). In general, apatite crystals that have been produced by biological systems exhibit crystal sizes less than that of synthesized HA crystals. These characteristics can be readily observed in trabecular bone, which consist of an anisotropic percolated macrolattice composed of HA nanocrystals as will be seen in the morphological characterization section 14.3.2. In an attempt to emulate cancellous bone, tissue engineering has developed scaffolds from materials like ceramics and polymers. These scaffolds are isotropic macrolattices without preferential structural orientation and microporosity. Some of the properties of bone have been reported as necessary for an ideal scaffold, such as biocompatibility, bioresorbability, mechanical properties, and pore size (Bose et al., 2012), but trabecular bone has other forgotten physicochemical properties, such as porosity and interconnectivity, which are necessary too.

14.2 Bone Composition

FIGURE 14.4 Hydroxyapatite monocrystals from bovine bone, obtained after calcination at 1000 C.

Fig. 14.4 shows the characteristic morphology of BIO-HA single crystals obtained after calcination of bovine bone at 1000 C. Quasi-hexagonal structures that have grown due to heating treatment are presented in this image. As was previously mentioned, BIO-HA is polycrystalline with nanometric sizes that form clusters in bone. This kind of single HA crystal does not appear in any bone.

14.2.5.1 Structural properties of BIO-HA X-ray diffraction is an excellent technique for identifying crystalline structures in any sample and specifically the crystalline percent of a sample when it has micrometric sizes. Biohydroxyapatite obtained through the calcination of bovine cortical bone at 1000 C (BIO-HA-1000) was achieved using the cortical bone of the same femur, and it was used as a reference for X-ray and Raman experiments (Ramirez-Gutierrez et al., 2016). Fig. 14.5A shows the X-ray diffraction pattern of this sample; the red lines correspond to the identification of HA using the JCPDF card No. 72 1243. These narrow peaks are characteristic of BIO-HA microcrystals. Fig. 14.5B shows a SEM image of BIO-HA single crystals where the presence of hexagonal structures and another minority spherical phase, which has been identified as magnesium oxide (MgO), are clear (Ramirez-Gutierrez et al., 2017). It is important to point out the differences between crystalline percentage and crystalline quality. Crystalline percentage is defined as:

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FIGURE 14.5 (A) X-ray diffraction pattern of BIO-HA single crystals. (B) BIO-HA single crystals.

%C 5

CA TA 2 N

(14.1)

where %C is crystalline percentage, CA is the crystalline area under the pattern, N is the characteristic electronic noise of the system, and TA is the total area. Here, it is important to recall that in order to determine this parameter, it is necessary to take into account the complete diffractogram. Giraldo-Betancur et al. (2013) studied the crystalline percentage of commercial BIO-HA used as medical implants, and this parameter changed from 60% to 80%. Crystalline quality (CQ) is calculated by studying the full width at haft maximum (FWHM) of any characteristic peak. The most intense peak is usually used

14.2 Bone Composition

for this calculation, but any peak can be characterized. It is necessary to obtain the FWHM (assuming Gaussian or other behaviors) in order to determine the crystalline quality of a sample; if the FWHM increases, it is an indication that the CQ decreases and vice versa. The CQ of bone can be obtained using the inverse of the FWHM value. Fig. 14.6 shows a characteristic X-ray diffraction pattern of raw powder bovine cortical bone. In the case of microcrystals, the crystalline area is taken from the base of the green line and the blue line that covers all diffracted peaks. The area below the green line corresponds to the background where there are contributions from the electronic noise of the system and the amorphous region of the material. However, this is not the case with HA from mammalian bones, which is nanocrystalline. The interphase between the nanocrystals acts like a diffraction grating when the radiation interacts with the material. This phenomenon produces patterns with broad peaks that are, in fact, curves instead of patterns. Ramirez-Gutierrez et al. (2016) studied the effect of the cooling rate on the structural properties of bovine cortical bone incinerated at 940 C and cooled in liquid nitrogen, water, furnace air, and air. Fig. 14.7A shows the [211] peak for all samples, indicating that the sample cooled in the furnace produced the BIO-HA with the best crystalline quality. Fig. 14.7B and C show the procedure used to calculate the crystallinity quality using the [211] diffracted peak; Fig. 14.7B corresponds to the changes in the FWHM value of each sample, and Fig. 14.7C shows the inverse of this parameter. This calculation was possible because the HA crystals were grown until they reached micrometric sizes using a calcination process.

FIGURE 14.6 X-ray diffraction pattern of defatted bovine bone powder.

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FIGURE 14.7 Diffraction peak [211] for hydroxyapatites, (A) FWHM of hydroxyapatite samples. (B) The CQ obtained for [211] peak.

The phase identification was obtained using the Powder Diffraction File (PDF) that is part of the any X-ray diffraction system library. In Appendix A, the PDF 09 0432 of hydroxyapatite is shown; other PDFs can be used for the identification of HAP, such as PDF 72 1243. These PDFs correspond to synthetic HA, and then the peak positions of BIO-HA indicate a shift due to the inclusion of other ions, such as Mg, N, K, among others, which distort the lattice.

14.2.5.2 Mineral composition of BIO-HA The ion traces of the following samples were quantified through inductively coupled plasma atomic emission spectroscopy (ICP-AES): bovine bone powder (BBP), BBP with a hydrothermal process (BBP-HTP), BIO-HA from BBP-HTP calcinated at 600 C, 800 C, and 1100 C (Bio-HA600, Bio-HA800, Bio-HA1100). Table 14.2 shows the mineral traces of the samples in parts per million; the major elemental components are sodium and magnesium.

Table 14.2 Mineral Composition of Bovine Bone, Hydroxyapatites, and Commercial Samples Mineral Content (ppm)

Sample Na

Mg

Al

K

Mn

Fe

Ni

Cu

Zn

Ba

Bovine bone powdera BBP-HTPa Bio-HA600a Bio-HA800a BioHA1100a Defatted Bio-HAb Bio-HA alkalineb Bio-HA calcinedb NISTb

6689.23

3048.49

33.98

414.89

—

9.89

—

—

90.17

—

5463.60 7493.14 8546.57 8111.94

2720.90 3818.75 4307.55 3936.94

21.73 32.27 44.81 28.35

147.45 211.08 225.00 169.64

— — — —

16.63 17.01 21.51 14.44

— — — —

— — — —

78.54 130.70 92.06 81.88

— — — —

—

—

6.6 6 0.21

345.27 6 5.06

—

31.36 6 0.1

—

1.07 6 0.76

79.46 6 0.98

294.08 6 4.74

—

—

6.66 6 0.32

—

—

83.33 6 14.5

—

5.22 6 2.01

135.81 6 0.25

241.99 6 1.56

—

—

20.01 6 1.27

228.14 6 6.84

1.25 6 0.04

42.71 6 0.45

—

9.05 6 0.22

139.38 6 6.05

265.58 6 13.15

—

—

—

17.44 6 0.11

645.45 6 17.18

5.8 6 0.54

2.7 6 0.01

180.87 6 3.38

240.76 6 5.16

Apafill Gb Biograft b SigmAldrichb Coralina b

— — —

— — —

270.53 6 4.45 103 6 2.9 6.62 6 0.92 5.72 6 0.6

— — —

59.73 6 1.97 — —

137.17 6 4.86 16.69 6 1.42 11.57 6 1.45

— — —

15.33 6 0.46 — —

24.88 6 1.16 119.05 6 2.64 3.66 6 0.08

150.3 6 5.02 1.62 6 0.07 —

—

—

14.29 6 0.28

—

—

25.98 6 0.33

—

3.98 6 0.88

6.94 6 0.26

8.49 6 0.11

a

This work. Giraldo-Betancur et al. (2013).

b

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Because the presence of these traces enhances bone regeneration, it is important to monitor their presence in the material. The small variations in mineral content indicate that these ions belong to the inorganic phase, and little losses can be attributed to the fusion temperature of some of them. Iron levels are low in comparison to the results reported in the literature; these low iron levels may be due to an adequate milling process without contaminants. The loss in mineral content after the hydrothermal process is due to the partial removal of the organic phase, which indicates that some of the found ions are not in the HA structure. Bone calcination at elevated temperatures does not produce a significant loss of substitutional ions. However, this lack of loss should not imply that this kind of process provides a suitable material for tissue engineering. HA obtained by calcination at adequate temperatures guarantee the absence of any organic compounds, but its morphology, porosity, and surface area are not adequate for medical purposes. Synthetic scaffolds do not have these ions in their composition; they are quite dissimilar from HA found in biological sources. In this section, it is worth noting that the chemical composition of bone can vary due to diet, gender, and age of the animal, the kind of bone, and the location of the studied bone, among other factors. Therefore, certain changes in mineral content values are not due to the conditioning process (cleaning and calcination), but to the heterogeneous content of the sample. In the case of HA from Sigma, it only has aluminum, iron, and zinc, which indicates that this HA was obtained by means of a chemical process (Giraldo-Betancur et al., 2013).

14.2.5.3 Thermal properties of BIO-HA The thermal properties of bone (cortical and trabecular) can be studied by means of differential scanning calorimetry (DSC) and thermogravimetric analysis (TGA). The objective of these techniques is to determine the thermal changes in bone when it is under the influence of a temperature program. Such investigations help to explain how HA is obtained from biowaste, which implies incineration at high temperatures. The aim of calcination should be the complete removal of the organic phase but without any modification of the physicochemical properties of the mineral phase. Thermal changes depend on different factors, such as the heating rate applied, which in the case of bone has been reported from 1 C to 50 C/min; the atmosphere where the sample was calcined; and the nature of the sample. It is well known that thermal changes during a thermal treatment are due to fat and protein (organic phase) degradation at temperatures below 650 C, while for temperatures above to this value, BIO-HA turns into other crystalline phases through a dehydroxylation process (London˜o-Restrepo et al., 2018). The thermogravimetric curve for DSC and its second derivative as a function of the temperature were obtained for the bovine bone powder sample using TG Q500 equipment (TA Instruments, USA). A 12.0 6 1.0 mg sample was placed in a platinum crucible (TA Instruments, USA). The sample was heated beginning at room temperature and up to 1200 C at a heating rate of 10 C/min; the measure

14.2 Bone Composition

FIGURE 14.8 DSC curve of bovine bone powder.

was carried out in a constant N2 flow. The TG data were processed with the use of a Universal Analysis 2000 TA-software. Fig. 14.8 shows the bone thermal degradation process of a sample without any defatting or deproteinizing process heated at 5 C/min, and its second derivative as a function of the temperature. The first part of the thermogram corresponds to the degradation of the organic matrix from room temperature until about 650 C, and the second part is associated directly with the inorganic phase (650 C 1100 C). According to the literature, some of these peaks can be identified as: Changes below 100 C correspond to dehydration of the sample (free water), while water is lost at 225 C (Heredia et al., 2013; Sofronia et al., 2014; London˜o-Restrepo et al., 2018). Changes from 220 C to 570 C correspond to degradation and collagen combustion processes (degradation and protein denaturalization or debranched). It is possible that the peak located at 770 C (I) is related to an endothermic process that takes place as a result of the recrystallization process of the HA polycrystals (London˜o-Restrepo et al., 2018). Decarboxylation is attributed to the peak at 870 C, and the dehydroxylation that turns HA into oxyapatite (OHA) is assigned to peaks at 900 C and 1120 C (Yoganand et al., 2011; London˜o-Restrepo et al., 2018). A more complete assignation of the peaks is shown in Table 14.3.

14.2.5.4 Methods to obtain HA and BIO-HA There are many ways to obtain hydroxyapatite; it can be synthesized through dry methods (solid-state and mechanochemical methods), wet methods (emulsion, chemical precipitation, hydrolysis, sol-gel, hydrothermal, and sonochemical

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Table 14.3 Thermal Events During Bone Calcination Temperature ( C)

Thermal Event

Reference

B200 290

Collagen denaturation and decomposition

622 800 1000 345, 425 (exothermic peaks) 900 1000

Dissociation of calcium carbonate Partial dehydration of hydroxyapatite Collagen degradation and combustion

Trębacz and Wójtowicz (2005) Kusrini and Sontang (2012)

217 665

Liberation of chemically bonded water, decomposition of the organic matrix and carbonates Decarbonization

835 (700 1100) 988 1050 1100 25 120 200 250 400 330 420 550 700 876 and 949 1041

Removal of remaining organic components

Decarboxylation Decomposition of oxyapatite to calcium phosphate Decomposition HAp-βTCP-αTCP Adsorbed water liberation Lattice water liberation HPO422 decomposition Phospholipids degradation Protein denaturalization Degradation of the primary protein structure Whitlockite formation Ca10(PO4)6OH - Ca10(PO4)6O Ca10(PO4)6O - Ca4P2O9 1 2Ca3(PO4)2

Heredia et al. (2013) Niakan et al. (2015) Sofronia et al. (2014)

LondoñoRestrepo et al. (2018)

methods), and by high-temperature processes (combustion, pyrolysis, or obtaining from biogenic sources) (Sadat-Shojai et al., 2013; Thuault et al., 2014; Me´ndezLozano et al., 2017). Only when HA is obtained from biowaste, like mammalian bones, is it named BIO-HA. Some properties of synthetic HA include stoichiometry, high purity, homogeneous composition, and nanometric sizes. Synthetic HA is ideal for applications related to adsorbents, sensors, catalysis, and chromatography for protein extraction and purification (Go´mez Ortega et al., 2004). Dry methods do not use liquid solvent, and they are used for mass production; large hydroxyapatite with an irregular shape is obtained as a result of these methods, while wet methods produce nanometric hydroxyapatite with regular morphologies. The formation of secondary calcium phosphate phases is presented as a disadvantage of the HA obtained through wet methods, but the mineral phase of mammalian bones has, in fact, these other phases (London˜o-Restrepo et al., 2018).

14.2 Bone Composition

Hydroxyapatite from mammalian bones is like human HA, which is not stoichiometric due to the presence of substitutional ions, like Na1, K1, Mg21, Ba21, F2, Cl2, that are bonded to the crystal lattice. BIO-HA has a deficient calcium content due to substitution by the carbonated group, sodium, potassium, magnesium, zinc, or calcium. It has been reported that processes that involve high temperatures are suitable for producing high purity homogeneous HA in a one step process, however, HA loses porosity and surface area after calcination at 700 C. It is worth noting that processes at high temperatures not only consist of bone calcination, they also involve the synthesis of HA from natural precursors such as algae, corals, and eggshell (Akram et al., 2014).

14.2.6 COLLAGEN Collagen has a supramolecular structure with periodicity and hierarchical levels. The primary structure of collagen is composed of a typical amino acid sequence that determines the collagen type; there are 21 different types of collagens with at least 42 polypeptide chains. Collagen type I is the main protein in bone and it is composed of three left α-helices with a length of about 300 nm and variable diameter from 4 to 20 nm. Each chain has about 1050 amino acid. Two of the chains are named α(1) because they have the same amino acid sequence and the other one is called α(2). The general sequence for collagen chains is (Gly-XY)n , where X is proline and Y is hydroxyproline (Olszta et al., 2007). Tropocollagen is aligned to form fibrils of about 200 nm, but these fibrils are not continuous. Collagen fibrils have gaps with lengths of about 40 nm where hydroxyapatite crystals are located; because there is a low concentration of proline and hydroxyproline, these zones have high structural flexibility. The high thermal stability of collagen is due to its structural arrangement and chemical bonds; in triple helix chains, there are hydrogen bonds between amino and carboxyl groups of amino acid chains. The main function of collagen is to give flexibility, elasticity, and tensile strength longitudinally. Biomechanical properties that collagen gives to bone depend on concentration, spatial orientation, and bond stability. Collagen incorporated into grafts can produce an immunologic response, and for this reason, its presence in implants should be avoided. Other components of organic phase are noncollagen proteins, such as osteocalcin (BPG), osteonectin (ON), osteopontin (OPN), thrombospondin (THBS1), bone morphogenetic proteins (BMPs), and bone sialoproteins (BSP), polysaccharides, lipids, cytokines, and primary bone cells: osteoprogenitors, osteoblasts, osteocytes, and osteoclasts (Campa Molina et al., 2007). Osteoprogenitor cells promote the formation of osteoblasts, these cells are responsible for the creation of new bone through collagen secretion that is thereafter coated with noncollagen proteins. Noncollagen proteins can retain minerals (mainly calcium and phosphate), and these aid blood flow. Osteocytes are necessary cells for maintaining the biomechanical properties of bone tissue; these cells are osteoblasts that have

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been trapped in the matrix during the mineralization process, and they take on a starry appearance. Finally, osteoclasts secrete substances that dissolve the inorganic bone phases for bone resorption. Phosphate and calcium ions are deposited into the collagen matrix to form an amorphous calcium phosphate that is subsequently turned into hydroxyapatite. As previously mentioned, the mineral phase has citrate, magnesium, and other minority ions that contribute to the mineral structure being more stable. These traces contribute to the life cycle of hard tissue because they influence some biochemical reactions related to bone metabolism (Akram et al., 2014). The mechanism of bone formation is still unclear, and the knowledge about it is limited. Minerals in bone provide rigidness and hardness due to packed crystals within the collagen gaps; hence, these properties depend on mineral quantity, packaging level, and the crystal ordering around the collagen fibers.

14.2.7 OSTEOCALCIN Osteocalcin (BGP: bone Gla protein) is a nonphosphorylated glycoprotein; rather, it is the main noncollagen protein in bone. BGP helps capture osteoblasts and osteoclast cells which play an essential role in bone formation and resorption during a specimen’s life, respectively. This protein also influences bone mineralization due to its high affinity with hydroxyapatite, but the reaction mechanism has not been clarified yet. BGP has negative charges on its surface, which coordinate five calcium ions in a particular configuration that is complementary to calcium ions in hydroxyapatite (Hoang et al., 2003). Human BGP has 49 amino acids, but it is not constant for diverse species: there is a range of 46 50 amino acids depending on the species (γ-carboxyglutamic acid (Gla) is in the amino acid sequence). According to the bone sample and its age, its composition can be made up of up to 20% noncollagen proteins in adult bones and about 2% total bone proteins. There is a 90% similarity between human and bovine osteocalcin for amino acid sequences, especially in the central zone of the protein where there is a high concentration of Gla residues. BGP exhibits a folded structure due to disulfide bridges of cysteine (Cys), that are present in BGP residues (Diaz, 1996). Through X-ray diffraction and nuclear magnetic resonance (NMR), it has been determined that BGP is a globular protein made up of three alpha-helices, a hydrophobic core, an N-terminal as β-sheet, and a C-terminal as β-sheet exposed. The amino acid sequence is NH2-Tyr-Leu-Tyr-Gln-Trp-Ley-Gly-Ala-Pro-Tyr-Pro-Asp-Pro-LeuGla-Pro-Arg-Arg-Gla-Val-Cys-Gla-Leu-Asn-Pro-Asp-Cys-Asp-Glu-Leu-Ala-AspHis-Ile-Gly-Phe-Gln-Glu-Ala-Tyr-Arg-Arg-Phe-Tyr-Gly-Pro-Val-COOH. The position of three Gla residues are the same for both sources: 17, 21, and 24.

14.2.8 WATER Water dwells within vascular channels that flow into collagen and mineral matrices. Due to the polarity of water, it can interact with both phases, then, water is

14.3 Study of Spongy Bone

bonded with the hydrophilic groups of collagen (glycine, hydroxyproline, carboxyl, and hydroxylysine) and with charged groups, such as phosphate and calcium ions. In addition, water is present as surface bonded water and lattice water (Behari, 2009). Surface bonded water is lost at low temperatures during a calcination process (,200 C) as was seen in the thermal degradation section, but lattice water which is bonded to the structure, requires a higher temperature to break the bonds (200 C 400 C). There is free water that fills the Haversian canals and the lacuno-canalicular system (see Fig. 14.1). This water provides viscoelastic properties to the bone, and is the cause of nutrient diffusion (Wilson et al., 2006). During mineralization and demineralization processes, ions need to be transported to and from osteoid sites. Therefore, this water serves as a means of transport (Behari, 2009). Wilson et al. (2006) reported water in two different crystal lattice environments in bone: lattice water and surface water. Lattice water is bonded to the crystal in vacant sites of carbonated hydroxyapatite, which provides structural stability through hydrogen bond formation between ions. The collagen and the mineral phase can be mechanically coupled due to the presence of the surface water; this water type serves as a cushion against mechanical stress. Surface water has a dual-action mechanism. First, the movement of the water provides a bone support against mechanical stress with less deformation, and second, this water protects collagen from shear under uniaxial stress by serving as a sacrificial layer.

14.2.9 FAT Lipids (2% 4%) in the organic phase are essential for bone metabolism and mineralization. The main lipid source are cells because they have a lipid membrane that regulates the input and output flux of components to the cells. There are lipids in the extracellular matrix—associated with collagen as extracellular matrix vesicles (ECMVs) where bone calcification actively occurs (Boskey, 2004). In particular, complex phospholipid acids have been identified in actively mineralizing tissues. These phospholipids are present in small quantities in mineralized tissue, and they can nucleate mineral formation and regulate the growth of mineral crystals. Also, there are other lipids in bone that are important for bone cell function: leptins can be essential for osteoblast function stimulation, which causes the formation of bone as well as osteoclast activity suppression or bone resorbing (Boskey, 2004). Regarding tissue engineering, it is important to remove this component from the bone to avoid interference with other organic and inorganic compounds of the bone.

14.3 STUDY OF SPONGY BONE The second part of this work is dedicated to the study of natural spongy bone obtained from bovine femur samples, which are fundamental for tissue

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engineering in order to understand the 3D structure of the sponge that affects the structural, morphological, and mechanical properties of this bone. The understanding of this system is still an unresolved problem for tissue engineering.

14.3.1 COLLECTION AND PREPARATION OF SAMPLES Spongy bone samples are usually obtained from biogenic sources, such as bovine and porcine. Because the femoral head region has a huge volume and is rich in spongy bone, it is commonly selected for sampling. Samples can be easily collected from slaughterhouse biowaste. In the case of bovine bone, the slaughter age is about 2 years old. This means that the bone has reached high mineral density and a determined chemical composition comparable to other vertebrate animals in the same age range. Spongy bones are nonisotropic systems, which means that their properties depend on the analyzed direction, especially the mechanical properties. To study spongy bones, it is recommended to differentiate the internal part of the femur head by assigning three different directions to the femur: X, Y, and Z, as shown in Fig. 14.9A, and it is stablished a coordinate system. Keeping this sign and direction convention, samples can be obtained with several geometries, such as cubes or cylinders, in order to study the mechanical properties of a bone as a function of direction. Another factor in studies of trabecular bone is the zone in the trabecular slice from which the cube is taken. With such samples, it is crucial to identify the zone where the cube was taken (Goldstein, 1987). The central zone of the slices is less dense than a section obtained from the boundary of the bone (Fig. 14.9B). Samples obtained from biowaste must be cleaned through the removal of soft tissue with surgical instruments, like a dissection cutter; this task can be facilitated by boiling the bone in distilled water. After that, a defatted process is necessary, and several routes have been followed. Some common chemical solvents used to defat bones are petroleum ether, acetone, xylol, hydrogen peroxide,

FIGURE 14.9 (A) Femur head; (B) Transversal slice of the femur head.

14.3 Study of Spongy Bone

sodium hydroxide, ethanol, and dihydrogen phosphate, among others (Lindahl, 1976; Vastel et al., 2007; Giraldo-Betancur et al., 2013; London˜o-Restrepo et al., 2016). Soxhlet extraction using petroleum ether at 30 C and alkaline treatment employing sodium hydroxide are widely reported processes for removing fat (Giraldo-Betancur et al., 2013), but these methods have not been entirely successful. Hydrothermal treatment using an autoclave at 1.2 atm for 1 hour has also been reported; sample-distilled water 1:5 w/v has been used, and this process must be carried out almost three times in order to remove all fat (London˜oRestrepo et al., 2018). Ultrasonic cleaning in acetone has been employed to clean bone, but this process causes internal damage to the structure of trabecular bone, which affects the mechanical properties of the specimen. Deproteinization is another important process to remove the organic phase from bone. Sodium hypochlorite, hydrogen peroxide, sodium hydroxide, and potassium hydroxide have been used to clean bone. All these processes are necessary to avoid rejection by host tissue.

14.3.2 MORPHOLOGICAL CHARACTERIZATION SEM is a powerful tool that can be used to characterize bone surface as well as to detect the presence of hydroxyapatite nanocrystals, but for this last task, a highresolution SEM along with a suitable cleaning process is necessary to visualize the crystals without gold sputtering. Morphological characterization provides significant information about porosity, crystal shape, and porous size. These three parameters are closely related to the success of grafts because all of them are necessary for biological activity. Porosity and porous size are correlated with blood irrigation and the presence of vessels through new bone created from osteoblast grafts. Meanwhile, crystal shape is related to possible damages to bone cells; it has been reported that needle-type crystal morphologies can destroy the cell membrane of bone cells and kill them. The morphology of hydroxyapatite nanocrystals in a bone can be influenced by factors such as the source of the sample, bone type, age, and position in the bone where the sample was taken. Images obtained by SEM have highlighted the interconnected character of trabecular bone as well as its tridimensional arrangement, which is nonisotropic. Macro- and microarchitectures can be studied in trabecular bone; macroarchitecture corresponds to trabeculae as a whole, while microarchitecture refers to the morphology of a trabecula. Due to anisotropic changes in macroarchitecture that result from the distinct functions of bones within the body, morphological characterization must consider the three aforementioned directions. Fig. 14.10 shows how trabecular bone has a complex three-dimensional structure with different arrangements for each X, Y, and Z direction, respectively. Longitudinal and transverse channels that percolate through the bone, with dimensions in the order of millimeters, can be observed in this figure. These pores provide the primary porosity of trabecular bone, which allows for blood irrigation. In the surface of these trabeculae, there are many

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FIGURE 14.10 Bone microarchitecture; primary porosity.

FIGURE 14.11 Bone microarchitecture; secondary and tertiary porosity.

microholes where blood vessels run through the bone (Ural and Vashishth, 2007). While osteogenesis benefits from microporosity, which provides a large superficial area allowing for osteoblast to attach, macroporosity enhances bone growth (Woodard et al., 2007). The secondary porosity of trabecular bone is related to lacunar cavities where osteocytes are located (see Fig. 14.11A). A more detailed analysis of the trabecular bone reveals that it has a tertiary porosity due to the presence of canaliculi and the space between HA nanocrystals (Fig. 14.11B). Osteocytes are responsible for the interchange of nutrients and waste through canaliculi. It is worth noting that trabecular bone pores have a characteristic spatial arrangement with a 60 degrees slope in the plane perpendicular to the Z direction where the charge is mainly supported. This fact can be attributed to the functional

14.3 Study of Spongy Bone

properties of bone as a mechanism to successfully support the mass and weight of the body. To design scaffolds, these complex architectures and porosities must be considered. Regarding microarchitecture, trabeculae have a narrow central zone with diameters in the range of 150 200 μm (Fig. 14.11C). Furthermore, trabeculae have a preferential growing habit as lamellae with about 20 μm in diameter (see Fig. 14.11D).

14.3.3 X-RAY TOMOGRAPHY Sections of bone and a smaller cubical piece extracted from it were scanned with a Nikon XT H 225ST industrial micro-CT system using a tungsten reflection target. With the improvements of scan resolution, X-ray CT is now a powerful nondestructive technique capable of resolving microestructures, such as the channel networks inside a trabecular bone. A schematic diagram for the CT process is shown in Fig. 14.12. The X-ray CT system configuration consists basically of an X-ray source tube and a detector, between which a specimen can be mounted on a rotary table. The rotary table, which sits between the X-ray source and the detector, rotates 360 degrees while the detector takes radiographic images of the specimen at different angular positions. The differences in contrast of radiographic images are generated by differences in X-ray attenuation that arise principally from differences in density within a specimen’s material composition. Also, X-rays traversing different paths through a specimen will emerge with different intensity attenuations, which are subsequently measured by the detector. In

FIGURE 14.12 CT process diagram.

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general, the attenuation of intensity of X-rays is theoretically described by an exponential attenuation law, the Beer Lambert law (Als-Nielsen and McMorrow, 2011; Leroy and Rancoita, 2011). Based on this principle, several radiographic images—also known as projection images (or simply “projections”)—are taken from a large number of different angular positions around a specimen on a 360 degrees rotation. By computing the attenuation in the radiation that has passed through a slice of a specimen over the length of the projection (i.e., line integrals), and with the aid of a reconstruction algorithm, it is possible to reconstruct a three-dimensional image (Deans, 2007). After the 3D reconstruction, the CT data offer the possibility of qualitative inspection and quantitative measurements of both the interior and exterior of the specimen under investigation. The 3D image resulting from a reconstruction algorithm represents a volumetric density map—depicted in gray levels (although false colors can be added to different sections)—of the mean X-ray absorption coefficients for the workpiece specimen along the crossing paths of X-rays. This map depends on both the density and composition of a specimen’s material and the energy of the X-rays passing through it. Each gray level offers information about what the X-rays have encountered in their paths along the entire line connecting the source and the respective detector element, thus, revealing external and internal structures of an object, and in CT data, their spatial location. The object’s surfaces can be extracted from the volumetric CT data, which is composed of voxels or “pixels” in three dimensions, using a precise surface determination algorithm that facillitates the specimen’s dimensional measurements (Villarraga-Go´mez et al., 2018). As a result, interior measurements can be obtained as well as the localization of structural material inclusion and the identification of tissue clusters not usually visible through traditional methods of nondestructive testing. Once the scans are complete, dimensioning of the specimen is performed for size and morphology determination, measurement of porosity distribution, and the location of bone tissues. In the work presented here, the measuring procedure included securing the sample with synthetic foam, which is transparent to X-rays, and using settings adjusted for imaging the sample with a standardized cone-beam CT setup operated with an X-ray tube voltage of 165 kV and a tube current of 100 μA. The number of radiographs taken per CT scan was 2880 over a full 360 degrees rotation (i.e., each radiograph was taken at steps of 0.125 degrees). All 2880 radiographs (or image projections) were used for 3D reconstruction. The exposure time for each projection was about 0.708 seconds and the voxel size resolution was 20 μm for the bone slices and 10 μm for the small cube sample. All scans were acquired using Inspect-X and reconstructed using CT Pro 3D (Nikon Metrology, USA), which uses an FDK (Feldkamp et al., 1984) type algorithm. This software takes the 2D projection images acquired by the X-ray detector and generates a 3D image as well as serial cross-sectional images of the bone sample. These consisted of matrices of 2000 3 2000 pixels, which collectively composed a volume of isotropic (20 μm)3 for the bone slices and (10 μm)3 for the small cube sample. Once the 3D reconstruction was complete, a local adaptive or dynamic gradient

14.3 Study of Spongy Bone

threshold, performed with the software package VGStudio Max 3.0 (Volume Graphics, GmbH), was applied for better surface determination.

14.3.3.1 Imaging By imaging the specimen with CT, the 3D reconstruction reveals all internal and external geometry, which enables the analysis of the specimen via clipping views or cross-sectional images with high spatial resolution. The cubical sample in this study was scanned with a resolution of approximately 10 μm (see Fig. 14.13). The cube’s three orthogonal planes intersect at a point that can be considered the origin (0,0,0) of a Cartesian system of coordinates that serves as a reference. These planes can be moved to any location in the volume, or additional tilted planes can be featured at any angle and positioned so that the 3D volume can be inspected slice by slice. From the 2D slices also shown in Fig. 14.13, it can be seen that bone microstructure is composed of an interconnected network of microporous channels that facilitate blood irrigation. 3D analysis of the dimensions (size, spacing, volume, etc.) and network arrangement of the channels provide novel insights related to the porosity as well as the structural and mechanical properties of bone. The 2D cross-sectional images of the cubical sample show marked differences in the cross-sectional network of channels and overall porosity, particularly anisotropy in the orientations as was shown in SEM images (see Fig. 14.10). For example, at Y 5 3.15 mm, some peculiar features can be seen in the bone that

FIGURE 14.13 2D imaging of trabecular bone cube.

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are also evident in the images obtained by SEM for macroarchitecture and primary porosity. As such, the CT-based analysis provides an efficient way to image and describe the 3D architecture inside the trabecular bone. Such analysis facilitates a better understanding of sponge, which can be applied to cases where scaffold design is needed. Furthermore, through this process, it is possible to quantitatively obtain information on topographical variations relative to cortical channels as a function of bone position. Fig. 14.14 shows a slice of the sponge (i.e., a portion of the whole sponge in the XZ plane). Porosity may be defined as the quality of being porous, or alternatively, as the ratio of a cumulative volume of interstices or voids in a material to the total volume. For visualization and pore network analysis, the software package VGStudio Max 3.0 from Volume Graphics, GmbH was used. The porosity analysis of the bone indicated that the overall porosity of the full bone specimen was approximately 60%, while the porosity of the small cubical section was close to 70%. These porosity values can be attributed to the differences in the sponge as a function of the bone position. The scan resolution used in the study (10 μm) was relatively high compared to the mean diameter of the cortical channels, which are typically in the range of 150 1600 μm; therefore, it can be concluded that effective imaging was achieved to compute the cortical porosity. However, if smaller channels or porosity with diameters well below 10 μm are expected, the analysis of these 3D microstructures requires considerably higher resolutions. As the spatial resolution of CT scan increases, the field of view decreases, and this ultimately compromises the size of the sample that can be scanned (samples smaller than 1 cm3 would be required).

FIGURE 14.14 Interconnected system in trabecular bone.

14.3 Study of Spongy Bone

14.3.4 STRUCTURAL PROPERTIES 14.3.4.1 Transmission electron microscopy Transmission electron microscopy (TEM) is a technique that allows for the morphology and size of crystals as well as their atomic arrangements to be seen. Based on this information and using a PDF, it is possible to determine crystalline directions. An electron beam is transmitted through a sample, and then an image is created through magnetic lenses. Images can be produced by scattered electrons, and in this case, a dark-field image is obtained. When an image is obtained by the interaction between a sample and electrons that overpass the sample, it is called bright field. The main advantage of this technique is its atomic resolution, allowing crystalline arrangements to be observed. To observe crystals from mammalian bones, it is necessary to develop a complete organic phase removal without modifying the mineral phase in bone powder. A Soxhlet extraction at 60 C using petroleum ether as a defatting process should be carried out. After that a deproteinizing with NaOCl is recommended, note that these kind of processes induce salts formation, then a multiwashing with hot distilled water and stirring is recommended. When a sample is completely clean, the HA powder should be put in an alcohol with rapid evaporation so as to be sonicated with an ultrasonic homogenizer to break up the particles in the sample. After that, about 2 μL of the diluted sample should be deposited on the specimen holder. TEM is a necessary technique for determining crystalline quality due to misunderstandings related to X-ray diffraction. It has been reported that HA from bones has a low crystalline percent and a low crystalline quality based on X-ray diffraction patterns due to some misinterpretations of the results and technique. Crystalline percent is commonly calculated as the ratio between the crystalline area (see Fig. 14.6), which has the contribution of the scattering in the case of nanocrystals as is the case of HA from bones, and the total area minus instrumental noise (see Eq. 14.1). This calculation can fail if the sample is composed of nanocrystals as was previously mentioned, but it also may be affected by the obtaining process. If the organic phase is not completely removed from the sample, the diffracted pattern will have a substantial amorphous contribution to the total area. Thus, this calculation provides a kind of ratio of concentration which can be modified; then, it is an error to say that HA from bones has low crystalline percent. The second error is related to crystal size: as was previously mentioned; as crystal size decreases to nanometric sizes, the pattern becomes to a curve with broad peaks. In this last case, the correction in Eq. 14.1 that corresponds to the contribution of the scattering by the effect of the crystal size could not be made yet. TEM images in dark field show HA crystals of cleaned trabecular bovine bone powder (Fig. 14.15A). Fig. 14.15B shows the atomic arrangement of these nanocrystals, and Fig. 14.15C exhibits some crystalline planes that belong to HA.

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FIGURE 14.15 TEM images for bovine trabecular bone.

Based on these images, it is possible to know that HA is the main crystalline phase in bone that is composed of polycrystals with irregular shapes in the order of nanosizes (6 9 nm diameter). Regarding crystalline quality, it is clear that there are some defects in the lattices because the substitutional ions. The crystal size for HA from bovine trabecular bone has been studied by TEM. Chen et al. (2011) reported ranges of size from 1.5 to 4 nm, but this value was assigned to an effect of the sample conditioning process because sizes of 30 200 nm length obtained by atomic force microscopy (AFM) were reported.

14.3.4.2 X-ray diffraction The structural properties of a bone can be studied by X-ray diffraction. This technique also requires that the sample be protein and fat-free to avoid the amorphous contribution of these elements to the pattern. This property depends on the macro and microarchitecture of the bone. Then, a complete analysis for trabecular bone must be carried out in the three spatial directions. In order to identify the crystalline phases in a bone, the International Organization for Standardization (ISO) has stablished some Powder Diffraction Files (PDF) in accordance with the International Center of Diffraction Data (ICCD). The numbers of these files are 09 0169 for β-tricalcium phosphate, 9 0169 for α-tricalcium phosphate, 9 432 or 72 1243 for apatite, 25 1137 or 70 1379 for tetracalcium phosphate, and 4 0777 or 82 1690 for calcium oxide. To determine the position of each peak in a pattern, a patron like lanthanum hexaboride powder (National Institute Standards and Technology (NIST)) should be used as an internal standard. As mentioned previously, trabecular bone has different architectures as a function of direction. Fig. 14.16 shows the X-ray diffraction curves for each direction in ascending order. These curves correspond to nanopolycrystalline hydroxyapatite crystals as was confirmed by comparison with PDF 09 0432. Usually, the FWHM is used as a parameter of crystalline quality, but in the case of nanocrystals, this parameter cannot be employed as a quality parameter because these kinds of crystals produce scattering, and as a result, the peaks get broad.

14.3 Study of Spongy Bone

FIGURE 14.16 X-ray diffraction patterns taken for the three studied directions: (A) YZ; (B) XZ; and (C) XY.

Regarding phase identification, only hydroxyapatite could be identified; this fact does not mean that HA is the unique crystalline phase in the sample—bones have several calcium phosphate phases, as well as calcium carbonate, but the amount of these phases is not enough to be identified by this technique. The X-ray curve for the Z-direction also reveals that the hexagonal structure of hydroxyapatite has preferential growth in the C direction. The ratio between I002/I211 is 0.487, 0.474, and 0.565 for X, Y, and Z direction, respectively. These ratios confirm the preferred orientation to the Z-direction. It is worth noting that there is a shift to the right of the peaks regarding the dashed lines, which correspond to synthetic and stoichiometric hydroxyapatite; this issue is related to the composition of the sample which distorts the lattice size. At this point, it is clear that the preferred orientation of hydroxyapatite crystals should be incorporated into the requirements for an ideal scaffold shown by O’Brien (2011) and Bose et al. (2012), as well as macro and microarchitecture.

14.3.5 VIBRATIONAL CHARACTERIZATION: RAMAN SPECTROSCOPY Raman spectroscopy provides information about chemical composition through the identification of rotational and vibrational modes in the functional groups of a molecule. This method is based on inelastic dispersed light when a monochromatic light beam interacts with matter. Some advantages of this technique are its nondestructive character and the small sample sizes necessary to develop an

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FIGURE 14.17 Raman bands for (A) trabecular bone and (B) for BIO-HAp calcined at 1000 C.

analysis. Raman spectrum of bovine trabecular bone exhibits bands that belong to the phosphate group, which corresponds to hydroxyapatite, but there is also a band for B-type carbonate, which confirms that hydroxyapatite in bone is carbonated (Fig. 14.17A). Bovine bone also has collagen and noncollagen proteins, which can be identified by the presence of bands for amide I, amide III, proline (Pro), hydroxyproline (Hyp), and phenylalanine (Phe). A complete composition of trabecular bone is presented in Table 14.4. Synthetic hydroxyapatites do not have all of these compounds because they are stoichiometric. The chemical composition of a scaffold should be another factor that requires consideration in order to obtain the ideal scaffold. In the case of bone characterization by Raman spectroscopy, only a qualitative identification of the components of the sample is possible because some bands are overlapped. Some processes to obtain hydroxyapatite from mammalian bones include calcinations at high temperatures above 900 C, but London˜o-Restrepo et al. (2018) showed how this kind of process turns hydroxyapatite into oxyapatite, after which the O-H band is not present (Fig. 14.17B).

14.3 Study of Spongy Bone

Table 14.4 Characteristic Raman Bands for Bone Material Functional Groups PO4

32

O H CO322

C H

Phospholipids

Amide I Amide III

Phenylalanine (Phe) Hydroxyproline (Hyp) Tryptophan (Trp) Other amino acids

Collagen/Lipid ADN

H 2O

Wavenumber (cm21)

References

ν1-960 ν2-430 ν3-1068 ν4-588 3232 (stretching) 754, 1068 (type B) 1102 (type A) 1416 (ν3 asym) δ-1450 ν-2883 ν-2938 ν-2971 1123 (ν(C C) trans), 1131 (fatty acids), 1745 (ν CQO) 1313 (δ(QCH)) 1625, 1663 1240 (β sheet) 1278 (α helix) 1338 (α helix) 618, 717, 1003, 1190, 1610

Penel et al. (2005)

1582 δ(CaC) 870 1530, 1576 740, 779 645, 857, 920 (proline)

670 (C—S; cytosine); 697 (C—S; methionine); 1206 (tyrosine) 1313 (CH3CH2 twisting mode of collagen/lipid) 682 (ring breathing mode), 725 (ring breathing mode), 823 (out of plane ring breathing) 1638 (intermolecular bending mode of water)

Movasaghi et al. (2007) Penel et al. (2005), Gong and Morris (2015) Movasaghi et al. (2007) France et al. (2014)

Movasaghi et al. (2007) Penel et al. (2005) Penel et al. (2005) Maiti et al. (2004), Makowski et al. (2013) Penel et al. (2005) Penel et al. (2005), Guangyong et al. (2011) Movasaghi et al. (2007) Shibata et al. (2013), Gong and Morris (2015) Penel et al. (2005), Movasaghi et al. (2007) Guangyong et al. (2011) Wopenka et al. (2008), Guangyong et al. (2011), Gong and Morris (2015) Movasaghi et al. (2007) Movasaghi et al. (2007) Movasaghi et al. (2007)

Movasaghi et al. (2007)

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14.3.6 MECHANICAL PROPERTIES Mechanical properties are one of the most relevant parameters with which to evaluate a scaffold. Information about elastic modulus, strength, structural stiffness, and ultimate stress can be obtained from a mechanical test (Wall and Board, 2014). Cylindrical and cubic specimens of bones are usually used in such tests. Here, it is essential to carry it out in the three spatial directions because the mechanical properties change as a function of the tested plane due to the anisotropic character of bone. Different results have been reported in the literature, which is due to multiple factors, such as the source of the bone, age, gender, localization, mineral content, organic content, and humidity, among others. Strain rate is one parameter that affects study results; when this rate increases, elastic modulus and fracture stress increase too (Meyers et al., 2008). When a material is under enough load its shape changes, this is known as deformation. The resulting curve (stress strain) of a compression test of a solid can be formed by three regimens before the fracture point: elastic, elastic-plastic, and plastic. Elastic deformation happens when the material returns to its original shape after the load is removed; a linear relationship between load and deformation is found in the stress strain curve. When the deformation is permanent, plastic deformation has been reached; in this case, there is a bond rupture between some atoms. The elastic-plastic regimen is commonly found in amorphous polymers, which is a combination of the two mentioned effects. According to O’Brien (2011), mechanical properties should be consistent with location where the graft is implanted and strong enough to resist the surgical process. It is clear that emulating the mechanical properties of bone is one of the biggest challenges for tissue engineering since its complex 3D structure has not been reproduced until now. Defatted and deproteinized bovine trabecular bone cubes were kept at 60% relative humidity and 23 C for a month. After that, compression tests for the three spatial directions of a bone cube were carried out with 0.05 mm/min deformation rate. Compressive strength, limit of proportionality, compression at rupture, and modulus of elasticity were analyzed. Fig. 14.18 shows the results for the compressive test for the bone cubes when a load is applied in the X, Y, and Z directions, respectively. In all cases, point “a” represents the proportional limit. It is clear that until point a, the trabecular bone has an elastic behavior, while from point a to c, it is a combined elastic-plastic regimen. In the Y direction, the trabecular bone is formed by cylindrical tubes with microholes, where molecular forces govern behavior. The compressive curve for the Z direction shows the highest breaking force for all samples, which is indicated by letter a 5 9.8 kN, and this is related to the fracture of the columns (i.e., the microstructure of the trabecular bone). Plastic deformation corresponds to the range from point b to c, which is associated with planes perpendicular to the columns. Then, mechanical properties depend on microarchitecture, density, trabeculae number, and crystalline structure. Due to its function within the body, trabecular bone has different responses

14.4 Synthetic Scaffolds Versus Trabecular Bone

FIGURE 14.18 Compression test carried out for the three studied directions: (A) YZ; (B) XZ; and (C) XY.

when under uniaxial compression for each direction. In most cases of synthetic scaffolds, their mechanical properties are isotropic. Table 14.5 shows the compressive force and Young’s module for some materials used in scaffold fabrication.

14.4 SYNTHETIC SCAFFOLDS VERSUS TRABECULAR BONE The mechanical properties of a trabecular bone depend on the complex hierarchical structure of bone, its chemical composition, and chemical interactions between atoms. The elaborate design of bone is the result of billions of years of evolution. The new age has allowed humans to design and create materials inspired by nature as is the case of synthetic scaffolds, which is still a challenge for material science and engineering and tissue engineering. Despite the complexity of bone structure, studies of synthetic scaffolds do not consider the architectonic design of scaffolds together with their physicochemical characterization. As previously mentioned, the materials used for implants should satisfy these requirements: biocompatibility, biodegradability, mechanical properties consistent with anatomical requirements (site), architecture, and manufacture

475

Table 14.5 Mechanical Properties of Bones Plane

Compressive Force

Modulus of Elasticity

Strain Rate

Type of Bone

Reference

YZ

Peak a: 3.09 MPa Peak b: 2.93 MPa Peak a: 2.45 MPa Peak a: 4.58 MPa Peak b: 5.25 MPa Peak c: 4.67 MPa

53.72 MPa 40.13 MPa 92.62 MPa 135.74 MPa 95.27 MPa 57.56 MPa 8 24 GPa 22.6 GPa 12.4 GPa 16.2 GPa 4.5 9 GPa 381.7 6 181.9 MPa 679 6 317 MPa

0.05 mm/s

BTB

This work

0.05 mm/s

BTB BTB

This work This work

1 3 1023 s21

Cortical human bone BCB

Meyers et al. (2008) Novitskaya et al. (2011)

0.1 mm/min 5 mm/min 1 3 1023 s21 6 3 1024 s21 0.33 mm/min

HA (α-TCP 1 pHA) BTB BTB HA scaffold β-TCP scaffold

Zhang et al. (2011) Kelly and McGarry (2012) Chen and McKittrick (2011) Woodard et al. (2007) Tarafder et al. (2013)

XZ XY

L R T

120 MPa 142 MPa 112 MPa 39 103 MPa 12.7 6 5.5 MPa 34.4 6 2.2 MPa 10.95 6 1.28 MPa

0.05 mm/s

BTB: Bovine Trabecular Bone; BCB: Bovine Cortical Bone; L: Longitudinal; R: Radial; T: Transverse.

14.4 Synthetic Scaffolds Versus Trabecular Bone

technology. The last one is a critical issue today because products created by tissue engineering should be able to be scaled-up to satisfy the demands of this growing industry. The primary function of bone scaffolds (natural or synthetic) is replacing and restoring any bone damage. Trabecular bone and synthetic scaffolds can exhibit limitations related to mechanical properties, biocompatibility, and immunogenic reaction, among others. Here, new trends are associated with an interdisciplinary study of the design of bio-inspired systems based on the understanding of the interaction of materials with tissue in order to introduce these materials into the human body. These are the so-called biomimetic materials. Biomaterials for scaffolding can be separated into two groups: natural biomaterials and synthetic biomaterials. The first one is related to the extraction or obtaining of materials from biogenic sources, for example, silk, collagen, gelatin, fibrin, chitosan, agarose, alginate, hyaluronan, and biogenic hydroxyapatite. The second one is related to the synthesis of polymers, such as poly(ethylene glycol) (PEG), poly(e-caprolactone) (PCL), poly (D, L-lactide) (PDLLA), poly(lactic-coglycolic acid) (PLGA), poly(glycolide) (PGA); and ceramics, such as zirconia, calcium phosphate, bioglass, alumina, and synthesized hydroxyapatite (Alaribe et al., 2016). Therefore, scaffold properties and functionality will be determined by the type of biomaterial, architecture, biocompatibility, and biodegradability. A porous preform is required to design synthetic scaffolds; this preform is of regular geometry (e.g., cylinders, cubes, hexagons, and spheres) for incrustation in a solid matrix. Fig. 14.19A D shows some common scaffolds built through regular porous preforms. Other architectures are related to the imitation of natural trabecular bone using random preforms (Fig. 14.19E) or reaction-diffusion models (Velasco et al., 2016). In the past years, additive manufacturing (3D printing) has become a trend. A collection of porosities, shapes, and sizes can be considered to generate a scaffold model. Their mechanical properties have been studied using numerical simulation such as finite element method (FEM) (Gallegos-Nieto et al., 2014) or mechanical tests. These kinds of scaffolds are not fully dense; their porosity is intended to give them biological functionality. Moreover, it is well known that the mechanical properties of scaffolds depend on their porosity regime. Even so, the compressive stress strain behavior of these systems is characterized by having three main regions: an elastic region, a plastic region (plastic plateau), and a densification region. As previously mentioned, bovine trabecular bone has an anisotropic mechanical behavior because of its scaffold geometry. Fig. 14.20 makes a comparison between a bovine trabecular bone cube compressed in the Z-direction, synthetic scaffolds built through regular preforms, and biomimetic scaffolds. The stress and strain curve is simulated by FEM in the case of cubic and spherical incrustations (Fig. 14.19A and D) for 70% porosity, which is approximate to the porosity of trabecular bone. Mechanical properties strongly depend on the porosity of the designed structure. These structures are relativity isotropic compared to natural bone or biomimetic systems, and exhibit three characteristic regions (elastic,

477

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CHAPTER 14 Study of microstructural, structural, mechanical

FIGURE 14.19 Designs of synthetic scaffolds made using regular preforms.

plastic plateau, densification). However, they present only a single deflection region (collapse of architecture) between the elastic region and the plastic plateau which is a sign of regular architecture. While hydroxyapatite-based biomimetic scaffolds (Alonso-Sierra et al., 2017) exhibit poor mechanical behavior compared with the other systems. This experiment evidences the role of the organic phase as reinforcement. Mechanical properties can be improved with the addition of a proteinic matrix, like collagen or gelatin. Finally, bovine trabecular bone has a complex behavior before the plastic region which is a contribution of its natural architecture. Briefly, Fig. 14.20 shows that some behaviors of the strain stress curves for scaffold systems are determined, to a large degree, by the architecture of the sample.

14.5 Conclusions and Perspective

FIGURE 14.20 Mechanical behavior comparison between natural trabecular bone, ceramic biomimetic scaffolds, and synthetic scaffolds made using regular preforms, such as spheres and cubes.

There are many scaffold manufacturing technologies, but the most popular today is related to additive manufacture or 3D printing. One proposal to improve synthetic scaffolds is to copy the architecture of trabecular bone, solving its architecture by X-ray tomography, and replicating it.

14.5 CONCLUSIONS AND PERSPECTIVE For many years, scaffold requirements have been limited to mechanical and biological properties, but it is clear that a natural scaffold, as is the case of bovine trabecular bone, involves many other properties that are necessary for the successful behavior of a scaffold within the body. The complex architecture of trabecular bone defines its kinds of porosity, which allow for blood irrigation, vascularization, and cell proliferation, and provides intercommunication sites for osteocytes. The structural properties of this bone demonstrated that hydroxyapatite crystals have a preferred orientation parallel to the axis where the mass is charged. Moreover, it was found that hydroxyapatite crystals have nanometric sizes which provide a large surface area to promote the interaction between scaffolds and organisms. It is imperative that the host organism that receives the scaffold be able to reproduce all the physicochemical properties of natural bone, for this purpose, the provided scaffold must have the previously mentioned properties. This

479

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CHAPTER 14 Study of microstructural, structural, mechanical

chapter does not only evidence the huge difference among natural bone and synthetic scaffold, it also provides a complete characterization of trabecular bone which can be used as a guideline for scaffold design which is one of the biggest challenge in tissue engineering.

ACKNOWLEDGMENTS The authors would like to thank Dra. Genoveva Hernandez-Padro´n, M. Alicia del RealLo´pez, and Dra. Beatriz Millan-Malo for their technical support for the Raman Spectroscopy, SEM images, and X-ray diffraction respectively. S. M. London˜o-Restrepo and C. F. Ramirez-Gutierrez acknowledge financial support from Consejo Nacional de Ciencia y Tecnologı´a (CONACyT). This project was financially supported by PAPIITUNAM (Project number IN112317).

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FURTHER READING Gamsjaeger, S., et al., 2014. Pediatric reference Raman data for material characteristics of iliac trabecular bone. Bone. 69, 89 97.

Appendix A

APPENDIX A Powder diffraction file of synthetic hydroxyapatite. Taken from: Hodge et al., 1938 PDF#09 0432: Ca5(PO4)3 (OH), Radiation 5 CuKα λ 5 1.54056 nm Hexagonal structure; Density (c) 5 3.160 g/cm3 d(A)

I(v)

h

k

l

2θ

d(A)

I(v)

h

k

l

2θ

8.170 5.260 4.720 4.070 3.880 3.510 3.440 3.170 3.080 2.814 2.778 2.720 2.631 2.528 2.296 2.262 2.228 2.148 2.134 2.065 2.040 2.000 1.943 1.890 1.871 1.841 1.806 1.780

4.0 3.0 2.0 7.0 7.0 2.0 33.0 11.0 16.0 100.0 61.0 62.0 27.0 7.0 10.0 25.0 3.0 13.0 5.0 11.0 3.0 8.0 42.0 24.0 9.0 61.0 31.0 19.0

1 1 1 2 1 2 0 1 2 2 1 3 2 3 2 3 2 3 3 1 4 2 2 3 3 2 3 4

0 0 1 0 1 0 0 0 1 1 1 0 0 0 1 1 2 1 0 1 0 0 2 1 2 1 2 1

0 1 0 0 1 1 2 2 0 1 2 0 2 1 2 0 1 1 2 3 0 3 2 2 0 3 1 0

10.820 16.841 18.785 21.819 22.902 25.354 25.879 28.126 28.966 31.733 32.196 32.902 34.048 35.480 39.204 39.818 40.452 42.029 42.318 43.804 44.369 45.305 46.711 48.103 48.623 49.468 50.493 51.283

1.754 1.722 1.684 1.644 1.611 1.587 1.542 1.530 1.503 1.474 1.465 1.452 1.452 1.433 1.407 1.407 1.348 1.316 1.316 1.306 1.306 1.280 1.265 1.265 1.257 1.249 1.235 1.221

26.0 33.0 7.0 17.0 14.0 7.0 11.0 11.0 19.0 23.0 8.0 25.0 25.0 18.0 8.0 8.0 6.0 11.0 11.0 9.0 9.0 15.0 7.0 7.0 20.0 2.0 25.0 21.0

4 0 1 3 3 5 4 3 2 5 5 3 3 5 4 4 5 4 4 2 5 4 3 6 2 4 5 5

0 0 0 2 1 0 2 3 1 0 1 2 0 1 2 1 1 0 3 0 2 2 2 0 1 3 1 2

2 4 4 2 3 1 0 1 4 2 0 3 4 1 2 3 2 4 1 5 0 3 4 2 5 2 3 2

52.100 53.143 54.440 55.879 57.128 58.073 59.938 60.457 61.660 63.011 63.443 64.078 64.078 65.031 66.386 66.386 69.699 71.651 71.561 72.286 72.286 73.995 75.022 75.022 75.583 76.154 77.175 78.227

485

CHAPTER

Laser processing of biopolymers for development of medical and high-tech devices

15

Nadya E. Stankova1, Petar A. Atanasov1, Nikolay N. Nedyalkov1, Konstantin Kolev2, Eugenia Valova2 and Stephan Armyanov2 1

Institute of Electronics, Bulgarian Academy of Sciences, Sofia, Bulgaria 2Institute of Physical Chemistry Rostislaw Kaischew, Bulgarian Academy of Sciences, Sofia, Bulgaria

15.1 INTRODUCTION Polydimethylsiloxane (PDMS) is an important technical polymer, which has various applications in medicine and the pharmacological industry due to its outstanding characteristics. It belongs to the group of silicones which have the common characteristic of a backbone composed by an alternate succession of silicon (Si) and oxygen (O) atoms joined together with strong, covalent interatomic bonds. The Si atoms are coupled with two adjacent O atoms and two organic radicals, that is, CaH or CaR (where R is an organic group). The various Si-based polymers only differ from each other by organic radicals, for example, methyl, vinyl, hydrocarbon, or other organic groups. The mechanical, physical, and chemical properties of these polymers can be controlled, depending on their application, by including different additives such as plasticizers, fillers, and crosslinkers. The specific volume and viscosity of the PDMS polymer can be driven by varying its molecular weight from a few hundreds to several hundred thousand g.mol21 and, thus, to control the product form - liquid at low molecular weights and solid gum at high molecular weights. Thus, PDMS polymer, can be classified as viscous fluid or solid. When reinforcement is performed then cross-linked solids are formed, known as elastomers. Elastomers with different hardness can be fabricated depending on their application subject. At force applied the material will react with elastic deformation, which is defined by the material mechanical properies. In general, PDMS elastomers possess high electrical insulation, thermal stability, and water repellency properties. PDMS is highly transparent to ultraviolet (UV), visible (VIS) and infra-red (IR) light, and starts to absorb below wavelengths of 280 nm (i.e., above 4.4 eV).

Materials for Biomedical Engineering: Hydrogels and Polymer-based Scaffolds. DOI: https://doi.org/10.1016/B978-0-12-816901-8.00015-8 © 2019 Elsevier Inc. All rights reserved.

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This polymer is characterized with high chemical and thermal stability, high durability, high dielectric breakdown, nontoxicity, and controllable flexible properties. Only strong acids or strong bases can depolymerize the siloxane chain. As a result, the PDMS is not very susceptible to oxidation or thermal degradation and it can be sterilized by heating. It is also well-known to possess superior elasticity and flexibility in mechanical applications, therefore, PDMS has become an excellent choice for biological studies and applications because of its nontoxicity and biocompatibility to cells and tissues, as well as its high permeability to gases (Kuo, 1999; Caprino and Macander, 1999). Thus, PDMS is an important and remarkable material for the development of microelectromechanical systems (MEMS) or long-term neural implants (without disruption of the soft tissues like the brain, spinal cord, and muscles) (Qina et al., 2014). Due to their chemical inertness, silicones, and especially PDMS, are perfect biocompatible and biostable materials for use as medical implants. Chemical coupling of foreign species to silicone requires opening its structure which involves an irreversible modification and interatomic bond breaking (Kolasinski, 2009; Laude et al., 2008). However, this cannot be done by mechanical or thermal treatment since silicone does not melt, sublimate, or evaporate, but rather transforms into a glassy and extremely fragile material at temperatures exceeding 230 C. PDMS is a flexible material which is excellent for microfabrication of MEMS devices (McDonald et al., 2000). One useful property of PDMS is that its surface can be chemically modified in order to obtain interesting interfacial properties (Makamba et al., 2003). For example, PDMS can be covalently functionalized through exposure to a reactive oxygen plasma whereby SiaCH3 groups on the surface transform into SiaOH ones. These silanol surfaces are easily transformed with alkoxysilanes yielding a different chemistry (Gelest, 2004; Hermanson et al., 1992). After polymerization, solid PDMS samples become strongly hydrophobic (McDonald et al., 2000). Thus, it is difficult to be wetted by polar solvents, for instance water. Various applications of PDMS require surface modification in order to change its morphology and chemical composition, namely to functionalize the surface. The most used method is via plasma oxidation in atmospheric air, oxygen, or argon plasma for changing of the chemical properties of the surface, that is, creation of superficial silanol (SiOH) groups (Vladkova et al., 2005). The recovery of the surface’s hydrophobicity after the plasma treatment is unavoidable regardless of the surrounding medium, that is, vacuum, air, or water (Hillborg et al., 2000). The temporal evolution of the contact angles between the deionized water and PDMS treated with O2 plasma is reported by Donzel et al. (2001). The interaction of photons (using VUV lamps) with polymers with an emphasis on UV laser ablation and surface modification was extensively described and reviewed by Lippert (2005). It is deduced that laser ablation of polymers is an established process in industrial applications, but the mechanisms of ablation are still controversial. Different polymers, such as poly(methyl) methacrylate,

15.1 Introduction

polyimide, and specially designed polymers are used to show that the mechanism is a mixture of photochemical and photothermal features which are closely related to the polymer structure and properties. In this chapter, different approaches to probe the ablation mechanisms and to improve ablation are discussed. Various trends for laser ablation, such as ultrafast ablation or VUV ablation are also provided. Surface modifications of polymers using VUV photons from lamps are discussed for oxidation and nitriding of PDMS and polyolefins. The mechanism of the PDMS surface oxidation is presented in detail. An important and strictly positioning alternative method for modifications of PDMS is the laser treatment. Ablation by laser pulses is a complicated mixture of photochemical and photothermal reactions where the ratio between them is a function of the polymer properties and the laser processing parameters (Urech and Lippert, 2007). The use of laser-pulsed radiation offers a useful approach to the processing and selective surface functionalization of silicone. The increase of Si in the EDS spectra compared to the native PDMS is observed to be caused by irradiation of the material by Nd:YAG laser (Dupas-Bruzek et al., 2009b). Laser treatment is also applied in order to facilitate selective metalization of the areas activated in this way (Dicara et al., 2003; Laude et al., 2003a; Laude et al., 2003b). After irradiation of PDMS with a Nd:YAG laser, a new peak at 516 cm21 into the μ-Raman spectrum appears which is assigned to mono- and/or polycrystalline nanostructured silicon (c-Si) (Palma et al., 1999; Dicara et al., 2003; Graubner et al., 2006; Dupas-Bruzek et al., 2009b). The intensity of this peak increases with the number of the laser pulses and fluence. Moreover, depending on the experimental conditions, in the μ-Raman spectra broad bands around 1346 and 1597 cm21 after irradiation with excimer laser with wavelength of 248 nm, and around 1320 and 1603 cm21 after irradiation with Nd:YAG laser with wavelength of 266 nm, appear. They can be assigned to D and G bands of carbon (C) which are related to defects and disorder, and to sp2 C-C bond vibrations, respectively, in all C structures (Graubner et al., 2006). Lippert et al. (1997) proposed that with the addition of dopants into the bulk it has become possible to sensitize most known classes of polymers for UV laser ablation at any desired wavelength, including fluoropolymers. Important features of dopant-induced ablation are the reduction of threshold energy fluence required for ablation and the enhancement of the etching rate by factors higher than ten. The investigated dopant/polymer systems are summarized and compared in the review. Based on the available information, a general scheme including relevant pathways is suggested, revealing that in each particular case the dominant mechanism depends on the specific system under study. Information on the changes of the PDMS structure caused by ns-laser pulses is typically obtained by μ-Raman measurements (Dicara et al., 2003). Photodecomposition of silicone results in volatile organic radicals, Si clusters, and substoichiometric silica foam that develop on a rather chaotic surface topography. This novel medium is further exploited by inducing selective

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auto-catalyzed metalization of the once-irradiated sample surface. Laser irradiation can induce breaking of the chemical bonds with minor heat effects and thus to cause chemical activation of the surface. Laser treatment can also be used in the ablation regime to micro- or nanostructure the polymer surface and thus to enhance the adhesion of the metal film subsequently deposited (Laude, 1997; Dini, 1997). Recently, “negative ablation” or swelling of PDMS samples processed by KrF excimer laser (λ 5 248 nm) by as low as 50 mJ energy pulses was reported by Kolev et al. (2016). The authors explained that this effect is caused by the low fluence and small repetition rate of the laser used. Moreover, the appearance of such defects in the form of cones in depth of the polymer sheet and changes of its chemistry are considered as a precursor for the ablation at higher than the ablation threshold fluence. Modification of PDMS films induced by UV irradiation is investigated via X-ray photoelectron spectroscopy (XPS) (Schnyder et al., 2003). The binding energies (BEs) of the Si 2p and O 1s increase and reach values corresponding to SiO2. After spin-coated films of PDMS exposure to microwave oxygen plasma, XPS data indicate that a significant quantity of the oxidized layer contained Si-bonded to three or four oxygen atoms (SiOx) (Hillborg et al., 2000). The degree of the conversion of organic to inorganic silicon (SiOx) increases with the increase of the microwave plasma dose. Similar results were obtained by Hillborg and Gedde (1998) when samples were exposed to a corona discharge. The PDMS surface can be enriched with oxygen and depleted of carbon by UV irradiation (Schnyder et al., 2003; Graubner et al., 2004; Lippert, 2005). It is also supposed that the methyl group CH3 can be replaced, not only by oxygen, but also by the hydroxyl group OH. In other groups of investigations, excimer laser pulses (e.g., KrF, ArF, and F2) are used for the fabrication of optical elements (Takao et al., 2004a, 2005; Okoshi et al., 2005), deposition of thin films (Okoshi et al., 2002a,b, Takao et al., 2004b), modification of optical properties (Okoshi et al., 2007a, 2009a,b), and surface processing or functionalization (Dicara et al., 2003; Ferreira et al., 2013; Okoshi et al., 2004, 2007b, 2009c; Takao et al., 2004b; Dupas-Bruzek et al., 2009a). For example, KrF excimer laser (λ 5 248 nm) processing of siloxane-based silicone rubber is used to photo-decompose the sample surface to change its surface relief (Dini, 1997; Laude, 1997; Dicara et al., 2003; Laude et al., 2008). Recently, superhydrophobicity is obtained and studied in the grid patterns produced by excimer laser (λ 5 248 nm) processing of PDMS (Farshchian et al., 2016). Processing with IR femtosecond (fs) laser pulses is also accomplished and microfabrication of PDMS has been investigated (Huang and Guo, 2009; Huang and Guo, 2010; Huang, 2010; Fukami et al., 2004). For example, Huang and Guo (2009) studied fs-laser ablation of PDMS induced by 900 fs at λ 5 1552 nm, both on a fixed spot and for the formation of a continuous line. They observed an

15.2 Structure and Raman Spectrum of Polydimethylsiloxane

ablation threshold of  4.6 J/cm2 and studied the influence of pulse overlapping rate and irradiation energy on line width, internal ablation interface depth, as well as ablation surface quality. Many of the investigations exploiting fs-laser pulses are restricted to the IR wavelength radiation, while studies concerning UV fs-laser ablation of PDMS are still scarce. In the framework of our investigations during the last 4 years, we have obtained sufficient progress and success in laser micro- or nanoprocessing and activation of the biocompatible polymer PDMS, formation of well-defined trenches and structures, and subsequent successful electroless metalization with nickel (Ni) and platinum (Pt), which was not found to depend significantly on the time interval after laser treatment. Laser treatment is accomplished by four laser systems lasting at nanosecond (ns), picosecond (ps), and femtosecond (fs) pulses in UV, VIS, and near-infrared (NIR) spectra. The influence of the wavelength, pulse duration, number of subsequent pulses, and laser fluence on the ablation depth and width of the trenches, the form and quality of the structures produced, as well as changes of the morphology and chemical composition of the surface were studied. The produced trenches and structures were analyzed via numerous high-resolution analytical techniques such as High-Resolution Scanning Electron Microscopy (HR SEM), μ-Raman spectroscopy, optical and 3D color laser microscopes, XRD, and XPS. Future investigations and possible applications are envisaged in the concluding section.

15.2 STRUCTURE AND RAMAN SPECTRUM OF POLYDIMETHYLSILOXANE The molecular structure of the PDMS consists of a SiaOaSi backbone to which methyl groups are attached through Si-C linkage, so giving this polymer the general formula [SiO(CH3)2]n, where n is an integer denoting the number of repeating monomer units. The atomic motion of the vibrational modes of the methyl groups determines the Raman spectra of PDMS. The Raman spectra of PDMS taken by two laser excitation wavelengths are depicted in Fig. 15.1 and the characteristic vibrational modes are summarized in Table 15.1. The typical μ-Raman spectrum of pristine PDMS-elastomeric material shows peaks, which characterize the various chemical bonds: 488 cm21 (SiaOaSi symmetric stretching); 685 cm21 (SiaCH3 symmetric rocking); 709 cm21 (SiaC symmetric stretching); 787 cm21 (CH3 asymmetric rocking 1 SiaC asymmetric stretching); 859 cm21 (CH3 symmetric rocking); 1262 cm21 (CH3 symmetric bending); 1411 cm21 (CH3 asymmetric bending); 2909 cm21 (CH3 symmetric stretching); and 2970 cm21 (CH3 asymmetric stretching) (Bae et al., 2005).

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CHAPTER 15 Laser processing of biopolymers

15,000 Intensity (arb.units)

492

Native PDMS

10,000

5000

0

400

800

1200

1600

2000

Raman shift

(cm–1)

2400

2800

FIGURE 15.1 Raman spectra of native PDMS-elastomer taken with laser excitation wavelengths: 514 and 785 nm. Reprinted from Stankova, N.E., Atanasov, P.A., Nedyalkov, N.N., Stoyanchov, T.R., Kolev, K.N., Valova, E.I., et al., 2015. Fs- and ns-laser processing of polydimethylsiloxane (PDMS) elastomer: comparative study. Appl. Surf. Sci. 336, 321 328; Atanasov, P.A., Stankova, N.E., Nedyalkov, N.N., Fukata, N., Hirsch, D., Rauschenbach, B., et al., 2016a. Fs-laser processing of medical grade polydimethylsiloxane (PDMS), Appl. Surf. Sci., 374, 229234.

Table 15.1 Summary of the Characteristic Vibrational Modes of PDMS Atomic Group

Vibration Mode

Raman Shift (cm21)

SiaOaSi SiaCH3 SiaC CH3 1 SiaC CH3 CH3 CH3 CH3 CH3

Symmetric stretching Symmetric rocking Symmetric stretching Asymmetric rocking 1 asymmetric stretching Symmetric rocking Symmetric bending Asymmetric bending Symmetric stretching Asymmetric stretching

488 687 708 787 862 1262 1412 2907 2965

15.3 EXPERIMENTAL AND ANALYTICAL TECHNIQUES PDMS-elastomer sheets (medical grade NuSil MED 4860) of different thicknesses (100 800 μm) are processed by four types of laser systems: (1) ns-laser—Q- switch Nd:YAG; (2) fs-lasers—Nd:glass and (3) Ti:sapphire; and (4) ps-laser—Nd:YVO4. The parameters of the laser systems are summarized in Table 15.2.

15.3 Experimental and Analytical Techniques

Table 15.2 Summary of the Laser Systems Used for PDMS Processing Wavelength Type Ns-laser

Ps-laser

Fs-laser 1

Fs-laser 2

UV

VIS

NIR

Average Fluence Range on the Surface

266, 355 nm 15 ns 10 Hz 355 nm 7 ps 500 kHz 263 nm 300 fs 33 Hz 266, 355 nm

532 nm 15 ns 10 Hz

1064 nm 15 ns 10 Hz

0.744.3 J/cm2 for UV 3.044.0 J/cm2 for VIS 10.0416.0 J/cm2 for NIR

-

-

0.0240.05 J/cm2

527 nm 300 fs 33 Hz 532 nm

1055 nm 900 fs 33 Hz 1064 nm

1.143.9 J/cm2 for UV 1.145.4 J/cm2 for VIS

35 fs 1000 Hz

35 fs 1000 Hz

35 fs 1000 Hz

0.1640.50 mJ/cm2 for UV, VIS, and NIR

Continuous linear trenches are formed by the overlapping of different numbers of consecutive laser spots focused on the surface of the material which is mounted on a stepper-motor x y table. In order to measure the optical properties, the sample is stationary and a series of consecutive laser pulses are focused on a single spot area. All PDMS samples are cleaned before laser treatment by following steps: washing in a detergent solution using an ultrasound bath; rinsing with deionized water; cleaning with ethanol in an ultrasound bath; and, finally, air stream drying. A variety of experimental techniques are applied to characterize the PDMS-elastomer (native, laser irradiated, and metalized tracks): optical spectrometry (Ocean Optics DH-2000 light source and Ocean Optics HR 4000 Spectrophotometer) to measure the optical transmission in the UV, VIS, and NIR ranges of the electromagnetic spectra; optical microscopy (Zeiss Opton, West Germany) to observe any visible permanent modifications of the surface and to measure the ablation depth; scanning electron microscopy (SEM) (Hitachi SU-70 with field emission gun or Zeiss Ultra55, Gemini) and SEM/FIB (Lyra/Tescan dual beam system) to assess the surface morphology of laser-treated or metalized tracks; VK-9700K 3D color laser microscope (KEYENCE) to view and analyze the laser-treated areas; three μ-Raman spectrometers with spectral resolution of 1 or 0.2 cm21, respectively, (Invia, Renishaw) using 514 or 785 nm excitation wavelengths and beam spot on a sample 5 3 5 μm for 100 3 objective, (LabRAM HR Evolution-Horiba Scientific) equipped with 100 3 magnified objective and laser source at λ 5 532 nm, and RMS-310 μ-Raman spectrometer (Photon Design) equipped with laser, operating at λ 5 532 nm having a beam spot on a sample of 1 3 1 μm for determination of the chemical composition. The μ-Raman spectra of the laser-treated PDMS surface are acquired from different sections at the bottom of the tracks and compared to the Raman signals of the native

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material. Scanning transmission electron microscopy (STEM) coupled with the facility for selected area electron diffraction (SAED) PHILIPS CM 20 Ultratwin are used to investigate the interface structural morphology of metalized tracks. The PHI model 1600 system equipped with an Omni Focus Lens III and a standard Mg Kα X-ray source at a high voltage of 15 kV and 300 W is used for XPS studies. The PHI data treatment and interpretation is made by applying Multipak 9 software.

15.4 OPTICAL PROPERTIES OF POLYDIMETHYLSILOXANE DURING NS-LASER TREATMENT The change of the optical properties of PDMS polymers during repeatedly nanosecond laser processing was studied by Stankova et al. (2016). The absorbance of the native elastomer PDMS decreases strongly with the wavelength as seen in Fig. 15.2. The values of the absorption coefficient and the corresponding penetration depths of the material are calulated according to the Beer-Lambert law. It must be noted that the scattering is neglected and only the linear absorption coeficient is assumed for the wavelengths used at the laser treatment; see Table 15.3. It is well-known that native medical grade silicones such as PDMS elastomer are optically transparent from the UV up to NIR region of the spectra. Laser treatment of such materials with wavelengths in this spectral region is possible because of the incubation effect which increases the optical absorption due to the accumulation of defects and cracks below the surface. Moreover, changes in the relief and chemical composition of the native material occur. It is supposed that the incubation in the polymers is a result of these intermolecular bond breaking 0.5

Absorbance (arb. units)

494

0.4

0.3

0.2

0.1

0.0

300

400

500

600

λ (nm)

FIGURE 15.2 Experimental absorbance spectra of pristine PDMS.

700

800

900

1000

15.4 Optical Properties of Polydimethylsiloxane

Table 15.3 Optical Parameters of Native PDMS Wavelength (nm) Absorption coefficient—α (cm21) Penetration depth (mm)

266 14.9 0.669

355 7.38 1.354

532 3.58 2.794

1064 2.86 3.502

Reprinted from Stankova, N.E., Atanasov, P.A., Nikov, R.G., Nikov, R.G., Nedyalkov, N.N., Stoyanchov, T.R., et al., 2016. Optical properties of polydimethylsiloxane (PDMS) during nanosecond laser processing, Appl. Surf. Sci., 374, 96 103.

chemical transformations which occur in the local area of the laser-treated zone only without changing the bulk material properties. These modifications depend simultaneously on the laser wavelength and fluence applied. Moreover, they appear after a certain number of pulses. This means that at a given wavelength and fluence, the absorption will increase during the laser irradiation with the number of pulses, that is, a certain number of pulses is required to accumulate local chemical transformations and to reach the ablation conditions. Consequently, the given absorbed fluences are valid only for the first laser pulse hitting the material. When continuous tracks are produced on the PDMS by consecutive laser pulses, an overlapping of each adjacent spot of the laser beam occurs. The moving speed of the x y table is related to the number of pulses necessary to perform ablation at an unit area and laser repetition rate. Thus, it is very important to know how the number of consecutive laser pulses, that is, the incubation process, affects the optical parameters and chemical characteristics of the native material. Practically only the first pulse hitting the surface interacts with the unchanged native material properties. In this respect, a specific experimental study is performed. At still stand position of the sample, an increasing number of laser pulses and different laser fluences are applied and every time an intermediate measurement of the optical transmittance is performed. Fig. 15.3 shows the changes of the optical properties of medical-grade PDMS interacting with different numbers of subsequent laser pulses at a given wavelength and fluence. For all wavelengths applied, the lowest number of pulses at which the incubation effect begins (i.e., the optical transmittance starts decreasing) is defined as eight pulses, but different laser energy is needed to deliver. The laser fluence (at fixed laser spot size) increases significantly from UV to NIR irradiation in order to cause visible damage with the same number of pulses of eight. It can be summarized that at every wavelength used, the number of initial pulses at which the optical transmittance begins to decrease and visible damages appear, that is, the transparent polymer starts to absorb the light efficiently, up to eight pulses with an increase of the laser fluences as follows: 1.0 J/cm2 for 266 nm, 2.5 J/cm2 for 355 nm, 10 J/cm2 for 532 nm, and 16 J/cm2 for 1064 nm. This means that the threshold laser fluence, which induces the incubation process after eight pulses depends on the wavelength. At the given wavelength and fluence, respectively, further increase of the number of pulses leads to a gradual decrease of the optical transmittance and at 22 and more pulses the material becomes almost optically

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CHAPTER 15 Laser processing of biopolymers

100

100 80 60 40 20 200 400 600 800 1000 λ (nm) CH3

PDMS chemical structure

Si O CH3

0.0

n

80

Transmittance (%)

1.5

60

8 pulses 11 pulses 16 pulses 22 pulses 33 pulses

λ = 266 nm

40

F = 1.0 J cm–2

20 0

400

600

λ (nm)

800

1000

400

(D)

100 80 8 pulses 11 pulses 16 pulses 22 pulses 33 pulses

60 λ = 355 nm F = 2.5 J cm–2

40

600

800

Wavelength (nm)

20

100

Transmittance (%)

200

(C)

(B)

3.0 Transmittance (%)

Absorbance (arb. units)

(A)

Transmittance (%)

0

80 60 8 pulses 11 pulses 16 pulses 22 pulses

λ = 532 nm

40

F = 10 J cm–2

20 0

400

600

800

400

Wavelength (nm) (E)

600

800

Wavelength (nm)

100

Transmittance (%)

496

80 λ = 1064 nm

60

–2

F = 13 J cm

11 pulses 14 pulses 16 pulses 22 pulses

40 20 0 400

600

800

Wavelength (nm)

FIGURE 15.3 Optical properties spectra: (A) the absorbance and transmittance of the medical grade pristine PDMS samples (thickness of 170 μm). Optical transmittance of the PDMS areas after laser processing with (B) 266 nm, (C) 355 nm, (D) 532 nm, (E) 1064 nm, and different number of pulses and fluences. Reprinted from Stankova, N.E., Atanasov, P.A., Nikov, R.G., Nikov, R.G., Nedyalkov, N.N., Stoyanchov, T.R., et al., 2016. Optical properties of polydimethylsiloxane (PDMS) during nanosecond laser processing, Appl. Surf. Sci., 374, 96 103.

opaque in the entire area treated by UV and VIS wavelengths. This means that the incubation saturates after a certain number of pulses, that is, the optical absorption of the material increases significantly for the laser ablation to occur

15.4 Optical Properties of Polydimethylsiloxane

100

Transmittance (%)

2

266 nm; 1.0 J/cm 2 355 nm; 2.5 J/cm 2 532 nm; 10.0 J/cm 2 1064 nm; 13.0 J/cm

80 60 40 20 0 10

20

30

Number of pulses (N)

FIGURE 15.4 Change of the transmittance with the number of consecutive laser pulses at given wavelength and laser fluence.

effectively. Under NIR irradiation, the incubation saturation occurs earlier, at 14 pulses, probably due to the higher penetration depth of this wavelength and higher energy dose delivered (16 J/cm2) in the laser-treated area. It is worth noting that at a wavelength of 1064 nm the material becomes optically opaque after 22 pulses and the laser fluence decreases up to 13 J/cm2 (in this case the incubation starts at a higher number of pulses emitted, i.e., 11 pulses). A summary of the changes of the transmittance on the number of laser pulses at given wavelengths and laser fluences are provided in Fig. 15.4. It should be noted that a further increase of the number of pulses after reaching the saturation of chemical transformations leads to drilling of the polymer. The number of pulses which cause drilling at UV irradiation is much higher (above 110 pulses at 266 and 355 nm) than at VIS and NIR wavelengths (above 22 pulses). A possible explanation could be that the incubation occurs below the surface in a depth higher than the UV light penetration depth. This behavior is a result of the lower absorption coefficient of the pristine PDMS in the VIS-NIR region in comparison with the UV region of the spectra, which results in higher penetration depth at longer wavelengths. As a result, the ablation depth measured is highest under NIR irradiation (B150 μm) and decreases gradually to 40 μm with reducing wavelength of up to 266 nm. Also, it should be considered that the contribution of the heat accumulation in the material at VIS and NIR irradiation is higher than at UV laser irradiation. Furthermore, much higher values of the laser fluence needed to induce incubation in the PDMS polymer by irradiation with VIS (λ 5 532 nm) and NIR (λ 5 1064 nm) lights could be explained by the lower photon energies, 2.3 and 1.2 eV, respectively, in comparison with the UV photon energies, such as 4.7 and

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3.5 eV for UV wavelengths of 266 and 355 nm, respectively. Despite of the low optical absorption (linear absorption coefficient; see Table 15.3) of the PDMS material to the photons with energy in the range of 5.041.165 eV (i.e., between 248 and 1064 nm wavelength range, respectively), UV photons are absorbed selectively and exclusively onto Si-C electron bonds with a binding energy of 3.3 eV. Thus, it is expected that organic radicals will be ejected as a result of the laser irradiation of the surface, while the backbone SiaO and CaH bonds, which have higher binding energy of 4.7 and 4.3 eV, respectively, could be broken with higher fluence or higher energy delivery (i.e., higher number of consecutive pulses). Obviously, much higher energy density should be delivered into the laser-treated zone in order the corresponding chemical bonds to be broken and the material to be ablated using VIS and NIR irradiation light. This assumption is confirmed by the μ-Raman spectroscopy measurements. In conclusion, we can deduce that the wavelength, laser fluence, number of pulses, and properties of the polymer are parameters which simultaneously affect the incubation and ablation processes.

15.5 FS-LASER NANOSTRUCTURING When the PDMS samples are processed by fs-laser pulses the surface inside the trenches becomes very rough with nanosized random structures. Moreover, the intensities of the peaks in the micro-Raman spectra at 488 cm21 (SiaOaSi stretching mode), 708 cm21 (SiaC symmetric stretching mode), and 2907 cm21 (CH3 symmetric stretching mode), respectively, strongly decrease with the increase of the laser fluence. Well-defined and good quality trenches are produced via UV, VIS, and NIR fs-laser treatment (Atanasov et al., 2016a). Fig. 15.5 depicts laser microscope images taken of the trenches. In all cases, an almost linear dependence of the trench depth on laser fluence is observed; see Fig. 15.6. An increase of the speed of the sample with respect to the processing laser beam results in a corresponding decrease of the depth due to the reduced number of overlapping laser pulses. At a fixed fluence, a larger depth for UV with respect to VIS pulses with the same pulse duration is observed, which agrees with the finding of Atanasov et al. (2014). This effect can be explained by the higher absorption of the material for UV with respect to VIS light. However, at similar fluence levels, NIR fs-pulses generated trenches obtain larger depth than those corresponding to VIS and UV. The different laser ablation efficiencies depends on the very low absorption of PDMS in the NIR that allows coupling photon energy at deeper depths in the sample with respect to lower wavelengths. The different laser-pulsed irradiation and possible incubation effects related to the laser photon wavelength can also play a role as reported by Graubner et al. (2002).

15.5 Fs-Laser Nanostructuring

FIGURE 15.5 Pictures of the trenches produced by UV, VIS, and NIR fs-laser pulses, respectively. (A) UV processing: laser energy 5 79 μJ, scanning speed 5 38 μm/s, depth of trench 5 15.5 μm, and width 5 62.2 μm; (B) VIS processing: laser energy 5 142 μJ, speed 5 38 μm/s, depth of trench 5 9.3 μm, and width 5 206.1 μm; and (C) NIR processing: laser energy 5 307 μJ, speed 5 95 μm/s, depth of trench 5 16.8 μm, and width 5 63.5 μm. Reprinted from Atanasov, P.A., Stankova, N.E., Nedyalkov, N.N., Fukata, N., Hirsch, D., Rauschenbach, B., et al., 2016a. Fs-laser processing of medical grade polydimethylsiloxane (PDMS), Appl. Surf. Sci., 374, 229 234.

The chemical changes induced by laser irradiation are analyzed at the bottom of the trenches by μ-Raman spectrometer and compared to that of the native material; Fig. 15.7 represents the Raman spectra. As illustrated, no significant changes to the general features of the Raman spectrum are observed when processing the PDMS with UV light (see Fig. 15.7A). However, the intensity of the peaks corresponding to the SiaOaSi, SiaC, and methyl stretching bonds gradually decreases with the laser fluence and number of the pulses, which is consistent with the findings of Atanasov et al. (2014) and

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CHAPTER 15 Laser processing of biopolymers

(A)

(B)

60 50

Trench depth, μm

Trench depth, μm

16

12

8

4

0.8 1.2

1.6 2.0 2.4 2.8

40 30 20 10

3.2 3.6 4.0

0.8 1.2 1.6 2.0 2.4 2.8 3.2 3.6 4.0 4.4

Laser fluence, J/cm2

Laser fluence, J/cm2

FIGURE 15.6 Dependence of the trench depth on the laser fluence for UV (dots), VIS (triangles), and NIR (quadrates) laser light, respectively. (A) 33 overlapping pulses; scanning speed 95 μm/s; (B) 83 overlapping pulses; scanning speed 38 μm/s. Reprinted from Atanasov, P.A., Stankova, N.E., Nedyalkov, N.N., Fukata, N., Hirsch, D., Rauschenbach, B., et al., 2016a. Fs-laser processing of medical grade polydimethylsiloxane (PDMS), Appl. Surf. Sci., 374, 229 234.

(A)

(B) Intensity, Arb. units

Intensity, Arb. units

3500 3000 2500 2000 1500

2500

c-Si 2000 1500 1000

1000 400

600

800 1000 1200 1400 1600 1800

400

600 800 1000 1200 1400 1600 1800

Raman shift (cm–1)

Raman shift (cm–1)

(C)

(D)

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32,000

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FIGURE 15.7 μ-Raman spectra taken from the bottom of the tracks produced by fs-laser 1: (A) UV; (B) VIS; (C) NIR laser light. (D) The Raman spectrum of the nonprocessed PDMS sample. The spectra correspond to the cases given in Fig. 15.5C. The inset in (C) illustrates the appearance of two amorphous carbon G and D bands. Reprinted from Atanasov, P.A., Stankova, N.E., Nedyalkov, N.N., Fukata, N., Hirsch, D., Rauschenbach, B., et al., 2016a. Fs-laser processing of medical grade polydimethylsiloxane (PDMS), Appl. Surf. Sci., 374, 229 234.

15.5 Fs-Laser Nanostructuring

Stankova et al. (2015). Instead, in the case of VIS or NIR laser processing a sharp peak between 512 and 518 cm21, ascribed to single crystalline silicon (c-Si) (Palma et al., 1999), appears in the Raman spectra (see Fig. 15.7B and C). The intensity of such peak increases with the number of subsequent overlapping pulses and laser fluence with respect to the decrease of the intensity of the peaks of the native material. This could be explained by breaking the corresponding chemical bonds and chemical transformation of the material caused by laser processing. Additionally, with the increase of the laser fluence and number of the overlapping NIR pulses, a strong carbonization appears (see the inset in Fig. 15.7C) accompanied by adjacent cracks of the material in the surrounding areas. The broad peaks between 1340 and 1360 cm21 and in the wavenumber range 1570 1600 cm21 are related to the D (breathing mode) and G (stretching mode) bands, respectively, of the sp2 bonds in amorphous carbon. They reveal the formation of highly disordered graphite structures. This decomposition of the organic material by breaking its chemical bonds and transformation and production of inorganic species (Si and C) as a result of the laser treatment is the base of the next technological operation: metalization of the trenches. Using fs-laser 2 with shorter pulse duration, that is, 35 fs (see Table 15.2), some specific features can be mentioned. In order to get well-defined trenches, a higher number of pulses (more than 2 orders of magnitude) compared to the case of 300 fs-laser are needed. Fig. 15.8A D depict some of the trenches’ profiles produced by 35 fs-laser at four different wavelengths and laser parameters (energy per pulse and number of subsequent pulses) viewed using a laser microscope. As illustrated, the width and the depth of the trenches depend on the laser processing parameters, that is, wavelength, fluence, and number of subsequent pulses. The best quality is obtained for all wavelengths applied at higher numbers of subsequent pulses. The lowest laser energy required for producing ablation

FIGURE 15.8 Profiles of the trenches produced by 35 fs laser: (A) UV (266 nm) laser light, 6 μJ and 1 3 104 subsequent laser pulses; (B) UV (355 nm), 6 μJ and 5 3 103 pulses; (C) VIS (532 nm), 25 μJ and 8 3 102 pulses; and (D) NIR (1064 nm), 14 μJ and 1 3 104 pulses.

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(ablation threshold) and formation of the well-defined trenches is obtained in case of UV lights, that is, 266 and 355 nm, respectively. This is predictable since the optical absorption of the material is highest in UV and decreases rapidly with wavelength increase (Stankova et al., 2016). It is worth noting in case of VIS irradiation when the pulse energy is at the level of the ablation threshold for PDMS or slightly above it, instead of producing a trench or ablate at the very central area, a swelling of the material is produced; see Fig. 15.9. This effect is also more pronounced if the number of subsequent laser pulses is relatively low (1024103). Fig. 15.10 presents some of the Raman spectra taken at the bottom of the trenches produced by VIS light of the fs-laser 2 at different numbers of subsequent pulses (curves 2 and 3). The spectrum of pristine PDMS is also presented for reference (curve 1). It is obvious that the intensities of the common peaks of PDMS decrease with the number of pulses, which is consistent with our previous investigations. Moreover, a photochemical transformation of the material occurs with the increase of the number of pulses applied. A second peak between 515 and 518 cm21 appears and can be attributed to single and/or polycrystalline silicon crystallites (c-Si) (Palma et al., 1999). This is due to the reduction and breaking of the Si-O-Si bonds and formation of Si crystallites. Generally, its intensity increases with increasing the laser fluence applied and becomes more dominant with the rise of the number of pulses, while the intensity of the 488 cm21 (SiaOaSi bonds) peak decreases. In addition, no visible wide peaks at

FIGURE 15.9 Trenches and their profiles produced by: (A) VIS (532 nm), 15 μJ and 1.4 3 103 pulses; and (B) VIS (532 nm), 20 μJ and 8 3 102 pulses.

15.6 Ps-Laser Processing

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FIGURE 15.10 μ-Raman spectra taken at the bottom of the trenches produced by fs-laser 2 (λ 5 532 nm). Laser energy per pulse is 25 μJ and the number of subsequent pulses is: 800—curve 2 and 2 3 1000—curve 3. Spectrum 1 is taken from the native PDMS. The inset depicts a more detailed curve 3.

1330 1336 cm21 and 1590 1615 cm21, respectively, can be observed in the Raman spectra. It is obvious that the energy delivered combined with very short fs-laser pulses cause preferentially breaking of the bonds with higher binding energy, whereas carbonization into the trenches is not observed. All these results are consistent with our previous findings (Atanasov et al., 2016a).

15.6 PS-LASER PROCESSING Preliminary experiments are accomplished via ps-laser. Fig. 15.11 depicts the LM view and profile of the track produced by ps-laser and μ-Raman spectra taken at three different points. As one can see, three zones are formed: c—ablation trench at the center; m—swelling zone, where the material forms hills from both sides of the trench; and n—slightly affected area in which the material does not change its properties. Both zones c and m are totally carbonized, that is, no other characteristic PDMS peaks are seen with the exception of wide and strong peaks at 1330 and 1580 cm21. These two peaks correspond to D and G modes and are attributed to the amorphous carbon, containing sp2 and sp3 bonds (Stankova et al., 2015, 2016). It is known that amorphous carbon is composed by a mixture of sp3, sp2, and sp1 sites. Obviously complete carbonization occurs. No other characteristic for PDMS peaks are seen (Fig. 15.11B).

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FIGURE 15.11 LM view and profile of the trench produced by UV ps-laser (355 nm), scanning speed— 0.1 mm/s and 5 3 105 pulses (A). (B) μ-Raman spectra taken at the three different zones: n, nonprocessed PDMS; m, swelling zone; and c, at the center of the trench.

15.7 COMPARISON BETWEEN FS- AND NS-LASER PROCESSING Fs-laser 1 and ns-laser processing of the PDMS-elastomer surface in air was investigated and the results were compared. The ablated trench depths depend on the overlapping of the consecutive pulses applied and are a function of the laser energy density as well. The ablation depth is evaluated by measuring and averaging five different points at the bottom of the trench with respect to the nonprocessed surface. Fig. 15.12A and B show the ablation depth for ns- and fs-laser processing as a function of the pulse numbers at a given fluence. As is illustrated, for all cases the ablation depth increases almost linearly with the number of pulses applied and has some tendency to saturate. Moreover, the ablation depth obtained in the case of ns-laser processing (Fig. 15.12A) is much higher compared to this one obtained at the fs-laser processing (Fig. 15.12B). The chemical composition in the trenches of the laser-processed PDMS surface is investigated via μ-Raman spectroscopy. The Raman spectra are taken and averaged from five different points at the bottom of each track (see Figs. 15.13A C and 15.14).

15.7 Comparison Between Fs- and Ns-Laser Processing

FIGURE 15.12 Ablation depth of the trenches produced by the number of subsequent pulses applied at different fluences: (A) ns-laser processing; and (B) fs-laser processing. Filled symbols are for UV light and open symbols for VIS light, respectively. Reprinted from Stankova, N.E., Atanasov, P.A., Nedyalkov, N.N., Stoyanchov, T.R., Kolev, K.N., Valova, E.I., et al., 2015. Fs- and ns-laser processing of polydimethylsiloxane (PDMS) elastomer: comparative study, Appl. Surf. Sci., 336, 321 328.

Significant differences in the Raman spectra of the UV, VIS, and NIR ns-laser-treated areas are observed compared with the native (untreated) PDMS surface (Fig. 15.13). Strong and sharp peaks between 513 and 518 cm21 appear at all processing parameters applied and are recognized as Si crystallites (caSi) and ascribed to mono and/or polycrystalline silicon (Palma et al., 1999). Its intensity rises considerably with the laser fluence and the number of pulses for all laser wavelengths used. While, the intensity of the SiaOaSi peak at 488 cm21, which is typical for the native PDMS sharply decreases. Therefore, the increase of the c-Si peak is most probably related to the corresponding breaking and, thus, reduction of the SiaOaSi bonds. A considerable drop of the intensities of the peaks at

505

FIGURE 15.13 Changes of the structure of the μ-Raman spectra caused by ns-laser processing: (A) λ 5 266 nm, (B) λ 5 532 nm, and (C) λ 5 1064 nm. The inset shows spectra obtained from some sections of the laser tracks at both wavelengths applied, respectively. The laser fluence applied is 4.3 J/cm2 in case of UV light, 4.0 J/cm2 in case of visible light and 13.0 J/cm2 in case of NIR light, respectively. Reprinted from Stankova, N.E., Atanasov, P.A., Nedyalkov, N.N., Stoyanchov, T.R., Kolev, K.N., Valova, E.I., et al., 2015. Fs- and ns-laser processing of polydimethylsiloxane (PDMS) elastomer: comparative study, Appl. Surf. Sci., 336, 321 328; Stankova, N.E., Atanasov, P.A., Nikov, R.G., Nikov, R.G., Nedyalkov, N.N., Stoyanchov, T.R., et al., 2016. Optical properties of polydimethylsiloxane (PDMS) during nanosecond laser processing, Appl. Surf. Sci., 374, 96 103.

15.7 Comparison Between Fs- and Ns-Laser Processing

FIGURE 15.14 μ-Raman spectra of native and fs-laser-processed PDMS-elastomer. Laser fluence is 3.9 J/cm2 in both cases and 83 subsequent laser pulses are applied. Reprinted from Stankova, N.E., Atanasov, P.A., Nedyalkov, N.N., Stoyanchov, T.R., Kolev, K.N., Valova, E.I., et al., 2015. Fs- and ns-laser processing of polydimethylsiloxane (PDMS) elastomer: comparative study, Appl. Surf. Sci., 336, 321 328.

685, 709, 787, and 859 cm21 with respect to the peaks corresponding to the native material is also observed in the Raman spectra at all parameter conditions applied under UV, VIS, and NIR irradiation. This could be due to the breaking of the SiaCH3 and SiaC bonds and also contributes to the formation of Si crystallites. Whereas under the fs-laser irradiation, this peak appears only in cases of VIS (Fig. 15.14) and NIR (Fig. 15.7C) wavelengths used (527 and 1055 nm) and is not as prominent. The Raman spectra, measured after UV, VIS, and NIR fs-laser treatment of the PDMS polymer reveal that the intensities corresponding to the SiaOaSi, SiaC, and SiaCH3 bonds also gradually decrease with enhancement of the laser fluence and the number of pulses. It worth noting that weak broad Raman band appears between 940 and 960 cm21 after ns-laser processing. Its intensity tends to increase with an increasing energy dose delivered in the laser-treated area, namely with increasing of the number of pulses or the laser energy (Fig. 15.13). This spectral band can be attributed to the microcrystalline SiaC bond (Ward et al., 2004). Furthermore, two new broad and quite pronounced peaks in the ranges of 1328 1345 and 1570 1615 cm21 are observed in the Raman spectra after ns- and fs-laser processing under certain parameter conditions. They are assigned to the D (breathing mode) and G (stretching mode) bands of amorphous carbon. However, in the case of fs-laser irradiation these two peaks appear only at NIR wavelength (Fig. 15.7C), at highest fluence, and higher number of overlapping pulses.

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In the case of ns-laser irradiation these two peaks are primarily observed at higher laser energy delivered by higher fluence and/or higher number of pulses under UV and VIS wavelengths irradiation. At NIR wavelength these peaks are much less expressed since the polymer is drilled with increasing laser energy delivered in the laser-processed area. It is worth noting that in the Raman spectra obtained from some sections of the laser tracks only D and G band peaks appeared at 1328 1336 and 1570 1580 cm21, respectively. Their signal is highly intensive and hinders the other Raman features. When the peaks appear simultaneously with the other Raman peaks (see Fig. 15.13A and B), their intensity decreases significantly and the position of the D band is slightly shifted toward higher wave numbers 1336 1342 and 1590 1600 cm21 (inset of Fig. 15.13B). The nature of the D and G vibration modes is complicated and depends highly on the laser wavelength and laser energy of the resonant excitation. The origin of the D peak is still a much-debated issue and the key point to understanding the D peak is screening the electron phonon interaction of graphite and amorphous carbon stages. The investigation showed that in perfect graphite this mode is forbidden and activates only in the presence of disorder. In case of amorphous carbon the appearance of D peak is related with structure disordering. It is well-known that amorphous carbon can have any mixture of sp3, sp2, and sp1 sites. But visible Raman spectroscopy is more sensitive to the sp2 sites since the cross section of the visible excitation is much higher for them (about 50 230 times) than this one for the sp3 sites. It is worth noting that the presence and the positions of the D (1328 1336 cm21) and G (1570 1580 cm21) peaks in the Raman spectra of the laser-treated PDMS polymer could be related to the presence of sp3 sites, which is characteristic to the diamond phase (Ferrari and Robertson, 2004; Ferrari, 2007). Finally, it could be concluded that the appearance of inorganic products like silicon and carbon is certain proof for actual chemical activation of the laserprocessed surface. It seems that the surface chemical transformation, that is, silicone decomposition, has a much more complex character of the laser fluence, pulse duration, and number of overlapping pulses than only of the wavelength. Obviously, the effective absorption of the PDMS-elastomer is high enough than that measured at UV, VIS, and NIR wavelengths (for native PDMS) and increases with increasing the laser energy delivered. Therefore, incubation occurs during the irradiation with every consecutive pulse with respect to the initial laser pulse and thus increases the absorption coefficient of the polymer. Most probably, both photothermal and photochemical reactions with ratios could contribute to the ablation mechanism depending on the laser beam parameters (fluence, pulse duration, and wavelength) as well as the polymer properties. The results of the investigation of the surface modifications due to micromachining of PDMS-elastomer by fs- (λ 5 263, 527, and 1055 nm) and ns-laser (266, 355, 532, and 1064 nm) irradiation can be summarized as follows. It was found by μ-Raman characterization that chemical activation by decomposition of

15.7 Comparison Between Fs- and Ns-Laser Processing

the silicone is induced after fs- and ns-laser treatment of the surface. It results in a strong decrease of the intensities peaks at 488, 685, 709, 787, 859, 2909, and 2970 cm21 which correspond to the SiaOaSi stretching, Si-CH3 symmetric rocking, SiaC symmetric stretching, CH3 asymmetric rocking 1 SiaC asymmetric stretching, CH3 symmetric rocking, CH3 symmetric, and CH3 asymmetric and stretching modes, respectively. Several different important features are evidenced, such as: 1. Si crystallites are formed into tracks fabricated by ns-laser treatment (λ 5 266, 355, 532, and 1064 nm), while in the case of fs-laser processing this is observed only at VIS and NIR wavelengths (527 and 1055 nm); 2. During the ns-laser and NIR fs-laser activation of the surface, amorphous carbon containing sp2 and sp3 sites are formed. Since the Si crystalline peak becomes more prominent, accordingly the intensity of the carbon bands decrease. The local chemical transformations are a complex function of the laser fluence, pulse duration, and the number of the pulses (incubation) rather than only of the wavelength. Optical microscopy and SEM reveal the following peculiarities: 1. The incubation process strart occuring below the PDMS surface after several pulses emitted on the laser-treated area without changing the properties of the surrounding bulk material. As a result, with increasing the laser ernergy and the number of pulses delivered, respectively, significant decrease of the optical transmittance of the processed surface is observed. 2. All the processed areas show a high-roughness morphology with the presence of pores and cavities of different sizes, which are much larger in the case of ns-laser and VIS and NIR fess-laser treatment. Whereas the UV fess-lasertreated surface is much more regular and uniform. 3. All laser-treated samples show the presence of nanometer-scale grains. The increase of the number of overlapping pulses leads to a more pronounced structure seemingly due to the melting and solidification of the material. Regarding the changes of the chemical composition of PDMS, that is, Si and C formation in the laser-processed area by higher fluence and larger number of overlapping pulses, a XRD at angle of 1 degree grazing incidence study is accomplished. For this purpose, continuous trenches with dimensions of 12 3 3 mm2 and different depths are prepared. The results from the XRD analyses are presented in Fig. 15.15. Fig. 15.15A depicting the XRD spectrum of the nonprocessed PDMS. The spectrum is similar to that reported by Ferreira et al. (2013) and Atanasov et al. (2016b). As illustrated in Fig. 15.15B, all the diffraction patterns are similar and consist of two halos: a first larger one, which is located at 12.3 degrees; and a second, smaller and broader one at 22.2 degrees. These results reveal that the microstructure of the samples is amorphous and the ns-laser processing does not change the properties of the basic material. Additionally, the peak of the material treated by

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CHAPTER 15 Laser processing of biopolymers

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FIGURE 15.15 XRD spectra of the (A) unprocessed; and (B) ns-laser-processed area of PDMS. Reprinted from Atanasov, P.A., Stankova, N.E., Nedyalkov, N.N., Stoyanchov, T.R., Nikov, Ru.G., Fukata, N., et al., 2016b. Properties of ns-laser processed polydimethylsiloxane (PDMS), J of Phys.: Conf. Ser., 700, 012023, 1 5.

UV (λ 5 355 nm) ns-laser pulses has about twice as high intensity compared to the other wavelengths applied, which is probably related to the better ordering of the material processed. More investigations are needed in this line.

15.8 XPS STUDY OF NS-LASER PROCESSING OF POLYDIMETHYLSILOXANE A detailed study of the influence of the laser irradiation on PDMS using XPS was presented by Armyanov et al. (2015). In order to enable XPS investigations, ns-laser (Table 15.2) is used to treat PDMS-elastomer sheets (medical grade, KCC-corporation, South Korea) (800 μm thick). UV, VIS and NIR wavelengths are applied. Continuous wider trenches (B500 μm) are formed by the moving velocity of the x y table and applying a number of the consecutive pulses (from 11 to 110) per laser beam spot on the material. In addition to the careful calibration, the condition to obtain reliable XPS data is to establish a line on an absolute basis enabling the rest of the lines to be positioned with respect to that one. For this purpose in the case of PDMS the carbon peak C 1s is adjusted at a binding energy of 284.38 eV (Beamson and Briggs, 2012; Louette et al., 2005), as it is shown in Fig. 15.16A. Thus, one can clearly reveal the differences in Si 2p and O 1s positions in Fig. 15.16B and C, respectively, whenever they appear. The peaks presented in Fig. 15.16 are symmetrical. However, the peak’s height and full width at half maximum (FWHM) varies in each case due to

FIGURE 15.16 Comparisons of XPS of PDMS treated by ns-laser radiation at different wavelength (in nm) and pulses number (p): (A) C 1s; (B) Si 2p; and (C) O 1s. Reprinted from Armyanov, S., Stankova, N.E., Atanasov, P.A., Valova, E., Kolev, K., Georgieva, J., et al., 2015. XPS and μ-Raman study of nanosecond-laser processing of polydimethylsiloxane (PDMS), Nucl. Instr. Meth. Phys. Res. B, 360, 30 35.

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geometrical reasons: shape, morphology, and depth of the trenches. This is due to the treatment regime (wavelength and number of pulses), specifying the ablation differences in areas, as demonstrated. The O 1s and Si 2p peaks of the sample treated by NIR irradiation 1064 nm and 66 pulses exhibit the highest shifts and they are suitable to assess the oxidation degree of silicon. The different degrees of silicon oxidation are demonstrated in Fig. 15.17 and Table 15.4, obtained by the deconvolution of Si 2p peak. For pure silicon a peak at 99.5 eV, corresponding to Si( OH)x, could appear (Alam et al., 2013) which, however, cannot be identified in Figs. 15.16 and 15.17. It looks like the chemical status of Si in PDMS is determined to a great extent first by its position in the polymer chain and, during the oxidation process, by its bonding with oxygen. Then, a very low influence of H through O on the BE of Si 2p could be displayed. A similar supposition was made for the absence of OH related peaks in Fourier Transform Infrared Spectroscopy of treated PDMS (Holgerson et al., 2005). Free SiOH could be quantitated by eliminating the water and siloxane interferences in a siloxane band, and applying dissolution of PDMS in carbon tetrachloride (Griffith, 1984).

FIGURE 15.17 Curve fitting of XPS spectra of Si 2p for treatment by ns-laser: λ 5 1064 nm, 66 pulses, and 6.5 J/cm2. The red line displays the original peak and the dashed line is the fitting curve. The orange line (1) corresponds to Si( O)4 103.3 eV; the green line (2) corresponds to Si( O)3 102.6 eV; and the blue line (3) corresponds to Si( O)2 101.8 eV, respectively. The areas under the curves are as follows (Table 15.5): Si( O)4, 43.2% Si ( O)3, 46.8 %, Si( O)2 10.0%. Reprinted from Armyanov, S., Stankova, N.E., Atanasov, P.A., Valova, E., Kolev, K., Georgieva, J., et al., 2015. XPS and μ-Raman study of nanosecond-laser processing of polydimethylsiloxane (PDMS), Nucl. Instr. Meth. Phys. Res. B, 360, 30 35.

15.8 XPS Study of Ns-Laser Processing of Polydimethylsiloxane

Table 15.4 Silicon Chemical Environments and the Corresponding Si 2p BE (see also Beamson and Briggs, 2012; Louette et al., 2005; Alexander et al., 1999; O’Hare et al., 2004; Wagner et al., 1979) Structure

Abbreviation Experimental BE (eV) Reference BE (eV) Calculated from fitting in Fig. 15.17 (%)

Si( O)4 103.3 103.3 43.2

Si( O)3 102.6 102.67 46.8

Si( O)2 101.8 101.79 10.0

Reprinted from Armyanov, S., Stankova, N.E., Atanasov, P.A., Valova, E., Kolev, K., Georgieva, J., et al., 2015. XPS and μ-Raman study of nanosecond-laser processing of polydimethylsiloxane (PDMS), Nucl. Instr. Meth. Phys. Res. B, 360, 30 35.

FIGURE 15.18 Curve fitting of XPS spectra of O 1s for treatment by ns-laser λ 5 1064 nm, 66 pulses, 6.5 J/cm2. The red line displays the original peak, the dashed line is the fitting curve. The areas under the curves (Table 15.6): orange (1)—Si( O)4 36.9%; green (2)—Si( O)3 44.6%; purple (3)—Si( O)2 14.4%; and dark orange (4)—Si( OH) 4%. Reprinted from Armyanov, S., Stankova, N.E., Atanasov, P.A., Valova, E., Kolev, K., Georgieva, J., et al., 2015. XPS and μ-Raman study of nanosecond-laser processing of polydimethylsiloxane (PDMS), Nucl. Instr. Meth. Phys. Res. B, 360, 30 35.

The curve fitting of the O 1s peak is presented in Fig. 15.18 and Table 15.5. For SiOH in as-received pure Si, BE is about 530.4 eV (Alexander et al., 1999). When Si is treated under humid conditions, the BE of Si(OH)x is about 1 eV higher. (Alexander et al., 1999 and references therein). For substances different from PDMS, the BE of O 1s is reported as follows: 532.20 eV for SiO1.8 (Pitts et al., 1986), 532.6 eV for SiO1.9 (Chao et al., 1986), and 533.20 eV for SiO2

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Table 15.5 Oxygen Chemical Environments and the Corresponding O 1s BE (see also Lippert, 2005; Hofmann, 2013; Beamson and Briggs, 2012; Louette et al., 2005; O’Connor, 2012) Structure

Abbreviation Experimental BE (eV) Calculated from fitting in Fig. 15.17 (%)

Si( O)4 533.0 36.9

Si( O)3 532.4 44.6

Si( O)2 531.6 14.4

Si( OH) 530.6 4.0

Reprinted from Armyanov, S., Stankova, N.E., Atanasov, P.A., Valova, E., Kolev, K., Georgieva, J., et al., 2015. XPS and μ-Raman study of nanosecond-laser processing of polydimethylsiloxane (PDMS), Nucl. Instr. Meth. Phys. Res. B, 360, 30 35.

(Finster et al., 1990). Thus, when the oxidation degree of Si in PDMS is increased, the peak shift to higher values of BE can be expected. The BE of the peaks after the deconvolution shown in Fig. 15.18 and Table 15.5 are close to those mentioned for the oxidation forms of the pure Si. Comparison of the areas under the three peaks in Fig. 15.17 and Table 15.4 with the respective three peaks in Fig. 15.18 and Table 15.5 shows that they are of the same order of magnitude. The presence of OH groups on the surface is responsible for the hydrophilic behavior of the laser-treated areas, being one of the major factors favoring electroless metalization. However, the photochemical conversion of surface methylsilane groups (Si CH3) to silanol groups (SiaOH) on the PDMS surface is responsible for the large increase in surface-free energy. Thus, after laser irradiation with time a subsequent degradation of the polymer and formation of SiOx was observed (Graubner et al., 2004). The detailed results of this degradation are provided in Figs. 15.17 and 15.18. The decreasing number of silanol groups on the PDMS surface is the reason for continuous decreasing of hydrophilicity and its rejuvenation to enable the metalization by electroless plating is a challenging task. A possible recovery of the surface after the laser ablation during 1-month lapse between the laser treatment and XPS measurements cannot be excluded. Similar phenomena after plasma treatment of PDMS were observed (Donzel et al., 2001). In Fig. 15.19 the μ-Raman spectra of the laser processed area with wavelength of 1064 nm and 66 pulses is compared with the pristine (native) PDMS. It is seen that in addition to already discussed peaks - D and G bands of the carbon, the sharpest peak of crystalline Si appears. It is worth to remind the difference in the sampling depth of x-ray photoelectron and μ-Raman spectroscopic measurements. XPS is surface-sensitive technique, which measures the elemental composition up to several nanometers in depth, whereas by the μ-Raman spectroscopy the information of the sample properties is received from several and more microns in depth. This is the reason for displaying crystalline Si inside the ablated area and Si oxides on the top of it.

15.9 Electroless Metallization Directly After the Laser Treatment

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FIGURE 15.19 μ-Raman spectra of pristine and ns-laser-processed PDMS-elastomer. Irradiation with λ 5 1064 nm, 66 pulses and laser fluence of 6.4 J/cm2 is applied. Reprinted from Armyanov, S., Stankova, N.E., Atanasov, P.A., Valova, E., Kolev, K., Georgieva, J., et al., 2015. XPS and μ-Raman study of nanosecond-laser processing of polydimethylsiloxane (PDMS), Nucl. Instr. Meth. Phys. Res. B, 360, 30 35.

15.9 ELECTROLESS METALLIZATION DIRECTLY AFTER THE LASER TREATMENT UV and VIS ns- and fs-laser surface processed PDMS films are successfully electroless metalized with Ni or Pt (Atanasov et al., 2014, 2016a; Armyanov et al., 2018; Stankova et al., 2015, 2018). The electroless deposition of Ni or Pt in the tracks is performed, excluding sensibilization and chemical activation usually preceding this process. In this way, a selective metalization is performed without application of masks or external templates. It is worth noting that this process was successfully accomplished for the first time after VIS laser treatment of the PDMS polymer. Strikingly, it was found that when both fs- or ns-laser were applied, the time interval between laser treatment and metalization was not a critical process parameter. After the laser treatment, the PDMS samples are prepared for metalization by rinsing with deionized water and air stream drying. Cleaning is an important step to avoid quick spontaneous decomposition of the electroless solution. Then electroless deposition of Pt or Ni is applied to the laser-processed tracks. The processes of sensitization by tin and activation by palladium, which usually precedes the electroless deposition of Pt or Ni on non-conducting materials, here are not applied. Hydrazine hydrate is used as a reducer. The pH of the plating baths is B12 and the deposition temperatures are 70 C for Pt and 80 C for Ni, respectively. The deposition lasts 25 minutes in a thermostatic vessel under rigorous

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stirring. After metalization, the samples are washed with deionized water and dried. It is worth noting that the Ni metallization is the best and cheaper analog of the classic Pt electroless plating process, based on the hydrazine hydrate reducer. The as-described metalization, similar to that presented in brief in Atanasov et al. (2014) and Stankova et al. (2015) is much simpler and not as sensitive with respect to the time interval between laser treatment and metalization as proposed by Laude (2005) and Dupas-Bruzek et al. (2009a). The electroless metalization with Ni is applied to samples processed with fs-laser 1 and ns-laser-processed PDMS. Table 15.6 summarizes the results of the electroless metalization of the fs-laser 1 processed PDMS samples with Ni. Good quality and uniform metalization is achieved in the following cases: UV light—for all fluences and larger number of overlapping pulses (83); VIS light—for all fluences and larger number of overlapping pulses (83); and NIR light—practically for both numbers of overlapping pulses and all fluences. Fig. 15.20A C presents examples of typical high resolution SEM pictures of metalized trenches after UV, VIS, and NIR fs-laser treatment. Uniform Ni coating is also deposited on UV ns-laser-treated surface of PDMS elastomer. SEM pictures prove that Ni crystallizes in “pappus” globules-like structures (Fig. 15.21A and B). The EDS elements map and EDS composition spectrum investigation of the metalized track confirm the content of Ni in the track (Fig. 15.21C and D). Additionally, trenches produced by UV ns-laser are metalized with Pt using hydrazine as a reducer at different concentrations; see Fig. 15.22. It was found that the surface morphology depends strongly on the concentration of hydrazine. The size of the metal nanospikes is ranged from several tens to several hundreds of nanometers and they tend to grow vertically.

15.10 NS-LASER PROCESSING IN DIFFERENT ENVIRONMENTS During ns-laser processing debris on the both sides of the trenches are formed. The debris are also metalized during the electroless deposition, which leads to fast depletion of the metal in the authocatalitic bath. In order to obtain free of debris adjacent zones or to prevent production of debris new aproaches of ns-laser processing of PDMS are performed. One is to cover the PDMS surface with a thin film of PMMA before the laser ablation in air and the second one (Stankova et al., 2018 Patent Application) is to perform the ablation in water environment. SEM pictures of Pt-metalized tracks by electroless plating of the track produced by laser ablation of PDMS covered with PMMA thin film is presented in Fig. 15.23. Here ns-laser with wavelength of 355 nm is used for

15.10 Ns-Laser Processing in Different Environments

Table 15.6 List of the Successfully Electroless Metalized With Ni Trenches Produced by Fs-Laser 1 Processing Number of Pulses

Energy per Pulse (μJ)

Fluence (J/cm2)

33

172 6 5

2.2

138 6 5

1.76

59 6 3

0.75

6

79 6 4

1.0

7

125 6 6

1.5

56 6 3

0.7

13

101 6 4

1.3

14

142 6 5

1.8

307 6 9

3.9

16

204 6 6

2.6

17

119 6 4

1.5

18

64 6 3

0.8

125 6 9

1.6

21

209 6 13

2.7

22

315 6 5

4.0

No.

λ (nm)

Pulse Duration (fs)

1

263

300

2 83

5

12

15

20

527

1055

300

900

83

33

83

Metalization Ni in form of “teeth” and “flowers” Not continuous with Ni “flowers” Uniform. Large and small Ni “teeth” Uniform. Large and small Ni “teeth” Uniform. Less dispersion in the size of Ni “teeth” Not continuous with big Ni “flowers” Almost continuous and uniform with Ni “flowers” and large Ni “teeth” Almost continuous with large Ni “teeth” and some Ni flowers Uniform with large Ni “teeth” Almost uniform with big and small Ni “teeth” Almost uniform with large and small Ni “teeth” and some “flowers” Almost uniform with grains with Ni “teeth” with different shape Partially covered with Ni “teeth” in places Partially covered with Ni “teeth” in places Partially covered with Ni “teeth” in places

Reprinted from Atanasov, P.A., Stankova, N.E., Nedyalkov, N.N., Fukata, N., Hirsch, D., Rauschenbach, B., et al., 2016a. Fs-laser processing of medical grade polydimethylsiloxane (PDMS), Appl. Surf. Sci., 374, 229234.

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FIGURE 15.20 SEM pictures of the electroless Ni metalized samples registered at the bottom of the trenches produced by fs-laser 1. (A) UV laser pulses at a fluence of 1.5 J/cm2. (B) VIS laser pulses at a fluence of 0.7 J/cm2. (C) NIR laser pulses at a fluence of 2.7 J/cm2. In all cases, the images refer to 83 overlapping pulses (scanning speed of 38 μm/s). Reprinted from Atanasov, P.A., Stankova, N.E., Nedyalkov, N.N., Fukata, N., Hirsch, D., Rauschenbach, B., et al., 2016a. Fs-laser processing of medical grade polydimethylsiloxane (PDMS), Appl. Surf. Sci., 374, 229234.

ablation. After the laser processing and before the metalization the PMMA film is removed. The trenches produced after ablation with ns-laser with 266 nm in double distilled water are shown in Fig. 15.24A. For comparison, Fig. 15.24B presents trenches produced at similar conditions, but in an air environment. The energy density for processing of the polymer depends on the height of the water column above the surface of the PDMS. As illustrated, the trenches produced in water are of quite good quality and no debris are produced in the vicinity around them. After the laser processing successful electroless Pt metalization is made. Fig. 15.25 depicts the metalized trenches prepared at similar conditions like those provided in Fig. 15.24. The quality of the metalized trenches is dramatically improved after laser processing of PDMS in water; see Fig. 15.25A. High efficiency of electroless metalization is obtained since Pt cover only the laser-treated area. Whereas, at electroless metalization of trenches obtained after laser ablation in air, the Pt is deposited predominantly on the debris around the trenches; see Fig. 15.25B).

15.10 Ns-Laser Processing in Different Environments (A)

(B)

(C)

(D)

FIGURE 15.21 SEM pictures indicating the Ni metalized tracks by electroless plating (A) and (B). EDS elemental mapping (C) and EDS spectrum of metalized PDMS-elastomer (D). UV ns-laser processing—fluence of 4.3 J/cm2 and 77 pulses. Reprinted from Stankova, N.E., Atanasov, P.A., Nedyalkov, N.N., Stoyanchov, T.R., Kolev, K.N., Valova, E.I., et al., 2015. Fs- and ns-laser processing of polydimethylsiloxane (PDMS) elastomer: comparative study, Appl. Surf. Sci., 336, 321 328.

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FIGURE 15.22 SEM pictures indicating the Pt-metalized tracks by electroless deposition. The elastomer is processed by 355 nm (A and B) and 266 nm (C and D) ns-laser. The hydrazine reducer at different concentration is applied.

FIGURE 15.23 SEM pictures of Pt metalized tracks by electroless plating. The track is produced by laser ablation of PDMS covered with thin PMMA film. 355 nm ns-laser is used for ablation.

15.11 Conclusion and Perspectives for Future Investigations

FIGURE 15.24 SEM pictures of the trenches, produced by ns-pulses at wavelength of 266 nm: (A) ablation in distilled water; and (B) ablation in air.

FIGURE 15.25 SEM pictures of the Pt metalized trenches, which previously are produced by ns-laser ablation at wavelength of 266 nm: (A) in water; and (B) in air environment.

15.11 CONCLUSION AND PERSPECTIVES FOR FUTURE INVESTIGATIONS Our results show promising prospects to implement such methods of laser-based micro- or nanofabrication and electroless metal plating of PDMS-elastomer devices used in MEMS, NEMS, and MEAs having applications for implantable neural interfacing technologies for monitoring and/or stimulation of neural activity. Direct laser writing on the optically transparent PDMS elastomers could be performed successfully by short and ultra-short pulsed laser irradiation in the UV, VIS, and NIR spectra with suitable exposure parameters like laser fluence and number of pulses for producing of high-definition multielectrode arrays.

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ACKNOWLEDGMENTS Financial support of the National Science Fund of Bulgaria under the project T02/24 entitled “New advanced method for processing nanocomposite materials for the creation of microsystems for medical and high-tech applications” is highly acknowledged. We thank Ekaterina Yordanova and Georgi Yankov (Institute of Solid State Physics, Bulgarian Academy of Sciences, Sofia, Bulgaria) for processing of the PDMS elastomer via the laser system “Fs-laser 2” (pulse duration of 35 fs, repetition frequency of 1000 Hz); M. Zamfirescu and B. St. Calin (Center for Advanced Laser Technologies (CETAL), National Institute for Lasers, Plasma, and Radiation Physics (INFLPR), Magurele, Romania) for processing of the PDMS elastomer via the laser system “Ps-laser” (pulse duration of 7 ps, repetition frequency of 500 kHz); Anastas Nikolov (Institute of Electronics, Bulgarian Academy of Sciences, Sofia, Bulgaria) for consultation regarding ns-laser ablation in water environment. We thank all the researchers-coauthors in the corresponding papers listed in the “References,” for laser processing of the PDMS elastomer via the laser systems “Ns-laser” and “Fs-laser 1,” for performing of the μ-Raman spectroscopy, SEM, XPS, XRD, STEM and SAED analyses, and for the Color 3D Laser Scanning Microscope measurements.

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References

Okoshi, M., Kuramatsu, M., Inoue, N., 2002a. Thin-film deposition by laser ablation of dimethylpolysiloxane. Appl. Surf. Sci. 197 198, 772 776. Okoshi, M., Kuramatsu, M., Inoue, N., 2002b. Pulsed laser deposition of SiO2 thin films with dimethylpolysiloxane targets. Jpn. J. Appl. Phys., Part 1 41, 1395 1399. Okoshi, M., Kimura, T., Takao, H., Inoue, N., Yamashita, T., 2004. Photochemical modification of silicone films using F2 laser for selective chemical etching. Jpn. J. Appl. Phys., Part 1 43 (6A), 3438 3442. Okoshi, M., Li, J., Herman, P.R., 2005. 157 nm F2-laser writing of silica optical waveguides in silicone rubber. Opt. Lett. 30, 2730 2732. Okoshi, M., Sekine, D., Inoue, N., Yamashita, T., 2007a. White-light emission from silicon rubber modified by 193 nm ArF excimer laser. Jpn. J. Appl. Phys., Part 2 46 (15), L356 L358. Okoshi, M., Cho, J., Inoue, N., 2007b. F2 laser photochemical welding of aligned silica microsphere to silicone rubber. Jpn. J. Appl. Phys., Part 1 46 (4A), 1516 1520. Okoshi, M., Iyono, M., Inoue, N., 2009a. Controllable change of photoluminescence spectra of silicone rubber modified by 193 nm ArF excimer laser. Jpn. J. Appl. Phys., Part 1 48, 12. 122503-1-5. Okoshi, M., Iyono, M., Inoue, N., 2009b. Photochemical surface modification of silicone rubber into photoluminescent material by 193 nm ArF excimer laser radiation. Jpn. J. Appl. Phys., Part 1 48, 10. 102301-1-7. Okoshi, M., Iyono, M., Inoue, N., Yamashita, T., 2009c. Photochemical welding of silica microsphere to silicone rubber by 193 nm ArF excimer laser. Appl. Surf. Sci. 255 (24), 9796 9799. Palma, C., Rossi, M.C., Sapia, C., Bemporad, E., 1999. Laser-induced crystallization of amorphous silicon-carbon alloys studied by Raman microscopy. Appl. Surf. Sci. 138 139, 24 28. Pitts, J.R., Thomas, T.M., Czanderna, A.W., Passler, M., 1986. XPS and ISS of submonolayer coverage of Ag on SiO2. Appl. Surf. Sci. 26, 107 120. Qina, Y., Howladera, M.M.R., Deena, M.J., Haddaraa, Y.M., Selvaganapathy, P.R., 2014. Polymer integration for packaging of implantable sensors. Sens. Actuators B 202, 758 778. Schnyder, B., Lippert, T., Koetz, R., Wokaun, A., Graubner, V.-M., Nuyken, O., 2003. Uv irradiation induced modification of PDMS films investigated by XPS and spectroscopic ellipsometry. Surf. Sci. 532 535, 1067 1071. Stankova, N.E., Atanasov, P.A., Nedyalkov, N.N., Stoyanchov, T.R., Kolev, K.N., Valova, E.I., et al., 2015. Fs- and ns-laser processing of polydimethylsiloxane (PDMS) elastomer: comparative study. Appl. Surf. Sci. 336, 321 328. Stankova, N.E., Atanasov, P.A., Nikov, R.G., Nikov, R.G., Nedyalkov, N.N., Stoyanchov, T.R., et al., 2016. Optical properties of polydimethylsiloxane (PDMS) during nanosecond laser processing. Appl. Surf. Sci. 374, 96 103. Stankova, P.A., Atanasov, N.N., Nedyalkov, Dr. Tatchev, K.N., Kolev, E.I., Valova, St. A., et al. 2018. Laser-induced surface modification of biopolymers micro/nanostructuring and functionalization, IOP Conf. Series: J. Physics: Conf. Series, 992, 012051. Stankova, N.E., Nikolov, A.S., Atanasov, P.A., Nedyalkov, N.N., 2018, Patent Application. Incoming number № 112728/03.05.2018, Method for structuring and activating of polymers with laser pulses.

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Takao, H., Okoshi, M., Miyagami, H., Inoue, N., 2004a. Fabrication of SiO2 microlenses on silicone rubber using a vacuum-ultraviolet F2 laser. IEEE J. Sel. Top. Quantum Electron. 10, 1426 1429. Takao, H., Okoshi, M., Inoue, N., 2004b. Swelling and modification of silicone surface by F2 laser irradiation. Appl. Phys. A 79, 1571 1576. Takao, H., Miyagami, H., Okoshi, M., Inoue, N., 2005. Microlenses fabricated on silicone rubber using F2 Laser, Jpn. J. Appl. Phys. Part 1 44 (4R), 1808 1811. Urech, L., Lippert, T., 2007. Designed polymers for ablation. In: Phipps, C. (Ed.), Laser ablation and its applications. Springer series in optical sciences, Springer, Boston, MA. 129, 281 297. Vladkova, T.G., Keranov, I.L., Dineff, P.D., Youroukov, S.Y., Avramova, I.A., Krasteva, N., et al., 2005. Plasma based Ar 1 beam assisted poly(dimethylsiloxane) surface modification. Nuclear Instrum. Methods Phys. Res. B 236, 552 562. Wagner, C.D., Riggs, W.M., Davis, L.E., Moulder, J.F., Muilenberg, G.E. (Eds.), 1979. Handbook of X-Ray Photoelectron Spectroscopy, A Reference Book of Standard Data For Use In X-Ray Photoelectron Spectroscopy. Perkin-Elmer Corporation, Physical Electronics Division 6509 Flying Cloud Drive Eden Prairie, Minnesota, USA. Ward, Y., Young, R.J., Shatwell, R.A., 2004. Application of Raman microscopy to the analysis of silicon carbide monofilaments. J. Mater. Sci. 39 (22), 6781 6790.

FURTHER READING Brochure of the Company Gelest, Silicon Compounds: Silanes and Silicones, 2004. Gelest, Inc., Morrisville, PA, 560. Ferrari, A.C., Robertson, J., 2001. Origin of the 1150 cm21 Raman mode in nanocrystalline diamond. Phys. Rev. B 63, 121405R. Takao, H., Okoshi, M., Inoue, N., 2002. Fabrication of SiO2-humps on silicone rubber using F2 laser. Jpn. J. Appl. Phys., Part 2 41 (10A), L1088 L1089.

Index Note: Page numbers followed by “f” and “t” refer to figures and tables, respectively.

A Absorption tomography, 8 Additive manufacturing (AM), 235 polyurethane-based structures obtained by. See Polyurethane-based structures African trypanosomiasis, 183 Agarose hydrogel, 43 Alginate, 403 417 characteristics of, 404t Alginate hydrogel, 43 44 Alkaline phosphatase (ALP), 377 Allografts, 441 α-tricalcium phosphate (α-TCP), 373t Alzheimer’s disease (AD) PLGA based therapies for, 179 Anticancer therapy poly(lactide-co-glycolide) (PLGA) scaffolds, 171 174 Antigen presenting cells (APCs), 181 Apatite, 447 448 Aromatic diisocyanates, 238 Arteriovenous loop, 26 Atomic force microscopy (AFM), 61 Autografts, 441 Autologous cells, 2 Azotobacter vinelandii, 221 223, 227 228

B Bacillus smithii, 127 Bacillus sp., 128 Bacterial cellulose, 282, 284 BActerial SYnthesized Cellulose (BASYC), 282 Barium sulfate, 26 BDO-BDI-BDO triblocks, 239 β-glycerophosphate, 409 β-tricalcium phosphate (β-TCP), 265, 273, 372 374, 373t, 419 420 β-tricalcium phosphate (β-TCP)-based porous scaffold, 372 374, 385 386 cell experimental evaluation of, 392f material characteristic evaluation of, 389f two phase structural scaffold, 385 386 characterization of, 385 386 fabrication of, 385 Bioabsorbable polymers, 98 Bioactive biomaterials, 39 40, 88 Bioactive glass, 273 274

Biocellulose (BC), 279 284 biomedical applications of, 281 284 antimicrobial biomaterials, 283 for bone tissue regeneration, 282 for cell culture, 282 283 drug delivery applications, 284 substitute biomaterials for, 281 282 fibril structure, 279 280 properties, 280 281 biocompatibility, 281 biodegradability, 281 hemocompatibility, 281 mechanical properties, 280 nontoxicity, 281 Bioceramic, 373t naturally produced, 14 15 Biocompatibility, 39 40, 99 100 Biocoral, 14 15 Biodegradable materials, 157 Biodegradable matrices, 297 299 Biodegradable polymers, classification of, 261f Biofabrication, 236 237 future trends in, 107 109 bioresponsive hydrogels, 108 109 electrospinning, 107 108 3D bioprinting rapid prototyping, 108 tissue engineering, biopolymer composites in, 109 Biohydroxyapatite (BIO-HA), 442, 447 459 methods to obtain, 457 459 mineral composition of, 454 456, 455t structural properties of, 451 454 thermal properties of, 456 457 Bioinert, 39 40 Biologically active molecules (BMP), 428 429 Biomaterials, 85 90 bioreabsorbable biopolymers, 89 90 poly(α-hydroxy acids). See Poly(α-hydroxy acids) polymeric biomaterials, 88 89 Biomedical engineering applications, specific properties for, 76 cytotoxicity, 76 degradability, 76 Biomimetics, 298 Bio-Oss, 274 276 Biopolymer matrices, 403 417 types of, 403 417

527

528

Index

Biopolymers, 259 260 Bioprinted scaffolds, 35 adherence and biocompatibility, 39 40 bioprinting, 36 degradation rates, 40 fibrous polymer-derived scaffolds, 52 geometry of scaffolds, 37 hydrogel-derived scaffolds, 40 43 agarose hydrogel, 43 alginate hydrogel, 43 44 cellulose hydrogel, 45 chitosan hydrogel, 44 45 fibrin hydrogel, 46 gelatin/collagen hydrogel, 46 48 hyaluronic acid hydrogel, 48 49 Matrigel hydrogel, 49 51 pore size, 38 39 porous polymer-derived scaffolds, 53 54 postbioprinting, 36 37 prebioprinting, 35 surface properties, 38 synthetic hydrogels, 51 52 Bioprinting, 36 Bioreabsorbable biopolymers, 89 90 Bioreabsorbable polymers for tissue engineering, 83 biocompatibility, 99 100 biofabrication, future trends in, 107 109 biopolymer composites in tissue engineering, 109 bioresponsive hydrogels, 108 109 electrospinning, 107 108 3D bioprinting rapid prototyping, 108 biomaterials, 85 90 bioreabsorbable biopolymers, 89 90 poly(α-hydroxy acids), 90 91 polymeric biomaterials, 88 89 poly(α-hydroxy acids), applications of, 104 106 medical applications, 105 106 nonmedical applications, 105 poly(α-hydroxy acids), copolymerization of, 97 98 poly(α-hydroxy acids), mechanisms of degradation of, 98 99 poly(α-hydroxy acids), toxicity of, 100 104 in vitro cytotoxicity tests, 100 102 in vitro hemocompatibility test, 102 103 in vivo biocompatibility tests, 103 104 poly(α-hydroxy acids) synthesis, 91 96 scaffolds, 85 tissue engineering, 83 84 Bioresorbable matrices, 302 304 Bioresorbable polyurethanes in biomedical devices, 237 239

Bioresponsive hydrogels, 108 109 Biosilicate, 273 274 Biphasic calcium phosphate (BCP), 273, 420 421 Blood brain barrier (BBB), 24, 176 177 Bloom index, 66 67 Bone calcination, 456 thermal events during, 458t Bone composition, 444 461 biohydroxyapatite (BIO-HA), 449 459 methods to obtain, 457 459 mineral composition of, 454 456 structural properties of, 451 454 thermal properties of, 456 457 bone porosity, 446 447 collagen, 459 460 cortical bone, 444 446 fat, 461 hydroxyapatite (HA), 447 448 methods to obtain, 457 459 osteocalcin, 460 trabecular bone, 446 water, 460 461 Bone Gla protein (BPG). See Osteocalcin Bone marrow-derived stem cells (BMSCs), 317 318 Bone microarchitecture, 6, 464f Bone morphogenetic protein-2 (BMP-2), 283 delivery of, 278 279 Bone morphogenetic proteins (BMPs), 1, 459 460 Bone porosity, 446 447, 447t Bone powder, 449, 450f Bone sialoproteins (BSP), 459 460 Bone structures, 11 15 Bone tissue engineered scaffold, general requirements of, 407 Bone tissue engineering, 14 15, 38 39, 162 PLGA scaffolds for, 162 171 Bone tissue regeneration biocellulose-based scaffolds for, 282 and human osteoblast cells, 223 227 Bone cartilage tissue construction, 397 398 Bovine bone powder (BBP), 454 Bovine trabecular bones. See Trabecular bones BrdU (5-bromo-2-deoxyuridine) assay, 39 40 Breast cancer, PLGA interventions in, 175t Brushite, 418 419 Bulk polymerization, 92 1,4-Butanediol diglycidyl ether (BDDGE), 303

C Calcium alginate, 423 Calcium phosphate particle loaded porous/ nonporous composites, 305 307

Index

Calcium phosphates, 271 273, 305 as essential part of composite materials, 417 421 Candida antarctica lipase B (CAL-B), 135 Candida boidinii, 127 Carboxymethylated chitosan (CM-chitosan), 64 Cardiac progenitor cells (CPCs), 21 Cartilage, 15 17 Cell culture, scaffolds for, 282 283 Cell-based regenerative medicine, 24 Cellulase, 281 Cellulose, 279 Cellulose based-hydrogel scaffolds, 45 Cellulose hydrogel, 45 Central and peripheral nervous system, 24 25 Central nervous system (CNS) delivery PLGA interventions in, 174 179, 178f Chemical hydrogels, 69 Chitin, chitosan extraction from, 65f Chitosan, 373t, 403 417 biodegradability of, 65 66 characteristics of, 404t extraction, from chitin, 65f physicochemical characteristics of, 66t Chitosan hydrogel, 44 45 Chitosan-based hydrogels, 61 biomedical engineering applications, specific properties for, 76 cytotoxicity, 76 degradability, 76 chemical hydrogels, 69 chitosan nature and main properties, 64 66 physical hydrogels, 67 69 potential applications and future trends of chitosan hydrogels, 76 77 property measurements, 74 76 active compound release assessment, 74 contact angle, 75 mechanical resistance, 74 swelling index, 75 thermal analysis, 76 viscosity (sol gel analysis), 74 75 structural analysis, 71 73 microstructural and spectroscopic analysis, 71 72 ultraviolet visible spectroscopy and Fouriertransform infrared spectroscopy, 72 73 Chitosan collagen composite hydrogel materials, 409 Chondroitin-4-sulfate, 46 47 Cold atmospheric plasma (CAP), 359 360 Collagen, 271 276, 373t, 403 417, 459 460 bioactive ceramic composites, 272 274 collagen-bioglass based composites, 273 274

collagen-HAP (CHAP) composites, 272 273 collagen TCP/BCP composites, 273 characteristics of, 405t commercially available collagen products, 275t medical applications of, 274 276 Collagen coatings, 38 Collagen fibrils, 304 Collagen hydrogel, 46 48 Collagen/hydroxyapatite composite scaffolds, 442 443 Collagen-apatite porosity (CAP), 444 446 Collagenases, 303 Collagen-based porous scaffold, 374 378 fabrication and characterization of particle distributed scaffold, 374 376 characterization of particle distributed scaffold, 375 376 fabrication of particle distributed scaffold, 374 375 in vitro cell experiment, 376 378 cell culture, 376 compression test, 376 377 evaluation of cell number and alkaline phosphatase activity, 377 gene expression analysis, 377 378 microstructural characterization, 377 statistics, 378 Collagen glycosaminoglycan (Col-GAG) scaffolds, 38 39, 46 47 Collagen-HA hybrid nanocomposite, 310 313 Colloidal gels, 61 Colon Cancer, PLGA interventions in, 175t Composite materials, 373t formation of, 421 427 Composite scaffolds, 297 applications in tissue engineering, 304 325 biodegradable matrices, 297 299 bioresorbable matrices, 302 304 for bone, 304 313 calcium phosphate particle loaded porous/ nonporous composites, 305 307 collagen-HA hybrid nanocomposite, 310 313 fiber-loaded composites, 307 310 for osteochondral defects, 313 325 gradient porous/nonporous composites, 320 323 magnetic bioinspired hybrid nanocomposites, 323 325 multilayer porous scaffolds, 315 319 Composites based on bioderived polymers, 259 biocellulose, 279 284 biomedical applications of, 281 284 fibril structure, 279 280 properties, 280 281

529

530

Index

Composites based on bioderived polymers (Continued) collagen, 271 276 bioactive ceramic composites, 272 274 medical applications of, 274 276 polyesters, 260 271 poly(lactic acid), 261 269 poly(lactic-co-glycolic acid) (PLGA) copolymers, 269 271 properties and medical applications, 262t silk fibroin (SF), 276 279 medical applications of, 278 279 processing of, 277 278 structure of, 276 277 Computer-aided design (CAD) techniques, 235 Contact angle, 75, 211 Corn steep liquor, 124 Cortical bone, 442, 444 446 Corynebacterium glutamicum, 125 Crosslinking methods, 42 43 Cryoprotectants, 249 250 Crystalline percentage, 451 452 Crystalline quality (CQ), 452 453 Cyclosporin A, 179 Cytokines, 459 460 Cytotoxicity, 76

D Degradability, 65 66, 76, 105, 281, 351 354 Degradation rates, 40 Dehydrative cross-linking (DHT), 303 304 Demineralized freeze-dried bone allografts (DFDBA), 274 276 Dexon, 162 Dibutyl tin dilaurate (DBTDL), 253 Dichloromethane, 174 Differential scanning calorimetry (DSC), 210, 456 457 Dimethyl sulfoxide (DMSO), 249 250 Direct ink writing (DIW), 240 241, 243 252 Double emulsion W/O/W method, 174 176 Double-nozzle low-temperature deposition manufacturing (DLDM), 240 241, 248 251, 250f Dry spinning, 169 Duchenne muscular dystrophy (DMD), 2

E Elastic modulus, 345 Electroactive polymers (EAP), 325 327 Electroless metallization, 491, 514 516 Electron microscopy (EM) methods, 4 Electrospinning, 107 108, 162, 169, 277

Electrospun biomimetic scaffolds, 203 bone tissue regeneration and human osteoblast cells, 221 227 cell viability and HEK293 cells, 221 223 methods of characterization, 208 212 contact angle, 211 Fourier-transformed infrared spectroscopy, 210 materials, 208 209 polarized optical microscopy, 211 212 scaffold fabrication, 209 210 scanning electron microscopy (SEM), 212 small-angle light-scattering (SALS), 211 thermal analysis, 210 X-ray scattering, 210 211 poly(β-hydroxybutyrate) (PHB) electrospun fibrous scaffolds, 212 221 aging, 218 220 scaffolds morphology, 212 217 sterilization methods and influence on physical properties, 220 221 wetting behavior, 217 218 Electrospun fibrous scaffolds, 205 207 Embden Meyerhof Parnas pathway, 118 Energy dispersive system (EDS), 64 Epoxy matrix, 204 Escherichia coli, 118 Extracellular matrix (ECM), 38 39, 83, 205 206, 371 372 Extracellular matrix vesicles (ECMVs), 461 Extrusion-based methods, 240 241, 243 253 direct ink writing, 243 252 fused deposition modeling, 252 253

F Fabrication techniques, 163 168 Fat, 461 Fatigue, 40 FeHA, 323 324 Fiber-forming techniques, 168 170 Fiber-loaded composites, 307 310 Fibrillar collagen, 406 Fibrin hydrogel, 46 Fibrin-collagen bioink, 46 47 Fibrinogen, 408 Fibronectin, 38 Fibrous nonwoven scaffolds, 212 Fibrous polymer-derived scaffolds, 52 Fibrous scaffolds, 37, 168 169, 208 Field emission scanning electron microscope (FESEM), 375, 382f, 393f Filopodia, 224 225 Filtered back projection (FBP), 6

Index

Finite element method (FEM), 477 Flexural modulus, 40, 345 Fluorescein-50-isothiocyanate (FITC) molecule, 284 Fourier transform infrared (FTIR), 72 73, 340 341 Fourier-transformed infrared spectroscopy, 210 FS- and NS-laser processing, comparison between, 504 510 FS-laser nanostructuring, 498 503 versus NS-laser processing, 504 510 Full-width at half maximum (FWHM), 211 Fused deposition modeling, 252 253

G Gelatin, 168, 403 417, 405t Gelatin methacrylate (GelMA), 51 Gelatin/collagen hydrogel, 46 48 Gene therapy, poly(lactide-co-glycolide) strategies for, 180 182 Gentamycin, 411 Glass-transition temperatures, 238 Glucosamine, 65 Glucosamine-glucosamine, 65 66 Glucosamine-N-acetyl-glucosamine, 65 66 Glycerol, 125, 249 250 Glycine-arginine-glycine-aspartic acid-serine (GRGDS), 43 Glycol chitosan coated PLGA nanoparticles, 182 Gradient porous/nonporous composites, 320 323 Green fluorescence protein (GFP), 180 Growth factors (GFs), 1, 273, 278 279, 317 318, 322

H HA/calcium alginate microsphere composite, 423 424 HA collagen composites, 312 313 HAMC hydrogel, 179 Haversian canals, 444 446 HBsAg loaded chitosan, 182 Heart, 20 24 Hemocompatibility, 281 Hexafluoro-2-propanol (HFP), 277 Hexafluoroisopropanol, 161 Holotomography (HT), 10 Human embryonic kidney 293 cells (HEK293), 221 222, 222f, 223f Human embryonic stem cells, 349 Human mesenchymal stem cells (hMSCs), 322 Human serum albumin (HSA), 358 359 Human umbilical vein endothelial cells (HUVECs), 46

Hyaluronan, 246 248 Hyaluronic acid (HyA), 408, 417 characteristics of, 404t Hyaluronic acid hydrogel, 48 49 Hydrogel scaffolds, 37 Hydrogelation, 277 Hydrogel-derived scaffolds, 40 43 agarose hydrogel, 43 alginate hydrogel, 43 44 cellulose hydrogel, 45 chitosan hydrogel, 44 45 fibrin hydrogel, 46 gelatin/collagen hydrogel, 46 48 hyaluronic acid hydrogel, 48 49 Matrigel hydrogel, 49 51 Hydrogels, 61, 66 67, 108 109, 170 Hydrophilic gels, 61 Hydrophilic polymer, 67 Hydrophobic nanocarriers, 177 178 3-Hydroxyalkanoates, 340, 348 Hydroxyapatite (HA), 306 307, 373t, 403, 410 411, 415, 441 442, 447 448 collagen/HA composites, 419 crystalline structure of, 448f fillers, 419 HA-CS-Alg composite scaffolds, 428f HA-Gel, 429 430 methods to obtain, 457 459 microparticles of, 415 mineral composition of, 455t stoichiometric HA, 419 synthetic HA, 418 Hydroxybutyric-CoA, 261

I In vitro cell experiment, 386 388 alizarin red S staining, 388 cell culture, 386 387 evaluation of cell number and alkaline phosphatase activity, 387 evaluation of mechanical characteristics, 387 gene expression analysis, 388 microstructural characterization, 387 results of, 390 394 statistics, 388 In vitro cytotoxicity tests, 100 102 In vitro hemocompatibility test, 102 103 In vivo biocompatibility tests, 103 104 bone implants, general tests for, 103 104 stents, general tests for, 104 Induced pluripotent stem cell-derived mesenchymal stem cells (iPSCMSCs), 273

531

532

Index

Inductively coupled plasma atomic emission spectroscopy (ICP-AES), 454 Injectable microparticles, 170 171 Inkjet printing, 240 243 Integrated organ printing, 251 252 Interfacial deposition method, 174 176 Inverse inkjet printing, 25

K KEYENCE, 493 494

L Lactic acid, 117 131 culture medium for fermentation of, 123 125 extraction and purification of, 130 131 factors that influence production of, 121 123 microorganisms involved in the production of, 127 130 production, by fermentation, 125 127 Lactic acid bacteria, 118, 127 catabolic pathways of, 119f Lactic acid enantiomers, chemical structures of, 118f Lactobacillus, 122 Lactobacillus delbrueckii, 127 Lactobacillus lactis, 122 Lactobacillus rhamnosus, 127 Lactonitrile, 119 “Lacunar-canalicular” porosity (LCP), 444 446 Lamellipodia, 224 225 Laser processing of biopolymers, 487 comparison between FS- and NS-laser processing, 504 510 electroless metallization directly after laser treatment, 515 516 experimental and analytical techniques, 492 494 FS-laser nanostructuring, 498 503 NS-laser processing in different environments, 516 520 perspectives for future investigations, 521 polydimethylsiloxane (PDMS) optical properties of, during NS-laser treatment, 494 498 structure and Raman spectrum of, 491 PS-laser processing, 503 XPS study of NS-laser processing of polydimethylsiloxane, 510 514 Laser-based bioprinting, 42, 51 52 Layer-by-layer system, 35 Lb. helveticus, 118 Lipids, 449, 459 460

Liquid-frozen deposition manufacturing (LFDM), 240 241, 246 248 Liver cancer, PLGA interventions in, 175t Lung cancer, PLGA interventions in, 175t Lysine ethyl ester diisocyanate (LDI), 253 Lysozyme, 65 66, 69, 283

M Macrodiols, 238 239 Macromolecules, 88 Magnetic bioinspired hybrid nanocomposites, 323 325 Mammalian bones, 442 Mass change, 394 395 Matrigel hydrogel, 49 51 Matrix biocomposite scaffolds, 283 Melt spinning, 168 169 Mesenchymal stem cells (MSCs), 46, 51 52, 371 374 Methacrylated gelatin, 42 43 Microelectromechanical systems (MEMS), 487 488 Microfil, 26 Microporous tricalcium phosphate (b-TCP) scaffolds, 53 54 Microspheres, 170 171 Microstructural and spectroscopic analysis, 71 72 Modified thermal compression molding-particulate leaching approach (MTCM-PL), 163 167 MTT (3-(4,5-dimethylthiazol-2-yl)-2,5diphenyltetrazolium bromide), 39 40 Multilayer porous scaffolds, 315 319 μ-Raman-spectroscopy, 497 498 Muscle tissue engineering, 18 24 heart, 20 24 skeletal muscles, 18 20 Muscle tendon unit (MTU), 251 252

N N-acetyl-glucosamine, 65 N-acetyl-glucosamine-N-acetyl-glucosamine, 65 66 Nanocellulose, 284 Nanoengineering of delivery systems, 156 Nanofiber scaffolds, 162 Nanohydroxyapatite (NanoHA), 278 NanoMatrix3D (NM3D), 142 Nanometric hydroxyapatite, 449 Natural bioresorbable polymers, 302 303 Natural materials, 300 301 Natural polymers, 52, 259 260, 373t, 403, 408 Natural-derived hydrogels, 40 41 Naturally produced bioceramics, 14 15

Index

New age therapeutic strategies, challenges in, 155 Ni coating, 516 Nonbiodegradable scaffolds, 40 Noncollagen proteins, 459 460 Normal Human Osteoblast (NHOst) cells, 223 224, 225f, 226, 228 NS-laser processing in different environments, 516 520 versus FS-laser processing, 504 510 NucleoSpin RNAII, 377 378

O O/W emulsification method, 170 171 Obtained composite materials biomedical applications of, 427 431 OC scaffolds, 322 OC tissue vascularization, 314 315 OCP/Gel composites, 428 Octacalcium phosphate (OCP), 420 Oil-in-water emulsification technique, 423 Oncogel, 174 Origanum vulgare L. essential oil (OEO), 138 139 Osseointegration, 103 Osteoblasts, 223, 407, 459 460 Osteocalcin, 381, 459 460 Osteochondral defects, composite scaffolds for, 313 325 gradient porous/nonporous composites, 320 323 magnetic bioinspired hybrid nanocomposites, 323 325 multilayer porous scaffolds, 315 319 Osteochondral tissue, 321f magnetic bioinspired hybrid nanocomposites for, 323 325 Osteoclasts, 407 Osteoconduction, 1 Osteocytes, 444 446, 459 460, 463 464 Osteogenesis, 1 Osteoinduction, 1, 103 Osteonectin (ON), 459 460 Osteopontin (OPN), 459 460 Osteoprogenitor cells, 459 460 Ovarian Cancer, PLGA interventions in, 175t

P Pancreatic Cancer, PLGA interventions in, 175t Parallel plate model, 6 Particle binding technology, 241, 253 Particle distributed scaffold, 374 characterization of, 375 376, 378 fabrication of, 374 375

PEG-PLGA-PEG, 170 Pellicle/membrane scaffolds, 283 Phase-contrast (PhC) effect, 8 PhC-microCT, 10 11 Photopolymerization, 42 43 Physical hydrogels, 67 69 advantages and disadvantages of, 69t Plant cellulose, 279 Plasma polymerization, 356 Plasma polymers, 342, 356 Plasma postirradiation grafting, 356 Plasma syn-irradiation, 356 Plasma treated and untreated thermoplastic biopolymers/biocomposites. See Polyhydroxyalkanoates (PHAs)Poly (lactic acid) (PLA) Plasma treated polymer surfaces, characterization techniques for, 356 358 Plasma treatment, 355 356 of PHAs, 361 363 of PLA, 358 361 PLGA-b-PEG-b-PLGA sol-gel injections, 174 Polarized optical microscopy, 211 212 Poloxamer 407, 51 52 Poly(3-hydroxybutyrate-co-3-hydroxyhexanoate) (PHBHHx) scaffolds, 207 208 Poly(3-hydroxybutyrate-co-3-hydroxyhexanoate) copolymer, 351 Poly(3-hydroxybutyrate-co-lactate), 135 Poly(acrylic acid), 63 64 Poly(butylene succinate) (PBS), 136 Poly(D,L-lactic acid) (PDLLA), 90, 341 Poly(D-lactic acid) (PDLA), 90 Poly(D-lactide) (PDLA), 341 Poly(glycolic acid) (PGA), 85, 89 91, 99, 105 106, 156, 259 260, 269, 307 309, 373t Poly(glycolide) diol (PGA diol), 238 239 Poly(hydroxybutyrate-co-hydroxyvalerate) (PHBV), 352 354 Poly(lactic acid) (PLA), 85, 89 90, 99, 105 106, 131 142, 156, 259 269, 339 340 application in tissue engineering, 350 351 applications, 137 140 biodegradability of, 351 354 biodegradation, biocompatibility, and toxicity, 141 142 chemical and physical properties, 132 133 fabrication, 263 264 kinds of polymers, copolymers, and their features, 135 137 market development, 140 141 medical applications, 267 269 drug delivery, 268 269

533

534

Index

Poly(lactic acid) (PLA) (Continued) orthopedic implants and fixation devices, 268 scaffolds for tissue engineering, 267 268 3D printing, 269 wound healing and stents, 267 plasma treatment of, 354 364 characterization techniques for, 356 358 disadvantages of, 363 364 plasma and plasma surface interactions, 355 356 processing, 264 266 cast film and sheet, 265 drying and extrusion, 264 265 foaming, 265 266 injection molding, 265 stretch blow molding, 265 thermoforming, 265 properties, 266 267, 344 350 biocompatibility, 349 mechanical properties, 267, 345 physical properties, 266 processability, 349 350 thermal properties, 266, 346 348 transparency, 348 349 structure of, 340 341 synthesis, 133 135, 134f, 263f, 341 344 chemical polymerization, 133 134 enzymatic polymerization, 134 135 Poly(lactic acid)/polycaprolactone (PLA/PCL), 136 Poly(lactic-co-glycolic acid) (PLGA) copolymers, 90, 269 271 medical applications, 271 properties, 270 271 synthesis, 269 270 Poly(lactide-co-glycolide) (PLGA), 136, 156 158 properties, 158f, 160 162 synthesis, 158 160, 159f Poly(lactide-co-glycolide) (PLGA) scaffolds, 36, 155 in anticancer therapy, 171 174, 175t for bone tissue engineering, 162 171, 164t, 167f fibrous scaffolds, 168 169 hydrogels, 170 injectable microparticles, 170 171 porous scaffolds, 163 168 future trends, 184 186 interventions in central nervous system delivery, 174 179, 178f miscellaneous PLGA therapeutics, 183 184 strategies for gene therapy and vaccine delivery, 180 182 Poly(L-lactic acid) (PLLA), 90 91, 341, 373t Poly(L-lactide-co-ε-caprolactone) (PLCL), 373t

Poly(propylene carbonate) (PPC), 261 Poly(sodium acrylate), 63 64 Poly(vinyl alcohol) (PVA), 64 Poly(α-hydroxy acids), 87, 90 92 copolymerization of, 97 98 mechanisms of degradation of, 98 99 PLA and PGA, 104 106 medical applications of, 105 106 nonmedical applications of, 105 synthesis, 91 96 toxicity of, 100 104 in vitro cytotoxicity tests, 100 102 in vitro hemocompatibility test, 102 103 in vivo biocompatibility tests, 103 104 Poly(α-hydroxy esters), 156 Poly(β-hydroxybutyrate) (PHB), 207 209 Poly(ε-caprolactone) (PCL), 85, 89 90, 259 260, 373t Polyacrylamide, 63 64 Polycaprolactone (PCL), 204, 316 Polycarbonate (PC), 205 Polydimethylsiloxane (PDMS), 487 488, 516 518 optical properties of, during NS-laser treatment, 494 498 structure and Raman spectrum of, 491 XPS study of NS-laser processing of, 490, 510 514 Poly-D-lysine, 38 Polyesters, 260 271 properties and medical applications, 262t Polyhydroxyalkanoates (PHAs), 207, 339 340 application in tissue engineering, 350 351 biodegradability of, 351 354 plasma treatment of, 354 364 characterization techniques for, 356 358 disadvantages of, 363 364 plasma and plasma surface interactions, 355 356 properties of, 344 350 biocompatibility, 349 mechanical properties, 345 processability, 349 350 thermal properties, 346 348 transparency, 348 349 structure of, 340 341 synthesis of, 341 344 Polylactic acid-coated polyglycolic acid (PGA/ PLA) material, 316 317 Poly-l-ornithine, 38 Polymeric biomaterials, 88 89 Polymeric nanoparticulate therapy, 171 172 Polymers and bioceramics, 408 Polymers as medical devices, 203 204

Index

Polysaccharide nanocomposites, 411 Polysaccharides, 65 67, 403, 404t, 459 460 Polyurethane-based structures, 235, 242t additive manufacturing for biomedical polyurethane processing, 240 241 extrusion-based methods, 240 241 inkjet printing, 240 particle binding, 241 additive manufacturing of composite polyurethanes, 241 253 extrusion-based methods, 243 253 inkjet printing, 241 243 particle binding, 253 bioresorbable polyurethanes in biomedical devices, 237 239 perspectives, 253 254 Polyvinyl alcohol (PVA), 64, 239 Polyvinyl chloride (PVC), 205 Polyvinylpyrrolidone, 63 64 Poly-α-hydroxyesters, 301 302 Poly-β-hydroxybutyric acid (PHB), 208, 226 227, 261 electrospun fibrous scaffolds, 212 221 Poly-ε-caprolactone/hydroxyapatite (PCL/HA), 316 317 Porogen leaching, 163, 277 Porogen technique based on spheric paraffin particles, 168 Porous collagen scaffolds with ceramic particles, 408 Porous composite scaffolds, 310 with a PCL matrix endowed with HA/MgHA crystals, 308f Porous HA-CS-Alg composite scaffolds, 427 Porous polymer-derived scaffolds, 53 54 Porous scaffolds, 37, 53 54, 163 168 Postbioprinting, 36 37 Powder Diffraction File (PDF), 454 PowderX, 210 211 Prebioprinting, 35 Primary bone cells, 459 460 Proflo, 124 Prostate Cancer, PLGA interventions in, 175t Protein based biopolymers for biomedical applications, 405t Proteins, 403, 405 Pseudo-holotomography, 26 27 PS-laser processing, 503 Pt metalized tracks by electroless deposition, 520f by electroless plating, 520f Puros Demineralized Bone Matrix (DBM) Putty, 274 276

Q Quercetin loaded PLGA TPGS nanoparticles, 183 184

R Radiopaque contrast agents, 26 Radon Transform, 5 6 Raman spectroscopy, 443 bone characterization by, 471 473, 472f, 473t Rapid prototyping (RP) technology, 320 Rat bone marrow mesenchymal stem cells (rMSC), 374, 376 Reinforcement systems, 298 Rhizopus oryzae, 127 Ring opening polymerization (ROP), 133, 160 Room temperature compression moldingparticulate leaching method (RTCM-PL), 167 Room temperature injection molding-particulate leaching method (RTIM-PL), 167 168

S Satisfactory vascularization, 25 Saturated poly(α-hydroxy esters), 162 163 Scaffold fabrication, 209 210 Scaffold structure and cell response, variational mechanical behavior between, 382 385, 394 396 Scaffold-MSC systems, 398 Scaffolds, geometry of, 37 Scanning electron microscopy (SEM), 61, 63 64, 212, 463 Scanning transmission electron microscopy (STEM), 493 494 Scherrer’s equation, 211 Schiff base, 69 Schizosaccharomyces pombe, 129 130 Selected area electron diffraction (SAED), 493 494 Series expansion methods, 6 Shape memory polymers (SMPs), 204 Silicon-stabilized TCP/MSC composites, 26 27 Silk fibroin (SF), 276 279 medical applications of, 278 279 delivery of bioactive molecules, 278 279 fixation devices, 279 for tissue engineering, 278 processing of, 277 278 electrospinning, 277 hydrogelation, 277 porogen leaching, 277 SF composites, 278

535

536

Index

Silk fibroin (SF) (Continued) 3D bioprinting, 278 structure of, 276 277 Silver nanoparticles, 283 Skeletal muscles, 18 20 Skeletal tissue engineering, 11 18 bone, 11 15 cartilage, 15 17 tendons, 17 18 Small interfering RNA (SiRNA), 172 173, 180 Small-angle light-scattering (SALS), 211 Smart polymeric coatings, 204 205 Sodium alginate, 414, 423 Sodium ammonium phosphate (SAP), 342 Solution spinning, 169 SOX9 genes, 180 181 Soy protein isolate (SPI), 350 SPIONs (superparamagnetic iron oxide nanoparticles), 323 324 Spongy bone. See Trabecular bones Sporolactobacillus, 127 128 Sporolactobacillus inulinus, 123 SR-microCT, 12, 20, 25 SR-PhC-microCT, 16, 23 25 Stimuli-sensitive polymeric nanocarriers, 171 172 Streptococcus bovis, 122 Streptococcus pyogenes, 122 Stroke, PLGA based therapies for, 179 Surface energy, 38 Surface eroding polymers, 301 302 Surface modification, 177 178 Surface modified pellicle scaffold, 283 Swelling index, 75 Synchrotron radiation (SR)-based microCT, 4 Synchrotron radiation-based high-resolution tomography, 1 central and peripheral nervous system, 24 25 conduction, induction, and cell transplantation in tissue engineering, 1 4 high-resolution tomography, innovative approaches to by synchrotron radiation, 7 11 muscle tissue engineering, 18 24 heart, 20 24 skeletal muscles, 18 20 skeletal tissue engineering, 11 18 bone, 11 15 cartilage, 15 17 tendons, 17 18 vascularization, 25 27 X-ray computed microtomography, 5 6 Synthetic biodegradable polyesters, 260 261 Synthetic biodegradable polymer, 259 260, 373t Synthetic hydrogels, 51 52

Synthetic polymers, 20, 52, 301 302 Synthetic scaffolds versus trabecular bone, 475 479 Synthetic-derived hydrogels, 40 41

T Tecoflex, 252 Tendons, 17 18 Tensile strength, 40 Thermal analysis, 76, 210 Thermal equilibrium, 355 356 Thermally induced phase separation (TIPS), 168, 235 236 Thermogravimetric analyses (TGA), 210, 347, 456 3D bioprinting, 37f, 38 39, 278 rapid prototyping (RP), 108 3D polyurethane structures, 240 Three-dimensional printing, 235 Thrombospondin (THBS1), 459 460 Tissue engineering, 3f, 83 84, 156 biopolymer composites in, 109 composite scaffolds in, 304 325 for bone, 304 313 for osteochondral defects, 313 325 conduction, induction, and cell transplantation in, 1 4 muscle tissue engineering. See Muscle tissue engineering SF scaffolds for, 278 skeletal tissue engineering. See Skeletal tissue engineering steps for, 84f Tomographic image reconstruction, 5 6 Trabecular bones, 442, 446 study of, 461 475 collection and preparation of samples, 462 463 mechanical properties, 474 475, 476t morphological characterization, 463 465 Raman spectroscopy, 471 473, 472f, 473t structural properties, 469 471 x-ray tomography, 465 468 synthetic scaffolds versus, 475 479 2D imaging of, 467f Transferrin receptors, 177 178 Transform-based methods, 6 Transforming growth factor β (TGFβ), 246 248 Transmission electron microscope (TEM), 61 63, 469 470 Transport of intensity equation (TIE), 10 Trimethylene carbonate (TMC), 141 142 Two phase structural scaffold, 374 characterization of, 385 386, 388 389 fabrication of, 385

Index

U

X

Ultraviolet (UV) radiation, 220 Ultraviolet visible spectroscopy, 72 73

Xenografts, 441 X-ray computed microtomography, 5 6 X-ray diffraction, 451, 470 471 X-ray microtomography (microCT), 5 X-ray photoelectron spectroscopy (XPS), 357 358, 490 X-ray scattering, 210 211 X-ray tomography, 465 468

V Vaccine delivery, poly(lactide-co-glycolide) strategies for, 180 182 Vascular endothelial growth factor (VEGF), 274 Vascular porosity (VP), 444 446 Vascularization, 25 27 Vicryl L-PLGA, 162 Viscosity (sol gel analysis), 74 75 Volkmann canals, 444 446

W Water, in bone, 460 461 Wet spinning, 169 Wettability, 38 Wide-angle X-ray scattering (WAXS) patterns, 210

Y Yield strength, 40 Young modulus, 40

Z Zirconium oxide, 429

537