Advances in Medical and Surgical Engineering 0128197129, 9780128197127

Table of contents :
Cover
ADVANCES IN
MEDICAL AND
SURGICAL
ENGINEERING
Copyright
Contributors
Chapter 1 - Introduction to advances in medical and surgical engineering
References
Chapter 2 - Engineering advances in promoting bone union
1 - Introduction
2 - Principles of bone union
2.1 - Secondary bone union
2.2 - Primary bone union
3 - Biological factors in bone union
3.1 - Engineering advances influencing the local environment of growth factors in bone union
4 - Mechanical factors in bone union
4.1 - Engineering advances influencing the mechanical environment in bone union
5 - Future advances to facilitate bone to bone union
References
Chapter 3 -
Engineering advances in promoting tendon to bone healing
1 - Introduction
2 - Principles of tendon to bone healing
3 - Factors affecting tendon-to-bone healing
3.1 - Biological factors affecting tendon-to-bone healing
3.1.1 - Engineering advances influencing the local environment of growth factors in tendon to bone healing
3.2 - Mechanical factors affecting tendon-to-bone healing
3.2.1 - Engineering advances influencing the mechanical environment of tendon to bone healing
4 - Future advances to facilitate tendon to bone healing
References
Chapter 4 - Engineering advances in reverse total shoulder arthroplasty
1 - Background of shoulder arthroplasty
1.1 - Anatomic total shoulder arthroplasty (TSA)
1.2 - Reverse total shoulder arthroplasty (RTSA)
2 - Design systems and their development
2.1 - Prosthesis stability
2.2 - Medializing and distalizing the center of rotation
2.3 - Deltoid muscle function
2.4 - Range of motion post RTSA
2.5 - Implant designs
3 - Clinical outcomes of shoulder arthroplasty
3.1 - Infection
3.2 - Scapular notching
3.3 - Instability
3.4 - Fracture
3.5 - Humeral loosening
3.6 - Glenoid loosening
3.7 - Nerve palsy
3.8 - Reduced rotation
4 - Advances in implant design and surgical techniques
4.1 - Scapular notching
4.2 - Inferior inclination of the glenosphere
4.3 - Inferior (eccentric) positioning of the glenosphere
4.4 - Increased glenosphere offset
4.5 - Alteration of the neck-shaft angle of the humeral component
4.6 - Surgical approach
4.7 - Managing glenoid defects
4.8 - Correct deltoid muscle tensioning
4.9 - Implant fixation
4.10 - Muscle transfers
5 - Future challenges
References
Chapter 5 - Engineering advances in knee arthroplasty
1 - Introduction
2 - Prosthetic joint infection
2.1 - Mechanism of bacterial adhere to implant surfaces
2.2 - Strategies to tackle prosthetic joint infection
2.3 - Engineering advances in order to reduce prosthetic joint infection
2.3.1 - Bio-inert implant surface modifications
2.3.2 - Bioactive implant surface modifications
2.3.2.1 - Bioactive mineral implant surface modifications
2.3.2.2 - Bioactive organic implant surface modifications
3 - Strategies to improve implant longevity
3.1 - Cementless fixation total knee arthroplasty
3.2 - Bisphosphonate coatings
3.3 - Biomolecule coatings
3.4 - Alternative bearing surfaces
3.4.1 - Ceramic metal femoral components
3.4.1.1 - Ceramic and ceramicised metal alloy femoral knee components
3.5 - Patient specific knee arthroplasty
4 - Metal hypersensitivity in total knee arthroplasty
4.1 - Pathophysiology
4.1.1 - Current concepts
4.2 - The all polyethylene tibial component
4.3 - Methods used to make implants metal hypersensitivity friendly
4.3.1 - Fully coated implants
4.3.2 - Partially coated implants
4.3.3 - Alternative alloy implants
4.4 - Controversy around the need for metal hypersensitivity implants
5 - Conclusion
References
Chapter 6 - Biology of cartilage
1 - Introduction
1.1 - Structure
1.2 - Function
1.3 - Cartilage aging
1.4 - Cartilage injuries and healing
1.4.1 - Causes of chondral injury
1.4.2 - Classification of chondral defects
1.4.3 - Current treatments and challenges in management
1.4.4 - Biological resurfacing: microfracture
1.4.5 - Biological resurfacing: autologous matrix-induced chondrogenesis (AMIC)
1.4.6 - Biological resurfacing: transplant
1.4.7 - Biological resurfacing: autologous chondrocyte implantation
1.5 - Future options: stem cells
1.6 - Future options: button replacements
1.7 - Conclusion
References
Chapter 7 - Mechanical circulatory support: an overview
1 - Introduction
1.1 - History of mechanical circulatory support
2 - Short term mechanical circulatory support
2.1 - Intra-aortic balloon counterpulsation
2.2 - Percutaneous left ventricular assist devices
2.3 - Right ventricular support
2.4 - Surgical short-term mechanical circulatory support
2.4.1 - Extracorporeal membrane oxygenation
2.4.1.1 - Veno-venous ECMO
2.4.1.2 - Veno-arterial ECMO
2.4.1.3 - Complications
2.4.1.4 - Clinical use
2.4.2 - Short-term ventricular assist devices
2.5 - Implantable left ventricular assist devices
2.6 - First generation devices
2.7 - Second generation devices
2.8 - Third generation devices
2.9 - Total artificial heart
2.10 - Future directions
References
Chapter 8 - Advances in transcatheter aortic valve implantation
1 - Introduction
2 - Evidence and current indications for TAVI
3 - Imaging work up for TAVI patient
4 - Approaches for TAVI access
5 - Transcatheter heart valves and delivery systems
5.1 - Balloon expandable valve
5.2 - Self-expandable valves
5.3 - Mechanically expandable valve
6 - Future perspectives
6.1 - Extended indications
6.2 - Future THV and delivery systems
6.3 - Advances in CT imaging and closure devices
7 - Conclusion
References
Chapter 9 - Advances in magnetic resonance imaging (MRI)
1 - Introduction
2 - A brief history of development of MRI [4,5]
3 - Advantages and disadvantages of MRI
3.1 - Advantages of MRI
3.2 - Disadvantages of MRI
4 - Basic physics of MRI [6–10]
5 - Commonly used MRI sequences [6,9,12,13]
5.1 - Spin echo (SE) sequences
5.2 - Gradient echo (GE) sequences
5.3 - Inversion recovery (IR) sequences
5.4 - Diffusion weighted imaging (DWI)
6 - Advances in general MRI [6,12,13,15]
6.1 - Advanced spin echo (SE) sequences
6.1.1 - Fast/Turbo SE (FSE/TSE) [17–20]
6.1.2 - Ultrafast SE [21–26]
6.2 - Advanced gradient echo (GE) sequences [27–30]
6.3 - Advanced inversion recovery sequence [31–33]
6.4 - Advanced diffusion weighted sequence [34,35]
6.5 - Echo planar imaging (EPI) [36,37]
6.6 - Functional MRI (fMRI) [38–41]
6.7 - Perfusion MRI [42–44]
6.8 - MR angiography (MRA) sequences [45]
6.9 - Cerebrospinal fluid (CSF) sensitive sequence [46]
6.10 - Susceptibility weighted imaging (SWI) [47,48]
6.11 - MR spectroscopy [49–51]
6.12 - Hybrid sequences
7 - Advances in musculoskeletal (MSK) MRI
7.1 - Imaging joints with prosthesis
7.2 - MR arthrography [57,58]
7.3 - Dynamic MSK MRI
7.4 - Differentiating benign from malignant bone tumors [62,63]
7.5 - Investigating joint soft tissues
7.6 - Utilising 3-dimensional isotropic voxels in MSK imaging
8 - Impact of MRI field strength on imaging
8.1 - Understanding MRI scanner’s magnetic field strength
8.2 - Comparing 1.5 T with 3.0 T magnetic fields in MRI scanners [72–81]
8.3 - MRI field strengths used in musculoskeletal (MSK) imaging
9 - Conclusions and future of MRI [88–113]
References
Chapter 10 - Technological advances in breast implants
1 - Types of breast implants
1.1 - Complications of breast implants
1.2 - Breast reconstruction and acellular dermal matrix (ADM)
1.3 - Future developments in breast implants
References
Chapter 11 - Importance of biomaterials in biomedical engineering
1 - Introduction
2 - Chitosan
2.1 - Introduction
2.2 - Dressing fabrication
2.3 - Hydrogel applications
2.4 - Wound management
2.5 - Implants
2.6 - Conclusion
3 - Hyaluronic acid
3.1 - Introduction
3.2 - Hydrogel applications
3.3 - Dressings
3.4 - Composites, scaffolds, and matrices
3.5 - Implants
3.6 - Conclusion
4 - Silk fibroin
4.1 - Introduction
4.2 - Silk fibroin as a biomaterial
4.3 - Silk fibroin in drug delivery
4.4 - Conclusion
5 - Conclusions
References
Chapter 12 - Visible light activated antimicrobial silver oxide thin films
1 - Introduction
2 - Theoretical background
2.1 - The fight against pathogenic microorganisms
2.1.1 - The use of metals in the fight against pathogenic bacteria
2.1.2 - The use of antibiotics to fight microbes
2.1.3 - Nanoparticle antimicrobials
2.2 - Silver chemistry
2.2.1 - Crystal field theory and coordination chemistry of silver
2.2.2 - Crystal field splitting
2.2.3 - Silver coordinating complexes
2.2.4 - Applications of silver complexes as antimicrobial agents
2.3 - Properties of silver oxides
2.4 - Bioinorganic chemistry of silver-based antimicrobials
2.5 - Bacteria cell structure
2.5.1 - Outer cell layer composition of microorganisms
2.5.2 - DNA structure
2.6 - Mechanisms of antimicrobial activity of silver
2.6.1 - Evidence of silver ion attack on Gram-positive and Gram-negative bacteria cell membrane
2.6.2 - Evidence of silver ion attack on DNA (DNA denaturing)
2.6.3 - Attack on proteins
2.7 - Photocatalysis and antimicrobial activity on the surfaces and generation of reactive oxygen species
2.7.1 - Visible light activated photocatalysts
2.7.2 - Nanostructured modification for photocatalytic activation in the visible spectrum
2.7.3 - The use of photocatalysts as antimicrobials
2.7.4 - Nanoparticle photocatalyst semiconductors
3 - Materials and methods
3.1 - RF sputtering and silver oxide thin film coating
3.1.1 - Substrates
3.1.2 - Aim of thin film deposition
3.2 - Characterization of the thin films
3.2.1 - Morphology and topography of the surfaces
3.2.2 - Energy available on the surfaces for interaction
3.2.3 - Structural characterization
3.2.3.1 - X-ray diffraction analysis
3.2.3.2 - Evaluation of crystallite size
3.2.3.3 - Radial distribution function
3.2.4 - Optical characterization
3.2.4.1 - Transmittance
3.2.4.2 - Absorbance
3.2.4.3 - Energy bandgap
3.2.5 - Vibrational spectroscopy investigation of silver oxides
3.2.5.1 - FT-IR spectroscopy
3.2.5.2 - Raman spectroscopy
3.2.6 - Silver ion release
3.2.7 - X-ray photoelectron spectroscopy
3.2.8 - Antimicrobial tests
4 - Major challenges overcome by using silver/silver oxide thin films
4.1 - Monolithic films and nanoclustering
4.1.1 - Visible light activated photocatalyst
4.1.2 - Silver ion release in water and saline from the thin films
4.1.3 - Produced small nanoparticles for antimicrobial efficiency
4.1.4 - Effective contact killing of bacteria on the surfaces
5 - Further research
6 - Conclusion
References
Chapter 13 - Corrosion and Mott-Schottky probe of chromium nitride coatings exposed to saline solution for engineering and biome...
1 - Introduction
2 - Chromium nitride coatings for biomedical implants
2.1 - The films deposition
3 - The films characterization
3.1 - Chemical phase identification
3.1.1 - X-ray photoelectron spectroscopy (XPS)
3.1.2 - Raman spectroscopy
3.2 - The films morphology and structural properties
3.2.1 - Scanning electron microscopy (SEM)
3.2.2 - X-ray diffraction (XRD)
3.3 - Corrosion resistance
3.3.1 - Open circuit potential
3.3.2 - Potentiodynamic polarization
3.3.3 - Electrochemical impedance spectroscopy (EIS)
3.3.4 - Mott-Schottky analysis
3.4 - Immune cells response
4 - Conclusions
References
Chapter 14 - Characterization of cochleate nanoparticles for delivery of the anti-asthma drug beclomethasone dipropionate
1 - Introduction
2 - Controlling the size of empty and BDP drug-filled SPC liposomes
3 - Zeta potential of empty and BDP drug-filled liposomes
4 - Structure and morphology of SPC liposomes/cochleates
5 - SPS liposome size and zeta potential
6 - Size and zeta potential of cochleates
7 - Conclusion
References
Chapter 15 - Advances in nasal drug delivery systems
1 - Historical background
2 - Why the nasal route?
3 - Anatomy of the nose
4 - Nasal delivery
5 - Mechanisms of drug transport following intranasal administration
6 - Factors affecting nasal drugs delivery
7 - Barriers interfering with nasal drug delivery
7.1 - Low drug bioavailability
7.2 - Mucociliary clearance
7.3 - Enzymatic degradation
8 - Dosage forms for intranasal administration
8.1 - Liquid formulations
8.2 - Nasal powders
8.3 - Nasal gels
9 - Factors affecting particle deposition in the nasal cavity
10 - Mucoadhesive drug delivery systems
10.1 - Chitosan
10.2 - Sodium alginate
11 - Microspheres as a drug delivery system
11.1 - Preparation of microspheres
12 - Liposomes
12.1 - Classification of liposomes
12.2 - Liposomes in nasal drug delivery
13 - Nasal drops
14 - Nasal sprays
15 - Delivery devices of powdered nasal formulations
16 - Conclusions
References
Chapter 16 - Carbon nanotubes drug delivery system for cancer treatment
1 - Introduction
2 - Carbon-based materials
3 - Allotropes of carbon
4 - CNTs: structures and properties
4.1 - Single walled CNTs
4.2 - Multi walled CNTs
5 - Synthesis of CNTs
5.1 - Arc discharge
5.2 - Laser ablation
5.3 - Vapor phase deposition
6 - Functionalization
6.1 - Functionalization of CNTs with Pluronic
6.2 - Non-covalent functionalization; surfactant and chain onto the CNT
7 - CNTs as carriers for cancer treatment
7.1 - The potential use of CNTs as a carrier
7.2 - CNTs of targeted drug delivery
8 - Toxicity of CNTs
9 - Conclusions
References
Chapter 17 - Advances in multi-functional super magnetic iron oxide nanoparticles in magnetic fluid hyperthermia for medical app...
1 - Introduction
2 - Physics of IONPs
3 - Magnetic fluid heating
4 - Applications of SPIONs for hyperthermia
5 - Biocompatibility of SPIONs
6 - Techniques for physico-chemical properties of SPION
7 - Multipurpose smart system for medical applications
8 - In vivo applications of magnetic hyperthermia
9 - Conclusions
References
Chapter 18 - Taxane anticancer formulations: challenges and achievements
1 - Chemical structure and pharmacology of taxanes
2 - Traditional taxane formulations
3 - Stability of taxane formulations
4 - Undesirable effects of taxanes
5 - Nanotechnology as an approach to reduce taxane instabilities
5.1 - Liposomes
5.2 - Nanoparticles
5.3 - Polymeric micelles
5.4 - Dendrimers
5.5 - Cyclodextrins
6 - Lipid nanoemulsions
7 - Conclusions
Acknowledgement
References
Chapter 19 - Biomechanics of the mandible and current evidence base for treatment of the fractured mandible
1 - Structure of the mandible
2 - Forces relevant to mono-cortical fixation
3 - Rationale for fixation
4 - How the fracture heals
5 - Discussion
6 - Summary
References
Chapter 20 - Dental implants—the preparation of enamel, dentin, and bone by machining
1 - Introduction
1.1 - The structure of teeth and bone
1.1.1 - Micromachining of medical materials
Medical micromachining
1.2 - Trepanning of bone and teeth
1.3 - Trepanning machine
2 - Conclusions
Acknowledgments
References
Index
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Citation preview

ADVANCES IN MEDICAL AND SURGICAL ENGINEERING Edited by

Waqar Ahmed School of Mathematics and Physics & Lincoln School of Medicine, College of Science, University of Lincoln, Lincoln, United Kingdom

David A. Phoenix Office of the Vice Chancellor, London South Bank University, London, United Kingdom

Mark J. Jackson Kansas State University, Salina, KS, United States

Charalambos P. Charalambous Department of Trauma and Orthopaedics, Blackpool Victoria Hospital, Blackpool Teaching Hospitals NHS Trust, Blackpool, United Kingdom; School of Medicine, University of Central Lancashire, Preston, United Kingdom



Academic Press is an imprint of Elsevier 125 London Wall, London EC2Y 5AS, United Kingdom 525 B Street, Suite 1650, San Diego, CA 92101, United States 50 Hampshire Street, 5th Floor, Cambridge, MA 02139, United States The Boulevard, Langford Lane, Kidlington, Oxford OX5 1GB, United Kingdom Copyright © 2020 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. Library of Congress Cataloging-in-Publication Data A catalog record for this book is available from the Library of Congress British Library Cataloguing-in-Publication Data A catalogue record for this book is available from the British Library ISBN: 978-0-12-819712-7 For information on all Academic Press publications visit our website at https://www.elsevier.com/books-and-journals

Publisher: Mara Conner Acquisitions Editor: Fiona Geraghty Editorial Project Manager: Mariana L. Kuhl Production Project Manager: Nirmala Arumugam Designer: Victoria Pearson Typeset by Thomson Digital

Contributors Banu Abdallah School of Pharmacy and Biomedical Sciences, University of Central Lancashire, United Kingdom

Manchester; Division of Cardiovascular Sciences, University of Manchester, Manchester, United Kingdom

Hesham K. Abdelaziz Lancashire Cardiac Centre, Blackpool Victoria Hospital, Blackpool, United Kingdom; Department of Cardiovascular Medicine: Ain Shams University, Cairo, Egypt

Abdelbary M.A. Elhissi  College of Pharmacy and Office of the Vice President (Research and Graduate Studies), Qatar University, Doha, Qatar Aisha Ghauri  School of Mathematics and Physics & Lincoln School of Medicine, College of Science, University of Lincoln, United Kingdom; College of Pharmacy and Office of VP (Research and Graduate Studies), Qatar University, Doha, Qatar

Waqar Ahmed  School of Mathematics and Physics & Lincoln School of Medicine, College of Science, University of Lincoln, Lincoln, United Kingdom Sanil H. Ajwani  Department of Orthopaedics, Blackpool Victoria Hospital, Blackpool, United Kingdom

Imran Ghauri School of Mathematics and Physics & Lincoln School of Medicine, College of Science, University of Lincoln, United Kingdom; College of Pharmacy and Office of VP (Research and Graduate Studies), Qatar University, Doha, Qatar

Ahmed Aljawadi Wythenshawe Hospital, Wythenshawe, Manchester, England Abraham Atta Ogwu  East Kazakhstan State Technical University, Ust-Kamenogorsk, Republic of Kazakhstan

Douglas Hammond  School of Dentistry and School of Engineering, University of Central Lancashire, Preston, United Kingdom

Paul Callan Department of Cardiothoracic Transplantation and Mechanical Circulatory Support, Wythenshawe Hospital, Manchester, United Kingdom

Israr U. Hassan  Dhofar University, Salalah, Oman Chahinez Houacine  School of Pharmacy and Biomedical Sciences, University of Central Lancashire, Preston, United Kingdom

Charalambos Panayiotou Charalambous  Department of Trauma and Orthopaedics, Blackpool Victoria Hospital, Blackpool Teaching Hospitals NHS Trust, Blackpool; School of Medicine, University of Central Lancashire, Preston, United Kingdom

Luke Hughes  Department of Trauma and Orthopaedics, Blackpool Victoria Hospital, Blackpool Teaching Hospitals NHS Trust, Blackpool, United Kingdom

Michalis Charalambous  The Parapet Breast Unit, King Edward VII Hospital, Windsor, United Kingdom

Nozad Rashid Hussein  College of Pharmacy, Hawler Medical University, Iraq Luke J. Hyde  Purdue University, West Lafayette, IN, United States

Raouf Daoud  Breast Unit, Frimley Park Hospital, Firmley, United Kingdom

Mark J. Jackson  Kansas State University, Salina, KS, United States

Ioannis Dimarakis Department of Cardiothoracic Transplantation and Mechanical Circulatory Support, Wythenshawe Hospital,

Christopher Jump  ST3 Trauma and Orthopaedics, North Western Deanery, United Kingdom

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Contributors

David A. Phoenix  Office of the Vice Chancellor, London South Bank University, London, United Kingdom

Isabella Karat  Breast Unit, Frimley Park Hospital, Firmley, United Kingdom Iftikhar Khan  School of Pharmacy and Biomolecular Sciences, Liverpool John Moores University, Liverpool, United Kingdom

Abdul Rahman Phull  Department of Biochemistry, Shah Abdul Latif University, Khairpur, Pakistan

Maire-Clare Killen  Royal Victoria Infirmary, Newcastle upon Tyne; Charalambos Charalambous, Blackpool Victoria Hospital, Blackpool, United Kingdom

Saeed Ur Rahman  Institute of Thin films, Sensors and Imaging, School of Engineering and Computing, University of the West of Scotland, Scotland

Rukhsana Mahmood  School of Pharmacy and Biomedical Sciences, University of Central Lancashire, Preston, United Kingdom

David H. Roberts  Department of Cardiovascular Medicine: Ain Shams University, Cairo, Egypt

Abdul Majid  Department of Biochemistry, Shah Abdul Latif University, Khairpur, Pakistan

Grant M. Robinson  Purdue University, West Lafayette, IN, United States

Wael Mati  Department of Radiology, Blackpool Victoria Hospital, Blackpool, United Kingdom

Htet Sein  University of Lincoln, Lincoln, United Kingdom

Mohammad Najlah  Pharmaceutical Research Group, School of Allied Health, Faculty of Health, Education, Medicine and Social Care, Anglia Ruskin University, United Kingdom

Tapas Sen  School of Physical Sciences and Computing, University of Central Lancashire, Preston, United Kingdom Khurram Shahzad  Department of Radiology, Blackpool Victoria Hospital, Blackpool, United Kingdom

Farah Naz  Department of Biochemistry, Shah Abdul Latif University, Khairpur, Pakistan Abraham A. Ogwu  East Kazakhstan State Technical University, Ust-Kamenogorsk, Republic of Kazakhstan

Nathaniel T. Tsendzughul  School of Computing, Engineering and Physical Sciences, University of the West of Scotland, High Street, Paisley Campus, Paisley, Scotland

Huner Kamal Omer College of Pharmacy, Hawler Medical University, Iraq

Asma Vali  School of Pharmacy and Biomedical Sciences, University of Central Lancashire, Preston, United Kingdom

Ishrat Parveen  School of Pharmacy and Biomedical Sciences, University of Central Lancashire, Preston, United Kingdom

Justin Whitty  School of Dentistry and School of Engineering, University of Central Lancashire, Preston, United Kingdom

Yogita Patel-Sen  School of Physical Sciences and Computing, University of Central Lancashire, Preston, United Kingdom

Sakib S. Yousaf  School of Pharmacy and Biomedical Sciences, University of Central Lancashire, Preston, United Kingdom

Paul Sutton Department of Orthopaedics, Northern General Hospital, Sheffield, United Kingdom



C H A P T E R

1

Introduction to advances in medical and surgical engineering Waqar Ahmeda, David A. Phoenixb, Mark J. Jacksonc, Charalambos Panayiotou Charalambousdd,e a

School of Mathematics and Physics & Lincoln School of Medicine, College of Science, University of Lincoln, Lincoln, United Kingdom; bOffice of the Vice Chancellor, London South Bank University, London, United Kingdom; cKansas State University, Salina, KS, United States; d Department of Trauma and Orthopaedics, Blackpool Victoria Hospital, Blackpool Teaching Hospitals NHS Trust, Blackpool, United Kingdom; eSchool of Medicine, University of Central Lancashire, Preston, United Kingdom Clinicians aim to do an accurate and speedy diagnosis based on clinical symptoms, clinical signs and relevant investigations that will allow them to instate the appropriate treatment. In intervening to tackle medical or surgical conditions one aims to an effective and long lasting treatment, minimizing harm and facilitating early recovery. The aims of medical and surgical interventions are not only to prolong life but also improve quality of life. Hence, there is a constant drive to develop minimally invasive interventions, with fewer side effects, develop implants that will outlive the patient. Central do what Physicians and Surgeons do is “First do no harm” [1,2]. Close co-operation between clinicians and engineers is essential to allow the former express their needs and the latter to improve awareness as to what technology is available to meet those clinical needs. Working together, new horizons are explored and new opportunities are unleashed. This book aims to describe engineering advances in various areas of medicine and surgery both with regards to interventions but also radiological diagnostics. Potential targets of future engineering advances are also discussed. With an increasing aging but at the same time active population, musculoskeletal conditions, traumatic and degenerative, are on the rise [3–7]. The orthopedic surgeon strives to ease patients’ pain, facilitate, and speed healing and maintain high function. The initial chapters are devoted to engineering advances in various aspects of orthopedic surgery, the management of fractures and soft tissue injuries, the biomechanical improvements of shoulder and Advances in Medical and Surgical Engineering. http://dx.doi.org/10.1016/B978-0-12-819712-7.00001-2 Copyright © 2020 Elsevier Inc. All rights reserved.

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1.  Introduction to advances in medical and surgical engineering

knee arthroplasty implants, and the role of joint preserving procedures in the management of chondral disruption of the knee. The subsequent two chapters address two of the latest engineering developments in the management of heart conditions—mechanical circulatory support devices for patients with heart failure and transcatheter aortic valve implantation, a minimally invasive procedure that avoids more extensive open heart surgery. Choosing an appropriate and effective treatment relies on accurate diagnosis. The working diagnosis based on clinical symptoms and signs is often further investigated using, among others, radiological modalities. The next chapter explores engineering advances in magnetic resonance imaging that allow the clinician to view the body structures in both static and dynamic ways. Reconstructive surgery of the breast for malignant and benign conditions is on the rise [8,9]. The development of reliable and durable breast implants that can closely match the native tissues is essential, and the next chapter addresses engineering advances in this area. The subsequent chapters discuss the use of biomaterials in biomedical applications with particular reference to naturally occurring polymers, as well as basic engineering advances in making implants more infection resistant using antimicrobial silver oxide films. Advances in chromium-based coatings for artificial joint replacements with particular reference to corrosion resistance and biocompatibility are discussed. Despite having effective treatments for certain chronic conditions such as asthma or migraine, engineering advances have aimed to improve the availability and duration of action of therapeutic agents, to help improve the quality of life of sufferers. There has been extensive interest in using nanotechnology for drug delivery with nanocarriers providing specific tissue or cell targeting. In line with this there has been wide exploration of alternative routes of introducing agents into the body. The next two chapters discuss advances in the use of cochleates, a new class of nanocarriers derived from multilamellar liposomes for the pulmonary drug delivery of steroids in the treatment of asthma, and explore drug introduction into the body via the nasal route. Along with advances in the management of chronic conditions, recent years have seen the development of novel strategies for targeting cancer. The next three chapters explore some of these advances, in particular the use of carbon nanotubes and magnetic nanoparticles in malignant tissue and cell targeting as well as the development of novel nanocarrier systems (based on lipid and polymeric materials) for the delivery of taxane agents, a promising group of naturally occurring cancer fighting substances. The final two chapters explore engineering advances in managing mandibular conditions, namely the use of intermaxillary fixation for mandibular fractures and dental implants for the treatment of conditions such as temporomandibular joint dysfunction [10]. This book will hopefully increase awareness both among clinicians as well as engineers of what is achievable and stimulate further collaboration between clinicians and industry to further improve patient care.

References [1] Sokol DK. “First do no harm” revisited. BMJ 2013;347:f6426. [2] Charalambous CP. Career skills for surgeons. Springer; 2017.



References 3

[3] Prieto-Alhambra D, Judge A, Javaid MK, Cooper C, Diez-Perez A, Arden NK. Incidence and risk factors for clinically diagnosed knee, hip and hand osteoarthritis: influences of age, gender and osteoarthritis affecting other joints. Ann Rheum Dis 2014;73(9):1659–64. [4] Deshpande BR, Katz JN, Solomon DH, et al. Number of persons with symptomatic knee osteoarthritis in the US: impact of race and ethnicity, age, sex, and obesity. Arthritis Care Res (Hoboken) 2016;68:1743–50. [5] Johnson VL, Hunter DJ. The epidemiology of osteoarthritis. Best Pract Res Clin Rheumatol 2014;28:5. [6] Friedman S Mendelson F D.A. SM. Epidemiology of fragility fractures. Clin Geriatr Med 2014;30(2):175–81. [7] Mall NA, Chalmers PN, Moric M, Tanaka MJ, Cole BJ, Bach BR Jr, Paletta GA Jr. Incidence and trends of anterior cruciate ligament reconstruction in the United States. Am J Sports Med 2014;42(10):2363–70. [8] Nègre G, Balcaen T, Dast S, Sinna R, Chazard E. Breast reconstruction in France, observational study of 140,904 cases of mastectomy for breast cancer. Ann Chir Plast Esthet 2019; pii: S0294–1260(19)30128–1. [9] Yang RL, Newman AS, Lin IC, Reinke CE, Karakousis GC, Czerniecki BJ, Wu LC, Kelz RR. Trends in immediate breast reconstruction across insurance groups after enactment of breast cancer legislation. Cancer 2013;119(13):2462–8. [10] Lomas J, Gurgenci T, Jackson C, Campbell D. Temporomandibular dysfunction. Aust J Gen Pract 2018;47(4): 212–5.



C H A P T E R

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Engineering advances in promoting bone union Luke Hughesa, Charalambos Panayiotou Charalambousa,b, Ahmed Aljawadic a

Department of Trauma and Orthopaedics, Blackpool Victoria Hospital, Blackpool Teaching Hospitals NHS Trust, Blackpool, United Kingdom; bSchool of Medicine, University of Central Lancashire, Preston, United Kingdom; cWythenshawe Hospital, Wythenshawe, Manchester, England

1 Introduction Fast and complete bone healing (union) is essential in allowing quick recovery and rehabilitation following bone fractures or surgical osteotomies. Patient local and systemic factors as well as exogenous factors (such as mechanism of injury, magnitude of force applied, or surgical procedure), may predispose certain bone disruptions to delayed union or even nonunion. High rates of non-union are recognized in fractures of the scaphoid and the femoral neck (intra-capsular hip fractures). Chapter 3 discusses the pathways of bone to bone and tendon to bone union, and describes engineering advances that aim to facilitate these processes.

2  Principles of bone union Bone union may occur via secondary or primary union and these are discussed next. Before moving on it is important to understand the following terms: 1. Osteoinduction refers to the stimulation of osteogenesis through the recruitment of osteogenic cells to the site of injury. 2. Osteoconduction refers to the process by which bone grows onto a surface or into a structure (scaffold or implant). The exact structure of the surface or structure at a nanolevel may facilitate or hinder such growth. 3. Osteointegration refers to the incorporation of an exogenous material within bone—so that it is firmly attached to the surrounding bone or even replaced with time by bone. Advances in Medical and Surgical Engineering. http://dx.doi.org/10.1016/B978-0-12-819712-7.00002-4 Copyright © 2020 Elsevier Inc. All rights reserved.

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2.  Engineering advances in promoting bone union

2.1  Secondary bone union Secondary or indirect bone union (Fig. 2.1) describes the process of bone union that occurs following most untreated fractures or fractures treated without absolute stability upon fixation, for example, compression hip screw for fractured neck of femur; intramedullary nailing for diaphyseal fractures (Figs. 2.2 and 2.3). Immediately following fracture, bone and soft tissue trauma results in bleeding and hematoma formation. This hematoma connects the bony ends and creates a template for subsequent callus formation. The hematoma contains macrophages and other inflammatory cells, which initiate an acute inflammatory response by secreting various molecules, including interleukin (IL)-1, IL-6, IL-11, IL-18, and tumor necrosis factor alpha (TNF-α) among others [1]. The inflammatory response peaks within 24 hours and is complete after 7 days [2]. During this time phagocytes remove debris, while fibroblasts and mesenchymal cells migrate to the fracture site. Fibroblasts proliferate and a fibrin rich granulation tissue forms [3]. It is the mechanical environment, which drives the differentiation of mesenchymal cells along chondroblastic and later osteoblastic lineages [4,5]. Initially there is instability and hence movement at the fracture site. Chondrocytes produce type 2 collagen and soft callus forms after 2 weeks. As soft callus forms and connects the fracture ends some stability is conferred [6]. Ossification of the soft callus and bridging of the defect by hard callus further increases stability and rigidity such to permit weight bearing [7]. Chondrocyte apoptosis and re-absorption of the cartilaginous callus allows for angiogenesis and vessel ingrowth [8]. Although the hard callus is a rigid structure it does not replicate the biomechanical properties of normal bone. In order to achieve this the hard callus under goes remodeling in accordance with Wolffs’ law, wherein cyclic loading induces cells to modify the internal architecture of the bone trabeculae, such to best resist the mechanical stresses acting upon it [9]. This process involves osteoclastic resorption of hard callus and osteoblastic deposition of laminar bone and persists long after clinical union.

2.2  Primary bone union When surgical fixation ensures direct bone on bone contact and absolute stability, primary rather than secondary bone union predominates (Figs. 2.4–2.6). Under these conditions cutting cones are formed at the end of the osteons closest to the fracture site (Fig. 2.7) [10]. The osteon is the fundamental unit of cortical bone. They are cylindrical structures about 0.2 mm in diameter [11]. Cutting cones consist of osteoclasts, which act to create a microscopic cavity that crosses the fracture site. Osteoblasts follow the osteoclasts and lay down new bone in these cavities resulting in bone union and restoration of the Haversian systems. The Haversian systems comprise a central canal surrounded by concentrically arranged lamellae of bone matrix. The central canal carries the bone’s blood supply and delivers osteoblastic factors [12]. The bridging osteons subsequently undergo remodeling. This process allows for fracture union without callus formation.



FIGURE 2.1  Secondary bone healing.

2  Principles of bone union



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2.  Engineering advances in promoting bone union

FIGURE 2.2  Radiograph of mid-diaphyseal hu-

FIGURE 2.3  Radiograph of mid-diaphyseal humeral

meral fracture.

fracture, treated with intra-medullary nail. Evidence of secondary bone healing with callous formation.

3  Biological factors in bone union Biological factors are important for osteoinduction. Platelets are involved in clotting and initial hematoma formation. When activated they release a number of growth factors. These include platelet derived growth factor (PDGF), fibroblast growth factor (FGF), vascular endothelial growth factor (VEGF), and transforming growth factor beta (TGFβ) [13]. PDGF acts to recruit various inflammatory and progenitor cells to the fracture site [14]. FGF stimulates proliferation of mesenchymal cells [15]. VEGF promotes neovascularization of the callus with new blood vessels delivering cells and nutrients to the site of bone formation [16]. Bone morphogenetic proteins (BMP) are members of the transforming growth factor beta (TGFβ) superfamily and have been shown to have a diverse role in bone development and repair. Of the 15 BMPs in humans BMP-2 and 7 have been most studied in the context of bone union. These 



3  Biological factors in bone union

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FIGURE 2.4  Radiograph of radial diaphyseal fracture.

act by binding to osteoprogenitor cells, increasing the transcription of osteoinductive genes such as RUNX2 to enhance osteoblast differentiation [17]. Bone union relies upon the delivery of nutrients, growth factors and cells to the site of injury. This necessitates a connection with the systemic circulation. However, some bones have a tenuous blood supply. Examples include the scaphoid (Fig. 2.8), femoral neck (Fig. 2.9) and talus. At these sites a fracture can disrupt the blood supply to the proximal segment of bone. If this does occur there is a lack of factors necessary to promote bone union and non-union or avascular necrosis will result. If fracture union is compromised by the anatomy then it is important to consider ways to locally increase the concentration of growth factors. This can be achieved through the use of autologous bone marrow, which contains progenitor cells with both osteogenic and angiogenic properties. These cells are self-regenerating and are able to produce both VEGF and BMPs [18]. For larger defects autologous bone graft is preferred, as this also provides structural support. Other advantages of using autologous grafts include their low cost, low risk of disease transmission or immunological rejection. However, these must be balanced against the risk of donor site morbidity. Platelet rich plasma (PRP) is obtained from a sample of the patient’s blood drawn at the time of treatment. This is subsequently mixed with an anticoagulant and cold centrifuged by adjusting the acceleration force platelets sediment from solution [19]. In vitro studies have shown the local injection of PRP can stimulate osteoprogenitor cell proliferation, increase extracellular matrix formation, and promote angiogenesis [13].

3.1  Engineering advances influencing the local environment of growth factors in bone union Another method of achieving local growth factor delivery includes engineering processes, which achieve the incorporation of growth factors to a synthetic graft or fixation implant (Figs. 2.10–2.12). Recent work has explored these processes. One study investigated the use of 

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2.  Engineering advances in promoting bone union

FIGURE 2.5  Fluoroscopy image of radial diaphyse-

FIGURE 2.6  Radiograph of radial diaphyseal frac-

al fracture treated with dynamic compression plating.

ture undergoing primary bone healing in the absence of callus formation.

recombinant human PDGF incorporated in a beta-tricalcium phosphate scaffold, in patients requiring hindfoot or ankle arthrodesis. It demonstrated comparable outcomes to autograft in promoting fusion at 52 weeks [20]. VEGF-loaded nanographene coated internal fixation screws have been used to treat canine femoral neck fractures. These were prepared by way of direct liquid phase exfoliation of the graphite and VEFG loading by physical adsorption. Subsequent ELISA, X-ray, microangiography, and histological evaluation revealed sustained release of VEGF, without burst release, increased speed of fracture union, new bone formation areas and revascularization [21]. 



3  Biological factors in bone union

FIGURE 2.7  Primary bone healing.



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2.  Engineering advances in promoting bone union

FIGURE 2.8  Radiograph of a fracture through the

FIGURE 2.9  Radiograph demonstrating an intra-

waist of the scaphoid.

capsular fracture through the neck of a femur.

Sustained local release of VEGF with improved revascularization of canine femoral heads has also been achieved using a poly-lactic acid/glycolic acid delivery system, with formation of microspheres suspended within fibrin glue [22]. When BMP-2 was incorporated into a silica xerogel-chitosan hybrid coating of a porous hydroxyapitate scaffold, this demonstrated superior osteoblastic cell responses and increased bone formation, when compared to a control group wherein the hydroxyapitate scaffold was coated in the hybrid without incorporation of BMP-2 [23]. The addition of 50 micrograms of BMP-2 and its slow release from poly d, l-lactide coated titanium intramedullary implants has been demonstrated to facilitate callus consolidation and improve biomechanical properties (maximal load to failure and torsional stiffness) compared to uncoated implants at 28 and 42 weeks in a rat model [24]. An alternative method is to use stem cells, which when attached to implants can differentiate in vivo and locally produce the desired growth factors. One study isolated and cultured mesenchymal stem cells from bone marrow, suspended them in fibrin glue and then sprayed these on the surface of implants prior to implantation. Subsequent radiological and histological analysis of the ovine model demonstrated increased bone mass and increased bone contact when compared to controls [25].

4  Mechanical factors in bone union Mechanical factors influencing bone union include pressure, stability, strain, and fluid/ solid velocity. All forms of fixation work to provide stability and reduce movement at the fracture site. The type of bone union depends upon the degree of stability achieved. 



4  Mechanical factors in bone union

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FIGURE 2.11  AP fluoroscopy of fractured neck of femur treated with cannulated screws.

FIGURE 2.10  Radiograph demonstrating compres-

FIGURE 2.12  Lateral fluoroscopy of fractured neck

sion screw used to trad scaphoid fracture.

of femur treated with cannulated screws.

When using compression plating, rigid stability may be achieved, with zero strain and velocity. The result is primary bone union with Haversian remodeling. This necessitates bone contact and compression and hence can only be used in simple fracture configurations that permit this. In the absence of bone contact rigid stability may impair bone union, as the cutting cones cannot cross the gap and callus does not form. Other methods of orthopedic stabilization include external fixation and intramedullary nailing. These provide relative stability and bone union is secondary by way of endochondral ossification. In the early stages of secondary bone union, the hematoma provides little structural support and there is significant movement at the fracture site (i.e., high strain and



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2.  Engineering advances in promoting bone union

velocity). Recruitment of fibroblasts allows fibrous tissue to form. When fibrous tissue connects the boney ends, it acts to increased stability, such that strain and velocity decrease. With moderate strain and velocity, we then have chondrocyte differentiation, cartilaginous tissue formation, and soft callus forms. This soft callous gradually crystallizes and its stiffness increases. Strain and velocity are further reduced allowing for osteoblastic differentiation, which leads to hard callus and eventually bone formation [26]. Without stabilization and sufficient immobilization bone cannot advance through the stages of union and non-union may occur. However, some movement is important for the bone remodeling process. Wolff’s law dictates that upon cyclic loading, a bone with gradually remodel such that it becomes more resistant to that load [27]. Studies have shown that there is a positive correlation between compression rate and bone union [28]. If a bone is subject to reduced loading or strain over a prolonged period it becomes less dense and weaker due to the lack of stimulus required for remodeling [29]. This is the case in disuse osteopenia or bone loss as a result of implant stress shielding of bone. The remodeling response of bone to loading is via mechanotransduction, a process through which forces are converted to biochemical cell signals. With the role of mechanical stimuli on bone heading and remodeling described that it is clear to see how we must achieve a balance whereby sufficient stability is conferred to allow for union while subsequent rehabilitation to load the bone is necessary to restore its full strength by remodeling along lines of force transmission.

4.1  Engineering advances influencing the mechanical environment in bone union Bone loss poses a problem. Large defects do not permit direct bone on bone contact without significant shortening or deformity. Furthermore, gross instability can inhibit progression through the stages of secondary bone union. Traditionally defects were filled with autologous bone graft, to provide structural support. Scaffolds have been designed as an alternative to bone grafting. More recent advances in scaffold design have allowed for improved osteoconduction. To achieve this, scaffolds must be highly porous with interconnected pores of a diameter of at least 100 µm to allow ingrowth of cells and vessels [30]. Studies have looked at modifying the surface properties of scaffolds to further facilitate osteoconduction. One study immobilized gelatin to the surface of poly alpha-hydroxy acid films and porous scaffolds. Subsequent cell culture demonstrated significantly improved cell attachment and proliferation [31]. Likewise, the immobilization of Arg-Gly-Asp peptide of polycaprolactone films has been shown to significantly improve bone marrow stromal cell adhesion [32]. It is important that the kinetics of scaffold degradation match tissue regeneration. In this way the scaffold will provide sufficient support to allow for progression through the stages of bone union, while allowing for appropriate loading of the newly formed bone, necessary for remodeling. The process of mechanotransduction may be influenced with the use of locally applied ultrasound (Figs. 2.13 and 2.14). Ultrasound applied at specific frequency and intensity may distort and stimulate osteoprogenitor cells, impacting on gene regulation and chondrocyte differentiation with increased production of mRNA levels for Cbfa1/Rnx2 and osteocalcin necessary for osteogenesis [33]. Ultrasound also increases prostaglandin and nitric oxide production by osteoblasts [34]. Prostaglandins have been demonstrated to stimulate bone union





5  Future advances to facilitate bone to bone union

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FIGURE 2.13  Exogen kit—charger, power pack, transducer, strap, and gel.

and accelerate union [35], while nitric oxide facilitates bone remodeling [36]. Engineering advances have allowed the development of portable ultrasound devices that may be utilized at home by patients. A meta-analysis of 13 clinical studies, including 737 patients, compared pulsed electromagnetic fields (PEMF) or low intensity pulsed ultrasound (LIPUS) to placebo for the management of acute fractures. It concluded that PEMF and LIPUS accelerates time to radiological and clinical union of acute fractures, when these are treated non-operatively or fractures of the upper limb [37]. A meta-analysis of 24 unique randomized controlled trials and 429 patients demonstrated that LIPUS treatment resulted in a mean reduction in radiographic union of 39.8 days, with the greatest reduction in union time seen in fractures with a long natural union tendency. However, this analysis was unable to determine if LIPUS was useful in the prevention of delayed unions or if it led to quicker functional recovery, evaluated by return to work or active duty [38].

5  Future advances to facilitate bone to bone union Ongoing research aims to determine the optimum environment for facilitating bone union in terms of concentration, proportion, and temporal expression of various growth factors. Furthermore, ongoing work aims to facilitate the local availability of such growth factors by optimizing their incorporation within or on the surface of synthetic grafts and fixation implants and fine-tuning their release profiles to provide optimum bioavailability. This may lead to the development and widespread use in clinical practice of “biological” fixation implants that may help promote bone union, especially in those fractures predisposed to nonunion (such as “biological” hip and scaphoid fixation screws).



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FIGURE 2.14  Diagram of exogen mechanism of action.

References [1] Gerstenfeld LC, Cullinane DM, Barnes GL, Graves DT, Einhorn TA. Fracture healing as a post-natal developmental process: molecular, spatial, and temporal aspects of its regulation. J Cell Biochem 2003;88(5):873–84. [2] Cho TJ, Gerstenfeld LC, Einhorn TA. Differential temporal expression of members of the transforming growth factor beta superfamily during murine fracture healing. J Bone Miner Res 2002;17(3):513–20. [3] Rahn B. Bone healing: histologic and physiologic concepts. New York: Thieme; Stuggart; 2002. [4] Olivares-Navarrete R, Lee EM, Smith K, et al. Substrate stiffness controls osteoblastic and chondrocytic differentiation of mesenchymal stem cells without exogenous stimuli. PLoS One 2017;12(1):e0170312. [5] Ghiasi MS, Chen J, Vaziri A, Rodriguez EK, Nazarian A. Bone fracture healing in mechanobiological modeling: a review of principles and methods. Bone Reports 2017;6:87–100. [6] Dimitriou R, Tsiridis E, Giannoudis PV. Current concepts of molecular aspects of bone healing. Injury 2005;36(12):1392–404.



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[7] Gerstenfeld LC, Alkhiary YM, Krall EA, et al. Three-dimensional reconstruction of fracture callus morphogenesis. J Histochem Cytochem 2006;54(11):1215–28. [8] Ai-Aql ZS, Alagl AS, Graves DT, Gerstenfeld LC, Einhorn TA. Molecular mechanisms controlling bone formation during fracture healing and distraction osteogenesis. J Dent Res 2008;87(2):107–18. [9] Frost HM. Wolff’s Law and bone’s structural adaptations to mechanical usage: an overview for clinicians. Angle Orthod 1994;64(3):175–88. [10] HDaH. Fracture biology and biomechanics. Philadelphia: WB Saunders; 1993. [11] Osteon. Encyclopaedia Britannica Online; 2018. [12] Einhorn TA. The cell and molecular biology of fracture healing. Clin Orthop Relat Res 1998;(Suppl. 355):S7–S21. [13] Alsousou J, Thompson M, Hulley P, Noble A, Willett K. The biology of platelet-rich plasma and its application in trauma and orthopaedic surgery: a review of the literature. J Bone Joint Surg Br 2009;91(8):987–96. [14] Hollinger JO, Hart CE, Hirsch SN, Lynch S, Friedlaender GE. Recombinant human platelet-derived growth factor: biology and clinical applications. J Bone Joint Surg Am 2008;90(Suppl. 1):48–54. [15] Itoh N, Ornitz DM. Evolution of the Fgf and Fgfr gene families. Trends Genet 2004;20(11):563–9. [16] Street J, Bao M, deGuzman L, et al. Vascular endothelial growth factor stimulates bone repair by promoting angiogenesis and bone turnover. Proc Natl Acad Sci USA 2002;99(15):9656–61. [17] Jang WG, Kim EJ, Kim DK, et al. BMP2 protein regulates osteocalcin expression via Runx2-mediated Atf6 gene transcription. J Biol Chem 2012;287(2):905–15. [18] Sacchetti B, Funari A, Michienzi S, et al. Self-renewing osteoprogenitors in bone marrow sinusoids can organize a hematopoietic microenvironment. Cell 2007;131(2):324–36. [19] Dhurat R, Sukesh M. Principles and methods of preparation of platelet-rich plasma: a review and author’s perspective. J Cutan Aesthet Surg 2014;7(4):189–97. [20] DiGiovanni CW, Lin SS, Baumhauer JF, et al. Recombinant human platelet-derived growth factor-BB and betatricalcium phosphate (rhPDGF-BB/beta-TCP): an alternative to autogenous bone graft. J Bone Joint Surg Am 2013;95(13):1184–92. [21] Li S, Yuan H, Pan J, et al. The treatment of femoral neck fracture using VEGF-loaded nanographene coated internal fixation screws. PLoS One 2017;12(11):e0187447. [22] Zhang L, Zhang L, Lan X, et al. Improvement in angiogenesis and osteogenesis with modified cannulated screws combined with VEGF/PLGA/fibrin glue in femoral neck fractures. J Mater Sci Mater Med 2014;25(4):1165–72. [23] Jun S-H, Lee E-J, Jang T-S, Kim H-E, Jang J-H, Koh Y-H. Bone morphogenic protein-2 (BMP-2) loaded hybrid coating on porous hydroxyapatite scaffolds for bone tissue engineering. J Mater Sci Mater Med 2013;(3):773–82. [24] Schmidmaier G, Wildemann B, Cromme F, Kandziora F, Haas NP, Raschke M. Bone morphogenetic protein-2 coating of titanium implants increases biomechanical strength and accelerates bone remodeling in fracture treatment: a biomechanical and histological study in rats. Bone 2002;30(6):816–22. [25] Kalia P, Blunn GW, Miller J, Bhalla A, Wiseman M, Coathup MJ. Do autologous mesenchymal stem cells augment bone growth and contact to massive bone tumor implants? Tissue Eng 2006;12(6):1617–26. [26] Prendergast PJ, Huiskes R, Soballe K. ESB Research Award 1996. Biophysical stimuli on cells during tissue differentiation at implant interfaces. J Biomech 1997;30(6):539–48. [27] Ruff C, Holt B, Trinkaus E. Who’s afraid of the big bad Wolff?: “Wolff’s law” and bone functional adaptation. Am J Phys Anthropol 2006;129(4):484–98. [28] Hente R, Fuchtmeier B, Schlegel U, Ernstberger A, Perren SM. The influence of cyclic compression and distraction on the healing of experimental tibial fractures. J Orthop Res 2004;22(4):709–15. [29] Wolff J. The law of bone remodeling. Translation of the German 1892 edition. Berlin Heidelberg New York: Springer; 1986. [30] Ma PX. Biomimetic materials for tissue engineering. Adv Drug Delivery Rev 2008;60(2):184–98. [31] Liu X, Won Y, Ma PX. Surface modification of interconnected porous scaffolds. J Biomed Mater Res Part A 2005;74(1):84–91. [32] Zhang H, Lin CY, Hollister SJ. The interaction between bone marrow stromal cells and RGD-modified threedimensional porous polycaprolactone scaffolds. Biomaterials 2009;30(25):4063–9. [33] Chen YJ, Wang CJ, Yang KD, et al. Pertussis toxin-sensitive Galphai protein and ERK-dependent pathways mediate ultrasound promotion of osteogenic transcription in human osteoblasts. FEBS Lett 2003;554(1-2):154–8. [34] Reher P, Harris M, Whiteman M, Hai HK, Meghji S. Ultrasound stimulates nitric oxide and prostaglandin E2 production by human osteoblasts. Bone 2002;31(1):236–41.



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[35] Li M, Thompson DD, Paralkar VM. Prostaglandin E(2) receptors in bone formation. Int Orthop 2007;31(6):767– 72. [36] Krausz A, Friedman AJ. Nitric oxide as a surgical adjuvant. Future Sci OA 2015;1(1):FSO56. [37] Hannemann PF, Mommers EH, Schots JP, Brink PR, Poeze M. The effects of low-intensity pulsed ultrasound and pulsed electromagnetic fields bone growth stimulation in acute fractures: a systematic review and metaanalysis of randomized controlled trials. Arch Orthop Trauma Surg 2014;134(8):1093–106. [38] Rutten S, van den Bekerom MP, Sierevelt IN, Nolte PA. Enhancement of bone-healing by low-intensity pulsed ultrasound: a systematic review. JBJS Rev 2016;4(3):pii.



C H A P T E R

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Engineering advances in promoting tendon to bone healing Luke Hughesa, Charalambos Panayiotou Charalambousa,b a

Department of Trauma and Orthopaedics, Blackpool Victoria Hospital, Blackpool Teaching Hospitals NHS Trust, Blackpool, United Kingdom; bSchool of Medicine, University of Central Lancashire, Preston, United Kingdom

1 Introduction Tendon disruptions may be described as traumatic occurring secondary to a substantial force or degenerative occurring secondary to the application of no or minimal force. Tendon tears are quite common with over 9000 rotator cuff repairs completed in the United Kingdom between 2016 and 2017 [1]. Tendon tears are often avulsions from the bony insertion rather than mid-substance tears. Hence, surgery often aims to restore this bony attachment by reattaching the avulsed tendon to its native site of insertion. Similarly, ligamentous tears are very common. In 2017, 2,122 anterior cruciate ligament (ACL) reconstructions were recorded on the UK’s national ligament registry [2]. Ligaments contribute to joint stability and their loss may lead to symptomatic joint instability that hinders day-to-day activities or impairs occupational and recreational function. Ligamentous injuries may be mid-substance tears or bony avulsions. Although in some cases, surgery may attempt to repair the torn ligament or reattach the avulsed ligament to its bony insertion, in many cases the damaged native ligament cannot be restored and needs to be substituted (reconstructed) using a tendinous graft (autograft or allograft). Although primary repairs of acute ACLs tears are becoming more popular, ACL reconstruction using a tendon graft remains the mainstay of surgical management of such tears. Hamstring, patellar tendon, and quadriceps autografts are among the most commonly used in ACL reconstruction. Rapid tendon to bone healing is essential in the surgical repair of tendon avulsions from their bony insertion as well as in tendon or ligament reconstructions using autografts or autologous grafts, to facilitate rehabilitation and improvement in function. In highly demanding athletes, speedy tendon healing and fast rehabilitation may allow early return to sport at a highly competitive level. Advances in Medical and Surgical Engineering. http://dx.doi.org/10.1016/B978-0-12-819712-7.00003-6 Copyright © 2020 Elsevier Inc. All rights reserved.

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2  Principles of tendon to bone healing The enthesis is the connective tissue between tendon or ligament and bone. With a complex structure to accommodate force transmission and dissipation [3], entheses are either described as fibrous or fibrocartilaginous. Fibrous entheses are formed when the tendon gives rise to Sharpey’s fibres. These are mineralized collagen fibers, which perforate and anchor directly into bone or its surrounding periosteum [4]. Fibrous entheses (Fig. 3.1) are found in tendons that attach to the diaphyses of long bones. Fibrocartilaginous entheses (Fig. 3.2) are more common than fibrous. They form where tendon was first attached to primordial cartilage, which is progressively replaced on its inner surface by bone. They are found at epiphyses and apophyses and consist of four distinct zones: 1. Fibrous connective tissue 2. Uncalcified fibrocartilage

FIGURE 3.1  Diagram detailing fibrous enthesis.





2  Principles of tendon to bone healing

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FIGURE 3.2  Diagram detailing fibrocartilaginous enthesis.

3. Calcified fibrocartilage 4. Bone These create a structurally continuous gradient from uncalcified tendon to calcified bone [5]. The tidemark is a basophilic line that separates the uncalcified and calcified zones of the enthesis. Torn entheses heal poorly without surgical intervention. Even when surgically fixed, the healing process cannot re-establish the native tendon-bone insertion site formed during embryological development [6–8]. The healing enthesis tends to be fibrous [9] and forms through a process of: 1. hemostasis (minutes) 2. inflammation (0–7 days) 3. repair (5–14 days) 4. remodeling (>14 days) [3,5,10] When considering the healing of a tendon graft within an osseous tunnel (as in ACL reconstruction), one can expect to find the tendon-bone interface initially comprised of highly cellular fibrovascular tissue. New bone then begins to fill the tunnel. This new bone starts to invade the fibrous interface and then the tendon. In the final stages of healing, collagen fibers connect the tendon to the surrounding bone and are aligned in the direction of pull of the musculotendinous unit, closely resembling Sharpey’s fibres and a fibrous enthesis [6]. However, histologically there is a lack of the normal transition zones and instead scar tissue



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3.  Engineering advances in promoting tendon to bone healing

between the tendon and bone. The abrupt transition of stiffness between the tendon and bone impacts on force transmission and the fibrovascular scar tissue has been shown to be mechanically weaker and more prone to failure [6,7,11].

3  Factors affecting tendon-to-bone healing As with bone healing, both biological and mechanical factors play a role in the normal structural development of the enthesis and in tendon healing [11]. An understanding of these factors can help determine strategies to encourage tendon-to-bone healing following trauma and surgery.

3.1  Biological factors affecting tendon-to-bone healing Insulin-like growth factors (IGF) 1 and 2, platelet-derived growth factor (PDGF) and fibroblast growth factor (FGF) are prominent in the early phases of inflammation and proliferation, stimulating fibroblast migration and proliferation [12–14]. Metalloproteinases (MMPs) are responsible for the degradation of extracellular matrix and are important for angiogenesis and remodeling. Their activity is regulated by tissue inhibitors of matrix metalloproteinases (TIMPs) [15]. Both MMPs and TIMPs have been shown to be of elevated concentration at the site of tendon injury [16,17] and their regulation appears to be an active process. Studies have demonstrated that altering the MMPs verses TIMPS balance can influence tendon-to-bone healing following surgical repair, with the addition of intraarticular MMP inhibitor (α2 macrogobulin) resulting in increased fibrocartilage formation, increased numbers of Sharpey’s fibers and higher loads to failure [18,19]. Transforming growth factor beta (TGFβ) has a key role in tendon development and remodeling. It acts to promote cell migration, collagen production, and fibronectin binding, while at the same time regulating cell proliferation and protease activity [20–22]. The application of exogenous TGFβ has been demonstrated to increase the formation of collagen fibers bridging the tendon-bone interface with improved mechanical properties on pull out testing [23]. Bone morphogenic proteins (BMP), members of the TNF-β super-family, play a role in orchestrating the remodeling of bone at the tendon bone interface. BMP-12, 13, and 14 are expressed at the enthesis during embryogenesis while elevated levels of BMP-2 and 7 have been demonstrated to be elevated during tendo-osseus integration. Application of BMP-2, 12, and 13 at the time of surgical repair has been demonstrated to increase fibrocartilage formation, with more Sharpey’s fibers and higher loads to failure [24–26]. 3.1.1  Engineering advances influencing the local environment of growth factors in tendon to bone healing Increasing the local concentration of growth factors will facilitate tendon to bone healing. Engineering advances aim to accomplish this through local injection of growth factors, cell implantation, gene transfer or incorporation onto implants. Following local injection, there is often rapid clearance of growth factors, with many growth factors having a short half-life, susceptibility to inactivation, dilution and metabolism. If cells can be recruited to produce growth factors or growth factors can be fixed to implants, a more sustained delivery may be achieved.





3  Factors affecting tendon-to-bone healing

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Surgically anchoring periosteum onto the surface of reattached tendon promotes healing, due to precursor cells within the periosteum [27] and the application of mesenchymal stem cells to tendon grafts encourages fibrocartilage formation [28]. These mesenchymal cells can be genetically modified to produce growth factors (including BMP-2) using viral vectors. In a study of ACL repair, this was demonstrated to increase the formation of cartilage-like cells, reduce the size of the tibial bone tunnel and significantly increase the ultimate load and stiffness levels [29]. Alternatively cells and growth factors can be incorporated into an engineered matrix. Hydrogels comprise a network of polymer chains with a high water. As such they form a gel like substance with excellent handling properties. The addition of cells, growth factors, and drugs is through simple mixing and hydrogel can be injected percutaneously to fill defects [30]. The main disadvantage of hydrogels is their lack of biomechanical strength and inability to re-establish tissue continuity [31]. Scaffolds can be used to provide biomechanical support to the healing tendon, while acting as a vehicle for cells and new tissue formation [32]. The attachment of cells and/or growth factors to hydrogels and scaffolds can accelerate healing and improve the biomechanical properties of the graft. Cells commonly mixed with hydrogels or fixed to scaffolds include mesenchymal cells derived from adipose (ADSCs), bone marrow (BMSCs), and tendon (TSPCs) tissue. ADSCs have the advantage of being in abundance and easy to isolate [33], however, they have a clear preference to adipogenesis in vivo [34]. BMSCs are the most widely used stem cell. They exhibit superior tenogenic differentiation capacity when compared to ADSCs, however, harvesting causes greater donor site morbidity [35], and they are associated with higher rates of ectopic ossification and adhesion formation [36]. Although TSPCs have the greatest tenogenic abilities, they are the most difficult to isolate and confer the greatest risk of donor site morbidity [37]. One strategy to overcome the problems posed by harvesting of TSPCs is to use ADSCs or BMSCs that have been pre-differentiated toward the tendon lineage with the help of growth factors. Growth factors have been demonstrated to influence the rate of proliferation and degree of terminal differentiation of mesenchymal cells [38–40]. When incorporated into hydrogels/scaffolds, growth factors can demonstrate sustained release over time. This has been demonstrated to prolong the survival of stem cells and promote their differentiation [41,42]. Fixation devices for tendon repair can be coated in substrates designed to facilitate tendon to bone healing. When comparing the intubation of a poly-L-lactide acid screw (PLLA), a PLLA/β-tricalciumphospate screw, and a poly-l-liactide-co-glycolic acid/βtricalciumphospate screw with human osteoblast-like cells, one study demonstrated that cell number and cell contract points were significantly increased on the composite materials when viewed under scanning electron microscope [43]. As such it can be concluded that β-tricalciumphospate offers good ultrastructural properties for cell adhesion. Growth factors can also be incorporated into the implant/coating (Figs. 3.3, 3.4). One example is the addition of BMP-2 to the hydroxyapitate coating of ACL fixation screws. By using tagged linkBMP-2 peptides it was possible to quantify binding efficiency and release kinetics in a sheep model [44]. This study demonstrated sustained release of linkBMP-2 over a 6-week period. Comparison of the experimental group with a control group revealed improved histological scores in the experimental group, with the recruitment of mesenchymal cells to the interface between screw and tendon.



24

3.  Engineering advances in promoting tendon to bone healing

FIGURE 3.3  AP radiograph of knee demonstrating

FIGURE 3.4  Lateral radiograph of knee demon-

fixation of hamstring graft for anterior cruciate ligament reconstruction, endo-button proximal and inference screw distal.

strating fixation of hamstring graft for anterior cruciate ligament reconstruction, endo-button proximal and inference screw distal.

3.2  Mechanical factors affecting tendon-to-bone healing It is recognized that controlled loading of tendons is necessary to facilitate healing. Under controlled static and cyclical loading, mesenchymal cells and fibroblasts demonstrate increased proliferation and collagen production, with cyclical loading producing the greatest effect [45– 47]. Furthermore, the impact of administering exogenous growth factors is dependent upon the mechanical environment [48,49]. Mechanical stretch influences integrin-dependant signaling, with cytoskeletal organization realigning mesenchymal cells. Hence not only there is increased collagen production but better organization of fibers along lines of force [50]. In the case of ACL reconstruction, aggressive loading results in the accumulation of macrophages and catabolic inflammatory mediators (including MMPs) that can be detrimental to the healing process [46,51–53], while complete immobilization results in inferior mechanical properties [51,54]. These principles can be applied to both the post-operative rehabilitation period, with a controlled physiotherapy program likely to facilitate tendon-to-bone healing and increased strength of the repair; and to implant production. When cell-seeded decellularized extracellular matrices are exposed to cyclical stretching they demonstrate higher ultimate stresses and elastic moduli [55,56]. This is because of cell mediated remodeling and increased collagen deposition in the direction of strain [57,58]. 



3  Factors affecting tendon-to-bone healing

25

FIGURE 3.5  Exogen kit—charger, power pack, transducer, strap, and gel.

3.2.1  Engineering advances influencing the mechanical environment of tendon to bone healing When considering the engineering of scaffolds for the augmentation of tendon repairs, these have been developed to provide additional structural support. The scaffold can be used as an overlay on the surgical repair and functions in parallel to mechanically reinforce the repair by load sharing. It is important that the stiffness of the scaffold is such to permit load sharing, providing reinforcement, while at the same time ensuring some load will be applied to the repair site, required for optimal biologic repair. The scaffold should provide mechanical support over a sufficient period to permit healing and functional repair, while at the same time slowly degrading to permit full loading of the graft with time [59]. Low intensity pulsed ultrasound (LIPUS) has been demonstrated to improve tendon to bone healing (Figs. 3.5–3.7). This is through manipulation of mechanotransduction pathways such as to increase new bone formation, facilitate remodeling of the bone and fibrocartilage layers, leading to significantly improved biomechanical properties. Walsh and coworkers investigated the effects of LIPUS on tendon-bone healing in an intra-articular sheep model using digital extensor tendon autograft. LIPUS was demonstrated to increase cellular activity, improve tendon to bone integration and vascularity. Furthermore, stiffness and peak load were greater compared to the control group at 26 weeks post-surgery [60]. Lu and coworkers investigated the effects of continuous LIPUS on tendon to bone healing following partial patellectomy in rabbits. Subsequent radiographic, histologic, and biomechanical evaluations revealed significantly earlier and greater bone formation, improved tissue intergration, higher load to failure, and ultimate strength in the experimental group compared to controls [61]. Extracorporeal shockwave therapy (ESWT) may also stimulate tendon to bone healing. Wang and coworkers [62] demonstrated that a single treatment at 6 weeks postop increased bone formation and improved bone mineralization in a rabbit model. Histologically, the experimental group demonstrated more advanced remodeling in terms of better alignment of collagen fibers and thicker, more mature regenerated fibrocartilage at 8 and 12 weeks. 

26

3.  Engineering advances in promoting tendon to bone healing

FIGURE 3.7  Close up of transducer and strap.

FIGURE 3.6  Exogen kit demonstrating battery pack, transducer and strap applied.

Biomechanical testing also revealed higher tensile load and strength in the shock wave group compared with controls. These findings were supported by Chow and coworkers [63], with low and high intensity ESWT showing comparative osteogenesis enhancement and biomechanical advantages over controls at 8 and 12 weeks. EWST has also been demonstrated to improve outcomes following ACL reconstruction in humans, with X-rays and MRI showing significantly smaller tibial tunnels. This correlates with improved functional outcomes scores [64]. Human tenocytes also respond to EWS such as to show increased viability, proliferation, and tendon specific marker expression, as well as release of anti-inflammatory cytokines [65].

4  Future advances to facilitate tendon to bone healing Ongoing research aims to determine the optimum concentration, combination, and temporal expressions of various growth factors in tendon to bone healing. Furthermore, ongoing work aims to enhance the incorporation of such growth factors within or on the surface of synthetic grafts and fixation implants, as well as allowing their sustained release in a manner 

References 27

that provides optimum bioavailability. In addition, future engineering concepts should seek to address the disadvantage of harvesting host tendon grafts, through the development of synthetic grafts substitute that provides sufficient strength, deliver a high and sustained concentration of growth factors, recruit native cells, promote tissue formation and gradually resorb. A “biological” graft meeting these parameters would allow for immediate rehabilitation, while gradually evolving to create a tendon formed from host tissue with capacity to remodel and repair.

References [1] Digital N. Hospital Admitted Patient Care Activity, 2016-17. 2017. [2] The UK National Ligament Registry. 4th Annual Report 2018. [3] Lu HH, Thomopoulos S. Functional attachment of soft tissues to bone: development, healing, and tissue engineering. Ann Review of biomedical engineering. 2013;15:201–26. [4] Doschak MR, Zernicke RF. Structure, function and adaptation of bone-tendon and bone-ligament complexes. J Musculoskelet Neuronal Interact 2005;5(1):35–40. [5] Benjamin M, Kumai T, Milz S, Boszczyk BM, Boszczyk AA, Ralphs JR. The skeletal attachment of tendons-tendon “entheses”. Comp Biochem Physiol A Mol Integr Physiol 2002;133(4):931–45. [6] Rodeo SA, Arnoczky SP, Torzilli PA, Hidaka C, Warren RF. Tendon-healing in a bone tunnel. A biomechanical and histological study in the dog. J Bone Joint Surg Am 1993;75(12):1795–803. [7] Angeline ME, Rodeo SA. Biologics in the management of rotator cuff surgery. Clin Sports Med 2012;31(4):645– 63. [8] Juneja SC, Veillette C. Defects in tendon, ligament, and enthesis in response to genetic alterations in key proteoglycans and glycoproteins: a review. Arthritis 2013;2013:154812. [9] Oguma H, Murakami G, Takahashi-Iwanaga H, Aoki M, Ishii S. Early anchoring collagen fibers at the bonetendon interface are conducted by woven bone formation: light microscope and scanning electron microscope observation using a canine model. J Orthop Res 2001;19(5):873–80. [10] Voleti PB, Buckley MR, Soslowsky LJ. Tendon healing: repair and regeneration. Ann Rev Biomed Eng 2012;14:47– 71. [11] Thomopoulos S, Genin GM, Galatz LM. The development and morphogenesis of the tendon-to-bone insertion—what development can teach us about healing. J Musculoskelet Neuronal Interact 2010;10(1):35–45. [12] Molloy T, Wang Y, Murrell G. The roles of growth factors in tendon and ligament healing. Sports Med 2003;33(5):381–94. [13] Sciore P, Boykiw R, Hart DA. Semiquantitative reverse transcription-polymerase chain reaction analysis of mRNA for growth factors and growth factor receptors from normal and healing rabbit medial collateral ligament tissue. J Orthop Res 1998;16(4):429–37. [14] Jones JI, Clemmons DR. Insulin-like growth factors and their binding proteins: biological actions. Endocr Rev 1995;16(1):3–34. [15] Visse R, Nagase H. Matrix metalloproteinases and tissue inhibitors of metalloproteinases: structure, function, and biochemistry. Circ Res 2003;92(8):827–39. [16] Holmbeck K, Bianco P, Chrysovergis K, Yamada S, Birkedal-Hansen H. MT1-MMP-dependent, apoptotic remodeling of unmineralized cartilage: a critical process in skeletal growth. J Cell Biol 2003;163(3):661–71. [17] Choi HR, Kondo S, Hirose K, Ishiguro N, Hasegawa Y, Iwata H. Expression and enzymatic activity of MMP-2 during healing process of the acute supraspinatus tendon tear in rabbits. J Orthop Res 2002;20(5):927–33. [18] Bedi A, Kovacevic D, Hettrich C, et al. The effect of matrix metalloproteinase inhibition on tendon-to-bone healing in a rotator cuff repair model. J Shoulder Elbow Surg 2010;19(3):384–91. [19] Demirag B, Sarisozen B, Ozer O, Kaplan T, Ozturk C. Enhancement of tendon-bone healing of anterior cruciate ligament grafts by blockage of matrix metalloproteinases. J Bone Joint Surg Am 2005;87(11):2401–3210. [20] Glass ZA, Schiele NR, Kuo CK. Informing tendon tissue engineering with embryonic development. J Biomech 2014;47(9):1964–8. [21] Bennett NT, Schultz GS. Growth factors and wound healing: biochemical properties of growth factors and their receptors. Am J Surg 1993;165(6):728–37.



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[22] Zhu X, Hu C, Zhang Y, Li L, Wang Z. Expression of cyclin-dependent kinase inhibitors, p21cip1 and p27kip1, during wound healing in rats. Wound Repair Regen. 2001;9(3):205–212. [23] Yamazaki S, Yasuda K, Tomita F, Tohyama H, Minami A. The effect of transforming growth factor-beta1 on intraosseous healing of flexor tendon autograft replacement of anterior cruciate ligament in dogs. Arthroscopy 2005;21(9):1034–41. [24] Martinek V, Latterman C, Usas A, et al. Enhancement of tendon-bone integration of anterior cruciate ligament grafts with bone morphogenetic protein-2 gene transfer: a histological and biomechanical study. J Bone Joint Surg Am 2002;84-a(7):1123–31. [25] Kovacevic D, Rodeo SA. Biological augmentation of rotator cuff tendon repair. Clin Orthop Relat Res 2008;466(3):622–33. [26] Helm GA, Li JZ, Alden TD, et al. A light and electron microscopic study of ectopic tendon and ligament formation induced by bone morphogenetic protein-13 adenoviral gene therapy. J Neurosurg 2001;95(2):298–307. [27] Chen C-H, Chen W-J, Shih C-H, Yang C-Y, Liu S-J, Lin P-Y. Enveloping the tendon graft with periosteum to enhance tendon-bone healing in a bone tunnel: a biomechanical and histologic study in rabbits. Arthroscopy 2003;19(3):290–6. [28] Lim JK, Hui J, Li L, Thambyah A, Goh J, Lee EH. Enhancement of tendon graft osteointegration using mesenchymal stem cells in a rabbit model of anterior cruciate ligament reconstruction. Arthroscopy 2004;20(9): 899–910. [29] Chen B, Li B, Qi Y-J, et al. Enhancement of tendon-to-bone healing after anterior cruciate ligament reconstruction using bone marrow-derived mesenchymal stem cells genetically modified with bFGF/BMP2. Sci Rep. 2016;6:25940. [30] Yamada T, Gotoh M, Nakama K, Mitsui Y, Higuchi F, Nagata K. Effects of hyaluronan on cell proliferation and mRNA expression of procollagens alpha 1 (I) and alpha 1 (III) in tendon-derived fibroblasts from patients with rotator cuff disease: an in vitro study. Am J Sports Med 2007;35(11):1870–6. [31] Garg T, Singh O, Arora S, Murthy R. Scaffold: a novel carrier for cell and drug delivery. Crit Rev Ther Drug Carrier Syst 2012;29(1):1–63. [32] Turner NJ, Badylak SF. Biologic scaffolds for musculotendinous tissue repair. Eur Cells Mater 2013;25:130–43. [33] Zarychta-Wisniewska W, Burdzinska A, Kulesza A, et al. Bmp-12 activates tenogenic pathway in human adipose stem cells and affects their immunomodulatory and secretory properties. BMC Cell Biol 2017;18(1):13. [34] Neo PY, See EY, Toh SL, Goh JC. Temporal profiling of the growth and multi-lineage potentiality of adipose tissue-derived mesenchymal stem cells cell-sheets. J Tissue Eng Regen Med 2016;10(7):564–79. [35] Zhao C, Chieh HF, Bakri K, et al. The effects of bone marrow stromal cell transplants on tendon healing in vitro. Med Eng Phys 2009;31(10):1271–5. [36] Hsieh CF, Alberton P, Loffredo-Verde E, et al. Scaffold-free Scleraxis-programmed tendon progenitors aid in significantly enhanced repair of full-size Achilles tendon rupture. Nanomedicine (London, England) 2016;11(9):1153–67. [37] Yan Z, Yin H, Nerlich M, Pfeifer CG, Docheva D. Boosting tendon repair: interplay of cells, growth factors and scaffold-free and gel-based carriers. J Exp Orthopaed 2018;5(1):1. [38] Halper J. Advances in the use of growth factors for treatment of disorders of soft tissues. Adv Exp Med Biol 2014;802:59–76. [39] Han P, Cui Q, Yang S, Wang H, Gao P, Li Z. Tumor necrosis factor-alpha and transforming growth factor-beta1 facilitate differentiation and proliferation of tendon-derived stem cells in vitro. Biotechnol Lett 2017;39(5):711–9. [40] Tokunaga T, Shukunami C, Okamoto N, et al. FGF-2 stimulates the growth of tenogenic progenitor cells to facilitate the generation of tenomodulin-positive tenocytes in a rat rotator cuff healing model. Am J Sports Med 2015;43(10):2411–22. [41] Jeon O, Powell C, Solorio LD, Krebs MD, Alsberg E. Affinity-based growth factor delivery using biodegradable, photocrosslinked heparin-alginate hydrogels. J Control Release 2011;154(3):258–66. [42] Liu G, Pareta RA, Wu R, et al. Skeletal myogenic differentiation of urine-derived stem cells and angiogenesis using microbeads loaded with growth factors. Biomaterials 2013;34(4):1311–26. [43] Bernstein A, Tecklenburg K, Sudkamp P, Mayr HO. Adhesion and proliferation of human osteoblast-like cells on different biodegradable implant materials used for graft fixation in ACL-reconstruction. Arch Orthop Trauma Surg 2012;132(11):1637–45. [44] Lu Y, Markel MD, Nemke B, Lee JS, Graf BK, Murphy WL. Influence of hydroxyapatite-coated and growth factor-releasing interference screws on tendon-bone healing in an ovine model. Arthroscopy 2009;25(12): 1427–34. e1. 

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[45] Kuo CK, Tuan RS. Mechanoactive tenogenic differentiation of human mesenchymal stem cells. Tissue engineering Part A 2008;14(10):1615–27. [46] Wang JH. Mechanobiology of tendon. J Biomech 2006;39(9):1563–82. [47] Garvin J, Qi J, Maloney M, Banes AJ. Novel system for engineering bioartificial tendons and application of mechanical load. Tissue Eng 2003;9(5):967–79. [48] Aspenberg P. Stimulation of tendon repair: mechanical loading, GDFs and platelets. A mini-review. Int Orthop 2007;31(6):783–9. [49] Virchenko O, Aspenberg P. How can one platelet injection after tendon injury lead to a stronger tendon after 4 weeks? Interplay between early regeneration and mechanical stimulation. Acta Orthop 2006;77(5):806–12. [50] Xu B, Song G, Ju Y, Li X, Song Y, Watanabe S. RhoA/Rock, cytoskeletal dynamics, and focal adhesion kinase are required for mechanical stretch-induced tenogenic differentiation of human mesenchymal stem cells. J Cell Physiol 2012;227(6):2722–9. [51] Killian ML, Cavinatto L, Galatz LM, Thomopoulos S. The role of mechanobiology in tendon healing. J Shoulder Elbow Surg 2012;21(2):228–37. [52] Thomopoulos S. The role of mechanobiology in the attachmentof tendon to bone: International Bone and Mineral Society; 2011. [53] Riley G. Tendinopathy—from basic science to treatment. Nat Clin Pract Rheumatol 2008;4(2):82–9. [54] Galatz LM, Charlton N, Das R, Kim HM, Havlioglu N, Thomopoulos S. Complete removal of load is detrimental to rotator cuff healing. J Shoulder Elbow Surg 2009;18(5):669–75. [55] Androjna C, Spragg RK, Derwin KA. Mechanical conditioning of cell-seeded small intestine submucosa: a potential tissue-engineering strategy for tendon repair. Tissue Eng 2007;13(2):233–43. [56] Angelidis IK, Thorfinn J, Connolly ID, Lindsey D, Pham HM, Chang J. Tissue engineering of flexor tendons: the effect of a tissue bioreactor on adipoderived stem cell-seeded and fibroblast-seeded tendon constructs. J Hand Surg Am 2010;35(9):1466–72. [57] Gilbert TW, Stewart-Akers AM, Sydeski J, Nguyen TD, Badylak SF, Woo SL. Gene expression by fibroblasts seeded on small intestinal submucosa and subjected to cyclic stretching. Tissue Eng 2007;13(6):1313–23. [58] Nguyen TD, Liang R, Woo SL, et al. Effects of cell seeding and cyclic stretch on the fiber remodeling in an extracellular matrix-derived bioscaffold. Tissue engineering Part A 2009;15(4):957–63. [59] Killian ML, Cavinatto L, Galatz LM, Thomopoulos S. The role of mechanobiology in tendon healing. J Shoulder Elbow Surg 2012;21(2):228–37. [60] Azuma Y, Ito M, Harada Y, Takagi H, Ohta T, Jingushi S. Low-intensity pulsed ultrasound accelerates rat femoral fracture healing by acting on the various cellular reactions in the fracture callus. J Bone Miner Res. 2009;16(4):671–680. [61] Lu H, Liu F, Chen H, et al. The effect of low-intensity pulsed ultrasound on bone-tendon junction healing: Initiating after inflammation stage. J Orthop Res 2016;34(10):1697–706. [62] Wang L, Qin L, Lu HB, et al. Extracorporeal shock wave therapy in treatment of delayed bone-tendon healing. Am J Sports Med 2008;36(2):340–7. [63] Chow DH, Suen PK, Fu LH, et al. Extracorporeal shockwave therapy for treatment of delayed tendon-bone insertion healing in a rabbit model: a dose-response study. Am J Sports Med 2012;40(12):2862–71. [64] Wang CJ, Ko JY, Chou WY, et al. Shockwave therapy improves anterior cruciate ligament reconstruction. J Surg Res 2014;188(1):110–8. [65] de Girolamo L, Stanco D, Galliera E, et al. Soft-focused extracorporeal shock waves increase the expression of tendon-specific markers and the release of anti-inflammatory cytokines in an adherent culture model of primary human tendon cells. Ultrasound Med Biol 2014;40(6):1204–15.



C H A P T E R

4

Engineering advances in reverse total shoulder arthroplasty Christopher Jumpa, Charalambos Panayiotou Charalambousb,c a

ST3 Trauma and Orthopaedics, North Western Deanery, United Kingdom; bDepartment of Trauma and Orthopaedics, Blackpool Victoria Hospital, Blackpool Teaching Hospitals NHS Trust, Blackpool, United Kingdom; cSchool of Medicine, University of Central Lancashire, Preston, United Kingdom

1  Background of shoulder arthroplasty 1.1  Anatomic total shoulder arthroplasty (TSA) Replacement of the shoulder joint is performed to treat arthritis or trauma to the glenohumeral (GH) joint with the aim of reducing pain and improving motion and function. Anatomic total shoulder arthroplasty (TSA) involves replacement of the glenoid and humeral head with a prosthesis which matches the native configuration of a “ball” on the humeral head which articulates with a “socket” on the glenoid (Fig. 4.1). Anatomic TSA requires an intact rotator cuff to provide stability and motion. The rotator cuff muscles attach around the humeral head and compress it into the glenoid. This creates a fulcrum on which the deltoid can act to elevate the arm [1,2]. Compression of the humeral head into the glenoid and its effect on glenohumeral joint stability has previously been explained as a ball (humeral head) sitting in a concavity (glenoid). This principle, termed concavity-compression states that the deeper the concavity the larger a displacing force needs to displace the ball for a certain compressive load (Fig. 4.2) [3]. With tears of the rotator cuff tendon, this compression is lost altering the biomechanical forces passing through the joint and the humeral head becomes unstable. In a glenohumeral joint with an intact rotator cuff the humeral head remains in a fixed position throughout its range of motion relative to the joint center line (a line perpendicular to the glenoid articular surface) [4]. Without an intact rotator cuff the humeral head is unstable relative to the center line which results in compromised deltoid function and increased wear to the articular surfaces. The humeral head tends to migrate superiorly causing impingement on the acromion resulting in erosion. Arthritis in the presence of rotator cuff deficiency is known as rotator cuff arthropathy [5]. Advances in Medical and Surgical Engineering. http://dx.doi.org/10.1016/B978-0-12-819712-7.00004-8 Copyright © 2020 Elsevier Inc. All rights reserved.

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4.  Engineering advances in reverse total shoulder arthroplasty

FIGURE 4.1  Anterior view of the anatomy of a right Total Shoulder Arthroplasty (TSA) matching the native configuration of a ‘ball’ on the humeral head, which articulates with a ‘socket’ on the glenoid.

FIGURE 4.2  Principle of concavity-compression. The deeper the concavity the larger a displacing force is needed to displace the ball (humeral head) for a certain compressive load. Based on Figure 8.5 from Chapter 8, Principles of glenohumeral stability. In: Matsen F, Lippitt S, DeBartolo S, eds. Shoulder Surgery: Principles and Procedures. Philadelphia: WB Saunders; 2004;83 with permission.

Performing an anatomic TSA in patients with a deficient rotator cuff results in an unstable shoulder with the humeral head continuing to move unrestrained relative to the glenoid. This may result in eccentric glenoid loading, early wear of the glenoid component, loosening, and a high failure rate [6]. Rotator cuff arthropathy patients have previously been treated with shoulder hemiarthroplasty to avoid these glenoid component complications. However, although these patients





1  Background of shoulder arthroplasty

33

may achieve reduced pain levels they tend to have limited shoulder function and suffer glenoid and acromial bone erosion [5,7–12]. Despite attempts at improving anatomic TSA designs through enhanced component constraint to limit the superior movement of the humeral relative to the glenoid component, these implants often failed due to excessive forces at the bone implant interface resulting in loosening [13]. The reverse total shoulder arthroplasty (RTSA) was designed to overcome these problems and gives certain mechanical benefits.

1.2  Reverse total shoulder arthroplasty (RTSA) RTSA reverses the normal anatomy of the shoulder with the “ball” (convex surface) positioned on the scapula and the “socket” (concave surface) on the proximal humerus (Fig. 4.3). Its configuration is designed to overcome the loss of rotator cuff function and optimize the action of the deltoid. Joint stability is increased by resistance to the shoulder moving superiorly when deltoid contracts. Compared to the anatomic TSA, the reverse implant moves the center of rotation of the joint medially which increases the moment arm of the deltoid muscle, enhancing its efficiency [14]. The original indication of RTSA was to treat rotator cuff arthropathy in elderly, low-demand patients. A minimum age of 65 years is currently

FIGURE 4.3  Anterior view of the anatomy of a right Reverse Total Shoulder Arthroplasty (RTSA) with the ‘ball’ (convex surface) (glenosphere) positioned on the scapula and the ‘socket’ (concave surface) (humeral cup) on the proximal humerus.



34

4.  Engineering advances in reverse total shoulder arthroplasty

TABLE 4.1  Indications for RTSA. Rotator cuff arthropathy Inflammatory arthritis with cuff tears Proximal humerus fractures and sequelae Reconstruction after resection for tumor Revision arthroplasty in cuff deficient patients Primary osteoarthritis in elderly patients with an intact rotator cuff Pseudoparalysed shoulder in absence of osteoarthritis

considered acceptable for performing RTSA [15]. With experience, the number of indications of RTSA has increased and the accepted age of patients undergoing RTSA has decreased. See Table 4.1 for the indications of RTSA.

2  Design systems and their development The first shoulder arthroplasty is thought to have been performed by Jules Emile Pean in 1893 to treat tuberculosis. It consisted of a prosthesis made of rubber, platinum, and leather, which had to be later removed due to postoperative infection [16]. In the 1950s Kruger used a humeral implant to treat aseptic necrosis of the proximal humerus [17]. From 1953 onward, Charles Neer continued to develop shoulder arthroplasty designs, starting with the Neer Mark I monoblock non-constrained vitallium humeral prosthesis [18]. This prosthesis was used to successfully treat seven patients with acute proximal humerus fracture-dislocations and five patients with an old proximal humerus fracture complicated by avascular necrosis. Neer’s second generation prosthesis featured a modular design which articulated using a Morse Taper principle. Neer’s subsequent prostheses were more anatomically designed with regards to the humeral head allowing adjustments to be made to offset, head diameter and version to match the native anatomy and hence improve function [19]. Due to poor outcomes and unpredictable function of the anatomic prostheses in patients with rotator cuff tear arthropathy, Neer developed a reverse prosthesis. Reversal of the glenoid and humeral implants originated due to the inability to have a glenoid implant large enough to prevent proximal migration of the prosthesis. The Mark I reverse prosthesis featured a large glenosphere with the aim of achieving a highly stable joint and prevent proximal humeral migration. However the size of the glenosphere prevented reattachment of the rotator cuff to the humerus [20]. The Neer Mark II reverse prosthesis featured a smaller glenosphere which allowed repair of the rotator cuff but resulted in reduced ROM and excessive constrain subsequently causing high rates of glenoid loosening [21]. To improve ROM and limit constrain, the Neer Mark III reverse prosthesis featured axial rotation between the humeral stem and diaphysis, however this resulted in an increased rate of complications including dislocation. Further work focused on the importance of the rotator cuff in order to achieve improved ROM. Several other designs were developed with the aim of overcoming a deficient rotator cuff, often featuring a third implant, which articulated





2  Design systems and their development

35

between the glenoid and humeral component. Unfortunately many of these designs failed to improve outcomes and were not developed further [22–27]. In 1987 Paul Grammont described the RTSA as a treatment for rotator cuff deficient patients with a novel design that allowed the deltoid muscle to compensate for the non-functional rotator cuff [28]. His design consisted of a polyethylene humeral cup and a cobalt-chromium-molybdenum glenosphere. His implant had four key design principles that are still followed with the latest designs. These included: 1. Inherent stability of the prosthesis 2. Center of rotation is medialized and distalized 3. Deltoid muscle lever arm is effective from initiation of movement 4. Glenosphere is large and the humeral cup is relatively small creating a semi-constrained joint

2.1  Prosthesis stability Glenohumeral joint stability can be assessed quantitatively using the stability ratio and balance stability angle. The stability ratio is the maximum distracting force that can be stabilized by a compressive force. In vitro the stability ratio is assessed by applying a compressive load and then applying an increasing displacing force until dislocation occurs. The forces before the point of dislocation are then compared to produce a stability ratio. A native GH joint has a ratio of 0.5, a TSA of 1.0, and a RTSA a ratio of greater than 2.0 [29,30]. The balance stability angle (BSA) is the angle between the center and edge of the glenoid articular surface in any direction [4]. In a native joint this surface area is reduced in conditions such as glenoid dysplasia resulting in a lower BSA and less stable joint (Fig. 4.4). An anatomic TSA has a maximum BSA of 30° before dislocation occurs [31]. When a TSA is in situ with an intact rotator cuff and deltoid, this is easily managed. However when the rotator cuff is deficient, the shoulder is less stable as contraction of the deltoid creates a superiorly directed force causing dislocation as opposed to abduction. The RTSA will allow a BSA of 45° due to its more stable design which features the glenosphere and humeral cup surfaces having the same degree of curvature and the humeral component having an increased concavity [32,33]. Further increasing the concavity of the humeral cup will result in increasing stability.

FIGURE 4.4  Balance stability angle (BSA) is the angle between the center of the glenoid articular surface and the edge of the glenoid articular surface in any direction.



36

4.  Engineering advances in reverse total shoulder arthroplasty

Other factors that increase the stability of the shoulder replacement include the compressive forces acting on the GH joint and use of an inferior offset. In the RTSA performed with a non-functional rotator cuff, the tensioning of deltoid has the biggest compressive effect across the GH joint.

2.2  Medializing and distalizing the center of rotation Grammont’s initial design consisted of a glenoid component forming two-thirds of a sphere and a humeral stem with a proximal concavity forming one-third of a sphere. By medializing the center of rotation by 10 mm and distalizing the center of rotation by 10 mm Grammont’s design increased the moment arm of the deltoid by 20% and 30%, respectively [34–36]. Distalizing the glenoid implant also has the advantage of reducing impingement of the humerus on the scapula neck during adduction. Medializing also has the benefit of recruiting more deltoid fibers to act as abductors [37]. Grammont’s next generation prosthesis introduced in 1991, the Delta III, further medialized the center of rotation by changing the two-thirds glenosphere to a hemisphere [37]. In a native GH joint the center of rotation changes constantly during movement although it lies close to the center of the humeral head. In RTSA the center of rotation is in a fixed, medialized location (Fig. 4.5). With movement of the shoulder a combination of compressive and shear forces continuously pass through this location. Contraction of deltoid results in superiorly directed shear forces to pass through the glenosphere bone-implant interface risking the stability of implant fixation. Medializing the center of rotation minimizes the lever arm for these shear forces and changes the direction of the culmination of forces at the center of rotation from being directed superiorly to medially into the scapula hence converting them from a shear force to a compression force resulting in a reduced risk of glenoid baseplate loosening [38,39].

FIGURE 4.5  Diagram showing native glenohumeral joint (A) and reverse total shoulder arthroplasty (B). RTSA results in medialization of the center of rotation (dot) compared to native joint. Horizontal arrow demonstrates distance of medialization. RTSA also results in distalization (dotted arrow). Based on Figure 2, Ackland et al. Prosthesis design and placement in reverse total shoulder arthroplasty, Journal of Orthopaedic Surgery and Research (2015) 10:101 with permission.





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2.3  Deltoid muscle function Medializing the center of rotation results in an increase in the moment arm of the deltoid by more than 20% [36,39,40]. Deltoid function is particularly improved in early abduction. This compensates for the lack of supraspinatus action. Distalizing the center of rotation increases tension in deltoid further improving its function [36,37,41]. At the same time it provides a greater impingement free ROM. Hence, RTSA overall reduces the forces needed for deltoid to contract and move the upper limb.

2.4  Range of motion post RTSA In the native joint, the anterior deltoid is predominantly a flexor, middle deltoid is an abductor, and posterior deltoid an extensor. The posterior and anterior deltoid subregions also contribute to external rotation and adduction, and internal rotation and adduction movements respectively. RTSA makes all three deltoid sections predominantly abductors, resulting in decreased ROM [42,43]. This along with a lack of rotation due to deficient rotator cuff muscles causes reduced rotation after RTSA [44]. These movements are crucial for performing activities of daily livings (ADLs) such as personal hygiene (reaching the bottom, washing hair) [37,45]. Medializing the center of rotation may also reduce subscapularis and teres minor function due to a reduction in muscle tension. This further compounds the lack of rotation post RTSA [46]. Grammont created a semi-constrained joint using a large diameter glenoid component hemisphere and relatively smaller humeral cup with the aim of producing a larger arc of motion [37,47]. A larger glenoid component results in a greater range of adduction prior to impingement but offers no advantage in the degree of rotational movement [48].

2.5  Implant designs In addition to improvements with implant design, options of implant fixation have been developed which give advantages in certain patient groups. Standard humeral implants may be cemented or cementless with press fit fixation. Cemented fixation gives the benefits of allowing better implant positioning as adjustments can be made during insertion, it provides instant stability and has lower risks of early stem migration and intraoperative humerus fracture [49]. Stemless implants are inserted into the humeral metaphysis without cement fixation. They are often used in the younger patient with adequate bone stock. Short stem implants are also available and along with stemless designs may preserve humeral bone which allows easier revision surgery in the future if needed. Cementless implants require a stable fixation at the time of insertion to allow bone ingrowth. They are coated in special materials with a specific porosity which allows osteointegration to occur. Many modern implant systems now feature modularity allowing an implant to be separated and recombined as needed which is advantageous in certain situations [50–52]. For example, when a proximal humerus fracture is treated with a hemiarthroplasty the tuberosities may suffer non-union or the rotator cuff may tear [53]. A modular system allows the hemiarthroplasty to be converted to RTSA without needing to revise the humeral stem. The articular surface aspect of the implant can be exchanged for a humeral cup which will then



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4.  Engineering advances in reverse total shoulder arthroplasty

articulate with an implanted glenosphere. Similarly, many anatomic TSA implant systems can now be converted to RTSA if the patient suffers a rotator cuff tear without needing to revise the humeral stem.

3  Clinical outcomes of shoulder arthroplasty Common anatomic TSA complications include implant loosening, glenoid wear, instability, rotator cuff tear, periprosthetic fracture, infection, hematoma, and nerve injury [54]. Long term follow-up studies report over 80% implant survival at 15 years with relief of pain achieved in 83% of patients [55,56]. Radiological lucency at the bone-cement interface of humeral and glenoid implants has been shown in 60%–75% of patients postoperatively after 10 years, but many of these are asymptomatic [55]. Complications after RTSA are similar to anatomic TSA as described earlier. In RTSA instability, periprosthetic fracture and infection are the commonest encountered complications [54,57,58]. These complications along with humeral and glenoid loosening are the commonest reasons for patients undergoing revision RTSA. Complications after revision RTSA are approximately 3 times higher than primary RTSA [57,59–61]. During the early years of RTSA initial complication rates were as high as 68%, with scapular notching one of the commonest complications [62–66]. As experience and understanding of RTSA biomechanics has improved and implant designs and surgical techniques have evolved, complications including notching have reduced in prevalence. More recent studies report complication rates after RTSA of 13%–20% [57,59,60,67–70]. RTSA implant survival is currently approximately 91% at 10 years follow-up [71,72]. Long-term follow-up studies have demonstrated RTSA results in significantly improved Constant-Murley scores, ROM, and pain levels [38,66,68,73–76]. The Constant-Murley score is a combined objective and subjective assessment of a patient’s pain level, activity level, abduction strength and ROM in flexion, abduction, external rotation, and internal rotation. It is widely recognized as a reliable and useful assessor of overall shoulder function. In a recent review of 14 studies [77], RTSA has been reported as improving ConstantMurley scores on average from 19.9 to 60.8 with follow-up ranging from 22 to 69.6 months [38,63,64,66,73–76,78–84]. After RTSA the most significant improvements in ROM occur in forward flexion with improvements reported up to 82°. Improvements in external rotation are generally poorer and reported inconsistently in the literature [38,63–66,73–75,79,81–89]. The majority of gains in ROM occur in the first 6 months then continue until 2 years postoperatively before plateauing [90,91]. Studies comparing a Grammont-style (medialized center of rotation) RTSA implant design against newer prostheses including a curved short stem design demonstrate significant improvements in Constant-Murley score and pain levels with both prostheses but with greater relative improvements after implantation of the more modern prostheses [92]. Similar improvements in postoperative ROM have been shown with both designs except for external rotation which improved greater after implantation of the modern design. This result was significant when the glenoid was lateralized with a humeral head bone autograft prior to 



3  Clinical outcomes of shoulder arthroplasty

39

glenoid baseplate fixation. Comparison of postoperative complications of new and earlier designs demonstrate significantly lower rates of notching and radiolucency at the glenoidbone interface with 5% of newer designs suffering notching and 39% of original designs after a mean follow-up of 32 months [92]. With increasing experience, RTSA has been used to manage increasingly complex conditions including sequelae of proximal humerus fractures including varus or valgus malalignment and osteonecrosis. Significant improvements in Constant-Murley scores, ROM, strength, pain levels, and patient satisfaction after treatment of these conditions has been demonstrated [93,94]. RTSA has also been used to treat nonunions after proximal humerus fracture and although significant improvements in ROM and Constant-Murley scores have been shown, high complication rates have also been reported with notching in 50% of patients, dislocation in 34% of patients, and 28% of patients requiring revision surgery [95]. Although these findings are from small studies, they demonstrate the requirement of continuous developments in surgical techniques and implant designs to allow RTSA to be used successfully to treat more challenging, complex conditions. Described further are some of the common complications of RTSA.

3.1 Infection Studies in the literature report up to a six-fold increase in infection rate in RTSA compared to anatomic TSA [33,96]. It is suggested to be caused by a combination of the large surface area of the prosthesis and the formation of a dead space under the acromion due to the lack of rotator cuff, which can accommodate hematoma formation, leading to subsequent infection. Formation of hematomas occurs in 20.6% of patients post RTSA [79]. Infection rate after primary RTSA has been reported in up to 15% of patients with Propionibacterium acnes, staphylococcus aureus, and staphylococcus epidermidis the commonest pathogens [97,98]. Routine drain usage has been recommended to minimize the risk of hematoma formation.

3.2  Scapular notching Scapular notching is a complication specific to RTSA. It occurs when adduction of the arm results in impingement between the scapular neck and medial aspect of the humeral prosthesis resulting in scapular bone erosion [66]. Impingement causes increased wear of the humeral polyethylene cup producing debris resulting in osteolysis causing additional notching leading to loosening of the implants [64,79,99–101]. Scapular notching affects between 50% and 96% of patients after RTSA and with longer term follow-up incidence and severity often increases [37,38,62–64,66,72,79,101–103]. 48% of cases are present at 1 year, 60% at 2 years, and 68% at 3 years. However, not all cases of notching progress with some stabilizing at 2 years follow-up [104]. Severe notching has been reported as affecting up to 45% of patients post RTSA and is more common when using a superolateral approach than a deltopectoral approach [103,104]. Incidence of notching increases with activity levels and preoperative diagnosis. Most studies in the literature show no relationship between notching and pain, function or postoperative Constant-Murley scores [63,79,83,102,104]. ROM and strength have been shown to be higher in the absence of notching. Progressive development of notching will result in complications including glenoid loosening. Incidence of



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4.  Engineering advances in reverse total shoulder arthroplasty

scapular notching has been reduced by improved implant design and surgical techniques as detailed further.

3.3 Instability Primary RTSA dislocation occurs in up to 14% of patients [38,60,63,79,88,100]. Dislocation rates of revision RTSA is approximately twice as common [60,63]. The majority of dislocations occurs within several months of surgery and are usually a result of surgical technique [105]. Multiple factors contribute to dislocation including implant malpositioning, glenosphere size, impingement, infection, axillary nerve injury, trauma and insufficiency of subscapularis, deltoid and pectoralis major muscles [63,106–109]. Implant malposition results in altered local muscle tensioning and function creating instability and dislocation.

3.4 Fracture Postoperative fractures of the acromion occur in 1%–4% [83,84]. This is thought to be caused by the subluxed humeral head eroding the acromion and due to increased tension of deltoid post RTSA leading to eventual fracture [110]. This results in significant loss of ROM and function, although with limited pain [84,111]. Postoperative humeral fractures occur in 1.6% of patients, commonly due to low energy falls [70]. Intraoperative fractures can occur with glenoid reaming and during the preparation of the proximal humerus.

3.5  Humeral loosening Humeral loosening occurs in up to 10% of patients [69,70]. In RTSA medialization of the center of rotation shifts stresses from the glenoid to the humeral implant [38,112,113]. Loss of bone at the proximal humerus, particularly the greater tuberosity is associated with increased risk of loosening [113–116]. This is common in certain indications of RTSA including fracture and post tumor resection.

3.6  Glenoid loosening Glenoid loosening is largely due to infection or technical error, occurring after 9% of RTSA procedures [69,71,117] A baseplate positioned too superior or with excessive superior inclination creates shear stresses at the bone-implant interface increasing the risk of loosening [109,118].

3.7  Nerve palsy Risk of nerve injury has been reported as 10 times higher in RTSA than conventional TSA, most commonly affecting the axillary nerve [119]. Insertion of RTSA results in increased strain on the nerve roots which can lead to nerve palsy [120]. This risk can be minimized by not at­ tempting to excessively lengthen the limb too much with the aim of tensioning the deltoid muscle. Lengthening by more than 2 cm results in a significantly increased risk of nerve in­ jury [119,121]. In the majority of cases nerve palsy is transient.





4  Advances in implant design and surgical techniques

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3.8  Reduced rotation RTSA results in reduced postoperative external rotation due to a reduction of the moment arms of teres minor and posterior deltoid [42]. The lack of effective rotator cuff muscles further compounds the loss of rotation after RTSA. From the literature, the majority of patients have no improvement in external rotation post RTSA, with some studies reporting up to one-third of patients having reduced external rotation [63]. Studies that do report increases in external rotation postoperatively report maximal improvements up to 33° [65,79,82].

4  Advances in implant design and surgical techniques Reported complication rates of RTSA are significantly greater than anatomic TSA. As a result, there have been great efforts to develop implant designs and also surgical techniques in order to minimize such complications and improve clinical outcomes for these patients.

4.1  Scapular notching Medializing the center of rotation gives the advantage of increasing the compressive forces and action of deltoid. However, when the arm is adducted the humeral component impinges on the scapular neck resulting in inferior scapular notching. Several techniques have been described to avoid this including: • • • • •

Inferior inclination of the glenosphere Inferior (eccentric) position of the glenosphere Increased glenosphere offset Alteration of the neck-shaft angle of the humeral component Surgical approach used

4.2  Inferior inclination of the glenosphere Increasing the inferior inclination of the glenosphere may increase the range of motion in both abduction and adduction before impingement occurs (Fig. 4.6) [14,101,102,122,123]. Inferiorly tilting the glenosphere by 15° has been shown to produce greater compressive

FIGURE 4.6  Diagram showing a glenosphere in a superior inclination (A), neutral (B) and inferior inclination (C) position.



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4.  Engineering advances in reverse total shoulder arthroplasty

forces and lesser tensile forces and micromotion compared to neutral and 15° superiorly inclined glenospheres [109,124]. Superiorly tilted glenospheres are associated with a higher degree of early glenoid implant loosening [125]. It is thought that these implants were fixed at the same inclination as the existing glenoid and due to increased superior glenoid wear from cuff arthropathy this resulted in a superiorly tilted glenosphere. The glenoid morphology should therefore be reviewed prior to fixation and glenoid reaming optimized, for example, a greater degree of inferior glenoid reaming, to prevent causing superior inclination [101].

4.3  Inferior (eccentric) positioning of the glenosphere In an anatomically positioned glenosphere abduction of the humerus may result in the humeral component impinging on the acromion and with adduction may notch on the scapula. An eccentric position places the glenosphere relatively inferiorly with the glenosphere overhanging the glenoid. This creates a larger space for the humeral component to move in both abduction and adduction therefore increasing the ROM before notching occurs (Fig. 4.7) [14,47,122,126,127]. It also has the benefit of not altering the center of rotation of the prosthesis. However, the angulation of the glenosphere during fixation must also be considered. In a concentric positioned glenosphere implanting at a neutral or inferior inclination will minimize shear forces at the bone-implant interface. In an eccentrically positioned glenosphere, inferior tilt has been shown to increase these shear forces affecting glenoid implant stability. These forces are reduced if the glenosphere is implanted in neutral inclination [124]. Eccentric positioning of the glenosphere has been shown to result in significant improvements in Constant-Murley scores and active ROM of flexion and external rotation [128]. Although some studies show no difference in the incidence of notching between eccentric and concentric glenospheres, at a 30 month follow-up of 47 patients, 5% of eccentrically positioned glenospheres had severe notching compared to 37% of concentric glenospheres [128].

FIGURE 4.7  Diagram showing a centered glenosphere (A) and eccentric glenosphere (B). An eccentric glenosphere (B) increases the range of motion in adduction and abduction before impingement occurs. Based on Figure 5, Nyffeler RW, Werner CM, Gerber C. Biomechanical relevance of glenoid component positioning in the reverse Delta III total shoulder prosthesis. J Shoulder Elbow Surg 2005;14:524-8 with permission. 



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FIGURE 4.8  Diagram showing a reverse total shoulder arthroplasty (A) and reverse total shoulder arthroplasty with lateral offset glenoid (B). The solid and hollow dots indicate the center of rotation of the reverse total shoulder arthroplasty and reverse total shoulder arthroplasty with lateral offset glenoid respectively. Based on Figure 2 from Ackland et al. Prosthesis design and placement in reverse total shoulder arthroplasty, Journal of Orthopaedic Surgery and Research (2015) 10:101 with permission.

4.4  Increased glenosphere offset Increased lateral offset also offers the benefit of a greater range of motion before impingement occurs (Fig. 4.8) [129]. By increasing lateral offset by 10 mm, ROM may increase up to 32° [130]. However, increased offset creates greater torque forces at the interface between the glenosphere and glenoid increasing the risk of loosening [37,66,131–133]. These forces can be minimized by inferiorly tilting the glenosphere. The lowest adverse forces at the boneimplant interface are achieved with either a neutral offset or lateral offset glenosphere positioned in inferior inclination [124].

4.5  Alteration of the neck-shaft angle of the humeral component The anatomical humeral neck-shaft angle is between 135° and 140°. It is measured by the angle between a line drawn along the axis of the humeral shaft and a line drawn along the axis of the anatomical neck of the humerus (Fig. 4.9). Grammont’s prosthesis had a neck-shaft

FIGURE 4.9  Diagram showing humeral neck-shaft angles of 140˚ (A) and 170˚ (B) demonstrating a reduction in adduction range of motion prior to impingement with an increased humeral neck shaft-angle of 170˚. Based on Figure 2A from Gutierrez S et al. Evaluation of abduction range of motion and avoidance of inferior scapular impingement in a reverse shoulder model. J Shoulder Elbow Surg 2008;17:608-15 with permission. 

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4.  Engineering advances in reverse total shoulder arthroplasty

angle of 155° to increase stability of the GH joint. This lowered the humerus, tensioning the deltoid, resulting in increased deltoid force and ROM. However, increasing the neck-shaft angle positions the medial aspect of the glenosphere closer to the scapula, reducing the adduction range of motion prior to impingement (Fig. 4.9) [130]. A systematic review of 38 studies (2,222 shoulders) comparing rates of notching between prostheses with neck-shaft angles of 135° and 155° demonstrated a significantly higher rate of notching with 155° prostheses [134]. A neck-shaft angle of less than 155° results in a significantly greater range of external rotation compared to implants with an angle of 155° [130].

4.6  Surgical approach Notching is more common and severe when using a superolateral surgical approach than a deltopectoral approach [104]. This is thought to occur due to the approach resulting in the baseplate being positioned on the glenoid more superiorly and with a greater degree of superior inclination [135]. Use of a deltopectoral approach may allow better visualization of the entire glenoid, particularly the inferior portion allowing the surgeon to implant the baseplate in the intended position. However, the incidence of instability has been reported to be lower with a superior approach than a deltopectoral approach [136]. This may be due to a difference in the status of the subscapularis tendon reattachment postoperatively. A study comparing patients with and without subscapularis reattachment in RTSA reported an instability rate of 0% in the reattached group and 9.2% in the non-reattached group, although further studies have suggested that subscapularis reattachment may not be essential [107].

4.7  Managing glenoid defects Glenoid erosion is common in patients requiring RTSA. Failure to correct a deficient glenoid prior to baseplate implantation leads to complications including instability, notching, incorrect base-plate position and inclination, and reduced function and implant longevity [125,137,138]. Surgical options to improve bony glenoid defects include eccentric reaming with autograft or allograft or use of specialized augmented glenoid baseplates [139]. A more recent option has been the use of angled bony-increased offset–reverse shoulder arthroplasty (BIO-RSA) [140]. This uses a trapezoidal bone graft matched to the defect taken from the native humeral head to replace the area of glenoid bone loss. This corrects the bony anatomy to normal, to which a baseplate with long peg and screws is implanted to create a lateralized center of rotation. This technique uses specialized instruments and preoperative templating to allow correction of individualized bony defects. Results of this technique at 3 year follow-up showed 94% of grafts had fused to the glenoid with no cases of severe notching [140]. Significant improvements in forward flexion, external rotation, internal rotation, and Constant-Murley scores were achieved similar to patients undergoing RTSA without complex glenoid bone deficiency [141–143]. Advantages of BIO-RSA over previous techniques include that it allows multiplanar deformity correction, has reduced costs compared to specialized baseplates, and use of humeral head autograft results in lower graft related morbidity [144–146].





4  Advances in implant design and surgical techniques

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4.8  Correct deltoid muscle tensioning One of the most important factors contributing to post RTSA function and stability is correct tensioning of the deltoid muscle [147,148]. Tensioning is predominantly affected by arm length. Arm length can be altered intraoperatively by several factors including the size of glenosphere used, position of glenosphere, inclination of glenosphere, type of humeral stem used, level of humeral cut, position of stem, and thickness of the polyethylene humeral insert. Arm lengthening resulting in appropriate deltoid tensioning correlates with active forward flexion [149]. Arm lengthening results in a significantly greater active forward flexion than arm shortening, although there may be no correlation with postoperative outcome scores. There is however a strong relationship between arm length and dislocation rates with some studies showing shortening to be present in all cases of RTSA dislocation [121]. A delicate balance is required to achieve adequate deltoid tensioning to prevent dislocation without causing nerve injury. This can be difficult in complex cases with a high dislocation risk including revision RTSA. Techniques used to reduce the risk of nerve palsy include use of a larger glenosphere, increasing the lateral offset of the glenosphere, and intraoperative nerve monitoring [150,151].

4.9  Implant fixation Screw fixation of the glenoid baseplate into the scapula is crucial to achieve optimum initial stability of the components [123,132,152]. This is affected by the position of screw insertion, the type of screw used, glenoid morphology, and the quality of the patient’s bone stock [153]. A well-fixed screw into the scapular pillar is most important in achieving stability as it is closest to the axis of the force applied to the limb when in use. Use of locking screws and a divergent screw trajectory provides greater stability of fixation and reduces the risk of loosening at the bone-implant interface [131]. As glenoid fixation is a cementless technique micromotion must be minimized to allow bone ingrowth to ensure long term stability.

4.10  Muscle transfers The use of muscle transfers during RTSA is used to optimize postoperative function in patients with rotator cuff arthropathy with significant external rotation weakness. The use of a latissimus dorsi transfer alongside RTSA was used first by Gerber to improve external rotation in patients with massive cuff tears [154–156]. This was performed through a separate posterior approach. Through the same deltopectoral approach Boileau [157] described the release of latissimus dorsi and teres major tendons from the floor and medial lip of the intertubercular groove of the humerus respectively. They were then passed around the posterior aspect of the humerus and fixed lateral to the intertubercular groove at the level of normal latissimus dorsi insertion. An alternative fixation site is the posterolateral aspect of the greater tuberosity at the level of teres minor insertion. Systematic reviews report significant improvements in external rotation, forward elevation and abduction after RTSA with latissimus dorsi transfer [158].



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4.  Engineering advances in reverse total shoulder arthroplasty

5  Future challenges There have been significant advances in the design and technique of shoulder arthroplasty over the last 50 years. Modern day implants continue to be designed around Grammont’s four key principles. Future work will involve developing these principles when designing new prostheses to find the perfect balance of implant longevity and function. The use of novel surgical techniques will further help to improve outcomes. For example, further work in muscle transfers may help to improve postoperative ROM after RTSA. Short-term results of RTSA are overall positive, however long term complication rates remain high. Further long term studies are required that differentiates the outcomes between the different indications for performing RTSA to help individualize complication risk assessment. Further work is needed to improve our understanding of the impact of different prostheses, implant positioning including version, inclination, medialization, offset and neck-shaft angle on the function of the muscles local to the GH joint. The biomechanics of the native and replacement shoulder are complex due to the large numbers of forces continuously acting across the joint. The GH joint particularly relies on muscles around the joint to provide stability. Individual muscle fibers continuously work to keep the joint stable, particularly when moving. These forces are difficult to recreate in current computer simulations due to their complexity. Further work in this area will increase our understanding of these forces and allow us to create biomechanical simulations which will allow translation into new implant designs and surgical techniques. The use of modern technology promises to help improve outcomes of shoulder arthroplasty. Computer assisted navigation systems aim to improve accuracy of glenoid implantation. Studies so far show improvements in accuracy compared to freehand techniques with up to 8° difference between the two methods in terms of glenoid implant position [159,160]. However, currently available systems may confer higher cost and result in a longer operating time. Utilization of patient specific instruments is another method that aims to improve positioning of implants. It uses computational methods to preoperatively build a template of the patient’s bony architecture which results in improvements in the positioning of screws and placement of the glenoid baseplate. This bony template can be recreated using 3-Dimensional printers to improve implant positioning by allowing the surgeon to preoperatively plan, visualize, and practice implant positioning. Initial studies demonstrate greater accuracy and consistency of glenoid positioning over standard techniques, although effect on functional outcome has not yet been assessed [161,162]. Computational models can also be used to create custom made implants to compensate for deformities or bone loss to give the optimum biomechanical functional outcome to patients. With continuing advancements and increased availability to surgeons these new techniques will help to improve positioning of components and functional outcomes and reduce complication rates in patients treated with RTSA.

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[128] Mizuno N, Denard PJ, Raiss P, Walch G. The clinical and radiographical results of reverse total shoulder arthroplasty with eccentric glenosphere. Int Orthop 2012;36(8):1647–53. [129] Kalouche I, Sevivas N, Wahegaonker A, Sauzieres P, Katz D, Valenti P. Reverse shoulder arthroplasty: does reduced medialisation improve radiological and clinical results? Acta Orthop Belgica 2009;75(2):158–66. [130] Gutierrez S, Comiskey CAt, Luo ZP, Pupello DR, Frankle MA. Range of impingement-free abduction and adduction deficit after reverse shoulder arthroplasty. Hierarchy of surgical and implant-design-related factors. J Bone Joint Surg Am 2008;90(12):2606–15. [131] Harman M, Frankle M, Vasey M, Banks S. Initial glenoid component fixation in “reverse” total shoulder arthroplasty: a biomechanical evaluation. J Shoulder Elbow Surg 2005;14(1 Suppl. S):162S–7S. [132] Hopkins AR, Hansen UN, Bull AM, Emery R, Amis AA. Fixation of the reversed shoulder prosthesis. J Shoulder Elbow Surg 2008;17(6):974–80. [133] Virani NA, Harman M, Li K, Levy J, Pupello DR, Frankle MA. In vitro and finite element analysis of glenoid bone/baseplate interaction in the reverse shoulder design. J Shoulder Elbow Surg 2008;17(3):509–21. [134] Erickson BJ, Frank RM, Harris JD, Mall N, Romeo AA. The influence of humeral head inclination in reverse total shoulder arthroplasty: a systematic review. J Shoulder Elbow Surg 2015;24(6):988–93. [135] Clouthier AL, Hetzler MA, Fedorak G, Bryant JT, Deluzio KJ, Bicknell RT. Factors affecting the stability of reverse shoulder arthroplasty: a biomechanical study. J Shoulder Elbow Surg 2013;22(4):439–44. [136] Walch G, Wall B, Mottier F. Complications and revision of the reverse prosthesis, a multicenter study of 457 cases. In: Walch G, Boileau P, Mole D, et al., editors. Reverse Shoulder Arthroplasty. France: Sauramps Medical: Montpellier; 2006. p. 335–52. [137] Hettrich CM, Permeswaran VN, Goetz JE, Anderson DD. Mechanical tradeoffs associated with glenosphere lateralization in reverse shoulder arthroplasty. J Shoulder Elbow Surg 2015;24(11):1774–81. [138] Valenti P, Sauzieres P, Katz D, Kalouche I, Kilinc AS. Do less medialized reverse shoulder prostheses increase motion and reduce notching? Clin Orthop Relat Res 2011;469(9):2550–7. [139] McFarland EG, Huri G, Hyun YS, Petersen SA, Srikumaran U. Reverse Total Shoulder Arthroplasty without Bone-Grafting for Severe Glenoid Bone Loss in Patients with Osteoarthritis and Intact Rotator Cuff. J Bone Joint Surg Am 2016;98(21):1801–7. [140] Boileau P, Morin-Salvo N, Gauci MO, Seeto BL, Chalmers PN, Holzer N, et al. Angled BIO-RSA (bony-increased offset-reverse shoulder arthroplasty): a solution for the management of glenoid bone loss and erosion. J Shoulder Elbow Surg 2017;26(12):2133–42. [141] Greiner S, Schmidt C, Herrmann S, Pauly S, Perka C. Clinical performance of lateralized versus non-lateralized reverse shoulder arthroplasty: a prospective randomized study. J Shoulder Elbow Surg 2015;24(9):1397–404. [142] Costantini O, Choi DS, Kontaxis A, Gulotta LV. The effects of progressive lateralization of the joint center of rotation of reverse total shoulder implants. J Shoulder Elbow Surg 2015;24(7):1120–8. [143] Athwal GS, MacDermid JC, Reddy KM, Marsh JP, Faber KJ, Drosdowech D. Does bony increased-offset reverse shoulder arthroplasty decrease scapular notching? J Shoulder Elbow Surg 2015;24(3):468–73. [144] Fowler BL, Dall BE, Rowe DE. Complications associated with harvesting autogenous iliac bone graft. Am J Orthop 1995;24(12):895–903. [145] Iannotti JP, Frangiamore SJ. Fate of large structural allograft for treatment of severe uncontained glenoid bone deficiency. J Shoulder Elbow Surg 2012;21(6):765–71. [146] Jones RB, Wright TW, Roche CP. Bone grafting the glenoid versus use of augmented glenoid baseplates with reverse shoulder arthroplasty. Bull Hospital Joint Dis 2015;73(Suppl. 1):S129–35. [147] Ladermann A, Walch G, Lubbeke A, Drake GN, Melis B, Bacle G, et al. Influence of arm lengthening in reverse shoulder arthroplasty. J Shoulder Elbow Surg 2012;21(3):336–41. [148] Ladermann A, Williams MD, Melis B, Hoffmeyer P, Walch G. Objective evaluation of lengthening in reverse shoulder arthroplasty. J Shoulder Elbow Surg 2009;18(4):588–95. [149] Jobin CM, Brown GD, Bahu MJ, Gardner TR, Bigliani LU, Levine WN, et al. Reverse total shoulder arthroplasty for cuff tear arthropathy: the clinical effect of deltoid lengthening and center of rotation medialization. J Shoulder Elbow Surg 2012;21(10):1269–77. [150] Nagda SH, Rogers KJ, Sestokas AK, Getz CL, Ramsey ML, Glaser DL, et al. Neer Award 2005: peripheral nerve function during shoulder arthroplasty using intraoperative nerve monitoring. J Shoulder Elbow Surg 2007;16(Suppl. 3):S2–8. [151] Boileau P, Moineau G, Roussanne Y, O’Shea K. Bony increased-offset reversed shoulder arthroplasty: minimizing scapular impingement while maximizing glenoid fixation. Clin Orthop Relat Res 2011;469(9):2558–67.



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[152] Parsons BO, Gruson KI, Accousti KJ, Klug RA, Flatow EL. Optimal rotation and screw positioning for initial glenosphere baseplate fixation in reverse shoulder arthroplasty. J Shoulder Elbow Surg 2009;18(6):886–91. [153] Chebli C, Huber P, Watling J, Bertelsen A, Bicknell RT, Matsen F 3rd. Factors affecting fixation of the glenoid component of a reverse total shoulder prothesis. J Shoulder Elbow Surg 2008;17(2):323–7. [154] Favre P, Loeb MD, Helmy N, Gerber C. Latissimus dorsi transfer to restore external rotation with reverse shoulder arthroplasty: a biomechanical study. J Shoulder Elbow Surg 2008;17(4):650–8. [155] Gerber C, Maquieira G, Espinosa N. Latissimus dorsi transfer for the treatment of irreparable rotator cuff tears. J Bone Joint Surg Am 2006;88(1):113–20. [156] Gerber C, Pennington SD, Lingenfelter EJ, Sukthankar A. Reverse delta-III total shoulder replacement combined with latissimus dorsi transfer. A preliminary report. J Bone Joint Surgery Am 2007;89(5):940–7. [157] Boileau P, Chuinard C, Roussanne Y, Neyton L, Trojani C. Modified latissimus dorsi and teres major transfer through a single delto-pectoral approach for external rotation deficit of the shoulder: as an isolated procedure or with a reverse arthroplasty. J Shoulder Elbow Surg 2007;16(6):671–82. [158] Sheth M, Ko JK, Namdari S. Reverse shoulder arthroplasty and latissimus dorsi tendon transfer. Am J Orthop 2017;46(5):E287–92. [159] Venne G, Rasquinha BJ, Pichora D, Ellis RE, Bicknell R. Comparing conventional and computer-assisted surgery baseplate and screw placement in reverse shoulder arthroplasty. J Shoulder Elbow Surg 2015;24(7):1112–9. [160] Berhouet J, Gulotta LV, Dines DM, Craig E, Warren RF, Choi D, et al. Preoperative planning for accurate glenoid component positioning in reverse shoulder arthroplasty. Orthop Traumatol Surg Res 2017;103(3):407–13. [161] Heylen S, Van Haver A, Vuylsteke K, Declercq G, Verborgt O. Patient-specific instrument guidance of glenoid component implantation reduces inclination variability in total and reverse shoulder arthroplasty. J Shoulder Elbow Surgery 2016;25(2):186–92. [162] Throckmorton TW, Gulotta LV, Bonnarens FO, Wright SA, Hartzell JL, Rozzi WB, et al. Patient-specific targeting guides compared with traditional instrumentation for glenoid component placement in shoulder arthroplasty: a multi-surgeon study in 70 arthritic cadaver specimens. J Shoulder Elbow Surg 2015;24(6):965–71.



C H A P T E R

5

Engineering advances in knee arthroplasty Sanil H. Ajwania, Paul Suttonb, Charalambos Panayiotou Charalambousa,c a

Department of Trauma and Orthopaedics, Blackpool Victoria Hospital, Blackpool Teaching Hospitals NHS Trust, Blackpool, United Kingdom; bDepartment of Orthopaedics, Northern General Hospital, Sheffield, United Kingdom; c School of Medicine, University of Central Lancashire, Preston, United Kingdom

1 Introduction Primary knee arthroplasty surgery is a commonly performed pain relieving procedure for patients with degenerative joint disease [1]. However, a substantial proportion of patients, having knee arthroplasty continue with pain, or are dissatisfied with their prosthesis even though there are no obvious clinical or mechanical causes for this. Increasingly clinicians are undertaking joint arthroplasty surgery in young demanding patients, who are most likely to be dissatisfied after total knee arthroplasty (TKA) [2,3]. There is a constant strive by knee surgeons to improve the success of knee arthroplasty by achieving better postoperative functional scores, decreasing postoperative pain, and increasing implant longevity (through a reduction in implant infection, wear, and loosening rates) [4,5]. Improvements in implant engineering may help knee surgeons in their strive for success. This chapter reviews recent engineering advances in knee arthroplasty implants, with specific reference to innovations aiming at reducing infection, loosening, and wear rates. In addition reference is made to the development of patient specific knee arthroplasty components, and to implant modifications that may counteract the possibility of metal hypersensitivity reactions.

2  Prosthetic joint infection Prosthetic joint infection (PJI) is a serious complication in orthopedic surgery. The reported risk of infection in TKR and THR is approximately 3% in primary arthroplasty, and up to 10% in revision surgery [6]. When it does occur it is a devastating complication which causes pain Advances in Medical and Surgical Engineering. http://dx.doi.org/10.1016/B978-0-12-819712-7.00005-X Copyright © 2020 Elsevier Inc. All rights reserved.

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and functional difficulty for the patient. Additionally, it often requires further surgery in the form of revision arthroplasty which increases patient morbidity and mortality and also has considerable socio-economic costs [7–9].

2.1  Mechanism of bacterial adhere to implant surfaces Bacteria are extremely versatile organisms that have developed very effective ways of colonizing and multiplying upon implant components [10]. This ability is centered upon numerous proteins in their cell walls that enable the bacterium to colonize implant surfaces. The molecules that allow the bacteria to bind to prosthetic surfaces are adhesive ligands found on the bacterial cell walls [11]. The point at which a bacterium colonizes the implant can occur at any point from surgery, from the intra-operative stage due to environmental contamination to many years post-surgery as a consequence of hematogenous spread [12–14]. Once on the implant, the bacteria are able to multiply and increase numbers and then produce a proteo-glycan rich coating that protects the colonized bacterium in a biofilm [15–18]. This biofilm essentially engulfs the bacterium protecting it from the host’s immune system and systemic antibiotics [15]. The success of the bacteria in implant colonization is characterized by a change in the phenotype of the bacterium from being a planktonic form to being in its biofilm form. The phases of adhesion of bacteria upon the implant surface can be divided into reversible and irreversible. An adhesion phase by planktonic bacteria is followed by gene expression for secretion of protective proteo-glycans. This process of forming the biofilm makes bacteria extremely resistant to host immune cells, antibodies, and antibiotic diffusion (Fig. 5.1). In addition to the bacteria’s ability to bind with the implant, the micro-environment of the implant also allows for a surrounding where bacteria can proliferate. Once the implant

FIGURE 5.1  Implant colonization.





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components are implanted the healing of the disrupted tissues results in the formation of avascular fibrous scar tissue [13,19]. This scar tissue to a degree encapsulates the implant and makes it difficult to be reached by the host’s immune system thus allowing for potential infection to proliferate [14]. This factor is particularly relevant in revision surgery. Gristina developed the concept of “race for the surface” theory [11], a simplified model of implant-related infection, whereby host and bacterial cells compete in determining the ultimate fate of the implant. The first to colonize the implant may result in successful osseointegration of the implant, or the possibility of PJI depending upon the biggest colonizing cell type. Occasionally, polybacterial types can proliferate upon an implant [14].

2.2  Strategies to tackle prosthetic joint infection Strategies to reduce infection in the peri-operative setting are key consideration for any arthroplasty surgeon. Key techniques that are commonly utilized to protect against prosthetic infection, include laminar flow theatre designs, use of exhaust suits, and strict control of personnel theatre transit [20,21]. Antibiotic administration prior to skin incision and the use of antibiotic impregnated cement has also been shown to be effective against decreasing bacterial load and PJI [21].

2.3  Engineering advances in order to reduce prosthetic joint infection As well as improving the sterility of the surgical environment when implanting prostheses, modification of arthroplasty implants may help reduce the risk of PJI [22]. These modifications aim to reduce the ability of bacteria to cause PJI in conjunction with the hosts own immune system and systemic antibiotics. Work has been done to develop coatings on the implant surface to competitively inhibit the bacteria’s ability to colonize the implants and allow the hosts cells to be the first at the surface [22,23]. Strategies have been developed to tackle infection with the aim to minimize bacterial adhesion, biofilm formation, and reduce bacterial load [23]. The types of techniques used can be loosely classified as bio-inert or bioactive surface modifications (Fig. 5.2). 2.3.1  Bio-inert implant surface modifications Bio-inert modifications of the implant surface rely upon using strategies that affect the physical properties of the implant to lessen the affinity of bacteria for the surface [23]. Surface characteristics of implants, like surface roughness [24–28], nano-structure [26,29], hydrophilicity [30,31], and surface potential (charge) [28, 32–34], are all physical properties that affect bacterial adhesion to the implants [35]. 2.3.2  Bioactive implant surface modifications Bioactive modifications are achieved via surface modifications that aim to have a bactericidal effect. These types of bioactive modifications include coating the implants with minerals such as silver, iodine, and selenium. Additionally, bioactive modifications include coating the implant with organic substances including antibiotics or antibacterial proteins and cytokines [23].



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FIGURE 5.2  Overview of the types of implant modification.

2.3.2.1  Bioactive mineral implant surface modifications

The use of minerals has been studied using silver [36–39], copper [40,41], and Zinc [42]. All of these metals have been shown to disrupt bacterial metabolic processes via different modalities [36,43]. There are concerns, however, with regards to toxicity of metal ions upon host cell activity. Research is being done aiming to reduce the toxicity effect of these metals by creating nano-structure coatings on the implants to maximize the therapeutic window of these metals while reducing their toxicity profile [29]. Non-metal minerals such as selenium and iodine, have also been investigated for their antimicrobial properties [44,45]. Selenium bound to the surface of titanium alloys has been shown to prevent S. aureus and S. epidermidis adhesion without affecting host tissue [44,46]. It works by creating superoxide radicals that inhibit bacterial growth and function [44]. Iodine has traditionally been used for its sterile properties in orthopedics for many years. However, new research has now looked at developing its role by coating titanium implants in iodine and assessing its results [45]. 2.3.2.2  Bioactive organic implant surface modifications

Organic bioactive modifications to the implant surfaces are being investigated utilizing antibiotics, antibacterial peptides, and cytokines [23]. The use of antibiotics that are either covalently bonded to the implant (non-eluting) or coated (eluting) are currently being investigated [47–50]. The success of applying antibiotics to the surface of implants does rely upon the bacterium being sensitive to the antibiotic used. The method of covalently bonding the antibiotic on to the implant does not easily allow for elution of the drug into the prosthesis surroundings. Conversely, antibiotics that are coated onto the prosthesis are able to elude around the component surroundings’ and hence attack bacterium that may not be directly





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adjacent to the implant [48,51]. Various antibiotics modifications are being developed that maybe applied to the implant in conjunction with hydroxyapatite in the form of an un-cemented prosthesis [52,53]. Antimicrobial peptides and cytokines are also being investigated for use as implant coatings. They have been shown to be effective against a myriad of pathogens. They mainly carry out their antimicrobial function by damaging bacterial metabolic processes as well as regulating host inflammatory responses [54–56].

3  Strategies to improve implant longevity Knee arthroplasty surgery is increasingly used in young, functionally high demand patients. Modern knee arthroplasty systems must be able to offer improved implant longevity in order for them to be used in this subset of patients. In order to tackle the issues that mainly affect longevity, engineering advances aim at addressing the problems of aseptic loosening, polyethylene wear, and implant mal-positioning (which may itself lead to accelerated eccentric wear of components). The longevity of knee arthroplasty implants may be influenced by the durability of the fixation to bone with the most widely used method being cementation. Current focus is on developing a more durable biological fixation to the bone. Traditional cobalt chrome (CoCr) and polyethylene components bearings have been associated with wear of the polyethylene bearing and need for early revision surgery in young patients. Furthermore, research is also looking at improving and developing alternative bearing surfaces with favorable wear characteristics.

3.1  Cementless fixation total knee arthroplasty The advantages of cementless TKA include the potential to preserve bone stock and achieve biological fixation of the implant to the bone [57]. Additionally, cementless TKA also obviates the generation of cement debris around the implant. Currently, the most commonly used coating is hydroxyapatite. It has been shown to have both osteoconductive and osteoinductive properties to help osseointegration [58–60]. Despite these properties there still remain concerns around the use of HA coating and its ability to fragment under physiological load and possibly generate third body wear particles [22,57].

3.2  Bisphosphonate coatings Coatings of the implant with a bisphosphonates help encourage osteoblast activity at the implant-surface while inhibiting osteoclastic activity [61]. This has the main benefit of increasing localized bone formation at the bone implant interface [62]. The use of a bisphosphonate in conjunction with hydroxyapatite improves bone implant contact, local bone mass, and bone mineral density [63]. The combined use of these agents may be ideal in patients with poor bone stock or osteoporosis [64,65]. The localized effects of the bisphosphonate would also theoretically reduce systemic side effects that are seen in patients that take them.



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3.3  Biomolecule coatings It is expected that the use of biomolecule coatings will improve implant surface osseointegration. The molecules that have been looked at in order to achieve this include growth factors [66], large proteins like collagen and chondroitin [67–71], and smaller molecules like protein peptides [72] and DNA [73]. Coating the implant surface with type 1 collagen and chondroitin, which acts as a biomimetic has been shown to accelerate osseointegration around implants [72]. Growth factor coatings have also been investigated which act as a chemoattractant to the host osteoprogenitor cells and induce blood vessel formation in order to promote implant osseointegration [66]. Furthermore, DNA molecules are also being studied, the DNA is incorporated into the implant coating and instructs hosts tissues and cells to create proteins to aid osseointegration [22,73].

3.4  Alternative bearing surfaces Conventional knee arthroplasty bearings use CoCr femoral components articulating with ultra high molecular weight polyethylene (UHMWPE) on the tibia [74]. The polyethylene on the tibial component can come in either modular or monoblock systems. This bearing combination has a long history of clinical success and modern polyethylene treatment techniques have improved wear characteristics, and reduced the incidence of subsequent aseptic loosening, which used to be common as a result of wear generated osteolysis [74,75]. Modern techniques aim to reduce the polyethylene wear for patients. Interestingly, a recent study by Kim and coworkers compared outcomes in patients who received a posterior cruciate-substituting total knee prosthesis with a conventional polyethylene tibial insert in one knee and the same prosthesis with a highly cross-linked polyethylene tibial insert in the contralateral knee. This study found no significant difference in clinical or radiological outcomes between the two groups at 5 years. Suggesting modern polyethylene treatment techniques seem to give reliable short term results [76]. 3.4.1  Ceramic metal femoral components Ceramic bearings in TKA have been developed as part of a continued drive to improve the wear characteristics of knee replacement systems [77]. At present there are three main ways of incorporating a ceramic femoral component into the knee replacement. 3.4.1.1  Ceramic and ceramicised metal alloy femoral knee components

Monoblock wholly ceramic femoral component development has been driven by the successes seen from modern ceramics in hip arthroplasty. The composite ceramic of the Biolox delta knee replacement by CeramTec is one such example of a monoblock ceramic femoral component (CeramTec, Plochngen, Germany) [78]. This ceramic is marketed as Zirconia toughened alumina femoral implant, with a titanium tray and polyethylene inlay. The use of monoblock ceramics has been a concern due to ceramic brittleness and the possibility of fracture during impaction of the prosthesis during implantation. This is more likely in the knee prosthesis as opposed to hip arthroplasty due to the complex geometry of the knee component. As a result of this, femoral impaction of the component with a hammer is not recommended [75,79].





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An alternative to wholly ceramic components is the use of ceramicised metals, such as Oxinium (Smith and Nephew, Memphis, USA) and the Columbus Advanced Surface (Bbraun & Aesculap, Tuttlingen, Germany) knees [75,80,81]. The Oxinum knee uses an alloy of niobium and oxidized zirconium, the surface of which is heated into a ceramic type material [82,83]. This process combines the strength and toughness of a metal with the wear properties of a ceramic at its surface. The advanced surface coating of the Columbus knee is formed by vapor deposition of zirconium nitride at the surface layer. Underneath this surface layer are layers of chrome nitride and an underlying CoCr alloy component. The fracture risk that is associated with a monoblock ceramic component is thought to be reduced with ceramicised metal implants. Additionally, the use of monoblock ceramic and ceramicised components may have a role in patients with metal hypersensitivity to nickel [75,84–86].

3.5  Patient specific knee arthroplasty A key technical consideration that affects outcomes of TKR is the implant size, position, and alignment of the components [87]. In order to help improve alignment accuracy and positioning, tools like patient specific instrumentation have been developed [88]. Disappointingly, current knee systems offer implants of limited sizes. The limitations in size may lead to mismatch between implant size and patient anatomy which can cause patient dissatisfaction. To address this patient specific components (matched to each patient’s unique anatomy) are also available that help to reduce issues caused by size mismatch. Patient specific knee arthroplasty jigs and components help restore the mechanical axis of the knee and also increase accurate implant positioning. The patient specific knee allows a surgeon to accurately recreate the coronal plane alignment of the knee and hence reduce the number of alignment outliers [87]. This is important as errors in surgical technique and component positioning may compromise the long term outcomes of surgery [1,87]. Patient specific systems recreate the anatomy of the knee digitally with a pre-operative CT or MRI scan of the knee joint [89]. This data is then used to recreate cutting jigs that are individualized to the patients’ specific anatomy. The jigs are then applied to the bone ends and act as a guide intra-operatively to aid cutting of the bone [90]. The purported advantages of patient specific implantation include improved implant position accuracy, reduced number of instrument trays meaning quicker theatre setup [88,91,92]. Other more contentious claims include reduced blood loss, operative time, length of hospital stay, and improved patient outcomes [93,94]. Disadvantages of this technology include the added costs to the procedure in obtaining the pre-operative scans and manufacture of the guides [87,92]. Also the claims that patient specific TKR results in lower overall wear remains to be seen in long term follow up clinical studies [95,96]. Current clinical evidence on the outcomes of patient specific arthroplasty seems to be mixed. Thienpont and coworkers carried out a systematic review and meta-analysis on the use of patient specific instrumentation. They concluded that patient specific instrumentation does not improve the accuracy of alignment of the components when compared with conventional instrumentation [97]. Similar results were also found in a systematic review done by Sassoon and coworkers. Sassoon and coworkers also found that patient specific



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instrumentation does require fewer surgical trays, but there is no decisive evidence that this technology improves surgical efficiency and cost effectiveness [96]. At present, long term outcomes of patient specific knee replacement are awaited, but current data seem to suggest there is limited clinical benefit in this technology, at least in the short term.

4  Metal hypersensitivity in total knee arthroplasty Post-operative pain and dissatisfaction occurs in 17%–54% of TKA patients and in many cases this is for unexplained reasons [98,99]. It has been suggested that a proportion of these patients may be suffering from implant related metal hypersensitivity [5,100–102]. The prevalence of cutaneous metal hypersensitivity to nickel, palladium, cobalt, and chrome is thought to occur in up to 20% of the general population [100]. The use of standard implants in patients with metal hypersensitivity has been linked to devastating complications such as accelerated aseptic loosening, deep localized inflammatory reactions, as well as ongoing pain [103–105].

4.1 Pathophysiology Orthopedic implants once implanted undergo both mechanical and chemical corrosion. As a result metal ions from the implants are released. These ions form complexes with host cells that surround the adjacent joint tissue. These protein-ion complexes combine with host tissue to form antigen presenting complexes that can activate the T helper cells of the immune system. Once activated these cells co-ordinate release of numerous inflammatory cytokines such as interleukins (IL), tumor necrosis factors alpha, (TNF-α,) interferons’, prostaglandins, and receptor activator of nuclear factors (RANK). These cytokines combine to become chemoattractants to immune cells and osteoclasts around the implants [100]. Activation of these cells and osteoclasts causes localized osteolysis, loosening and pain for the patient (Fig. 5.3). 4.1.1  Current concepts The majority of standard TKA systems available worldwide are made with cobalt and chromium alloy. This alloy contains a mixture of metals including nickel. In order to tackle the issue of metal hypersensitivity, companies have developed implants that are suitable for use in patients that suffer from this form of hypersensitivity (Fig. 5.4).

4.2  The all polyethylene tibial component A monoblock polyethylene tibial component is considered safe to use in patients that have got metal hypersensitivity. A monoblock tibia has been shown to have similar clinical outcomes to modular tibial components [106]. Additionally, the monoblock components may reduce exposure to metal allergens when used in this subclass of patient. All polyethylene components also have some inherent disadvantages such as a lack of modularity (limiting intra-operative options), no option for liner removal in the setting of acute irrigation and debridement for infection, and no option for late liner exchange [107].





4  Metal hypersensitivity in total knee arthroplasty

FIGURE 5.3  Pathophysiology of loosening in patients with metal hypersensitivity (100).



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FIGURE 5.4  Types of metal hypersensitivity related implants that are currently available.

4.3  Methods used to make implants metal hypersensitivity friendly 4.3.1  Fully coated implants Currently, metal hypersensitivity “friendly” implants fall into two categories, those that are coated implants and those made of alternative alloy to CoCr [108,109]. The majority of implant companies make a standard cobalt chrome implant coated with a superficial hypersensitivity “friendly” layer (usually titanium nitride or zirconia nitride) which encapsulates the prosthesis. These implants can be made custom made or be available off the shelf depending on the implant and manufacturer [109]. Typically, these hypersensitivity implants are associated with increased costs. The advantage of coating implants is it allows retention of the tribological properties of cobalt chrome such as strength, and durability [108,109]. However, worryingly, this method of coating the implant could be affected by asperities and scratching that occur to the prosthesis during implantation or during the lifetime of the implant. If such asperities were to occur then it would expose the patient to the underlying metal and lead to a potential hypersensitivity reaction. Recently, Theinpoint and coworkers [110], looked at standard CoCr implants and those coated with titanium niobium nitride and found no difference in clinical, radiological, or patient related outcomes between the two groups [110]. 4.3.2  Partially coated implants The majority of implants are completely encapsulated with a hypersensitivity “friendly” coating, which includes both the articulating and non-articulating surfaces of the implant (the part in contact with bone). However, a partially coated CoCr implant with titanium nitride is also obtainable. The partial coating is only applied to the articulating surface. A partial coating is only applied by some manufactures as they believe that coating the non-articulating surface can impair cementation. This type of hypersensitivity friendly implant is concerning



References 65

as exposure of the allergen over time could occur if cement lysis were to occur or if there are any cement mantle deficits when implanting the prosthesis [109]. 4.3.3  Alternative alloy implants Another method of manufacturing hypersensitivity “friendly” implants is by developing prostheses made entirely of non-cobalt chrome alloys. These implants’ are made entirely from titanium, ceramic, and ceramicised metal alloys [75,79,83,109,111]. These implants eliminate the risk of allergen exposure to nickel, cobalt, and chrome due to asperities and long term wear [105,112,113]. A potential disadvantage of titanium implants is reduced strength compared to CoCr alloy. The disadvantages of a purely ceramic component, is the brittle property of the ceramic implant and difficulties with impaction of the prosthesis during surgery. Ceramicised metal components are new and the early outcomes report excellent results with excellent revision rates and acceptable wear characteristics when compared to conventional implants [114,115].

4.4  Controversy around the need for metal hypersensitivity implants Metal hypersensitivity “friendly” implants are designed to help surgeons manage patients with metal hypersensitivity. There is however no strong evidence for the type of implants best to use in patients that have mild local skin reactions to nickel, cobalt, or chromium [116–118]. Guidelines and expert consensus studies do recommend that conventional implants be used in most patients with mild local cutaneous metal hypersensitivity reactions reported by patients or determined by patch testing [112,119–121]. Conversely, when there is a history of severe local cutaneous metal hypersensitivity reactions, or generalized systemic reactions it has been suggested that patients should be patch tested and appropriate hypersensitivity “friendly” implants utilized [86,100,119,120,122].

5 Conclusion Knee replacement surgery is a life transforming procedure for the majority of patients. However, it does currently have some issues which continue to affect patient satisfaction. Appreciation of the problems we currently face in knee replacement surgery can guide targeted areas of research to address our current difficulties. With advances in implant engineering aiming to reduce some of the issues and complications highlighted earlier.

References [1] Lombardi A Jr, Berend K, Adams J. Why knee replacements fail in 2013. Bone Joint J 2014;96(11 Suppl. A):101–4. [2] Price AJ, Longino D, Rees J, Rout R, Pandit H, Javaid K, et al. Are pain and function better measures of outcome than revision rates after TKR in the younger patient? The Knee 2010;17(3):196–199. [3] Williams D, Price A, Beard D, Hadfield S, Arden N, Murray D, et al. The effects of age on patient-reported outcome measures in total knee replacements. Bone Joint J 2013;95(1):38–44. [4] Mandalia V, Eyres K, Schranz P, Toms A. Evaluation of patients with a painful total knee replacement. Bone Joint J 2008;90(3):265–71. [5] Nam D, Li K, Riegler V, Barrack RL. Patient-reported metal allergy: a risk factor for poor outcomes after total joint arthroplasty? J Arthroplasty 2016;31(9):1910–5. 

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[6] Gallo J, Holinka M, Moucha CS. Antibacterial surface treatment for orthopaedic implants. Int J Mol Sci 2014;15(8):13849–80. [7] Alijanipour P, Parvizi J. Infection post-total knee replacement: current concepts. Curr Rev Musculoskelet Med 2014;7(2):96–102. [8] Lentino JR. Prosthetic joint infections: bane of orthopedists, challenge for infectious disease specialists. Clin Infect Dis 2003;36(9):1157–61. [9] Zmistowski B, Karam JA, Durinka JB, Casper DS, Parvizi J. Periprosthetic joint infection increases the risk of one-year mortality. J Bone Joint Surg Am 2013;95(24):2177–84. [10] Toms AD, Davidson D, Masri BA, Duncan CP. The management of peri-prosthetic infection in total joint arthroplasty. J Bone Joint Surg Br 2006;88(2):149–55. [11] Gristina AG, Naylor P, Myrvik Q. Infections from biomaterials and implants: a race for the surface. Med Prog Technol 1988;14(3-4):205–24. [12] Zimmerli W, Lew PD, Waldvogel FA. Pathogenesis of foreign body infection. Evidence for a local granulocyte defect. J Clin Invest 1984;73(4):1191–200. [13] Higgins DM, Basaraba RJ, Hohnbaum AC, Lee EJ, Grainger DW, Gonzalez-Juarrero M. Localized immunosuppressive environment in the foreign body response to implanted biomaterials. Am J Pathol 2009;175(1):161–70. [14] Tande AJ, Patel R. Prosthetic joint infection. Clin Microbiol Rev 2014;27(2):302–45. [15] Costerton JW, Stewart PS, Greenberg EP. Bacterial biofilms: a common cause of persistent infections. Science 1999;284(5418):1318–22. [16] Stoodley P, Ehrlich GD, Sedghizadeh PP, Hall-Stoodley L, Baratz ME, Altman DT, et al. Orthopaedic biofilm infections. Curr Orthop Pract 2011;22(6):558–63. [17] An YH, Friedman RJ. Concise review of mechanisms of bacterial adhesion to biomaterial surfaces. J Biomed Mater Res 1998;43(3):338–48. [18] Katsikogianni M, Missirlis YF. Concise review of mechanisms of bacterial adhesion to biomaterials and of techniques used in estimating bacteria-material interactions. Eur Cell Mater 2004;8:37–57. [19] Franz S, Rammelt S, Scharnweber D, Simon JC. Immune responses to implants - a review of the implications for the design of immunomodulatory biomaterials. Biomaterials 2011;32(28):6692–709. [20] Humphreys H. Surgical site infection, ultraclean ventilated operating theatres and prosthetic joint surgery: where now? J Hosp Infect 2012;81(2):71–2. [21] An YH, Friedman RJ. Prevention of sepsis in total joint arthroplasty. J Hosp Infect 1996;33(2):93–108. [22] Goodman SB, Yao Z, Keeney M, Yang F. The future of biologic coatings for orthopaedic implants. Biomaterials 2013;34(13):3174–83. [23] Romanò CL, Scarponi S, Gallazzi E, Romanò D, Drago L. Antibacterial coating of implants in orthopaedics and trauma: a classification proposal in an evolving panorama. J Orthop Surg Res 2015;10(1):1. [24] Del Curto B, Brunella MF, Giordano C, Pedeferri MP, Valtulina V, Visai L, et al. Decreased bacterial adhesion to surface-treated titanium. Int J Artif Organs 2005;28(7):718–30. [25] Zhang F, Zhang Z, Zhu X, Kang ET, Neoh KG. Silk-functionalized titanium surfaces for enhancing osteoblast functions and reducing bacterial adhesion. Biomaterials 2008;29(36):4751–9. [26] Oh S, Moon KS, Lee SH. Effect of RGD peptide-coated TiO 2 nanotubes on the attachment, proliferation, and functionality of bone-related cells. J Nanomater 2013;2013:9. [27] Braem A, Van Mellaert L, Mattheys T, Hofmans D, De Waelheyns E, Geris L, et al. Staphylococcal biofilm growth on smooth and porous titanium coatings for biomedical applications. J Biomed Mater Res A 2014;102(1):215–24. [28] Kaper HJ, Busscher HJ, Norde W. Characterization of poly(ethylene oxide) brushes on glass surfaces and adhesion of Staphylococcus epidermidis. J Biomater Sci Polym Ed 2003;14(4):313–24. [29] Pelgrift RY, Friedman AJ. Nanotechnology as a therapeutic tool to combat microbial resistance. Adv Drug deliv Rev 2013;65(13):1803–15. [30] Yu JC, Ho W, Lin J, Yip H, Wong PK. Photocatalytic activity, antibacterial effect, and photoinduced hydrophilicity of TiO2 films coated on a stainless steel substrate. Environ Sci Technol 2003;37(10):2296–301. [31] Gallardo-Moreno AM, Pacha-Olivenza MA, Saldana L, Perez-Giraldo C, Bruque JM, Vilaboa N, et al. In vitro biocompatibility and bacterial adhesion of physico-chemically modified Ti6Al4V surface by means of UV irradiation. Acta Biomater 2009;5(1):181–92. [32] Harris LG, Tosatti S, Wieland M, Textor M, Richards RG. Staphylococcus aureus adhesion to titanium oxide surfaces coated with non-functionalized and peptide-functionalized poly(L-lysine)-grafted-poly(ethylene glycol) copolymers. Biomaterials 2004;25(18):4135–48.



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C H A P T E R

6

Biology of cartilage Maire-Clare Killen Royal Victoria Infirmary, Newcastle upon Tyne; Charalambos Charalambous, Blackpool Victoria Hospital, Blackpool, United Kingdom

1 Introduction There are several types of cartilage, which are classified as fibro-cartilage, elastic, fibro-elastic, and hyaline. Joint surfaces are covered with hyaline cartilage; this is a highly specialized connective tissue which provides a low friction gliding surface, acts as a shock absorber and minimizes pressures on subchondral bone. Hyaline cartilage it is an avascular, aneural, alymphogenic, almost non-immunogenic structure and is nourished entirely by the surrounding synovial fluid. It therefore has a very limited capacity for intrinsic healing and repair. In contrast to most other tissues, cartilage has only one cell type, the chondrocyte. These articular chondrocytes are derived from mesenchymal stem cells and are bound together by an abundant extracelluar matrix. They are responsible for synthesizing the components of the extracellular matrix, as well as maintenance of the extracellular matrix metabolism. The extracellular matrix is mostly made up of water, which is held in place by the negative charge of proteoglycans. Collagen fibers form a meshwork, giving cartilage its high tensile strength. The extracellular matrix has numerous functions, including: 1. protection of chondrocytes from mechanical loading; 2. storage of cytokines and growth factors required for chondrocytes; 3. controls the type, concentration, and rate of diffusion of the nutrients to chondrocytes; 4. acts as a signal transducer for the cells. The constituents of articular cartilage and their main functions are summarized in Table 6.1.

1.1 Structure The structure of articular cartilage can be divided into four zones, from superficial (zone 1) to deep (zone IV), shown in Fig. 6.1. The composition, structure, and function of chondrocytes changes depending on the depth from the surface of the cartilage. Advances in Medical and Surgical Engineering. http://dx.doi.org/10.1016/B978-0-12-819712-7.00006-1 Copyright © 2020 Elsevier Inc. All rights reserved.

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TABLE 6.1  Overview of the components of articular cartilage. Component

Function

Chondrocytes (1%–5% wet weight)

Synthesis of matrix components (collagen, proteoglycans, enzymes). Regulate exctracellular matrix metabolism

Extracellular matrix Water (60%–85% of wet weight)

Allows load dependent deformation Nutrition Lubrication

Collagen (10%–20% of wet weight, 40%–70% dry weight) • Type ll (almost exclusively) • Also types VI, IX, X, XI

Tensile strength

Proteoglycans and glycosaminoglycans (10%–20% wet weight) • Aggrecan (predominant proteoglycan) • Hyaluronan • Decorin, byglycan, fibromodulin, syndecan, lumican, superficial zone protein

Compressive strength Maintain fluid and electrolyte balance

Glycoproteins • Cartilage oligomeric protein (COMP), laminin, lubricin, chondro-adherin • Cartilage matrix protein (CMP), cartilage matrix glycoprotein (CMGP) Chondronectic, fibronectin, anchorin,

Act to bind various components of the matrix and chondrocyte surface

Degradative enzymes • Matrix metalloproteinases

Degradation of collagen and proteoglycan aggregates as part of normal matrix turnover

Extracellular ions

FIGURE 6.1  Layers of articular cartilage: chondrocytes are arranged horizontally in the superficial layer, contrasted with the vertical orientation in the deep layer. Source: From Ref. [1].





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Zone 1: Superficial gliding zone The superficial zone is the thinnest of all the layers. It is made up of flattened ellipsoid chondrocytes and collagen fibrils, which lie parallel to the articular surface, providing tensile strength and giving resistance to shear forces. This layer also functions as a barrier to large macromolecules, protecting cartilage from the synovial fluid immune system. Chondrocytes in this zone synthesize high concentrations of collagen and low concentrations of proteoglycans; this zone therefore has the highest water content. This zone is covered by a clear film of small collagen fibrils and a cellular layer of flattened chondrocytes known as the “lamina splendens.” This provides an ultimate gliding surface to the cartilage. Disruption of the superficial zone results in alteration of the mechanical properties of cartilage and is the first to show degenerative changes in the development of osteoarthritis. Zone 2: Middle transitional zone In the middle, transitional zone, the density of cells is lower. It is predominately made of spheroid shaped cells, embedded in an abundant extracellular matrix. Large diameter collagen fibers are arranged obliquely and there is a higher concentration of proteoglycan. This zone acts to provide a transition between the shearing forces of the superficial layer and the compressive forces in the deeper zones. Zone 3: Deep (radial) zone The deep zone is the largest part of the articular cartilage. The cell density is lowest in this zone; chondrocytes are spheroidal in shape and are arranged perpendicular to the surface. This zone contains the largest diameter of collagen fibrils and has the highest concentration of proteoglycans. The perpendicular arrangement of the collagen fibrils distributes load and resists compression. The chondrocytes in zone two and three are involved in production of all the components of the extracellular matrix. The tidemark represents the boundary between the uncalcified and calcified cartilage. It is a layer which is free of cells and identifies the transition to less resilient subchondral bone. The tidemark tends to progress toward the surface with age. Zone 4: Calcified cartilage zone The calcified zone is anchored to the subchondral bone by hydroxyapatite crystals. This layer provides a barrier to diffusion from blood vessels supplying the bone. It contains a small number of cells embedded in a calcified matrix, which demonstrate a very low metabolic activity.

1.2 Function The two main functions or articular cartilage are joint lubrication and shock absorption, it has unique viscoelastic properties to aid this function. Cartilage provides a smooth, lubricated surface for low friction articulation and to facilitate load transmission to the underlying subchondral bone. Articular cartilage has a unique ability to withstand high repetitive loads, demonstrating little or no evidence of damage or degenerative change. Joint lubrication There are several types of lubrication that have been described in relation to cartilage, which act to results in minimal joint wear: • Elastohydrodynamic lubrication: this is the main mode of lubrication during dynamic joint movement. This occurs when pressure in the fluid film causes elastic deformation 

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6.  Biology of cartilage

of the joint surfaces, increasing the surface area and reducing escape of fluid between surfaces as they glide over each other. Boundary lubrication: this involves a single layer of lubricant molecule (e.g., the glycoprotein lubricin), absorbed on each surface of the joint, partially separating the joint surfaces and preventing direct articular contact. This is most important at rest or under load. Boosted lubrication (fluid-entrapment): the solvent part of the lubricant enters the cartilage which then leaves behind pools of concentrated hyaluronic acid to act as a lubricant. Hydrodynamic: this occurs when two surfaces are at an angle and sliding over each other, with a resulting fluid layer separating the two surfaces. Weeping or self-lubrication: when articular cartilage is compressed, fluid shifts out of the cartilage, separating the surfaces by hydrostatic pressure [2].

These types of lubrication were first described as engineering terms. As joints are non-rigid structures it is thought that modified forms of these types of lubrication occur during joint movement. When movement is initiated, boundary lubrication occurs where there is close contact between the joint surfaces, followed by fluid-film lubrication elsewhere. Shock absorption Cartilage is 10 times more effective at absorbing shock compared to bone and therefore acts to protect bone by distributing the load. With respect to its viscoelasticity, cartilage has two properties, termed creep and stress relaxation which occur through macromolecular and water movement: • Creep: under a constant compressive stress, cartilage will initially deform rapidly, allowing a large increase in surface area to dissipate forces, followed by a slower deformation until an equilibrium value is reached. • Stress relaxation: under constant deformation, high initial stress is followed by progressively decreasing stress to the level required to maintain the deformation. Although cartilage is freely permeable to water, the frictional drag of macromolecules will hinder the movement of water when compressive forces are high. This results in reduced flow of water, increasing the cartilage stiffness, therefore allowing greater resistance to higher loads. The complex arrangement of collagen fibers, collagen fiber cross-links (Fig. 6.1) and the collagen-proteoglycan interactions allows cartilage to be anisotropic; it can exhibit different mechanical properties depending on the direction in which is being loaded, contributing significantly to its shear-resistant properties [2].

1.3  Cartilage aging The overall structure and composition of both chondrocytes and extracellular matrix changes with age. There are changes in the distribution of chondrocytes, but the total number of chondrocytes remains essentially the same. Chondrocytes begin to dissipate in the superficial region, whereas the deeper layers have an increased number of cells. With increasing age, there is an overall reduction in the water content of the extracellular matrix, with a corresponding increase in compressive stiffness. This may result in increasing forces in the underlying subchondral bone, as the cartilage loses its ability to undergo reversible deformation. 



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TABLE 6.2  Alterations in the structure of cartilage seen with both aging and osteoarthritis. Component

Aging

Osteoarthritis

Water content

↓

↑ (90% compared to 65%–80%)

Modulus/stiffness

↑ (less elastic)

↓ (more elastic)

Chondrocytes

↓ (increase in size)

Cells cluster in late OA

Collagen

↔

↓

Proteoglyan synthesis

↓

↑

Proteoglycan degradation

↓

↑↑

Glycosaminoglycans Chondroitin sulfate (4-and 6-)

↓

↑

Glycosaminoglycans Keratan

↑

↓

Advanced glycosylation end produces (AGE)

↑

Accumulation of AGE thought to lead to knee and ankle OA

A summary of the changes in composition of cartilage with both aging and osteoarthritis is shown in Table 6.2.

1.4  Cartilage injuries and healing Injury to articular cartilage can lead to significant musculoskeletal morbidity and result in premature arthritis if not addressed appropriately. The unique and complex structure of articular cartilage makes treatment and repair or restoration of defects challenging. Localized cartilage injuries can broadly be divided into superficial lacerations, which occur superficially to the tidemark, and deep lacerations which occur through the tidemark. Attempts at cartilage repair are made by the cartilage itself and by surrounding tissue, the mechanism of which depends on the location of the injury. In superficial or partial-thickness lacerations, the cartilage makes minimal attempts to repair itself, with little assistance from the surrounding tissue. Chondrocytes proliferate and there is migration of a small number of synovial fibroblasts toward and into the defect, but no actual healing takes place. In full-thickness injuries, there is a marked response from the underlying subchondral bone with hematoma and migration of large numbers of stem cells. However, these undifferentiated marrow mesenchymal stem cells are unable to reconstitute tissue resembling the original hyaline cartilage and instead produce type one collagen. This results in the formation of fibrocartilaginous reparative tissue which is both mechanically suboptimal and inadequate for durable structural bonding with the uninjured surrounding hyaline cartilage. 1.4.1  Causes of chondral injury Chondral injuries can broadly be divided into focal and degenerative lesions. Focal defects are well demarcated defects, which usually occur following trauma, whereas degenerative



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defects tend to be poorly demarcated and occur secondary to joint malalignment, ligamentous instability, meniscal tears, or arthritis. Trauma is the most common cause of focal chondral injury; shearing forces create a stress fracture through the cartilage matrix and can extend into subchondral bone. Dislocation of the patella leads to the formation of an osteochondral fracture via this mechanism in some cases and is responsible for 40%–50% of osteochondral lesions around the femoral condyles. There are several other intra-articular pathologies which can result in focal chondral lesions, including osteochondritis dissecans, spontaneous or secondary osteonecrosis. In the absence of a clear history of trauma to the affected joint, these differential diagnoses should be considered. Osteochondritis dissecans (OD) is a condition affecting articular cartilage and subchondral bone. It results in varying pathological abnormalities, beginning with softening of the articular cartilage, early cartilage separation and partial or, in some cases, complete separation of an osteochondral fragment. It usually affects the posterolateral aspect of the medial femoral condyle. OD often affects adolescents and young adults [3]. Osteonecrosis of the knee can be broadly divided into four groups: spontaneous osteonecrosis of the knee, secondary, post-arthroscopic and post-traumatic osteonecrosis. Spontaneous osteonecrosis is the most common type of osteonecrosis and affects an older age group, in comparison to secondary osteonecrosis, which is associated with other underlying medical problems or medication. Alcohol, steroid use, sickle cell disease, myeloproliferative disorders, and renal disease have all been implicated in the development of secondary osteonecrosis [4]. Post-arthroscopic osteonecrosis is thought to be the rarest form, with an onset following arthroscopic knee surgery, more specifically, following meniscectomy. In post-traumatic osteonecrosis, as the name suggests, there is a history of trauma or surgery preceding symptom onset, leading to bone death, usually in an isolated area of the knee. It can sometimes be difficult to differentiate these conditions from the history and clinical examination alone and radiological findings can also be misleadingly similar. Certain characteristics may help to differentiate these conditions from one another and are summarized in Table 6.3 [5,6,7]. 1.4.2  Classification of chondral defects There are numerous classification systems for assessing and grading chondral damage. The Outerbridge classification was initially introduced as a simple grading system for Chondromalacia patellae but has since been extrapolated and is now in use for joints throughout the body. It grades lesions from 1 to 4 in order of increasing severity according to their macroscopic appearance (Table 6.4). Cartilage injuries can also be graded on their microscopic appearance according to the international cartilage repair society score (ICRS), which ranges from grade 0 (normal) to grade 4 (severely abnormal) (Fig. 6.2) [9]. 1.4.3  Current treatments and challenges in management Once trauma or disease provokes and intra-articular destructive process, adult articular cartilage has a limited ability to spontaneously heal, particularly for larger defects [10]. This is due to the inability of chondrocytes to migrate to the site of injury, its avascular nature and the absence of a fibrin clot scaffold.





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TABLE 6.3  Differential diagnoses to be considered for chondral lesions. Spontaneous osteonecrosis

Osteochondritis dissecans

Secondary osteonecrosis

Post-arthroscopic osteonecrosis

Patient age group

>55 years

Young-middle age

80% bilateral

Unilateral

Clinical features

Usually sudden onset of server pain

Insidious onset Prior knee trauma in 40%

Usually insidious onset of pain

Acute onset pain, often after initial post-op improvement

Radiological features

Dependent on stage: can progress from sclerotic lesion to condylar collapse

Well circumscribed area of subchondral bone separated from femoral condyle by crescentshaped radiolucent line.

Lesions much larger and area of osteonecrosis more diffuse than in SONK

Similar to spontaneous osteonecrosis

TABLE 6.4  The Outerbridge classification for grading chondral damage [8]. Outerbridge grade

Description

I

Softening and swelling of the cartilage

II

Fragmentation and fissuring in and area ≤1/2 inch in diameter

III

Fragmentation and fissuring in and area ≥1/2 inch in diameter

IV

Erosion of cartilage down to subchondral bone

Several treatment modalities have been described, depending on the site and size of the lesion. They represent a wide spectrum, varying from non-operative measures to joint-preserving and joint replacement surgery. Although numerous options exist for the treatment of cartilage defects, most options are directed at symptomatic relief, but do not recreate the normal joint mechanics to allow long-term healing. These methods are have shown success to



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6.  Biology of cartilage

FIGURE 6.2  International cartilage repair society score for classifying chondral injuries. Source: Images from https://cartilage.org/content/uploads/2014/10/ICRS_evaluation.pdf.

varying degrees, but they have their limitations, particularly as the properties of the repaired region are not as robust as the original cartilage and often lack integration with the surrounding, uninjured cartilage [11]. Numerous surgical treatments have been described for the treatment of focal chondral lesions, including arthroscopy with drilling, perforation, and decompression of the lesion; resurfacing the lesion either with microfracture and using autograft or allografts for transplant 



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TABLE 6 5  Current treatment options for chondral defects. Treatment

Description

Benefits

Limitation

Non-operative

Analgesia, physiotherapy

Avoids risks of surgery

Symptom relief only

Arthroscopic chondroplasty

Resection of loose cartilage to reduce mechanical symptoms

Simple procedure

Symptom relief only

Arthroscopic microfracture

Small intra-articular holes made with pick to release cells and encourage fibrocartilage formation

Minimally invasive, no graft needed Used for lesions 4% or logistic EuroSCORE I >10%), the decision between SAVR and TAVI should be made by the Heart Team according to the individual patient characteristics with TAVI being favored in elderly patients suitable for trans-femoral access (class IB).

3  Imaging work up for TAVI patient Multimodality imaging is needed for preprocedural planning and intraoperative decision making given the complex 3-dimensional (3D) anatomy of the aortic valve, sinuses, and annulus. Imaging guidance helps prevent suboptimal valve deployment, which is associated with an increased risk of complications such as paravalvular regurgitation, aortic injury, heart block, and embolization of the valve prosthesis [15]. This typically include multidetector computed tomography (MDCT) that gives detailed information on the aortic annulus, aortic root of valve sizing but also provide detailed information on the coronary arteries, leaflet morphology, calcification, and identification of other challenging anatomical features. In addition, it evaluates the entire thoraco-abdominal aorta, major thoracic arterial vasculature, carotids, and ilio-femoral vasculature and guide alternative access approaches to the transfemoral such as subclavian, transapical, direct aortic, carotid, or even transvenous access approaches. Echocardiography, transthoracic echo (TTE) and/or transoesophageal echo (TOE), is essential to confirm the diagnosis and assess the severity of AS, asses the left ventricular function and dimensions, and exclude other valvular abnormalities. Furthermore, when used intraprocedural, TOE helps to guide the implantation procedure and in early detection of the complications. Invasive coronary angiography is also required to assess the coronary artery disease and the need for percutaneous coronary intervention that can be done as a staged procedure prior to the TAVI.



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4  Approaches for TAVI access Various routes have been described for gaining access to the aortic valve for the implantation procedure. The femoral artery has been the most popular access site. Although originally requiring a surgical cut-down, most experienced groups now utilize a percutaneous puncture and suture pre-closure technique, avoiding the need for open surgical access. Current consensus and guideline strongly favors transfemoral arterial access as the preferred default approach for TAVI, whenever possible, as it is associated with better outcomes [14]. Vascular access sheaths are typically described in terms of their inner diameter in French size (3 × diameter in millimeters). The outer diameters are slightly larger by as much as 2–4 French (0.67–1.33 mm) depending on the material used. A relatively compliant non-diseased artery can generally accommodate a sheath slightly larger than its internal diameter. However, many patients have small or diseased femoral arteries. On occasions, an open surgical retroperitoneal approach is utilized to gain access to the larger iliac artery in patients with femoral disease. Other alternative approaches are available in case of difficulty in gaining aces through the transfemoral route. The transaxillary (sometimes referred to as subclavian) access, has gained popularity with a surgical cut down in the majority of the cases. A transapical approach, with direct access to the left ventricle through an intercostal thoracotomy can also be considered as a viable option that offer a direct pathway to the aortic valve and easier antegrade crossing of the valve. However, risks include myocardial injury, bleeding, mitral injury, and postoperative respiratory compromise and thoracotomy pain. A transaortic approach requiring a mini-thoracotomy and aortotomy is another access route that could potentially reduce the risk associated with transapical route. More recently, a transcaval access has been described by creating an iatrogenic abdominal aorto-caval fistula. The introducer sheath is advanced from the femoral vein through the inferior vena cava into the adjoining infrarenal abdominal aorta at a preselected target site. When the introducer sheath is removed, the caval-aortic access tract is closed by implanting a nitinol occluder device [16].

5  Transcatheter heart valves and delivery systems Since Alain Cribier performed the first transcatheter aortic valve implantation (TAVI) in an inoperable patient in 2002, transcatheter valve intervention has become an established therapy for patients with severe symptomatic AS. It has been widely recognized that TAVI is the treatment of choice for severe aortic stenosis in inoperable patients [17,18] and as a reasonable alternative to SAVR in patients with intermediate and high surgical risk [19,20]. The first prototype transcatheter heart valve (THV) designed by Cribier was a stainless steel stent that contained a trileaflet valve made up of bovine pericardium. After a few years, this prototype evolved into the Cribier-Edwards valve (Edwards Lifesciences), and transfemoral or transapical approaches were used instead of the original trans-septal technique [21]. Simultaneously, Medtronic has developed another device, the self-expandable CoreValve which is made of a nitinol frame containing a porcine pericardial valve. These two devices, which after a few years obtained CE Mark and US Food and Drug Administration (FDA) approval have the most clinical experience and published data to date. 



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5  Transcatheter heart valves and delivery systems

In the last 15 years, there have been great advancements in the TAVI technology that enabled transforming a challenging intervention into a standardized and safe procedure. Multiple modifications have been made to reduce the profile of the delivery catheter, minimize the rate of para-valvular leak (PVL), and facilitate deployment and accurate positioning as well as some additional features such as the ability of repositioning and retrieval of some of the devices. All these aim to attain better clinical outcome and reduce TAVI-related complications. To date, a significantly expanding number of THV are approved by CE and FDA. These devices are categorized into balloon-expandable, self-expanding, and mechanically expandable valves (Table 8.1). A weighted meta-analysis of 30 studies including 5923 patients showed that the rate of permanent pacemaker is lowest with the Portico and accurate Neo devices (9%) and highest with the Lotus (32%), the rate of disabling stroke is lowest with Sapien 3 (15) and Jenavalve (2.4%) and highest with Accurate Neo and Evolute R (3.5%). The rate of more than mild Aortic regurgitation is lowest with Accurate Neo and Lotus devices (1%) and highest with TABLE 8.1  Different types and characteristics of TAVI devices. Device name

Valve structure

Access route, delivery system, valve size

Reference access vessel diameter

Fully Repositionable retrievable

Sapien 3

Bovine pericardial tissue valve; balloon– expandable cobalt chromium frame

TF: Edward sheath: 14F (20,23,26 mm) 16F (29 mm) TA,TAo: cerititude sheath: 18F(20,23,26 mm) 26F (29 mm)

>5 mm (except for 29 mm: > 5.5 mm)

No

No

Evolut R

Porcine pericardial tissue TF, Tao, TSc; EnVeo valve; self-expanding sheath nitinol frame 14F (23,26, 29 mm) 16 F (34 mm)

>5 mm (except for 34 mm: > 5.5 mm)

Yes

Yes

Portico

Bovine pericardial tissue valve; self-expanding nitinol frame

≥6 mm

Yes

Yes

Acurate Neo

Porcine pericardial tissue TF: 18F (small, medium, valve; self-expanding large) nitinol frame TA: 28F (small, medium, large)

≥6 mm

No

No

—

Yes

No

TF, Tao, TSc; 18F( 23, 25 mm) 19F(27, 29 mm)

Jenavalve Porcine pericardial tissue TA: sheathless 32F valve; self-expanding (23,25,27 mm) nitinol frame Lotus

Bovine pericardial tissue valve; self-expanding braided nitinol frame

TF : 18F (23mm) 20F (25, 27 mm)

≥6 mm (except for 25 and 27 mm; >6.5 mm)

Yes

Yes

Allegra

Bovine pericardial tissue valve; self-expanding nitinol frame

TF: 18F (23, 27, 31 mm)

≥6 mm

Yes

Yes

Abbreviation: TF, trans femoral; TA, trans apical; Tao, trans aortic; TSc, trans subclavian.



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8.  Advances in transcatheter aortic valve implantation

Evolute R (5.5%), the rate of vascular complications is lowest with Lotus (2.4%) and higest with Jenavalve (6%), and the rate of 30-days mortality is lowest with Sapien and Evolute R (around 2%) and higest with Jena valve (11%) [22].

5.1  Balloon expandable valve The Edwards SAPIEN valve is the only commercially available balloon- expandable valve to date. It was initially made from bovine tissue and polyester supported with a stainless steel mesh frame (Fig. 8.1A). It was available in a diameter of 23 mm or 26 mm. The delivery system was called Retroflex3 and its diameter varied from 22 French (Fr) up to 24 Fr. Over the years some modifications led to the development of the most recent SAPIEN XT (Fig. 8.1B) and SAPIEN 3 valves (Fig. 8.1C). All SAPIEN valves require rapid ventricular pacing through a temporary transvenous pacing wire to stabilize the deployment process and allow for accurate valve positioning. Compared to prior generation, the SAPIEN XT consist of bovine pericardium mounted within a cobalt-chromium frame with thinner struts and more open cell design. These modifications allow for a lower crimped profile and a smaller sheath diameter, while maintaining the valve’s radial stiffness and structural integrity. The available sizes are 20 mm, 23 mm, 26 mm, and 29 mm with respective diameter of 16, 16, 18, and 20 Fr delivery sheaths for Novaflexe transfemoral (TF) access. For the transapical (TA) access, the Ascendra delivery system was used [23]. SAPIEN 3 THV represent the latest generation of this family. Compared to Sapien XT, it is covered by an outer polyethylene terephthalate cuff to enhance para-valvular sealing and the delivery system has been made ultra-low profile (the transfemoral Commander system, 14 Fr eSheath for the 20 mm, 23 mm and 26 mm valves, and 16 Fr eSheath for the 29 mm valve) [24]. In addition, the delivery system has been modified to enable advancing or retracting the valve several millimeters up or down within the annulus without pushing or pulling on the entire delivery system. With the use of these new delivery system for trans-femoral access, the loading of the valve occurs inside of the descending aorta which allows a reduction of the sheath diameter (minimum TF diameter of 5.5 mm). For transapical implantation, the Certitude is the new corresponding delivery system that also features a smaller nose cone.

FIGURE 8.1  The evolution of Edwards Sapien THV. (A) Sapien Valve; (B) Sapien XT; (C) Sapien 3. Source: Reproduced with permission of Edwards lifesciences LCC, California, United States.





5  Transcatheter heart valves and delivery systems

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The wide range of valve size (20–29 mm) allow for the treatment of patients with an aortic annulus of up to 27 mm and annular size range of 273 to 683 mm2 by Computed Tomography (CT) area.

5.2  Self-expandable valves I. The Medtronic CoreValve revalving system: The CoreValve (Fig. 8.2A) was the first generation valve for the TF approach or other retrograde access options (trans-aortic, trans-subclavian). The valve consists of porcine pericardium leaflets mounted within a nitinol frame. The Accutrack delivery system (18 Fr) was available for transfemoral implantation of prostheses of size 26 mm, 29 mm, and 31 mm and allows for the treatment of patients with an aortic annulus diameter of up to 29 mm.

FIGURE 8.2  The Medtronic CoreValve family. (A) CoreValve; (B) Evolute R; (C) Evolute PRO. Source: Reproduced with permission of Medtronic Inc, Minnesota, United States.



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8.  Advances in transcatheter aortic valve implantation

Once positioned within the diseased native valve the delivery catheter is withdrawn, releasing the THV. This valve features a long multi-staged frame which not only anchors within the aortic annulus, but also extends superiorly to anchor in the supracoronary aorta. The supra-annular valve design maximizes leaflet cooptation and promotes single digit gradients and large effective orifice area. The current generation, the Evolut R device (Fig. 8.2B), is currently available in four device sizes of 23, 26, 29, and 34 mm, allowing the treatment of native valves with a perimeter of 56.5–94.2 mm and annulus diameter up to 30 mm as assessed by CT-scan [25]. The Evolut R provides several modifications to improve anatomical fit, annular sealing, and durability. In particular, the device is designed to enable recapturability and repositionability with the option to recapture and reposition up to three times before reaching the “point of no recapture.” The Evolut R frame is tailored to reduce the overall height, while preserving the height of the pericardial skirt (13 mm) with an extended skirt of the inflow tract to provide a seal against PVR. In addition, cell geometry has been redesigned to achieve optimized radial force. The new delivery system, the EnVeo R, integrates an InLine which slides against the capsule to allow vascular access that is the equivalent of a 14-F system (16 F for the Evolut R 34 mm). This means that the Evolut R system is now indicated to treat minimum access vessels of ≥ 5 mm (Evolut R 23, 26, 29 mm) and ≥ 5.5 mm (Evolut R 34 mm). Furthermore, positioning accuracy is aided by the EnVeo R delivery system’s 1:1 delivery response. The Evolut PRO is the latest-generation of this family (Fig. 8.2C), which recently obtained approval from the FDA, and consists of the same platform as the Evolut R but incorporating an outer porcine pericardial tissue wrap to the first 1½ inflow cells to improve sealing further between the device and the native aortic annulus and to minimize PVL. It is indicated for vessels down to 5.5 mm and its EnVeo R system with the InLine sheath allows for a delivery profile as low as 16-Fr equivalent. It is currently available in 23, 26, and 29 mm sizes. It was tested by the investigators of the Evolut PRO clinical study (N = 60). At that 30-day followup, the Evolut Pro system resulted in high rates of survival (98.3%) and low rates of disabling strokes (1.7%). None to trace PVL was observed in 72.4% of patients while the remaining 27.6% experienced mild PVL. There were no patients with moderate or severe PVL, and lower rates of permanent pacemaker implantation (10%) as compared to previous iterations of selfexpandable valve [26]. II. PORTICO valve The Portico valve (Fig. 8.3) is composed of nitinol frame, bovine leaflets, and a porcine pericardial sealing cuff. There are multiple specific features for this valve which includes (1) the large cell area allow easy engagement of the coronary ostia after implantation and minimizes the risk of PVL by allowing valve tissue to conform around calcific nodules at the annulus; (2) low placement of leaflets within the frame allows for minimal protrusion into the left ventricular outflow tract meant to reduce the incidence of AV-blocks; (3) the proximal stent part is covered by a tissue cuff designed to minimize paravalvular leaks; (4) during implantation, the valve can be re-sheathed and re-positioned or even fully retrieved prior to full deployment; and (5) the valve is functioning during the deployment process to maintain hemodynamic and enable accurate stable positioning. A balloon aortic valvuloplasty is mandatory before deployment of this THV. The 23- and 25-mm valves are loaded onto an 18-F delivery system, whereas the 27- and 29-mm valves are loaded onto a 19-F delivery system [22].





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FIGURE 8.3  The Portico device. Source: Reproduced with permission of Abbott Inc., Illinois, United States.

The Portico valve can be delivered by transfemoral access, and when used with the SoloPath recollapsible introducer (St. Jude Medical, Inc.), has a low 13.5-F insertion profile. III. Symetis Acurate NEO valve (Boston Scientific) The Acurate Neo aortic bioprosthesis (Fig. 8.4) is a second-generation self-expanding valve composed of porcine pericardium tissue sewn onto supra- annular nitinol frame that is covered both externally and internally by an antileak porcine pericardium skirt. The device has multiple unique features including three stabilization arches to enable axial self-alignment to aortic annulus, a top crown for capping the native leaflets and provide coronary clearance,

FIGURE 8.4  The Acurate Neo THV. Source: Reproduced with permission of Boston Scientific Corporation, Massachusetts, United States.



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8.  Advances in transcatheter aortic valve implantation

and a bottom crown with minimal protrusion into the left ventricular outflow tract (LVOT) to minimize the conduction system disturbance, anti-leak skirt that provide sealing against PVL leak and the tissue is treated with BioFix anti-calcification technology. The prosthesis can be implanted through both the transapical (28 F) and the transfemoral (15–18 F) routes using a simple two-step deployment (top-down) and stable positioning. The Acurate Neo comes in three different sizes: small (21–23-mm aortic annulus), medium (23–25-mm aortic annulus), and large (25–27-mm aortic annulus). When compared to S3, the Accurate Neo has comparable 30 days safety and efficacy outcome and significantly reduced risk of new permanent pacemaker implantation and elevated trans-valvular gradients and overall lower Mean Gradients, but higher moderate/severe PVL rate [27]. IV. Direct flow Medical (Direct Flow Medical Inc.) The concept of the Direct Flow transfemoral transcatheter prosthesis relies on two inflatable rings for anchoring the valve at the annular level [22]. This makes the valve fully retrievable even after the device had been fully deployed. However, the radial forces of this design are lower than metallic based frame. Thus, in heavily calcified native aortic valves problems occurred. This valve is no longer commercially available. V. The JenaValve The Jena valve (JenaValve Technology, Inc.) consists of a porcine valve mounted on a low-profile, self-expanding nitinol frame. The low profile design and absence of stent mesh in the area of the coronary ostia mitigate the risk of coronary obstruction. The valve is designed to be repositionable and retrievable. It is available in three different sizes (23, 25, and 27 mm) for implantation in aortic annuli that are 21–27 mm in diameter [22]. A sheathless 32-F delivery system is used for a three-step deployment procedure through the transapical route. What make this valve unique is that the implantation relies on active clip fixation of the native aortic leaflets independent of the extent of native valve calcification providing secure anchorage. In fact, the JenaValve is the only TAVI device to have obtained CE Mark approval for non-calcified aortic regurgitation. It must be noted that the CE Mark for the JenaValve has expired in 2016, and the company stopped commercial distribution in June 2016. Currently, newer generation of JenaValve Pericardial THV is under development and assessment in trials with both transapical (22 F) and transfemoral (19 F) delivery systems (see Section 6). VI. ALLEGRA valve The Allegra THV (NVT AG) consists of a supra-annular nitinol frame and bovine pericardium (annular skirt and leaflets). The annular portion of the frame is covered with a 12 mm sealing skirt to minimize the PVL [22]. Unique features include six radiopaque gold markers to allow for perfect visibility and precise implantation and T- bars at the top of the prosthesis allowing for safe anchoring to the catheter. The valve is available in three sizes (23, 27, and 31 mm) to fit for annular diameter from 19 to 28 mm. The stent frame uses a variable cell size design to allow for axially tailored radial force distribution with higher force in the annular sealing section of the valve for secure anchoring. The upper section of the stent frame has larger cells to allow for flexure of the stent frame and accommodation of conformational changes during the cardiac cycle to mitigate leaflet stresses. The transfemoral delivery system incorporates an 18-F cartridge and a 15-F catheter shaft. The grip uses a “squeeze-to-release” mechanism allowing for a stepwise controlled





5  Transcatheter heart valves and delivery systems

113

implantation avoiding any rotation. The PermaFlow principal ensures accurate and controlled release of the valve in three steps and allow for stable hemodynamics at any time during the release process.

5.3  Mechanically expandable valve The LOTUS valve system (Boston Scientific Corporation) (Fig. 8.5) consists of a bovine pericardial tissue supported on a braided nitinol frame. It is mainly delivered by TF route but direct aortic and trans-axillary approaches are other alternative. Positioning of the prosthesis within the aortic root is facilitated by a central radio-opaque marker. The frame is covered with an adaptive seal at the inflow segment that minimize the PVL by conforming to the irregular anatomical surface. This THV is currently available in three sizes, 23, 25, and 27 mm, covering a range of annulus diameters from 19 to 27 mm. The 23-mm model can be delivered through an 18-F sheath (small), while the 25- and 27-mm valves require a 20-F sheath (large). This is the only TAVI device that it is fully recapturable and repositionable even after the valve has been fully deployed. The Lotus valve is associated with the lowest rate of PVL and similar mortality, stroke, and vascular complication rate when compared to other valves [28–30]. The REPRISE 3 trial showed that the mechanically expanded Lotus valve was not inferior in term of primary safety end point or the primary effectiveness end point to the selfexpanding CoreValve/Evolut R valves [31]. Nevertheless, the high rate of conduction disturbances requiring pacemaker implantation with this valve (approximately 30%) remained a concern.

FIGURE 8.5  The Lotus valve technology. Source: Reproduced with permission of Boston Scientific Corporation, Massachusetts, United States.



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8.  Advances in transcatheter aortic valve implantation

Newer Lotus Edge device is designed to mitigate this problem (see Section 6). In early 2017, Boston Scientific announced a voluntary removal of all Lotus THV devices from global commercial and clinical sites. The action is a response to reports of the premature release of a pin connecting the Lotus valve to the delivery system. It is expected that it will be back on shelves in Europe and the United States in 2019.

6  Future perspectives 6.1  Extended indications Two large recently published trials, the Medtronic arm using the CoreValve Evolut R System (NCT02701283) and the Edwards Sapien arm using the Edwards Sapien 3 valve system (Partner 3 trial, NCT02675114), that tested the safety and efficacy of TAVI in low-risk patients showed that TAVI is safe and as effective as SAVR. If approved by the FDA, it will open up TAVI as a treatment option to a much larger pool of patients with aortic stenosis. Data on TAVI are still limited for younger population  10 years are still lacking albeit there are no major issues of structural valve degeneration have been reported as yet. As mentioned earlier, Bicuspid aortic valve represents a challenge for TAVI given the complex anatomy. Hence, it remains to be determined whether newer generation THVs and newer technologies with THVs specific for bicuspid aortic valves may prevent paravalvular aortic regurgitation. Specific clinical trials among patients with bicuspid AS and newer generation THVs are warranted. To date, the role of TAVI for native aortic regurgitation treatment is marginal. The absence of aortic annular and aortic valve leaflet calcification in patients with pure aortic regurgitation is a known risk factor for THV embolization and migration. Even though, TAVI has been infrequently used in the treatment of native severe aortic regurgitation in in-operable cases [32]. For this off-label indication, the self-expandable valve with better radial strength is used to avoid valve dislocation in the absence of calcium to anchor the valve in place. New device technologies (J-Valve; JieCheng Medical Technology Co., Ltd., Suzhou, China), specifically designed to be implanted in non-calcific aortic valves, are under preclinical and clinical investigation and will probably extend the indications for TAVI in this population [33].

6.2  Future THV and delivery systems Medtronic incorporation plan to launch Evolute PRO 34 mm valve and the new EnVeo PRO Seamless tracking delivery system for easily trackability in tortuous calcified vessels. Also, future valve generations as Evolute NG for improved visualization, controlled release and lower profile and the Horizon valve for concentric deployment, enhanced sealing, and superior hemodynamics are in development.





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FIGURE 8.6  Sapien 3 Ultra valve. Source: Reproduced with permission of Edwards lifesciences LCC, California, United States.

Sapien 3 Ultra system (Fig. 8.6) featuring Sapien 3 leaflet and design with textured 3D PET skirt design which is 40% taller to allow for more contact surface with the native anatomy. The delivery system has a unique on-balloon design to remove the need for valve alignment, utilizes 14F Axela sheath for all valve sizes with 5.5 mm vessel indication, and has Seamless sheath design allowing for transient expansion and active contraction. The Edwards CENTERA valve (Fig. 8.7) is a self- expanding, repositionable, and retrievable valve that can be delivered through a low-profile, 14-French, motorized delivery system for TF use only. It has received CE Mark in early 2017 based on CENTERA-EU Trial that showed low incidence of cardiovascular mortality (4.6%), disabling stroke (4.1%), permanent pacemaker (6.0%), and no moderate or severe total aortic regurgitation [34]. It has a unique feature of being packaged with the valve fully pre-attached to the delivery system, which facilitates simple and rapid device preparation.

FIGURE 8.7  The Centera valve. Source: Reproduced with permission of Edwards lifesciences LCC, California, United States.



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FIGURE 8.8  The new Lotus edge valve. Source: Reproduced with permission of Boston Scientific Corporation, Massachusetts, United States.

The second-generation JenaValve Pericardial THV is currently under clinical evaluation with both transapical (22 F) and transfemoral (19 F) delivery systems. It features a dual annular sealing mechanism and locator engagement with the native valve cusps that is intended to anatomically mitigate the risk of low placement into the LVOT. The newer-generation LOTUS Edge device (Fig. 8.8) are developed with the hope of reducing the frequency of conduction disturbances by featuring proprietary Depth Guard technology, designed to reduce LVOT interaction and PPM rate by minimizing the depth of the valve during deployment. Also, it features an additional radio-opaque marker for one-view locking and the delivery catheter is made more flexible with lower profile. There will be also 2 additional sizes; 21 and 29 mm. Boston Scientific plans to relaunch the Lotus Edge valve in Europe by 2019. Finally, the Symetis Acurate Neo Advanced seal which features modified skirt material to reduce the PVL are recruiting patients for CE mark trial.

6.3  Advances in CT imaging and closure devices Advances in TAVI technologies are not limited only to the implantation devices but also are extended to include the imaging and closure devices. With regard to imaging, the incorporation of CT with advanced software such as the VesselNavigator and the HeartNavigator 3 (Philips Company, Amsterdam, The Netherlands) provide insightful planning and guidance for the procedure. The VesselNavigator allows the reuse of 3D vascular anatomical





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information from existing CT datasets as a 3D roadmap overlay on a live X-ray image providing an intuitive and continuous 3D roadmap to guide through vasculature during the entire procedure. The HeartNavigator 3 enable automatic “view” planning to get the best view, enhanced device selection to get the best fit, automatic 3D segmentation and aortic measurement to save time and avoid errors, automatic calcification indicator to enhance the preparation, live image overlay for better orientation, and real time guidance moving in-sync with C-arm. The development of percutaneous closure device such as PerQseal (Vivasure Medical Limited, Galway, Ireland) designed to enable closure of large arteriotomies up to 24 F using an intravascular patch (resorbable synthetic polymer) and the percutaneous MANTA vascular closure device (VCD) (Essential Medical Inc., Malvern, Pennsylvania, Unites States) which is a collagen-based technology dedicated to the closure of large bore arteriotomies up to 25 F will reduce the bleeding rate and time to achieve hemostasis.

7 Conclusion The new generation TAVI devices are proving to be considerably safer and more efficient than their ancestors. Ongoing advancement in this technology is expected to allow for treatment of increasing number of patients with different clinical and anatomical scenario.

References [1] Cribier A, Savin T, Saoudi N, et al. Percutaneous transluminal valvuloplasty of acquired aortic stenosis in elderly patients: an alternative to valve replacement? Lancet 1986;1(8472):63–7. [2] Bernard Y, Etievent J, Mourand JL, et al. Long-term results of percutaneous aortic valvuloplasty compared with aortic valve replacement in patients more than 75 years old. J Am Coll Cardiol 1992;20(4):796–801. [3] Percutaneous balloon aortic valvuloplasty. Acute and 30-day follow-up results in 674 patients from the NHLBI Balloon Valvuloplasty Registry. Circulation 1991;84(6):2383–97. [4] Kapadia SR, Leon MB, Makkar RR, et al. PARTNER trial investigators. 5-year outcomes of transcatheter aortic valve replacement compared with standard treatment for patients with inoperable aortic stenosis (PARTNER 1): a randomized controlled trial. Lancet. 2015;385(9986):2485–91. [5] Baron SJ, Arnold SV, Reynolds MR, US CoreValve investigator, et al. Durability of quality of life benefits of transcatheter aortic valve replacement: long-term results from the CoreValve US extreme risk trial. Am Heart J 2017;194:39–48. [6] Mack MJ, Leon MB2, Smith CR, PARTNER 1 trial investigator, et al. 5-year outcomes of transcatheter aortic valve replacement or surgical aortic valve replacement for high surgical risk patients with aortic stenosis (PARTNER 1): a randomised controlled trial. Lancet 2015;385(9986):2477–84. [7] Deeb GM, Reardon MJ, MD., Chetcuti S, et al. For the CoreValve US Clinical Investigators. 3-year outcomes in high-risk patients who underwent surgical or transcatheter aortic valve replacement. J Am Coll Cardiol 2016;67:2565–74. [8] Chakos A, Wilson-Smith A, Arora S, et al. Long term outcomes of transcatheter aortic valve implantation (TAVI): a systematic review of 5-year survival and beyond. Ann Cardiothorac Surg 2017;6(5):432–43. [9] Reardon MJ, Van Mieghem NM, Popma JJ, SURTAVI Investigators, et al. Surgical or transcatheter aortic-valve replacement in intermediate-risk patients. N Engl J Med 2017;376(14):1321–31. [10] Leon MB, Smith CR, Mack MJ, PARTNER 2 Investigator, et al. Transcatheter or surgical aortic-valve replacement in intermediate-risk patients. N Engl J Med 2016;374(17):1609–20.



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[11] Villablanca P, Mohananey D, Katarina N, et al. Comparison of local versus general anesthesia in patients undergoing transcatheter aortic valve replacement: a meta-analysis. Catheter Cardiovas Interven 2017;91(2). 10.1002/ ccd.27207. [12] Webb JG, Wood DA, Ye J, et al. Transcatheter valve-in-valve implantation for failed bioprosthetic heart valves. Circulation 2010;121(16):1848–57. [13] Mylotte D, Lefevre T, Søndergaard L, et al. Transcatheter aortic valve replacement in bicuspid aortic valve disease. J Am Coll Cardiol 2014;64:2330–9. [14] Baumgartner H, Falk V, Bax JJ, et al. The Task Force for the Management of Valvular Heart Disease of the European Society of Cardiology (ESC) and the European Association for Cardio-Thoracic Surgery (EACTS). 2017 ESC/EACTS Guidelines for the management of valvular heart disease. Eur Heart J 2017;38:2739–3279. [15] Binder RK, Webb JG, Willson AB, et al. The impact of integration of a multidetector computedtomography annulus area sizing algorithm on outcomes of transcatheter aortic valve replacement: a prospective, multicenter, controlled trial. J Am Coll Cardiol 2013;62:431–8. [16] Lederman RJ, Babaliaros VC, Greenbaum AB. How to perform transcaval access and closure for transcatheter aortic valve implantation. Catheter Cardiovasc Interven 2015;86(7):1242–54. doi:10.1002/ccd.26141. [17] Leon MB, Smith CR, Mack M, PARTNER Trial Investigators, et al. Transcatheter aortic-valve implantation for aortic stenosis in patients who cannot undergo surgery. N Engl J Med 2010;363(17):1597–607. [18] Popma JJ, Adams DH, Reardon MJ, CoreValve United States Clinical Investigators, et al. Transcatheter aortic valve replacement using a self-expanding bioprosthesis in patients with severe aortic stenosis at extreme risk for surgery. J Am Coll Cardiol 2014;63(19):1972–81. [19] Smith CR, Leon MB, Mack MJ, PARTNER Trial Investigators, et al. Transcatheter versus surgical aortic-valve replacement in high-risk patients. N Engl J Med 2011;364(23):2187–98. [20] Adams DH, Popma JJ, Reardon MJ, U.S. CoreValve Clinical Investigators, et al. Transcatheter aortic-valve replacement with a self-expanding prosthesis. N Engl J Med 2014;370(19):1790–8. [21] Cribier AG. The odyssey of TAVR from concept to clinical reality. Tex Heart Inst J 2014;41:125–30. [22] Torado D, Picci A, Barbanti M. Current TAVR Devices. Technical characteristics and evidence to date for FDAand CE Mark-approved valves. Cardiac Intervent Today. 2017;11. [23] Collas V, Philipsen T, Rodrigus I, et al. Transcatheter aortic valve implanation: review and current state of art. EMJ Int Cardiol 2014;1:52–61. [24] Binder RK, Rodés-Cabau J, Wood DA, et al. Transcatheter aortic valve replacement with the SAPIEN 3: a new balloon-expandable transcatheter heart valve. JACC Cardiovasc Interv 2013;6(3):293–300. [25] Manoharan G, Walton AS, Brecker SJ, et al. Treatment of symptomatic severe aortic stenosis with a novel resheathable supra-annular self-expanding transcatheter aortic valve system. JACC Cardiovasc Interv 2015;8:1359–67. [26] Forrest JK. 30-day safety and echocardiographic outcomes following transcatheter aortic valve replacement with the self-expanding repositionable Evolut PRO system. Washington: American College of Cardiology Annual Scientific Session; 2017. pp. 17–19. [27] Husser O, Kim WK, Pellegrini C, et al. Multicenter comparison of novel self-expanding versus balloon-expandable transcatheter heart valves. JACC Cardiovasc Interv 2017;10(20):2078–87. [28] Meredith IT, Worthley SG, Whitbourn RJ, et al. Transfemoral aortic valve replacement with the repositionable Lotus Valve System in high surgical risk patients: the REPRISE I study. EuroIntervention 2014;9:1264–70. [29] Meredith Am IT, Walters DL, Dumonteil N, et al. Transcatheter aortic valve replacement for severe symptomatic aortic stenosis using a repositionable valve system: 30-day primary endpoint results from the REPRISE II study. J Am Coll Cardiol 2014;64:1339–48. [30] Falk V, Modine T, Brecker S, et al. Post-market evaluation of a fully repositionable and retrievable aortic valve in 500 patients treated in routine clinical practice: interim results from the RESPOND study. Paper presented at: PCR London Valves; September 2015; London, United Kingdom. [31] Feldman TE, Reardon MJ, Rajagopal V, et al. Effect of mechanically expanded vs self-expanding transcatheter aortic valve replacement on mortality and major adverse clinical events in high-risk patients with aortic stenosis: the REPRISE III Randomized Clinical Trial. JAMA 2018;319(1):27–37.



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[32] Anna F, Raffaele P, George CM, Siontis, et al. Transcatheter aortic valve replacement for the treatment of pure native aortic valve regurgitation: a systematic review. JACC Cardiovascular Interventions 2016;9(22):2308–17. [33] Wei L, Liu H, Zhu L, et al. A new transcatheter aortic valve replacement system for predominant aortic regurgitation implantation of the J-valve and early outcome. JACC Cardiovasc Interv 2015;8:1831–41. [34] Reichenspurner H, Schaefer A, Schäfer U, Tchétché D, et al. Self-expanding transcatheter aortic valve system for symptomatic high-risk patients with severe aortic stenosis. J Am Coll Cardiol 2017;70(25):3127–36. doi: 10.1016/j.jacc.2017.10.060.



C H A P T E R

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Advances in magnetic resonance imaging (MRI) Khurram Shahzad, Wael Mati Department of Radiology, Blackpool Victoria Hospital, Blackpool, United Kingdom

1 Introduction Discovery of X-rays in 1895 by Wilhelm Conrad Röentgen created an immense interest and excitement for the scientists and physicians alike. Within 6 months, X-rays were being used to treat the patients by assessing fractures and localizing bullets in battlefields. In 1896, first X-ray department of the world was opened at Glasgow [1]. However, unfortunately due to their easy availability and no usage restriction, the general public was using X-rays for completely different reasons such as coin operated X-ray booths to take images as a souvenir similar to a modern day photograph, shoe stores offering customers to see their feet inside shoes with X-rays for a better fit [2], X-ray machines used at parties and in theatres during performances as part of the play, X-ray shops where anyone could have an X-ray to assess all kinds of illnesses. Members of the public were also getting exposed to radiation by fake physicians and individuals using radioactive radium dissolved in water and sold as a super drug which lead to some high-profile deaths. This was despite that not long after their discovery, it was realized that X-rays were not all that safe. In 1896, initial reports started to surface about the harmful effects of over exposure from X-rays, which led to damage hands and fingers. In the next 10 years, tissue damage caused by radiation was recognized and widely reported. It was not only the general public who were exposed to the harmful effects from over exposure to X-rays but also physicians as there was no concept at the time of using protective equipment when working with X-rays regularly [2,3]. It was the result of earlier mentioned concerns among physicians and scientists that International Commission on Radiological Protection (ICRP) was established in Stockholm in 1928 which issues guidelines for radiation protection of doctors and general public on a regular basis and provides guidance rules for Radiology departments throughout the United Kingdom. There was also awareness of the need for researchers to develop other means of Advances in Medical and Surgical Engineering. http://dx.doi.org/10.1016/B978-0-12-819712-7.00009-7 Copyright © 2020 Elsevier Inc. All rights reserved.

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imaging human bodies without exposing them to radiation and this was one of the reasons that led to the development of magnetic resonance imaging (MRI).

2  A brief history of development of MRI [4,5] Even though Nikola Tesla discovered rotating magnetic field in 1882, it was not until the 1930s and 40s when Professor Isidor Isaac Rabi, Felix Bloch, and Edward Mills Purcell observed and discovered the phenomenon of nuclear magnetic resonance (NMR) which is the underlying principle for MRI. In 1950, Herman Carr created a one-dimensional MR image and in 1956 Tesla was proclaimed as a unit on which magnetic field strength of all future MRIs was based. Stronger the magnetic field, higher the Tesla unit and stronger the signals creating MR image. Up until 1970, MRI was mainly used for chemical and physical analysis. In 1973, Paul Lauterbur produced the first NMR image of a test tube. In 1975, Richard Ernst proposed the use of phase and frequency encoding and Fourier transformation for acquisition of MR images. Initial progress in human MRI scanning was made by Raymond Vahan Damadian who built the first MRI scanner and created the first human thorax MRI of his post-doctoral student Larry Minkoff on 3rd of July 1977, an imaging process which took nearly 5 hours. He also produced an MRI image of a human body with cancer in 1978 to show the different signal characteristics produced by normal and cancerous tissues. During the same time in 1970s, Paul Lauterbur used Herman Carr’s methods from 1950 and devised a way to produce 2-dimensional and 3-dimensional MRI images, and Sir Peter Mansfield of the University of Nottingham developed the technique of echo planar sequence for rapid imaging. In 1979, Richard Likes followed by Ljunggren and Tweig in 1983 introduced the concept of k-space which helps gather data for MRI acquisition. In 1986, Le Bihan described the concept of diffusion weighted imaging in MRI. Initial attempts at cardiac MRI imaging were unsuccessful due to problems with cardiac motion and respiration. Real time MRI imaging of heart was developed in 1987 with further improvements made with the use of ECG gating, faster scans, and breath holding techniques. In 1990s, functional MRI brain was developed which is an oxygen sensitive sequence to depict levels of activity in different areas of brain. Since 2000, further significant progress has been made to develop MRI sequences for body imaging including imaging of fetuses.

3  Advantages and disadvantages of MRI 3.1  Advantages of MRI The obvious advantage of using MRI is lack of exposure to ionizing radiations unlike Xrays (and CT scans). It also produces superior soft tissue contrast which not only helps to delineate soft tissue anatomy but further research and progress has meant that tissues can be characterized based on their perfusion, diffusion, and function. Vascular structures can be imaged without needing contrast medium which is particularly useful in patients with poor renal function. 



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3.2  Disadvantages of MRI MRI scanners are enclosed spaces which can be troublesome for patients with claustrophobia. They take relatively longer to perform (compared to CT scans) which makes them vulnerable to movement artefact or natural phenomenon such as swallowing, breathing, etc. MRI is not as good in evaluating bones or calcium deposition in tissues where CT scans are superior. Manufacturers are continuously working to address the earlier and have made significant progress in these areas which will be discussed later in this chapter. Any metal can affect MRI image quality from artefacts (this can also happen with CT scans) as well as can cause safety concerns due to the magnetic field. However, more and more manufacturers are producing MRI safe implants. There are various websites that allow one to check the safety of MRI in metallic implants one of which commonly used is www. mrisafety.com.

4  Basic physics of MRI [6–10] MRI scanners work by using the magnetic properties of hydrogen atoms in human body to create an image. Luckily nature has privileged us with abundant amounts of hydrogen atoms in the form of body water and fat. Percentage of body weight due to water decreases with increasing age, for example, approximate percentage body weight due to water in a fetus is 90%, infant 74%, child 60%, teenage and adult male 59%, teenage female 56%, adult female 50%, 56% in a male over 50 years, and 47% in a female over 50 years. However, in obese patients, total amount of water in body as part of weight percentage may be slightly lower than the above figures due to relatively more adipose tissue which contains water only in the range of 10%–15% which is even lower than what bone can contain (20%). Human body fat can vary greatly from 6% to 13% in athletic males, 14% to 20% in athletic females, and 25% to 32 % in obese males and females, respectively. On average, the mean percentage body fat range for average male is 18%–24% and for female 25%–31% [11]. Furthermore, weight gain in old age from lack of possible exercise and other contributing factors lead to further proportional increase in total body fat content. Hydrogen atoms with their single proton act as small magnetic bars that are randomly spinning around their own axis at any given time. Strong magnetic fields such as those used in an MRI scan can align these protons to the magnetic field, either along or opposite to the direction of magnetic field. Additional energy is sent to body in the form of radiofrequency (RF) waves which disturb this alignment resulting in some of the protons to start spinning out of sync. When the RF waves are turned off, the protons relax and return to their original position realigning with the main magnetic field. By doing so, an echo is produced which is detected by receiver coils and recorded in an arbitrary space known as the k-space. Gradients for spatial encoding are used to decide how the k-space will be filled. The earlier-mentioned process is repeated many times until all required data is recorded which is then used to generate an MR image. There are two types of relaxation which occurs in the tissues; magnetic vector/longitudinal relaxation which is known as T1 relaxation and axial spin/transverse relaxation known as T2. As different body tissues have different quantities of water and fat, they relax at different 

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rates and the protons take different times to come back to their resting position. Therefore, if receiver coils stop receiving the signal at a set time, different tissues will be at different stages of relaxation and produce different signal intensity resulting in an image with certain amount of tissue contrast. In MRI, a particular pulse sequence combined with a particular magnetic field gradient produces a certain tissue signal and contrast to highlight different tissues of the body. All MRI sequences use the same principle of T1 and T2 relaxation but with different pulse sequence and magnetic field gradients. As most diseases manifest by increasing the affected tissue water content, they tend to have different relaxation times to normal tissue and as a result different signal and imaging appearance.

5  Commonly used MRI sequences [6,9,12,13] 5.1  Spin echo (SE) sequences These were the first sequences discovered by Erwin Hahn in 1949 and are still widely used in modern day practice as baseline. A 90 degrees RF pulse known as the flip angle is sent to dephase protons which have achieved homogeneity with the magnetic field followed by a further 180 degrees RF pulse, which rephases protons and put them in the opposite direction of spinning. Protons then try to return to their resting position of homogeneity with the magnetic field and as they do that, an echo is given which is used to produce an image. Different tissues have different relaxation times and therefore the echo strength produced from them at a certain point varies. The time at which this echo is produced is known as echo time or TE. To repeat the whole process again, a further pulse is sent and the time at which it is sent is known as the repetition time or TR. As we are only interested in a particular tissue for a particular disease, as soon as that tissue has relaxed enough to produce echo (TE), further pulse can be sent at a particular time (TR). Each TE fills one line in the K-space. By altering the TE and TR, T1 weighted, T2 weighted or proton density (PD) sequences can be produced. T1 weighted imaging (T1WI) is acquired by using short TE and TR times. It is useful for illustrating the anatomy of tissues, for demonstrating fat, blood (hyper acute and chronic), and lesions containing proteinaceous material as well as in contrast imaging all of which give a high signal (bright). In comparison fluid is low signal (dark) on T1WI. T2 weighted imaging (T2WI) is acquired by using longer TE and TR times. It is useful for assessing fluid which is of high signal (bright) whereas fat is of intermediate signal. Fat signal can be suppressed (dark) by using fat suppression sequences that are discussed later in this chapter. Proton density weighted imaging (PDWI) is acquired by using short TE and long TR times. It therefore uses some features of T1WI and some of T2WI with the aim to assess density of protons in a particular tissue. PDWI was previously used mainly in brain imaging which has been overtaken by other sequences. However, PDWI is still frequently used in musculoskele­ tal imaging as it provides good distinction between fluid, hyaline cartilage, and fibrocartilage. In PDWI, fluid and fat have high signal (bright), muscle and hyaline cartilage intermediate signal (grey), and fibrocartilage has a low signal (dark).





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5.2  Gradient echo (GE) sequences Flip angle used in gradient echo (GE) sequences is between 10 degrees and 80 degrees rather than the 90 degrees used in SE sequences and without the 180 degrees refocusing pulse which allows for quick scan times. Larger flip angles are used to produce predominantly T1WI and smaller flip angles for predominantly T2WI. Images produced are susceptible to magnetic field inhomogeneities due to lack of the 180 degrees refocusing pulse leading to signal loss. However, this phenomenon can be utilized to diagnose degraded blood products which due to the presence of hemosiderin are susceptible to signal loss and areas of hemorrhage will appear dark on these sequences. Image sequence produced in this way is known as T2* weighted. However, any metal objects in the imaging plain may also lead to signal loss due to magnetic field susceptibility which is not desired.

5.3  Inversion recovery (IR) sequences In an inversion recover (IR) sequence, an initial 180 degrees RF pulse is applied to flip the longitudinal magnetization in to the opposite direction. As the longitudinal magnetization returns to its initial value, it passes through a point where signal from a particular tissue will be zero. If a 90 degrees RF pulse is applied at this particular point, the tissue with zero signal will not produce an echo and will be suppressed which is the main concept behind these sequences. During the early use of IR sequences between 1980 and 1985, excellent images with T1 weighted contrast were produced but could take up to 15–20 minutes to perform. However, since late 1990s, techniques have been developed where IR sequences can be combined with other faster sequences to achieve images quicker. This has helped to bring the imaging time for an IR sequence down to the order of 5–10 minutes. Some of the uses of IR sequences include suppressing fat using short tau inversion recovery (STIR) sequence to assess lesions which contain fat to differentiate them from blood and protein rich fluids (which also appear bright on T1 weighted images) and suppressing fluid using the fluid attenuated inversion recovery (FLAIR) sequence to assess tissue edema. Main advantages of these sequences include relative insensitivity to magnetic field inhomogeneity, less image compromise from metallic prosthesis due to relatively low susceptibility to metals and they can be performed in any magnetic field strength. Main limitation of a STIR sequence is the inability to use it with MRI contrast agents, which shorten T1 and thus the signal in tissues which take up contrast may be nullified. It also suppresses signal from any tissues which have same T1 value as fat [14].

5.4  Diffusion weighted imaging (DWI) This type of imaging utilizes the concept of whether there is random Brownian motion of water molecules in different tissues. Images produced are either DWI or apparent diffusion coefficient (ADC) weighted. Broadly speaking, the cellular membrane is responsible for maintaining a homogenous intracellular environment. Damage to this will lead to excessive amounts of extracellular water moving intracellularly which restricts movement of water molecules. Diffusion restriction will lead to high signal (bright) on DWI sequence and loss of



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signal (dark) on ADC sequence. The ADC map or images are needed to compare with DWI images to avoid the risk of wrong image interpretation from T2 shine through, which produces artefactual high signal on DWI images. Diffusion weighted imaging has achieved a major role in day to day clinical practice. Some of the clinical applications include stroke imaging, differentiation of abscess from necrotic tumors, differentiation of epidermoid cysts from arachnoid cysts, and the assessment of diffuse axonal injury.

6  Advances in general MRI [6,12,13,15] Even though Raymond Damadian produced the first ever-human MRI image of a thorax [16], initially MRI was mainly used for brain and spinal cord imaging with each scan taking around an hour to complete. There has been vast research in to identifying ways to improve MRI scans by devising particular sequences with the aim to reduce scan duration and to get more specific information on different body parts. This has reduced scan duration in some cases to a few minutes, for example, for brain imaging. With the advent of analogue to digital converters in MRI, whole body scans which are particularly useful for detecting skeletal metastases have become easier to perform without the loss of spatial resolution. Reduction in scan duration has not only increased patient throughput but also reduced the chance of image degradation from movement artefact due to patient discomfort from having to lie still for long periods of time during the scan. There are multiple advanced sequences now in clinical use by different manufacturers which are discussed further. It is worth noting that various manufacturers name their sequences differently even though the final image tissue contrast produced may be same. For example, a Steady State GE (discussed later in the chapter) sequence produced by Philips is named Fast Field Echo (FFE), by Siemens as Fast Imaging with Steady-state Precession (FISP), by GE as Multi-Planar Gradient Recalled (MPGR), and by Toshiba as Field Echo (FE).

6.1  Advanced spin echo (SE) sequences 6.1.1  Fast/Turbo SE (FSE/TSE) [17–20] For this sequence, multiple 180 degrees inversion pulses are applied following each 90 degrees excitation pulse producing multiple echoes for image formation. This is done by cancelling phase encoding after each echo and applying a different phase encoding gradient. In this way, an echo train is produced which is then used to fill the k-space faster and make image acquisition quicker. Multiple 180 degrees pulses help to correct for any magnetic field inhomogeneities. Uses for this sequence include investigating of brain diseases, meniscal tears in knee, liver lesions, etc. However, long echo train leads to reduced signal to noise ratio (SNR) particularly for echoes obtained toward the end along with variation in T2 weighting of the image due to longer TE. To help with this, centre of k-space which is responsible for image contrast is filled first with initial echoes which have desirable TE and also the echo train length is restricted to a reasonable number to give sufficient high spatial frequency information.





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6.1.2  Ultrafast SE [21–26] This is another variant of the Fast SE sequence which is also known as SS-FSE (single shot fast SE) or HASTE (Half-Fourier Acquisition Single-shot Turbo Spin Echo). In this technique, whole K-space is filled during a single 90 degrees RF pulse in comparison to multiple excitation RF pulses needed for FSE/TSE. There is no repetition of the RF pulse and therefore no TR which is considered to be infinite. As the time at which echoes are collected increases, images become very heavily T2 weighted. However, some prior inversion pulses can be employed to change image contrast appear closer to T1 weighted imaging to a certain degree. There may be only half of K-space image acquisition with the other half mirror imaged using K-space symmetry properties (e.g., HASTE sequence). Even though this sequence can reduce imaging times considerably, long echo train lengths lead to low SNR and reduction in spatial resolution. Due to fast acquisition, these images are less sensitive to magnetic field inhomogeneities. Any static fluid can be imaged exquisitely using this technique and therefore these sequences are commonly used for imaging biliary (magnetic resonance cholangiopancreatography— MRCP) and urogenital systems. Other uses include; brain and fetal imaging, scout images, head or body scanning in children or uncooperative patients, fetal imaging, abdominal, liver and chest imaging (due to their low sensitivity to movement), myelography, sialography, and non-contrast MR angiography.

6.2  Advanced gradient echo (GE) sequences [27–30] Further progress in GE sequences has been made in the form of spoiled and steady state GE sequences. In spoiled GE sequence, different flip angles are used to vary the T1 weighting of an image and different echo times are used to vary the T2* weighting. There is a further variation to this sequence in the form of ultrafast GE sequence. These sequences enable fast imaging using breath holding techniques, subtraction imaging using pre- and post-contrast images and in and out of phase imaging to elicit fat in certain tissues. In steady state GE sequence, some residual transverse magnetization is conserved which is then used to obtain signal and create contrast among different tissues as required and according to the type of sequence employed. Other variant of this include the double echo steady state (DESS) sequence and the constructive interference steady state (CISS) sequence. These advanced GE sequences have found uses in cardiac, abdominal, neurological including CSF-cisternography for assessing cranial nerves at skull base and fetal imaging. In addition, some of these sequences can be used for MR angiography and to evaluate cartilage and meniscal lesions.

6.3  Advanced inversion recovery sequence [31–33] There is a relatively newer inversion recovery sequence developed for the assessment of neurological diseases called the double inversion recovery or DIR. In this sequence, two inversion times are used which suppress both cerebrospinal fluid (CSF) and white matter signals highlighting any focal cortical dysplasia or demyelinating lesions such as those in multiple sclerosis.



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6.4  Advanced diffusion weighted sequence [34,35] A relatively newer MRI sequence utilizing diffusion weighting is diffusion tensor imaging (DTI) which is also known as MRI tractography. This type of imaging makes use of the phenomenon of anisotropy which means that water molecules diffuse more freely along nerve fibers rather than across them. Different nerve pathways in the brain are established using anisotropy and then color coded to further clarify their direction. In addition, the micro-structural information about biological tissue obtained from DTI can also be utilized in investigating other brain pathologies such as ischaemia, inflammatory illnesses, demyelinating illnesses, infection, and tumors as well as pre-surgical mapping of white matter pathways. Other use suggested is in musculoskeletal imaging where it can be employed for depiction of normal anatomy as well as to investigate possible muscle and nerve abnormalities. Limitation for this particular sequence is low SNR, which can be increased by increasing number of averages obtained but this also increases the overall scan duration.

6.5  Echo planar imaging (EPI) [36,37] EPI is a technique which produces multiple echoes from a single RF pulse by using rephasing gradients rather than the repeated 180 degrees pulses used in SE sequences. This particular feature is utilized in brain, cardiac, and abdomen (with breath hold technique) imaging. EPI has revolutionized MRI imaging as it reduces scanning duration significantly. It takes approximately 20 milliseconds to obtain one slice and therefore a whole MRI scan may be performed in less than 1 minute. EPI helps to obtain T2* images much more quickly and has made functional and perfusion imaging possible. Furthermore, EPI can be used with SE, GE, inversion recovery or diffusion weighted sequences.

6.6  Functional MRI (fMRI) [38–41] Main use of this sequence is to assess regions of the brain that control different activities. While a particular activity is performed repeatedly such as flicking the fingers, blood flow to certain parts of the brain controlling that movement increases. Therefore, the ratio of oxyhemoglobin and deoxyhemoglobin in that particular region is different from rest of the brain and that helps to identify which part of the brain is responsible for carrying out a certain function. T2* weighted images are used to obtain fMRI and the earlier-mentioned principle is called blood oxygenation level dependent (BOLD) contrast principle. fMRI can be used to differentiate normal brain tissue from diseased tissue, for example, tumors and to establish pre-operatively which brain region controls which function. Patients are therefore able to make better informed decisions knowing possible functional repercussions following brain surgery. It is also being researched for use in cognitive and neuropsychological studies, and to investigate renal illnesses.

6.7  Perfusion MRI [42–44] Using EPI principles, perfusion MRI can be performed which is mainly used to assess how well a tissue is perfused by blood. It is used in stroke cases to assess whether there is any ischemic penumbra which would mean that improving blood flow to the region may result 



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in salvage of some neuronal function. It is also utilized to assess the vascularity of tumors to establish if there is any high grade component in a diffuse low grade appearing glioma and for investigation of neurodegenerative conditions, for example, Alzheimer’s disease.

6.8  MR angiography (MRA) sequences [45] This type of imaging relies on whether protons in blood entering or leaving an image slice have been exposed to the RF pulse. Some of the MR angiography sequences used in clinical practice are described as: MRA sequences

Clinical applications and limitations

Time of flight (TOF)

To visualize fast flowing blood but needs long image acquisition time especially for 3-dimensional TOF images and unable to depict vessels in which flow has reversed, for example, vertebral artery flow reversal secondary to subclavian steal syndrome. To assess for flow velocity and direction as signal is not dependent on direction of blood flow but poor for visualization of turbulent flow. For quick assessment of thoracic or abdominal vasculature but there is loss of signal with fast flow. Uses gadolinium to increase vascular signal but has associated risks from venepuncture, allergic reaction, renal effects, and nephrogenic systemic fibrosis (NSF), which is specific to gadolinium contrast used in MRI.

Phase contrast Fresh blood imaging (FBI) Contrast-enhanced MRA

6.9  Cerebrospinal fluid (CSF) sensitive sequence [46] Time resolved 2-Dimensional phase contrast imaging with velocity encoding is the most widely used CSF sensitive sequence. Two RF pulses are applied in opposite direction, which cancel each other with no signal returned from stationery CSF but normal signal obtained from flowing CSF. As the name suggests, this type of sequencing is mainly used to assess CSF abnormalities related to aqueduct stenosis, Chiari malformation, and normal pressure hydrocephalus.

6.10  Susceptibility weighted imaging (SWI) [47,48] Principle for this type of GE sequence is that tissues with paramagnetic (deoxyhemoglobin, hemosiderin, and ferritin), diamagnetic (bone and dystrophic calcifications), or ferromagnetic (metallic foreign bodies and implants) properties interact with local magnetic field resulting in their distortion and loss of signal. Most commonly SWI is used to localize small quantities of blood, calcium, or iron in tissues. Particular application is in neurological imaging for diagnosing multiple sclerosis, cerebrovascular accidents, tumors, trauma, and vascular malformations.

6.11  MR spectroscopy [49–51] In this sequence, no actual images are produced as such but instead different metabolites produced by tissues during both physiological and pathological processes are charted 

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on a graph to assess the area of interest. Two main types of sequences in clinical practice which use this principle are stimulated echo acquisition mode (STEAM) and point resolved spectroscopy (PRESS). There are a number of metabolites which can be assessed depending on the body organ being investigated, some of which include choline, creatine, N-Acetyl-Aspartate, acetate, and lactate. Clinical applications include assessment of tumors (e.g., brain, breast, and prostate), ischaemia, infections, inflammations, and metabolic pathologies mainly of brain. MR Spectroscopy needs a very homogenous and strong magnetic field to distinguish between different metabolite peaks. Stronger the magnetic field, better the chance of detecting a small metabolite change. If there were magnetic field inhomogeneities, this would result in spreading of the peaks which would skew the data to be analyzed.

6.12  Hybrid sequences These are a mixture of fast SE and GE sequences. Specific absorption rate (SAR) is the energy deposited in to tissues by RF pulses during an MRI scan and measured in watts per kilogram. Even though this is not significant in most sequences, the main advantage of hybrid sequences is reducing the quantity of this SAR energy deposited in tissues which may be useful in a certain group of patients such as those with impaired thermoregulatory mechanism (e.g., fever, obesity, old age, pregnancy, and cardiac impairment). Therefore, this sequence can be particularly useful in higher magnetic field strength scanners without too high SAR for patients.

7  Advances in musculoskeletal (MSK) MRI 7.1  Imaging joints with prosthesis There is substantial ongoing work to advance musculoskeletal (MSK) imaging further using MRI scans as discussed earlier. A major challenge is the imaging of joints post metallic prostheses insertion. These can be imaged using X-rays but the images are not detailed enough due to their 2-dimensional nature. A CT scan may provide more detailed 3-dimensional images but streak artefact from metal obscures the image areas and does not show soft tissue pathologies. Although MRI scans are good in investigating soft tissues due to the excellent contrast between them, they are still not immune to the artefacts from metallic prostheses which are mostly due to susceptibility from magnetic field inhomogeneity adjacent to the prostheses. There is a technique of pre-polarized MRI which uses two magnets in the scanner consisting of homogenous low-field readout magnet and inhomogenous high-field polarizing magnet [52,53]. However, the use of this technique clinically is limited due to lack of enough appropriate specialized scanning hardware [54]. Magnetic field inhomogeneity also leads to inadequate fat suppression adjacent to the prostheses. A new sequence developed to overcome this problem is the iterative decomposition of water and fat with echo asymmetry and least-squares estimation (IDEAL) fast SE sequence which provides better homogenous fat suppression near metallic implants [55,56].





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7.2  MR arthrography [57,58] MR arthrography (direct or indirect) is a well-established technique to evaluate joint illnesses. In direct MR arthrography, gadolinium contrast is injected in to the joint under fluoroscopic guidance followed by an MRI scan. The injected contrast causes distension of the joint to evaluate any minor changes; however, this approach is locally invasive with its inherent associated risks and requires skilled person and fluoroscopy for contrast injection. The other method gaining popularity is the indirect MR arthrography which still requires gadolinium contrast but this is injected peripherally intravenously. As there will be wider distribution of contrast in the body, a higher amount is required. The blood supply and therefore contrast concentration in the area of interest may be increased by exercising that particular region. However, this may be limited by the underlying condition such as infection, inflammation, joint or tendon injury affecting the area of interest and hence limiting the ability to localized exercise.

7.3  Dynamic MSK MRI MSK imaging is dependent on joint position at the time of imaging such as the position of muscles relative to the bony structures and the physical strains on the joint (e.g., differences between flexion and extension, lying down and standing). These factors could result in misdiagnosis of a condition simply due to alteration in the appearance of abnormal structures during imaging [59]. Kinematic MR imaging is a type of dynamic imaging gaining popularity where cine images are created during joint movements that reproduce the patient’s symptoms. Relative position of bones, soft tissues, and joints is then assessed at the point of symptom reporting to better understand the underlying pathology and plan treatment. However, the main limitation of this technique is with imaging of joints which may be moving much faster than it is possible to obtain the MRI scan. Another improvement in MSK imaging, particularly spine is the availability of MRI scanners which can tilt and it is possible to scan patients upright. This enables diagnosis of pathologies which may be occult in recumbent position [60,61].

7.4  Differentiating benign from malignant bone tumors [62,63] MR Spectroscopy is beginning to find its use in MSK imaging. Even though MRI scans have high sensitivity to diagnose bone malignancy, they have poor specificity necessitating biopsies or other investigations. With the use of MR Spectroscopy, a diagnosis of benign or malignant bone lesion can be reached with more certainty. Choline is the main marker used in MSK imaging which is involved in cell membrane metabolism. As the cell turnover increases in malignant lesions, the amount of choline in affected tissues also increase and this can be measured with MR Spectroscopy. It is however worth noting that choline is not a specific metabolite for MSK malignancy, and increased choline levels may be seen in other malignancies, demyelinating illnesses, and gliosis.



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7.5  Investigating joint soft tissues Conventional T2 weighted imaging produces little to low signal from joint soft tissues due to their short relaxation times compared to the long echo time (TE). This fact can be used with ultra-short TE to evaluate periosteum, cortical bone, menisci, tendons, and ligaments as these structures produce signal whereas the tissues with long T2 relaxation time do not and appear dark [64–66]. Diffusion tensor imaging (DTI) along with tractography are relatively promising tools which provide information about direction and orientation of complex muscle fibers. They can also help to investigate muscle, nerves, and articular cartilage diseases at early stage by detecting changes in the collagen fibers arrangement [35,53,67]. T2 mapping, T1rho, and sodium imaging (using sodium-23 [23Na]) are some of the rapidly evolving compositional MRI techniques which allow quantitative evaluation of cartilage matrix to help establish early diagnosis of osteoarthritis [68,69].

7.6  Utilising 3-dimensional isotropic voxels in MSK imaging Previously MSK MRI scans used 2-dimensional multislice fast SE sequences which provided anisotropic voxels. These produced thick slices and slice gaps with partial volume artefact and the inability to reformat them. This in turn meant longer scan durations if tissues needed to be viewed in different planes. The relatively recent introduction of 3D isotropic voxels has helped MSK imaging greatly in that the partial volume artefact is reduced by thin continuous slices which can be obtained comparatively quickly and can be reformatted later in any plane for tissue assessment [53].

8  Impact of MRI field strength on imaging 8.1  Understanding MRI scanner’s magnetic field strength Tesla (T) is the universally accepted unit used to demonstrate the strength of a magnetic field. Another unit used commonly to describe a magnetic field strength which is smaller compared to a Tesla is Gauss with 1 T equalling 10,000 gauss. For example, magnetic field strength of earth at the surface is 0.25–0.60 gauss and at its core it is measured to be approximately 25 gauss, a typical refrigerator magnet has an approximate magnetic field strength of 50–100 gauss and that of a junkyard electromagnet to lift cars is approximately 10,000 gauss (1 T). MRI scanners used in medical practice have a magnetic field strength ranging from 0.2 to 3.0 T however in human research, MRI scanners with magnetic field strength as high as 9.4 T have been used [8,70]. Small animal imagers exist with magnetic field strengths of up to 17 T [71]. Largest pulsed field created in a lab which destroyed the magnetic equipment but not the lab was 730 T [12].

8.2  Comparing 1.5 T with 3.0 T magnetic fields in MRI scanners [72–81] High magnetic field strength MRI scanners (3.0 T) are finding their way in clinical practice. In 2011, 72% of MRI scanners in Europe and 76% in the United States were 1.5 T in strength





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whereas 3.0 T scanners comprised 11% in Europe and 10% in the United States. In the same year, 24% of new MRI scanners sold in Europe were of 3.0 T strength. Although only a few 3.0 T scanners are currently in medical practice in the United Kingdom, following review is intended to highlight their benefits and limitations as they are possibly going to increase in number in the future. As explained earlier, SNR increases with higher magnetic field strengths. This particular trait of high magnetic field strength MRI scanners, for example, 3.0 T can be used to improve image quality or can be used to reduce scan time as the SNR same as from a 1.5 T can be achieved in shorter time. Even though 3.0 T scanners deposit higher energy in tissues (SAR) from stronger RF pulses when compared to a relatively lower magnetic field strength scanner, for example, 1.5 T, the reduction in scan time largely cancels this effect out. Parallel imaging is a technique employed to reduce scan duration as the number of RF pulses needed to form an image is reduced. This may lead to loss of signal which can be compensated with the higher SNR from 3.0 T scanners. Multiple channels are used in parallel imaging to collect data simultaneously reducing scan durations, for example, technically an 8-channel phased array coil is able to collect all data required for image formation 8 times quicker, however, in practice the maximum number of RF channels is limited due to practical considerations. 3.0 T scanners may require less contrast due to their greater sensitivity to gadolinium, better quantify metabolites measured in MR Spectroscopy by detecting even small changes, increased contrast to noise ratio (CNR) is used in fMRI using the BOLD imaging sequence and better fat saturation can be achieved. Imaging sensitivity for white matter tract pathologies such as demyelination in multiple sclerosis is improved due to better tissue visualization from higher SNR. In prostate imaging, 1.5 T scanners need to use a surface coil on prostate inserted endorectally to attain adequate tissue detail which can be uncomfortable for the patient. This can be avoided in 3.0 T scanners due to higher signal production without the need for a surface coil. Breast cancer lesion detection and characterization is also better with 3.0 T scanners due to higher SNR and better image detail. Cardiac imaging is not so good with 3.0 T scanners due to the dielectric artefacts which are more pronounced on higher field strengths. However, higher spatial and temporal resolution from increased SNR and decreased imaging time respectively can help with cardiac imaging. As the magnetic field strength increases, tissues take longer for T1 relaxation and conventional T1 SE sequencing may not be reproducible. However, this principle is employed in MR angiography where an increased T1 relaxation time of solid tissue on a background of relatively constant T1 relaxation of blood gives high quality images due to improvement in the blood signal against background solid tissue contrast. Signal from blood is also used to image vessels using an arterial spin labeling (ASL) technique. Accuracy of cerebral perfusion measurement increases with 3.0 T scanners due to increased number of sampling points. Susceptibility artefacts on 3.0 T scanners are also enhanced which can be used to detect even small amounts of blood degradation products in subtle hemorrhage. 3.0 T scanners may not be compatible with certain metallic implants which are only tested on 1.5 T scanners. There is a higher risk of skin irritation or burning if equipment malfunction were to happen. Any artefacts such as due to respiration, vascular pulsation, wrap around artefacts common to echo planar imaging are magnified in higher magnetic fields. However, manufacturers of 3.0 T scanners are continuously developing special imaging sequences with different parameters to overcome the problems described earlier.



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We do not aim to infer that 1.5 T MRI scanners are inferior to 3.0 T scanners in anyway but to inform the reader that 3.0 T scanners are also available and may have a role in imaging of certain specific body parts for the reasons described earlier.

8.3  MRI field strengths used in musculoskeletal (MSK) imaging When an MRI scanner is built, it is the specific amount of electric current applied to coils, which determines its magnetic field strength which cannot be altered later to vary the magnetic field strength. Also, once installed, the coils cannot be altered within a scanner. Most MRI scanners used in the United Kingdom are 1.5 T in strength however, in the United States, MRI scanners with a magnetic field strength as low as 0.2 T are also used for axial MSK imaging. On the other hand, higher filed strength scanners of 3.0 T are being evaluated in MSK imaging for higher SNR which can reduce imaging times and increased spatial resolution [82]. One of the issues with higher strength magnetic fields is that metallic implants which are commonly used in orthopedic patients and have been tested to be safe in a lower strength magnetic field may not be safe in higher magnetic field strength due to greater pull from the magnet or not being tested in higher field strength. Multiple published studies comparing low/intermediate strength magnetic fields (0.2/1.0 T, respectively) to relatively higher strength magnetic fields (1.5 T) for MSK imaging have shown comparable sensitivity and specificity [83–85]. Benefits of using 0.2 T scanners are that they do not require as much space to install as the larger commonly used 1.5 T scanners, are less expensive, can be open space rather than like the tunnel shape in 1.5 T scanners and are less physically restrictive making them better suited for patients with claustrophobia. Furthermore, the resolution in 0.2 T scanners is comparable to investigate certain joint conditions such as rheumatoid arthritis in which small bone erosive changes are among the earliest signs [86,87]. One of the limitations of 0.2 T scanners used to be its inability to image larger joints such as the shoulder or hip due to the confined imaging space but this problem has been overcome by manufacturers building open space MRI scanners. The images may be noisy and unclear due to poor SNR which generally increases with higher magnetic field strengths. SNR can be increased by using longer imaging duration but this may risk image degradation from movement artefact. Also, low field strength scanners struggle to perform adequate fat suppression due to less spectral separation between water and fat which is very valuable and commonly employed technique in MSK imaging to diagnose bone edema from injury. 0.2 T scanners provide low diagnostic value from MR arthrography due to the suboptimal contrast enhancement detection of tissues with gadolinium. Overall, although high field strength magnets produce nicer images which are more aesthetic to the eye, low field strength magnets can possibly produce good quality images with comparable diagnostic accuracy and certain other benefits as discussed earlier.

9  Conclusions and future of MRI [88–113] With advanced imaging sequences being developed for different physiological and pathological conditions, MRI is likely to become mainstay of cross sectional imaging of patients similar to CT scans. It is not only clinical practice which stands to benefit from technological



References 135

advances in MRI but also biomedical and neurosciences research. Although much of MRI based imaging progress has previously been around brain imaging and neurosciences; structural, functional, perfusion, and spectroscopy sequences are being evaluated in whole body imaging including that for investigation of liver, inflammatory bowel disease, and genitourinary pathologies. With faster techniques and technological advances, whole body MRI scans for staging purposes for certain cancers to detect nodal and bone metastasis are now possible. MRI has also been evaluated to help with image guided procedures including but not limited to; ablation of uterine fibroids, multiple MSK procedures including palliation of bone metastasis and for treatment of brain, breast, cardiac, prostate, and gastrointestinal diseases. Functional MR scans are being routinely used in some neurosurgery procedures for brain mapping pre-operatively and also intra-operatively to better guide the surgeon in what to expect during surgery and to preserve important functional centers. With the MRI software becoming advanced, faster scans are possible which enable cardiac and respiratory system imaging which was previously hampered by movement artefacts. Software is now available which allows achievement of different soft tissue contrast ratios in brain from one image acquisition meaning that from a single MRI scan sequence T1, T2, PD, STIR, T1 FLAIR, T2 FLAIR, dual inversion recovery (IR), and phase sensitive IR weighting can be formulated afterward. Noise from the MRI scanners can be quite distressing for patients. Manufacturers are offering newer scanners with reductions in sound intensity and pressure to improve patient comfort. There are also now scanners available which provide an ambient experience through lighting, sound, and videos which help with patient’s claustrophobia and production of a movement free scan. Scanner sizes and power needed to run them are decreasing which means that they can be installed in smaller spaces and lower cost to run, respectively. The options for progressing in MRI techniques are vast and the authors concur that the future for this technology looks promising. We believe that we are only at the beginning of discovering the true potentials, benefits and uses of MRI and are currently far from reaching its maximum potential.

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C H A P T E R

10

Technological advances in breast implants Michalis Charalambousa, Raouf Daoudb, Isabella Karatb a

The Parapet Breast Unit, King Edward VII Hospital, Windsor, United Kingdom; b Breast Unit, Frimley Park Hospital, Firmley, United Kingdom

A breast implant is an internal prosthesis, which is most commonly used either in breast reconstruction following a mastectomy, or aesthetically in breast augmentation. Breast implants may also be used to correct congenital deformities and defects of the chest wall. The first silicone breast implant was developed in 1961 by American plastic surgeons Thomas Cronin and Frank Gerow, and the world’s first augmentation mammoplasty was performed the subsequent year. In 1964, the French company “Laboratories Arion” manufactured the world’s first saline-filled breast implant.

1  Types of breast implants At present, there are two main types of breast implants: silicone and saline filled breast implants. 1. Silicone breast implants Since the early 1960s when the first silicone breast implant was developed, there have been five generations of silicone implants, with each generation defined by specific manufacturing and technical developments. The first generation silicone implants were developed in the 1960s and were produced until the early 1970s. These included the original Cronin-Gerow breast implant which had a teardrop shape, and consisted of a silicone, rubber envelop sac, and was filled with viscous silicone gel. The main complications of these early implants included capsular contracture and implant rupture. The second generation of silicone breast implants was also developed by Cronin and was in use between 1972 and 1986. This new breast implant featured three main Advances in Medical and Surgical Engineering. http://dx.doi.org/10.1016/B978-0-12-819712-7.00010-3 Copyright © 2020 Elsevier Inc. All rights reserved.

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technological advances. The first development included a thinner-gauge device-shell, and a filler gel of low-cohesion silicone, which improved the functionality and verisimilitude (size, look, and feel) of the implant. However, these implants suffered high rates of rupture and silicone escape (also known as “silicone gel bleed,” where filler silicone gel escapes through an intact implant shell). The second technological development was a special coating of the implant made of polyurethane foam, which reduces the rate of capsular contracture, through an inflammatory reaction that inhibits the development of a fibrous collagen capsule around the breast implant. However, the polyurethane foam coating had a carcinogenic by-product, namely 2,4-toluenediamine (TDA), and was briefly discontinued due to its potential health risk [1]. Subsequently, the United States Food and Drug Administration (FDA) reviewed the available medical evidence, and concluded that the potential risk of TDA-induced breast cancer is extremely low [2]. At present, polyurethane-coated breast implants are in medical use in Europe and South America, but not in the United States. The third development was a double-lumen breast implant, composed of an inner silicone implant contained within an outer saline implant. The main advantages of this technological development included improved cosmetic outcomes, and the possibility of adjusting the volume of the implant postoperatively. However, this more complex breast implant design suffered an increased rate of device-failure compared to single-lumen implants. The third and fourth generations of implants were developed in the 1980s and early 1990s, and included sequential advances in implant manufacturing technology. The filler gel of these implants was thicker, enhancing cohesion, and the implant shell was coated with elastomer, reducing silicone escape. Moreover, new breast implant models were developed, including anatomical models which provide a more natural breast look, and round models, with both models having either a smooth or a textured surface. Textured implants have a uniformly textured surface that reduces the risk of implant rotation. The fifth generation of breast implants has been in development since the mid-1990s, and includes implants made of a high-strength, highly cohesive silicone gel that reduces the risk of silicone escape. They are commonly known as “gummy bear” implants, as their consistency is similar to gummy bear candies. Due to the consistency of their filler silicone gel, they can maintain their shape and reduce the risk of rippling, although they have a slightly firmer feel. 2. Saline breast implants Saline breast implants have been in development since 1964. They are made of a thicker, silicone elastomer shell, which is vulcanized at room temperature, and are filled with sterile normal saline solution (0.9% w/v NaCl). Saline breast implants can be used for aesthetic breast augmentation, as they are inserted deflated, requiring a smaller incision, and are filled with saline once in the surgically created implant pocket. They are also commonly used in breast reconstruction following a mastectomy, where the overlying breast skin envelop requires gradual expansion (hence the term “tissue expander”). This is achieved by the gradual insertion of sterile saline solution into the implant over several weeks post-operatively, via an attached port that is placed subcutaneously during the operation, so that it can be readily accessible for saline





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insertion. Some of the latest models of tissue expanders have their port embedded within the implant, and incorporate the use of a magnet to correctly locate the port for saline insertion.

Photo showing a tissue expander, a type of adjustable saline breast implant, with its associated port and connection parts, while being prepared to be used for implant based reconstruction following a skin sparing mastectomy.

Compared to silicone gel breast implants, saline implants can produce cosmetically acceptable results, with increased breast size, smoother hemisphere contour, and a more realistic breast texture. However, they are more likely to cause aesthetic problems, especially in women with little breast tissue, and following a mastectomy. These include rippling and wrinkling of the skin of the breast envelop, as well as accelerated stretch of the lower breast pole. Saline breast implants are also generally more noticeable to the eye and to the touch, resulting in an overall inferior cosmetic result, compared to silicone implants [3].

1.1  Complications of breast implants There are potential complications associated with insertion of breast implants. These include general complications associated with surgery, such as wound problems, seroma, hematoma, adverse effects to anesthesia, atelectasis, and thromboembolic disease; as well as complications specific to implant insertion, such as chronic breast pain, paresthesia of the overlying skin, visible skin wrinkling, breast asymmetry, capsular contracture, implant rupture, silicone escape, infection, rotation of the implant, and breast contour changes such as symmastia (“kissing” of the two breasts at the midline). Moreover, complications and reoperations related to breast implant surgery may cause unfavorable scarring in up to 7% of patients [4]. 1. Implant rupture Breast implants have a limited life span, and although they may retain their mechanical integrity for decades after insertion, they may occasionally undergo rupture. The possible mechanisms of implant rupture include local trauma, damage during insertion, and chemical degradation of the implant shell. 

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Rupture of saline implants results to quick leakage of saline, causing deflation of the implant, which can be easily explanted. Rupture of silicone implants does not usually result in deflation of the implant, but the filled silicone gel can escape either just inside the implant capsule (intracapsular rupture) or migrate outside of it (extracapsular rupture), typically causing locoregional symptoms such as silicone granulomas and axillary lymphadenopathy. The rate of implant rupture for saline implants is 3%–5% at 3 years and 7%–10% at 10 years [5], while for silicone implants is 1% at 6 years [6]. 2. Capsular contracture The human body’s normal immune response to the presence of a foreign body such as a breast implant, is to try and isolate it by encapsulating it with tightly woven collagen fibers. Capsular contracture is usually a painful complication, and occurs when the collagen-fiber capsule thickens, resulting in compression and distortion of the implant and/or the breast tissue. The cause of capsular contracture is unclear, but the common risk factors include bacterial contamination, implant rupture, silicone escape, post-operative hematoma, as well as radiotherapy to the chest wall following a mastectomy in a breast cancer patient. There are several surgical precautionary measures that have been shown to reduce the incidence of capsular contracture, including sub-pectoral placement of the implant, use of textured implants, minimal handling and skin contact of the implant during implantation, and irrigation of the implant pocket with antibiotic solution. Capsular contractures can be quite painful, requiring open revisional surgery, including capsulotomy (surgical release of capsule), capsulectomy (excision of capsule), and removal or exchange of implant. 3. Revisional surgery Breast implants have a limited life span and may undergo complications, necessitating revisional surgery. The most common indications include unsatisfactory esthetic outcome, implant rupture, and capsular contraction. 4. Systemic illness There have been claims that breast implants are associated with an increased risk of several systemic illnesses, including connective tissue disorders, rheumatism, and autoimmune diseases. However, over the last 20 years, several independent systematic reviews have investigated these claims and all concluded that there is not any scientific evidence to support any of these claims [7,8]. 5. Anaplastic large cell lymphoma Textured breast implants are associated with a rare form of lymphoma, known as breast-implant-associated anaplastic large cell lymphoma (BIA-ALCL). The current lifetime risk of BIA-ALCL is unknown, but it was originally estimated to be in the region between 1:70,000 and 1:500,000 women with breast implants [9]. However, latest figures have estimated this risk to be at 1:1,000 to 1:10,000 for textured implants [10]. ALCL usually presents many years after implant insertion, with breast swelling or new onset seroma. It is typically diagnosed by cytological studies and the “CD30” marker. It normally carries a good prognosis and does not require systemic therapy.





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Treatment involves explantation of the implant and capsulectomy, although occasionally chemotherapy may also be required.

1.2  Breast reconstruction and acellular dermal matrix (ADM) In modern implant-based breast reconstruction, it has become common to use acellular dermal matrices (ADMs) in combination with a silicone implant or a saline tissue expander. ADMs used in breast reconstruction are derived from the dermis of either donated cadaveric human skin (AlloDerm), porcine skin (Strattice), or fetal bovine skin (SurgiMend). ADMs are soft connective tissue allografts generated by a decellularization process, which preserves the intact extracellular dermal matrix, which upon implantation, acts as a scaffold for donorside cells to facilitate subsequent incorporation and revascularization.

Photo showing SurgiMend, an ADM which is commonly used in implant based breast reconstruction.

In implant based breast reconstruction, ADMs are used mainly lower pole coverage and the shaping of the implant, resulting in a better esthetic outcome with better inframammary fold definition, greater breast projection, and a more natural look. Concerns have been raised over the rate of complications following implant based reconstructions using an ADM, with initial studies showing higher rates of postoperative seroma, infection and skin necrosis with the use of ADM. However, more recent studies have shown that by careful patient selection, and improving surgical technique and perioperative management, implant based breast reconstruction with an ADM, can result in improved esthetic outcomes without any difference in complication rates [11].



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Photo showing a tissue expander covered by ADM, ready to be used for implant based reconstruction.

1.3  Future developments in breast implants Silicone breast implant models are in continuous development, aiming for better esthetic results, and fewer complications. In particular, new designs need to be more resistant to infections, have a very low long-term rupture rate, minimize the development of capsular contracture even after radiotherapy, and when used in breast reconstruction, result in a breast that looks, moves, and feels as natural as possible. Currently, there are various technological advances in development, promising to revolutionize breast implant reconstruction. A new breast implant model (B-LITE) employs the use of microsphere technology to disperse tiny, inert hollow borosilicate beads throughout the filler silicone, resulting in a lighter implant for a given volume. Researchers are also developing a breast implant using 3-dimensional printing (MATTISSE), which is made of a bio-absorbable shell, filled with the patient’s own adipose cells, derived from abdominal fat. Upon implantation, the shell of this breast implant will be absorbed by the body, leaving only the implanted fat tissue in the breast.

References [1] National Toxicology Program. Bioassay of 2,4-diaminotoluene for possible carcinogenicity. Natl Cancer Inst Carcinog Tech Rep Ser 1979;162:1–139. [2] U.S. Food and Drug Administration. FDA breast implant consumer handbook 2004: timeline of breast implant activities. Silver Spring, MD: U.S. Food and Drug Administration; 2013. [retrieved 10 March 2019]. Available from: http://www.fda.gov/MedicalDevices/ProductsandMedicalProcedures/ImplantsandProsthetics/ BreastImplants/ucm064242.htm.



References 147

[3] Aesthet Surg J. 2010 Jul-Aug;30(4):557-70. Breast implants: saline or silicone? Spear SL, Jespersen MR. [4] Important Information for Women About Breast Augmentation with INAMED Silicone-Filled Breast Implants (PDF). Retrieved 10 March 2019. [5] Walker PS, Walls B, Murphy DK. Natrelle saline-filled breast implants: a prospective 10-year study. Aesthet Surg J 2009;29(1):19–25. [6] Hedén P, Boné B, Murphy DK, Slicton A, Walker PS. Style 410 cohesive silicone breast implants: safety and effectiveness at 5 to 9 years after implantation. Plast Reconstr Surg 2006;118(6):1281–7. [7] Tugwell P, Wells G, Peterson J, Welch V, Page J, Davison C, et al. Do silicone breast implants cause rheumatologic disorders? A systematic review for a court-appointed national science panel. Arthritis Rheum 2001;44(11): 2477–2484. [8] Balk EM, Earley A, Avendano EA, Raman G. Long-term health outcomes in women with silicone gel breast implants: a systematic review. Ann Intern Med. 2016;164(3):164–75. doi: 10.7326/M15-1169. [9] Implant-associated ALCL Facts | The MD Anderson Foundation“. www.mdanderson.org. Retrieved 10 March 2019 [10] Clemens M. Breast implant associated anaplastic large cell lymphoma (BIA-ALCL) Archived 2017-03-26 at the Wayback Machine; 2017. [11] Vardanian AJ, Clayton JL, Roostaeian J, et al. Comparison of implant-based immediate breast reconstruction with and without acellular dermal matrix. Plast Reconstr Surg 2011;128:403e–10e.



C H A P T E R

11

Importance of biomaterials in biomedical engineering Sakib S. Yousafa, Chahinez Houacinea, Iftikhar Khanb, Waqar Ahmedc, Mark J. Jacksond a

School of Pharmacy and Biomedical Sciences, University of Central Lancashire, Preston, United Kingdom; bSchool of Pharmacy and Biomolecular Sciences, Liverpool John Moores University, Liverpool, United Kingdom; c School of Mathematics and Physics & Lincoln School of Medicine, College of Science, University of Lincoln, Lincoln, United Kingdom; d Kansas State University, Salina, KS, United States

1 Introduction While there are many advances in the field of bioengineering, these are intimately dependent upon the availability of amenable materials, that is, biomaterials; without which bioengineering would be significantly limited. Biomaterials may be described as “Any substance or combination of substances, other than drugs, synthetic or natural in origin, which can be used for any period of time; augmenting or replacing partially or totally any tissue, organ or function of the body, in order to maintain or improve the quality of life of the individual” [1]. In response to increasing demands within the pharmaceutical and medical industries, the value of the biomaterial field is set to rise, at a compound annual growth rate of 6/7% between 2014 and 2019, yielding a $109.5 billion global market in 2019 [2]. The usage of biomaterials is certainly not a new concept, with the ancient Egyptians and Romans using gold and ivory in the treatment of cranial defects, and in the 1900s primitive implementations such as the usage of placenta in World War II [1]. Though metals and ceramics may also be identified as biomaterials, there is a dedicated body of literature surrounding synthetic and naturally derived polymers and their usage as biomaterials. It is structural similarities between the polymers present in tissues such as polypeptides with synthetic polymers which may account for their wide use as biomaterials, being related more closely, when compared to metal or ceramic biomaterials [3]. Advances in Medical and Surgical Engineering. http://dx.doi.org/10.1016/B978-0-12-819712-7.00011-5 Copyright © 2020 Elsevier Inc. All rights reserved.

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Despite there being many synthetic polymers utilized as biomaterials, there are many naturally derived polymers, which are also utilized. Unlike many synthetic alternatives, naturally derived polymers are able to facilitate cellular processes (i.e., adhesion, migration, proliferation, and differentiation) indicating potential for usage as part of drug delivery systems and medical devices [4]. Meeting the physical and cellular requirements of damaged tissues is a particular driving force for the development and amelioration of biomaterials [5-9]. This in turn is responsible for an elevated demand on the properties of biomaterials, that is, biodegradability, cost-effectiveness, non-antigenicity, and biocompatibility. The ability to ameliorate the aforementioned properties has allowed for the divergence of biomaterial utilization, that is, films, foams, gels, implants scaffolds, and microparticles. However, the application of biomaterials made from synthetic polymers is limited by issues surrounding biocompatibility; acting as propellant for research pertaining to biomaterials formed from naturally occurring polymers. This book chapter focus upon the applications of naturally occurring polymers (i.e., chitosan, collagen, hyaluronic acid, and silk fibroin) as biomaterials utilized in novel research.

2 Chitosan 2.1 Introduction Chitosan is a mucopolysaccharide present in the exoskeleton of crustaceans, mollusks, and insects; as well as the cells walls of fungal cells. It is derived from chitin, prepared via deactylation and depolymerization of chitin in solid form under alkaline conditions or through enzymatic hydrolysis in the presence of chitin deacetylase. The molecule is recognized for its biocompatibility, biodegradability, and low toxicity; consequently it is considered an adaptable polymer, which may be utilized in a variety of formulations, including: gels, films, nano/micro-particles and beads, etc. Moreover, it is considered to possess low immunogenicity and have the following biological properties which are associated with the primary amine functional group possessed by the molecule: low immunogenicity, high biocompatibility and biodegradability. The biological properties of chitosan include; mucco-adhesion, controlled drug delivery, in situ gelation, permeation enhancement, colon targeting, and efflux pump inhibition. It is further abundant on a global scale, holding second place to cellulose with respect to production through biosynthesis [10-16]. Optical examination of chitosan conducted by Firdous et al. [17] also acknowledged it as a strong biopolymer, capable of forming dense and biocompatible materials suitable for bioactive coatings applied to orthopedic implants. As a consequence of the materials structural similarity to that of corticol bone, it has been proposed as biocompatible material suitable for use in corticol bone cartilage of humans [17]. Being cationic in nature the versatility of chitosan as a biomaterial may be associated with active amino groups (i.e., an amino polysaccharide) on the polymer which potentially act as a site for a multitude of molecular group attachments. Essentially, it is a co-polymer comprised of N-acetyl-D-glucosamine and D-glucosamine units with one amine (NH2) and two hydroxyl (OH) groups (Fig. 11.1) [18,19]. The polymer also varies substantially in molecular weight, ranging from 300 to >1000 Kd. Deacetylation also varies from 30% to 95% with solubility in water increasing when deacetylation reaches 50% under acidic conditions (i.e., amino





2 Chitosan

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FIGURE 11.1  Molecular structures of cellulose, chitin, and chitosan.

groups protonate, solubilizing the molecules). Solubilization in acidic conditions results in the D-glucosmaine unit becoming positively charged, this effect makes the polymer molecule one of the very few pseudo natural cationic polymers present [15].

2.2  Dressing fabrication As naturally derived polymer chitosan has significant applications in the fabrications as dressings. For example, chitosan has been used in synergy with regenerative cellulose, silver nanoparticles, and the broad-spectrum aminoglycoside gentamicin, in the preparation of wound dressings, in recent research conducted by Ahamed et al. [20]. Upon utilization in rat subjects when compared to an untreated control, faster wound healing was observable when chitosan dressings were utilized. Notably, complete wound closure was observed within 25 days of chitosan dressing treated wounds, when compared to control group wounds, which took in excess of 30 days to reach wound closure. However, when compared to treatment with the antibiotic ointment soframycin, or dressings which did not contain gentamicin, a similar degree of wound closure was observed, indicating comparable efficacy. This suggested that gentamicin did not play as strong a role in the wound closure as the other dressing constituents [20] (Fig. 11.2).

FIGURE 11.2  Percentage contraction of wounds in rats treated with standard Soframycin, RC-Ch-Ag and RCCh-Ag-G. Source: Permission requested online Elsevier.



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Chitosan is also an example of a biopolymer which is also useful as a hemostastic agent (i.e., brings about the cessation of bleeding and restricts blood flow within compromised blood vessels). It is thus of key importance in wound management and hence utilized in wound dressings. At present, chitosan originated haemostatic agents have been utilized in post-extraction bleeding control in patients. Moreover, when combined carrageenan, the hemostatic capability of the polymer was enhanced, attributed to carboxymethylation and crosslinking [21]. Chitosan is also believed to facilitate wound healing via its own depolymerization, which releases the molecule N-acetyl-β-d-glucosamine resulting in fibroblast proliferation. The polymers monomers enable wound healing by sequential deposition of collagen, in turn stimulating and increasing the degree of hyaluronic synthesis at the wound site. Additionally, the polymer supplies a cellulosic matrix for regeneration of skin tissues and is also involved in the activation of macrophages which impede abnormal growth activity, enabling acceleration of full thickness wound healing, precluding scarring [22]. In the form of burn wounds dressings (hydrogels) chitosan has been combined with marine peptides extracted from tilapia [23]. When examined by Ouyang et al. [23] peptides alone were found to elicit a low antibacterial effect in comparison to when combined with chitosan in the form of a hydrogel, where a significant antibacterial effect was observed. The developed formulation further promoted cell proliferation and migration, when comapred to chitosan or peptides alone. Wounds healed to completion within 21 days, total regeneration of the epithelium was observed by 14 days and collagen fibre deposition within 7 days. In summary, chitosan marine peptide hydrogels were proposed to facilitate cell migration and promote skin tissue regenration in wound healing [23]. A number of biomaterials have also been investigated for their potential as scaffolds (i.e., man-made structured on which cells are seeded). These are required to possess the capacity of supporting the fabrication of 3D tissues, as well as facilitating restoration of organ functionality. Chitosan’s ability to form interconnected porous structures allows for cell seeding into the structure as well as the required degree of nutrients was noted in recent research [24]. As chitosan fabricated structures are structurally similar to glycosaminoglycans which are found within the extracellular matrix; and possess the ability to form scaffolds with comparative mechanical and morphological properties to the protein collagen, they are deemed an appropriate alternative in tissue engineering applications [25-27]. Chitosan’s ability to facilitate wound healing in the form of dressings has been demonstrated by the aforementioned pieces of research; impacting on various cellular processes and offering similarity to compounds which are able to enhance wound healing. The polymers ability impact wound healing is not restricted to the fabricated form of dressings and may also be formulated as hydrogels.

2.3  Hydrogel applications Hydrogels are constituted of three-dimensional polymer networks (utilizing polymers including chitosan), containing copious quantities of water. They typically have a soft and rubbery consistency; in addition to having low interfacial tension when in contact with biological fluids native to human tissues [28]. The elevated water content provides the basis of hydrogel biocompatibility and high viscoelasticity which in turn restricts surrounding





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tissue damage when implanted [18]. Chitosan’s biocompatibility is however limited by its low mechanical strength and as a consequence its application as a hydrogel, in spite of this, there are numerous pieces of research relating to its use as a hydrogel [29]. Additionally, chitosan purification has been suggested as a potential solution to inadequate mechanical strength elicited by chitosan based hydrogels [30]. In recent research Bi et al. [30] developed an enhanced dissolution method for chitosan purification involving the usage of freezethawing and sodium hydroxide. The developed formulation exhibited desirable swelling characteristics with a uniform polymeric structure. Moreover, the formed hydrogel was highly biocompatible and possessed high mechanical strength which was associated with the formation of woven networks which were highly uniform/homogenous. There are a plethora of chitosan-based hydrogels which are broadly recognized, including: physically associated chitosan gels, cross-linked by co-ordination complex and chemically cross-linked chitosan hydrogels [31]. The manufacture of physical hydrogels relies upon reversible non-covalent interactions/exchanges among polymeric chains (e.g., electrostatic interactions/hydrogen bonding) [32]. Furthermore, a number of factors may influence these interactions, such as: pH, concentrations and temperature; these may detrimentally influence stability (i.e. reversible gellation). Thus, amelioration of the abovementioned parameters may bring about an increase or reduction in the number of interactions between the polymers and as a consequence swelling variation of the hydrogels (i.e., a reduction in the number interactions forming a softer gel and higher number of interactions resulting in the converse) [31]. Contrastingly, chemically cross-linked chitosan hydrogels are developed via covalent bonding between the polymeric chains. The developed hydrogels intrinsically offer greater stability when compared to the physically associated chitosan gels, as gelation is irreversible in these. A drawback of this approach is the necessitation of chemical amelioration of chitosan’s primary structure, which provides the potential alter the intrinsic characteristics of the molecule and generate toxic remnant reactants or catalyst traces [31]. Despite poor mechanical properties offered by chitosan hydrogels the variety in form of chitosan hydrogels (i.e., physical and chemically crosslinked) offer scope for further research in their development and use in wound management. Whilst gels and dressings of chitosan are being purposed for this use, chitosan has wider applications in the area of wound management.

2.4  Wound management Biomaterials are also considered as a potential approach in reducing infection risks, for example, when developed in the form of coatings of implants, medical devices, and as a component of wound dressings; potentially providing antimicrobial properties which inhibit microbial growth [33]. The polymer has been used successfully in the development of films used in antimicrobial dressings that enhance healing in skin ulcers. Research conducted by Escárcega-Galaz et al. [34] demonstrated developed chitosan films to be have a high retention capacity (i.e., ability to absorb fluids), resistance to breakage and high flexibility. In terms of antimicrobial properties, the developed chitosan films were active against a number of microorganisms; Klebsiella pneumoniae and Pseudomonas aeruginosa (Fig. 11.3); indicating the films suitability as a suitable option in the treatment of cutaneous ulcers [34].



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FIGURE 11.3  Antimicrobial sensitivity test of chitosan films with (A) Klebsiella pneumoniae and (B) Pseudomonas aeruginosa. Source: Permission requested online Elsevier.

Chitosan has also been used in the fabrication of membranes imbued with antibacterial agents (silver sulfadiazine (AgSD) or tetracycline hydrochloride (TH) when using a combination of chitosan floccule suspensions and solvent evaporation as a casting method [35]. Ma et al. [35] attributed the mechanism for membrane formation to bonding at a molecular level between the plasticizer and polymer. Plasticizer concentration was observed to notably enhance swelling rate, enhance tensile strength, water vapor permeability, and membrane wettability. The fabricated membranes exhibited good stability following enzyme degradation, in addition to crucially offering sustained release of both tetracycline and silver sulfadiazine membranes with notable activity against E. coli and S. aureus bacteria; further indicating suitability of the membrane in wound healing [35]. The combination of chitosan with metal ions has also been observed to yield biomaterials which exhibit antimicrobial properties [36]. Research conducted by Gritsch et al. [36] developed copper (II)-chitosan complexes via in situ precipitation, purposed as a biomaterial independent of antibiotics. Samples that contained copper, elicited a significant antibacterial effect when compared to the chitosan control. Additionally, inhibition of both Gram (−) and Gram (+) bacteria was observed, in as little as 1 hour of exposure (Fig. 11.4) [36]. Chitosan has also been incorporated into sponges to facilitate wound healing. Hu et al. [37] formulated composite sponges from physical mixtures of chitosan and hydroxybutyl chitosan, using freeze-drying. The developed sponges were extremely porous (∼85% porosity), which allowed for the successful absorption of large quantities of blood/wound exudate. As well as the ability absorb high volumes of exudate, the high porosity was also demonstrated to augment cell viability and proliferation, with negligible cytotoxicity within the mouse fiber cell line L929. Promotion of wound healing was also exhibited by the sponge indicating antibacterial activity and scope for utilization as a material for wound dressing [37]. The amalgamation of chitosan with alternative components such as metal ions to control bacterial growth is an exciting application of the polymer as its usage in the form of sponges





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FIGURE 11.4  The bacterial growth assessment of Escherichia coli (top) and Staphylococcus carnosus (bottom) on pure chitosan and copper doped chitosan shows clear inhibition due to the presence of copper. All CuChiX samples exhibit a statistically significant decrease in bacteria growth, both Gram-positive and -negative, after 9 h from inoculation. Source: Permission requested from Elsevier.

to absorb exudate, but there are further applications of the polymer in terms of implants which touch upon some of the same properties; such as antibacterial activity, as discussed in the succeeding section

2.5 Implants Whilst titanium and its associated alloys are widely used in medical implants, issues such as bacterial infection remains a common cause associated with implant failure and compromised osseointegration. Recent research conducted by Gilabert-Chirivella et al. [38] coated titanium surfaces, through etching titanium disks using 3D printing and coating with chitosan;



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removing the requirement of covalently bonded polymer to the metal implant. Implants were stable for up to a period of 5 days when stored in water. Moreover, cell adhesion assay using MC3T3-E1 (a cell line used to test biomaterials for bone implants) and C2C12 (a cell line used for investigating muscle growth, proliferation and differentiation) demonstrated an osteoblast preference for titanium regions, but with no notable preference for myoblasts; suggesting overall good biocompatibility. A major cause of failure of dental implants is attributed to biofilm formation, which is associated with a number of dental pathogens [39]. A recent piece of research by Divakar et al. [39] conjugated silver with chitosan, purposed as material for the coating of titanium dental implants. The prepared particles exhibited an inhibitory effect against two dental pathogens S. mutans and P. gingivalis, prevention the formation and subsequent adhesion of biofilm. The developed coatings were directly compared in terms of antimicrobial activity with the antibiotic agents Gentamicin and Ceftazidime, the particles exhibiting a lower minimum inhibitory concentration; indicating superior activity (Table 11.1). Additionally, when combined with the antibiotics, the particles elicited a significant decrease in minimum inhibitory concentration, as well as an increased zone of inhibition for both control and test organisms; identifying the particles as appropriate for both mono and combination therapy.

TABLE 11.1  Antimicrobial testing profile results. Test

Disk diffusion (mm)

MIC (µg/mL)

Control

Bacterial species

S. mutans

P. gingivalis

Staphylococcus aureus ATCC 25923

E. coli 25922

Staphylococcus aureus ATCC 29213

Blank disc

0

0

0

0

0

Chitosan

17 ± 3

19 ± 2

19 ± 2

20 ± 3

19 ± 3

Ag-Chitosan

19 ± 1

21 ± 2

20 ± 2

21 ± 1

24 ± 1

Gentamicin

18 ± 2

20 ± 2

18 ± 2

20 ± 2

NA

Ceftazidime

17 ± 2

24 ± 2

17 ± 2

24 ± 2

NA

Ag.chitosan + Gentamicin

19 ± 3

21 ± 2

20 ± 2

22 ± 3

24 ± 3

Ag.chitosan + Ceftazidime

22 ± 1

27 ± 2

24 ± 2

26 ± 1

27 ± 1

Chitosan

1

1

NA

1

0.5

Ag-Chitosan

0.5

8

NA

0.5

8

Gentamicin

0.5

0.5

NA

0.5

0.5

Ceftazidime

0.25

8

NA

0.25

8

Ag.chitosan + Gentamicin

0.12

0.12

0.25

0.12

0.12

Ag.chitosan + Ceftazidime

0.06

1

4

0.25

2

Permission requested from Elsevier.





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Treatment of vitreoretinal diseases necessitates repeated intravitreal injections of methotrexate (a chemotherapeutic agent) [40]. A recent piece of work conducted by Manna et al. [40] demonstrated that the usage of poly lactic acid and chitosan-based methotrexate micro implants were able to sustain release for up to a period of circa 1 month in vitro and in vivo. The coating of micro implants using modified forms of the copolymer poly-(lactic-co-glycolic) acid (comprised of poly-glycolic acid and poly-lactic acid) were observed to increase the longevity of the implants to three to five months. Adjusting the ratio of the co-polymer, that is, by increasing the poly lactic acid composition, release of methotrexate from the implant initially was retarded, in addition to delaying swelling and biodegradation of the implant [40]. Devices in contact with bodily fluids, specifically blood, are connected with clinical issues such as thrombosis and calcification. Thin polymeric coatings have been implemented as an approach to regulating the interaction between devices and blood. In recent research, sulfonated chitosan was observed to decrease thrombosis and calcification [41]. Research carried out by Campelo et al. [41] also examined sulfonated chitosan applied to stainless steel surfaces. These elicited higher roughness and hydrophilicity when compared to pure chitosan. Sulfonated chitosan was also observed to negate platelet activation (limiting clot formation). In addition to reducing calcification; indicating the materials appropriateness as a biomaterial which is able to address clinical problems associated with blood interacting implants. Chitosan has also been employed in the development of chitosan-carbon nanotube implants, in the form of tubular hydrogels supplemented with calcium ions for use in tissue engineering. This use was based on the formulation constituent’s ability to augment axonal outgrowth. The fabricated implants were comparable in mechanical properties to peripheral tissue. Additionally, in vitro testing indicated biocompatibility with no notable cytotoxicity of the implants (Fig. 11.5) [42].

FIGURE 11.5  Results of cytotoxicity tests conducted for chitosan-carbon nanotube implants enriched with calcium ions on (A) L929 fibroblasts and (B) mHippoE-18 hippocampal cells. Diagrams present mean percentages of viable cells calculated for four technical repeats performed in triplicates. Source: Permission requested from Elsevier.



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Chitosan has been successfully combined with collagen to functionalize silane titanium dioxide coatings in order to enhance the biocompatibility of magnesium alloys used in bone repair applications [43]. Córdoba et al. [43] demonstrated that the utilization of biopolymers could significantly impact upon the composition and corrosion products of magnesium alloys; forming the corrosion products calcium carbon trioxide and magnesium carbon trioxide, when compared directly to pure silane tiantium oxide coatings. In summary, modification of the magnesium alloys was indicative of superior osseointegration. The usage of chitosan to enhance the integration, longevity and biocompatibility of implants are but a handful of ways in which the polymers properties have been harnessed as a biomaterial.

2.6 Conclusion Despite limitations pertaining to the mechanical properties of chitosan it is widely utilized as a biomaterial, with particular advances in wound healing, whether in the form of dressings, hydrogels or miscellaneous applications such as scaffolds used in tissue engineering. Overcoming barriers or properties associated with a given polymer may commonly be overcome by combination with alternative polymers which is touched upon the forthcoming sections.

3  Hyaluronic acid 3.1 Introduction Hyaluronic acid is a naturally occurring linear polysaccharide comprised of alternating units of b-1,4-linked D-glucuronic acid and (b-1,3) N-acetyl-D-glucosamine20,21 [44,45]. It may also be described as a high molecular weight glycosaminoglycan [46]. Its extraction is possible from umbilical cords, synovial fluid, vitreous humor or from the cartilage of rooster combs [47]. Biologically important, hyaluronic acid is involved a number of physiological functions, including; tissue and extracellular water absorption and lubrication [48,49]. Moreover, as the polymer is non-toxic, non-allergenic and biocompatible, it has a number of applications, including; skin moisturization, tissue regeneration, inflammation moderation, and wound healing [48,50-54]. On a cellular level, hyaluronic acid has been observed to influence cell functionality through binding to cell receptors. This binding gives rise to a plethora of effects, for example, raised expression of pro-inflammatory cytokines, cell migration, cell proliferation and development of granulized tissue matrix; all of which are critical to wound healing. Additionally, due to the polymers viscoelastic and hygroscopic properties, it is believed to also impact upon cell behaviors (i.e., proliferation, differentiation, adhesion, and migration [45,55-57]. The role of hyaluronic acid in wound healing has been explored in several in vitro and in vivo pieces of research, identifying the polysaccharide as facilitating mesenchymal and epithelial cell migration, differentiation, enhanced angiogenesis, and collagen deposition [51,58-60].





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3.2  Hydrogel applications While hydrogels from hyaluronic acid are effective as biomaterials for soft tissue regeneration, they are also commonly brittle [46]. In spite of this drawback they polymer is widely utilized as a hydrogel [61-64]. For example, in a recent piece of research by Chen et al. [64] examined hyaluronic acidbased hydrogel microparticles with covalent semi-interpenetrating and conventional 3D molecule networks. These were investigated as an alternative to the use of macroscopic hydrogels which are utilized commonly as injectable scaffolds, however, were identified with the associated risk of significant blood vessel obstruction, restricting their clinical use. The formulated hydrogels in this research were observed to demonstrate a superior safety profile (in terms of dispersity, cell safety, hemocompatability) in comparison to commercially available fillers. Moreover, the formulated particles demonstrated functionality with respect to the entrapment and subsequent release of hydrophilic agents [64]. This is not the first example of injectable hyaluronic hydrogels. Research conducted by Zhang et al. [61] examined the role of hyaluronic hydrogels in soft tissue repair, by combining the polymer with collagen (human), cross-linking to construct a 3D network (Fig. 11.6). Through histocompatibility analysis, the formed hydrogel was notably lower in toxicity then hyaluronic acid based gels alone. Moreover, following further investigation, it was found that the hydrogel was capable of facilitating baby hamster kidney cell growth, sustaining sufficient cell viability in addition to adhesion. Aside from the aforementioned properties, the hydrogel was associated with a protracted degradation time, anti-enzyme ability and marginal inflammatory response at 1, 2, and 4 week intervals; indicating the gel’s suitability as a biomaterial for implementation in soft tissue-filling and repair [61].

FIGURE 11.6  Schematic diagram of crosslinking of Hyaluronic acid/Human like collagen hydrogels [61]. Source: Permission requested from Elsevier.



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TABLE 11.2  Nonlinear least-squares curve fit of the relationship between mean size of granulation tissue area (cm2) and time (days) after treatment [66]. Treatment

Granulation tissue area–Day 3 (cm2)

Initial rate of decrease of granulation tissue area

Days to 50% original size

Control

6.99 ± 0.26

0.145 ± 0.014

28.17 ± 0.86a

Single gel

6.64 ± 0.28

0.153 ± 0.018

27.64 ± 0.97a

Multiple gel

6.96 ± 0.27

0.135 ± 0.013

27.39 ± 0.94a

Multiple film

6.84 ± 0.16

0.151 ± 0.009

25.26 ± 0.53b

Abbreviation: SEM, standard error of the mean. Mean ± SEM. Letters indicate significant differences between groups. Permission requested from Elsevier.

Hyaluronic acid has also been further utilized as scaffolding in the promotion of survival of transplanted stem cells, as part of cell-based therapy of neurological disorders, where unprotected cell survival is particularly poor [65]. Liang et al. [65] utilized hyaluronic acid based hydrogels for the survival of human ReNcells and human glial-restricted precursor (GRP) cells, which were implanted into the brains of immunocompetent or immunodeficient mice. When using the hydrogels as scaffolding for these cells, survival substantially improved, even in the presence of inflammatory responses which were observed over a two week period following injection in immunocompetent mice brains; potentially indicating the need for further optimization [65]. In a recent piece of research hyaluronic acid biomaterials (gel and film; single and multiple applications) were investigated in horses with regards to wound healing [66]. Crosslinked hyaluronic gels and films were compared with respect to wound healing capability via single and multiple applications. Healing of wounds was determined with respect to wound size, tissue repair quality, and also healing rate. Upon comparison to the formulated gels, films were noted to be superior in terms of wound healing, as wound size reduction was achieved at a significantly faster rate with marked healing (i.e., a stronger epithelium) (Table 11.2) [66]. The usage of hyaluronic acid hydrogels as scaffolding, with intent for utilization in modified drug delivery and potential tissue engineering, was more recently examined by [62]. The hydrogels (containing dexamethasone) were formulated by crosslinking using Diels-Alder chemistry offering modified release. Thermoresponsive properties exhibited by the hydrogel were attributed to the formation process (i.e., Diels-Alder chemistry), suggesting competitiveness as a preparation technique. Upon cell culture, the developed hydrogel exhibited no toxicity, as well as offering entrapment and delivery of an andipogenic factor; indicating formulation suitability for implementation in soft tissue engineering [62]. Crosslinking of hyaluranon has also been observed to be beneficial in the treatment of corneal blindness. When crosslinked with hydrazole in the form of a hydrogel, hyaluranon was used to facilitate the delivery of human adipose stem cells which are used as part of cell therapy to regenerate damaged corneal stromal tissue [67]. Porcine corneas were utilized in order to investigate cell delivery in the repair of stromal defects. The formulated hydrogels demonstrated significant cell viability following encapsulation. Moreover, the addition of the protein collagen to the hydrogels was noted to preserve cell activity (metabolic) and influence hydrogel mechanical properties (i.e., increased stiffness), whilst also reducing the gel 



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swelling; additionally the gels were acknowledged as being clinically viable drug delivery systems for corneal defects [67]. A recent study conducted by Wu et al. [68] investigated the synthesis of hyaluronic acid supplemented fibrin hydrogels designed as microenvironments in the area of cartilage tissue regeneration. In order to examine cell viability, chondral defect regeneration; and chondrogenic gene expression; human adipose-derived stem cells were grown in hyaluronic acid supplemented fibrin hydrogels; as well as in simple fibrin hydrogels. These were subsequently further tested in vitro and ex vivo on osteochondral core explants, to determine chondral defect regeneration. Increased swelling ratio in combination with smooth surface was noted for the hyaluronic acid enriched fibrin hydrogels. This was attributed to the higher number of interconnected pores present in the gel when compared to pure fibrin gels. This was accompanied by an increase chondrogenic gene expression and in turn tissue development. The usage of hyaluronic acid to facilitate rapid wound healing and cellular growth in hamster kidney cells, are exciting applications of the polymer. However like chitosan, the polymers ability to promote wound healing is not restricted to use as hydrogels but may also be effective in the form of dressings as the subsequent section discusses.

3.3 Dressings Hyaluronic acid has been examined for its effectiveness in preventing adhesion following spinal surgery (i.e., postlaminectomey) [69]. Implementation of hyaluronic acid sheets in combination with animal model of lumbar laminectomy in rabbit, were tested for the prevention of postlaminectomy. Upon usage of the hyaluronic acid sheets, a number of observations were made, these included; a reduction in the number of inflammatory cells at the site of laminectomy; as well as suppression of the enlargement of the dura (fibrous membrane covering the central nervous system). The usage of the hyaluronic acid sheets were identified as a robust membrane barrier which also exhibited anti-inflammatory action [69]. With large skin wounds being characterized by a significant loss of dermis; hyaluronic acid exhibits effect at a variety of stages of wound healing, such as: re-epithelialization, granulation and even inflammation. However, esterification of the natural polymer into less soluble derivatives with improved structural properties is often necessitated due to the polymers poor physical properties, for example, solubility and rapid degradation. The improved properties of the polymers make them suitable candidates for deep burn treatment in the form of dressings [70]. Hyaluronic acid has been observed to be concentrated in the extracellular matrix of connective tissues [71]. A recent study conducted by Grundelova et al. [71] examined modifications of hyaluronan-based material to absorb liquid without compromising the materials mechanical properties using N-(3-dimethylaminopropyl-N′-ethylcarbodiimide hydrochloride (EDC) and N-hydroxysuccinimide (NHS). The modified films exhibited a significantly large increase in elasticity with films stretching up 130% before breaking; while also possessing the ability to absorb large volumes of water without detriment to the film mechanical properties. Combinations of hyaluranon and chitosan have also been examined in the form of films designed for transdermal drug delivery [72]. Recent research carried out by Bigucci et al. [72] targeted the poor bioavailability (25%) of Thiocolchiside (a compound which offers anti-inflammatory properties) using films developed from combinations of hyaluranon and



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chitosan [73]. Among the plethora of combinations of chitosan and hyaluranon (i.e., different ratio films), films which were fabricated in a 2:2 ratio of the polymers combined were associated with significantly lower water uptake than the counterparts investigated. This was in addition to the lowest drug diffusion through the skin, illustrating the impact which the ratio of polymers with distinctly different properties may have on resultant film characteristics. Healing of diabetic foot ulcers at sufficient rate and to completion is an arduous task, particularly in diabetic patients where wound healing environments can offer unfavorable conditions. When used as dressing material, hyaluronic acid has been observed to facilitate the healing of diabetic foot ulcers [74]. Research conducted by Lee et al. [74] investigated hyaluronic acid in the fabrication of dressings across a three-month study which investigated ulcer healing to completion; compared to orthodox dressings. Toward the latter end of the three month period, it was noted in patients treated with the fabricated dressings, that significantly faster ulcer healing rate was elicited with no adverse effects attributed to the dressing [74]. Hyaluroanon has been combined with chitosan in the formation of layer-by-layer nanofilms under varying pH and Ionic strength of polyelectrolyte solutions. The prepared films were examined for antibacterial properties against Xylella fastisiosa, Gram-negative bacteria which has significant agricultural impact. The developed films were noted to offer prolonged antimicrobial resistance, retarding bacterial growth for a period of up to 8 days (Fig. 11.7) [75]. The usage of hyaluronic acid as particles in novel dressings impregnated with platelet lysate and vancomycin hydrochloride has also been demonstrated in recent piece of research by Rossi et al. [76]. The formulated particles coated with calcium alginate were transferred to an alginate matrix. Loading capacity of the platelet lysate hyaluronic acid particles onto

FIGURE 11.7  Widefield fluorescence microscopy images of X. fastidiosa grown on nanofilm and Si samples for 2, 4, and 8 days. The inset images show the exposed Si surface, with larger number of adhered cells, on the corresponding PEM sample. Scale bars of images (including insets): 40 m [75]. Source: Permission requested from Elsevier.





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the dressings was found to be impacted by calcium chloride concentration. The relationship between both calcium chloride and loading capacity was correlated positively, with increasing calcium chloride concentration, loading capacity increased proportionally; this effect was associated with elevated crosslinking. The fabricated skin ulcer dressings were noted to be sufficiently strong to endure packaging and handling. Additionally, the dressings were able to absorb large volumes of exudate, with formation of a shielding gel surrounding the ulcer. Healing was also augmented by loading and delivering of an antibiotic compound dually to the ulcer throughout the healing phases [76]. Combinations of hyaluronic acid, collagen and chitosan have additionally been utilized to form films imbibed with the antibacterial agent gentamicin sulfate [77]. Interestingly, films fabricated by Michalska-Sionkowska et al. [77] were observed to retard Gram-negative and positive bacterial growth. Incorporated gentamicin was observed to impact upon the physical properties of the manufactured films, with film smoothness increasing substantially in the absence of gentamicin, potentially suggesting interaction between the compound and the polymers. Contrastingly, in the presence of the compound, film water and oxygen permeability was found to improve. While hyaluronic acid alone has been demonstrated to beneficial in terms of wound healing its combination with other naturally-derived polymers such as chitosan and antibiotics have demonstrated its versatility as a polymer and provided scope for amelioration of polymer properties for adaptation to purpose.

3.4  Composites, scaffolds, and matrices Copolymerization of hyaluornic acid and chitosan has been utilized in the preparation of interconnected macroporous cryogel networks, designed for the purpose of tissue-engineering scaffolding [78]. Research conducted by Kutlusoy et al. [78] formulated cryogels which employed the use of glutaraldehyde as a cross linking agent in combination with the chitosan and hyaluronic acid in a myriad of ratios (i.e., 0, 10, 20, 30, and 50 wt% hyaluronic acid). The formulated gels were highly porous (90%-95% porosity) (Table 11.2). Moreover, chitosan’s mechanical properties were noted to improve significantly following incorporation of hyaluronic acid, ensuing the formulation of a malleable and robust blend, which was biodegradable and exhibited rapid swelling behavior. Upon cytoxicity analysis using MTT assay, cell viability results were indicative of no significant cytotoxicity associated with the cryogels, when fibroblast and SAOS-2 cells [78] (Table 11.3). A piece of research conducted by Sanad and Abdel-Bar [79] investigated hyaluronic acidchitosan combinations in the formulation of composite sponge scaffolds purposed for holding of lipid nanoparticles with andrographolide, developed for wound healing. Subsequent to successful formulation of angrapholide lipid particles in addition to chitosan-hyaluronon/ andrographolide nanocomposite scaffolds; the scaffolds were examined in terms of porosity and swelling ratio, offering desirable properties in terms of both parameters. The sponge scaffolds were demonstrated to enhance wound healing when tested in rats in vivo, no scar was visible with tissue quality also being enhanced (Fig. 11.8). Additionally, controlled release andrographolide was achieved from the sponge particles for up to a period of 72 hours. The augmented healing noted was associated with the synergistic action of chitosan and hyaluronic acid, that is, anti-inflammatory and anti-oxidant.



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11.  Importance of biomaterials in biomedical engineering

TABLE 11.3  Swelling ratio, porosity, and pore diameter of cryogels in equilibrium swelling. Swelling ratio

Porosity

Pore diameter

Elastic modulus (E)

Samples

(%)

(%)

(µm)

(kPa)

4G2C

1250

94

90-150

90

6G2C

1200

95

90-150

-

10G2C

850

97

90-150

-

4G2C10HY

1900

94

150-200

32

4G2C20HY

2000

92

150-200

34

4G2C30HY

2400

91

150-200

37

4G2C50HY

3450

90

150-200

40

Permission requested from Elsevier.

FIGURE 11.8  Photographs of treated wounds at different time intervals. Source: Permission requested from Elsevier.

While chitosan possesses ample biocompatibility, the polymers brittleness is restrictive of it usage in fabricating wound healing materials. Blending with other polymers (e.g., hyaluronic acid) has been deemed a potential remedy to this problem [80-82]. As a crucial component of skin extracellular matrix, the harnessing of hyaluronic acid in wound healing is of great interest. It is known to facilitate a moist environment which protects the wound by maintaining adequate moisture levels, promoting healing [79]. Research conducted by [79] formulated composite scaffolds made from chitosan-hyaluronic acid impregnated with andrographolide lipid nano carriers. The developed sponge offered controlled drug release, which augmented wound healing and brought about a reduction in scar formation. Histological progress was also enhanced in comparison to the control. These effects were associated with synergistic action of both polymers (e.g., anti-inflammatory and antioxidant effect of chitosan and hyaluronic acid) when combined with the andrographolide. The design and fabrication of bioactive scaffolds offering an artificial environment comparative to real biological environments is a critical area of interest in biomaterial science [83]. 



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165

A recent study conducted by Garcia-Fuentes et al. [83], recognized that whilst hyaluronic acid and silk fibroin are utilized widely in biomaterials individually, the combination of both to co-utilize desirable characteristics of the polymers (i.e., mechanical properties from silk fibroin and biological properties from hyaluranon) could be useful when fabricating scaffolds. Upon combination to form scaffolds and following seeding and culturing of stem cells (mesenchymal) for 3 weeks; the polymer blend scaffolds were found to be superior in terms of cellular in growth when compared to pure silk fibroin scaffolds. Moreover, greater efficiency in tissue formation upon usage of tissue-inductive stimuli was observed; suggesting the polymer blend scaffolds appropriateness as a basis for mesenchymal stem cell research. The abovementioned research is not the only example of the combination of hyaluronic acid with other polymers to yield modified properties. Recent research conducted by Sionkowska et al. [84] fabricated two-dimensional composites using combinations of hyaluronic acid, chitosan and collagen. Structural, biological and mechanical properties were found to vary with the percentage of hyaluronic acid incorporated (i.e., 1, 2, and 5%); with mechanical and thermal stability improving without impact upon cell viability. Electrostatic and hydrogen bonding between the incorporated hyaluronic acid with collagen and chitosan was deemed to improve elasticity and porosity (i.e., increasing); potentially increasing cell growth and proliferation. More recently, hyaluronic acid has been employed in the development of porous foams allowing for the drug delivery of NSAIDs (non-steroidal anti-inflammatory drugs) such as ketoprofen, when using cationic surfactants [85]. Roig et al. [85] developed foams were observed to release ketoprofen when exposed to phosphate buffer markedly better when hyaluronic acid was incorporated, with the presence of cationic surfactant further enhancing the release profile; thus indicating the systems suitability being biocompatible, biodegradable, and augmenting the solubility of lipid-soluble drugs for delivery [85]. The aforementioned pieces of research further highlight the scope offered from the combination of polymers to modify the properties of resultant biomaterials for use. However, much like chitosan, hyaluronic acid also has applications in the area of implants.

3.5 Implants A leading cause of titanium-based implant failure is their loosening following implantation. A recent study carried out by Ao et al. [86] formulated novel collagen/hyaluronic acid coatings applied in a layer-by-layer manner upon titanium implants. The resultant coatings exhibited superior biological properties in comparison to alternative coatings. Moreover, the presence of the coatings resulted in enhanced osseointegration (Fig, 11.9), suggesting potential to negate the slackening of titanium implants through covalent immobilization of the titanium coatings [86] In the treatment of symptoms associated with osteoarthritis (i.e., joint pain), hyaluronic acid has been observed to be effective [87]. A reduction in the lubrication of joints results in a plethora of changes, including: inflammation and increased wear and tear; which dually accelerate cartilage degradation. Research conducted by Corvelli et al. [87] found that hyaluronic acid was able to reduce the mean static and kinetic friction of cartilage samples (75% and 70% correspondingly). These values were on concordance with the properties of recently isolated synovial fluid in joints [87].



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11.  Importance of biomaterials in biomedical engineering

FIGURE 11.9  Histological morphology of the interface between implants and bone tissue after 3 months implantation in rabbit femur condyles. Source: Permission requested from Elsevier.

The initial stages of osteoarthritis are associated with changes in viscoelasticity of the synovial fluid which may cause a reduction in the protection offered by cartilage. The aforementioned effect is attributed to a decrease in concentration and molecular weight of hyaluronic acid. Hyaluronic acid injections have been posed to assuage the drop in elasticity associated with osteoarthritis and delay surgical procedures [88]. Henrotin et al. [88] investigated the effectiveness of intra-articular injections of hyaluronic acid and chondroitin in patients suffering from knee osteoarthritis. Clinical effectiveness was determined by gauging pain in patients who received the injection thrice weekly. The results demonstrated a noteworthy drop in mean pain; indicating a strong therapeutic effect. The varied applications of hyaluronic acid in arthritic disease states are stimulating, with enhancement of lubricity in joints being a unique point associated with the polymer.

3.6 Conclusion Much like chitosan, hyaluronic acid has been demonstrated to be particularly useful in the area of wound management, being formulated as hydrogels, dressings, and films to facilitate healing. It has further been beneficial in the treatment of arthritic conditions, particularly in the form of injections in patient suffering from knee osteoarthritis.

4  Silk fibroin 4.1 Introduction Principally there are two families of insects which are linked to the spinning of silk fiber, Bombycidae and Saturndiae. Silk proteins are a key component of cocoon filaments and resemble the polymers keratin, collagen and fibroin; these proteins are secreted by silk gland cells in silk worms. Silk fiber proteins predominantly comprise this secretion accounting for 75%; with the remaining 25% is accounted for by sericin; which is a crucial constituent influencing the quality and quantity of total secreted fiber ([89,90]). For example, sericin protein is noted for a number of key properties, including: UV protection, antibacterial, chemoprotective, and antioxidant. 



167

4  Silk fibroin

The fiber is applied in an array of industries (i.e., medical, textile, and industrial) due to specific structural properties, that is, long, thin, light, soft, ability to offer insulation, water absorbency, thermal stability, and also its ability to bond with dyes [91]. While there are many synthetic polymers available on the market, the natural properties of the fiber combined with the absence of strict processing parameters, makes it use desirable. Despite the lack of requirement for strict processing, it is notable that research conducted by Phillips et al. [92] identified the introduction of ionic liquid to be of great importance for the production of silk fibres which enhanced optical and mechanical properties, thus suggesting capacity for amelioration of the polymer properties [92].

4.2  Silk fibroin as a biomaterial Being documented as a biomaterial by the FDA (US Food and Drug Administration) in 1993, silk fibroin is also posed to have superior mechanical properties to many comparable natural biopolymers as well as being biodegradable and amenable having many applications as a biomaterial (Table 11.4) [93]. The aforementioned properties are attributed to the polymers distinctive physiochemical properties. Sericin being an amorphous protein and glue like, acts as a binding agent to help maintain the integrity of the silk fibers [94,95]. The application of silk-based biomaterials is widely documented, examples include: bone and bone tissue regeneration (3D scaffolding) [96-99], cartilage repair [100], implants in treatment of refractory epilepsy [101] and peripheral nerve repair (nerve conduits) [102]. While there are many applications of silk fibroin, there is significant amount literature dedicated to the polymers fabrication into nanoparticles for drug delivery.

TABLE 11.4  Application of silk fibroin in tissues and cell as biomaterial from Ref. [120]. Application

Form/material

References

Wound dressings

Film

[121]

Sponge

[122]

Sponge

[123,124]

Film

[125,126]

Hydrogel

[127,128]

Film

[129,130]

Porous sponge

[131,132]

Hydrogel

[133]

Ligament tissue engineering

Fiber

[134]

Tendon tissue engineering

Fibers

[135]

Hepatic tissue engineering

Films

[136]

Connective tissue

Non-woven mats

[137]

Endothelial and blood vessel

Non-woven mats

[138]

Anti-thrombogenesis

Films

[139]

Bone tissue engineering

Cartilage tissue engineering



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11.  Importance of biomaterials in biomedical engineering

Silk fibers derived from B. mori silk worms are typically utilized for biomedical applications, commonly as sutures. The fibers clinical applicability is marred by biological reactions to the protein, thus inferring reduced biocompatibility. Thus, undesirable effects such as hypersensitivity have been observed, the origins of these have been determined to be from the sericin in derived fibers. Moreover, while biodegradability of the material has been identified as an issue, literature is indicative that silk proteins are susceptible to proteolytic degradation in vivo and will also steadily absorb [103,104]. Although fibroin is incorporated into medical sutures, the prevalence of synthetic materials which do not suffer from the same drawbacks (i.e., poor biocompatibility, increased incidence of adverse reactions and sensitization), have essentially brought about a drop in preference in terms of its use. However, research carried out by Altman et al. [103], identified sericin (associated with Type I allergic reactions, due to upregulated IgEs) impurities in the employed materials as being the origin of the reactions observed. Consequently, used of fibroin as a biocompatible materials for medical applications has increased, with core fibroin protein observed not to elicit allergic reactions when incorporated in sutures.

4.3  Silk fibroin in drug delivery Silk fibroin has also been used as a material for the formation of matrices in the delivery of therapeutic agents (protein-based); offering preservation of potency and amelioration with respect to: concentration of morphology, crystallinity, silk fibroin, and molecular weight, all influencing release kinetics from silk fibroin matrices. There have many approaches modify release from silk fibroin matrices, which have included the coating of micro or nanoparticles with silk fibroin films; or alternatively incorporation of nano-/microparticles containing therapeutic compounds. A primary application of silk fibroin delivery systems is tissue regeneration, for example, silk fibroin scaffolds impregnated with growth factor were successfully used in tissue engineering of bone and cartilage, in addition to for nerve and vascular regeneration devices oral, ocular, and trans-mucosal drug delivery and wound healing products [105]. Silk fibroin has also been used in the lipid vesicle modification. For instance, liposomes and microparticles coated with silk fibroin facilitated prolonged release from the vesicles in addition to cellular recognition. Research carried out by [106] examined alginate and poly(lactic-co-glycolic acid) microparticles coated in silk fibroin using a layer-by-layer assembly technique. The coated microparticles exhibited lowered rates of decomposition when compared to uncoated particles. In addition to delaying decomposition, the coatings were able to sustain release of horse radish peroxidase from particles made from poly(lactic-coglycolic acid) and tetramethylrhodamine-conjugated bovine serum albumin (Rh-BSA) from alginate microparticles. The modified released demonstrated by SF coated particles was suggestive of superiority over alternative systems which rely on pure poly(lactic-co-glycolic acid) particles, which exhibit burst style release, rapid peaking of drug plasma levels, and increase the likelihood of toxicity [106,107]. As silk fibroin possesses a large number of active amino groups and tyrosine residues which offer potential as a site for conjugation with active molecules or growth factors [108]. Amalgamation of bio-conjugates of protein/polypeptides compounds with silk fibroin nanoparticles has been achieved through covalent attachment to the particle surface, allowing





4  Silk fibroin

169

for potential amelioration of function. For example, coupled with insulin covalently, silk nanoparticles exhibited resistance trypsin digestion, as well as thermal stability and enhancement of the in vitro stability of the insulin [109-111]. Research conducted by Gobin et al. [112] examined the use of silk fibroin coated liposomes in the delivery of the natural anticancer agent Emodin which is used in refractory tumors as a selective receptor tyrosine kinase inhibitor [112]. Delivery and potency of the compound was enhanced using the coated liposomes through localized delivery, increasing the specificity of cell recognition and residence time when compared to uncoated liposomes loaded with Emodin [112]. The mechanism of release of Emodin from the coated liposomes was determined to be by diffusion as opposed to a combination of swelling and diffusion observed by the uncoated liposomes. Lack of swelling observed with silk fibroin Emodin liposome may be attributed to the steric hindrance from the large silk fibroin molecules [112]. Moreover, silk fibroin coated liposomes have exhibited enhanced specificity, adhesive targeting, and efficacy against breast cancer cells. Enhanced cellular uptake as a result of the silk fibroin coating, resulted in elevated concentrations of drug in cells as well as hijacking of multiple pathways contributing to further growth inhibition [113]. Coatings of silk fibroin on nanoparticles have also been beneficial in the treatment of breast cancer cells using curcumin as a therapeutic compound. Higher intracellular uptake was observed with prolonged availability of curcumin from the silk fibroin coated nanoparticles. [114]. The polymer has also been utilized as a material to construct nanoparticles as opposed to mere coatings. Research conducted by Kundu et al. [108] and Huang et al. [115] were able to form stable, spherical, and negatively charged silk fibroin (extracted from a non-mulberry tasar silk worm) particles within a size range of 150-170 nm. The particles were successfully utilized in the sustained delivery of vascular endothelial growth factor with a biphasic release rate pattern noted for the vascular endothelial growth factor from the silk fibroin nanoparticles (i.e., an initial linear burst (circa 35%) succeeded by a sustained release for up to five (1% per day) [108,115]. Further examples of research include work conducted by Myung et al. [116] who examined the fluorescent dye rhodamine B encapsulated in spherical silk fibroin nanoparticles, which were found to exhibit enhanced stability suggesting potential application as device for molecular imaging and bioassays [116]. Owing to silk fibroins amphiphilic nature it is able to augment the loading of hydrophobic and hydrophilic into nanospheres derived from the polymer. Loading and release of the therapeutic compound may be ameliorated based on the degree of crystallinity within hydrophobic regions of the polymer or charge states within hydrophilic regions of the polymer [117]. The degree of crystallinity of silk fibroin determines its in vivo degradation rate, which in turn controls the drug release rate of bioactive molecules which may be loaded onto the silk fibroin structures [98]. Silk sericin is also deemed to possess a number of properties desirable in biomaterials, including: antibacterial action, oxidation, and UV resistance. Typically, small sericin pepties ( Ag-S >> Ag-Cl > Ag-N >> Ag-O (12.1) Silver is used both in heterogeneous oxidation processes and homogeneous silver-mediated and catalyzed reactions as well as the oxidation of primary and secondary alcohols to aldehydes and ketones because it is a mild oxidizing agent [6]. Silver exhibits four oxidation states; the most common state is Ag (I). Its coordination number varies from 2 to 6 due to its spherical symmetric configuration. The dz2 and s orbital of silver are close in energy allowing extensive hybridization to take place [19,49]. Silver (I) is stable in aqueous solution as [Ag(H2O)2]+, but the Ag+ ion has a low affinity for the hard aqua ligand giving a tendency for its crystalline inorganic salts to be anhydrous. Silver complexes containing N donor ligands readily form aqueous solutions with large formation constants observed for complexes of the type [AgL]+ and [AgL2]+ due to their preference for linear coordination [17]. The most important classes of O-donor ligands for complexation with silver (I) include carboxylates, crown ethers, and calixarenes [2,37]. Silver (I) also forms numerous complexes with the donor atoms S, Se, P, and As and binds to peptides and proteins with a preference for the thioether, sulfur, and nitrogen atoms in imidazole [2]. 

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12.  Visible light activated antimicrobial silver oxide thin films

The other oxidation states of silver include Ag(II) which has an electronic configuration of 4d9 and forms many complexes. The complexes are usually square planar and paramagnetic. Ag(III) oxidation state is very uncommon and forms highly unstable compounds, while the oxidation of fused Ag2O with persulfate give a mixed oxide AgIAgIIO2 [50]. Another oxidation state of silver is Ag(IV). 2.2.1  Crystal field theory and coordination chemistry of silver A coordination complex or metal complex is made of a metal atom or ion and a surrounding array of bound molecules or ions known as ligands or complexing agents. Each ligand makes available in the system a free pair of electrons and is a Lewis base, while the metal located in the coordinating center is a Lewis acid. The bonding between the metal base and the ligand is hinged on a Lewis base/Lewis acid interaction and is sometimes referred to as donor bond or coordinate covalent bond [51]. The nature of the bond between metal and ligand is stronger than intermolecular forces because they form directional bonds between the metal ion and the ligand but are weaker than covalent bonds and ionic bonds. The crystal field theory explains that the interaction between transition metals and ligands arises from the attraction between the positively charged metal cations and the negative charge on the non-bonding electron ligands. Crystal field theory is a model that describes the breaking of degeneracies of electronic orbital states usually those of the d or f orbitals due to the static electric field produced by a surrounding charge distribution brought about by neighboring anions resulting to a change in magnetic properties and color. Energy changes are introduced into the five degenerate d-orbitals surrounded by an array of point charges from the ligands [52]. As the ligand gets nearer the metal ion, the electrons from the ligands get closer to some of the d-orbitals and further away from others causing a loss of degeneracy. Metals and ligands are point charges in the complex. The metal is the positive cation, while the ligand is the negatively charged ion or a neutral molecule with negative polarity at its end. Metals used here are mostly transition metals such as copper, iron, gold, silver, etc. Examples of simple ligands and their ligand names include water (aqua H2O), ammonia (ammine NH3), chloride (chloro Cl−), azide (azido N3−), bromide (bromo Br−), cyanide (cyno CN−), hydroxide (hydroxo OH−), [51,52] etc. 2.2.2  Crystal field splitting The electrons in the ligands and those of the d-orbitals of the metal repel each other in a coor­ dinating complex. The d-electrons closer to the ligands will have higher energy than those further away, and this results in the d-orbitals splitting in energy [51]. The higher the metal’s oxidation state, the larger the splitting and if the effect of the ligands is substantial, there will be a more considerable difference between the high and low energy d-groups [51]. In the octahedral complex, six ligands form an octahedron around a metal ion. The d-orbitals split symmetrically into two sets with an energy difference (the crystal field splitting parameter devoted by ∆oct) where the dxy, dxz, and dyz are lower in energy than the d x 2 − y 2 and dx2 which have higher energy [51]. This is because the first orbitals are further from the ligands and experience less repulsion than the second group. The lower energy orbitals are collectively referred to as t2g and the two higher energy orbitals as eg, [51]. In tetrahedral complexes, four ligands form a tetrahedron around a metal ion. In this crystal field splitting, the d-orbitals again split into two groups with an energy difference of ∆tet and the lower energy orbitals are dzy and the higher energy orbitals





2  Theoretical background

183

are dxy, dxz, and dyz [51]. The energy splitting here is lower than that in octahedral complexes because the ligand electrons are not oriented directly toward the d-orbitals. The metal is located at the center of an imaginary x, y, and z coordinate system with the electrons arrange along the axes. The electrons are initially degenerate when they are far from the metal center but become non-degenerate as they approach the center of the x-, y-, and z-axis. The attraction between the positively charged metal cation and negative charge on the non-bonding electrons of the ligands is responsible for the interaction between transition metals and the ligands [51]. Energy changes occur in the five-degenerate d-orbitals surrounded by an array of point charges from the ligands. As the ligand get closer the metal ion, electrons from the ligand will get closer to some of the d-orbitals and further away from each other causing a loss of degeneracy. Repulsion therefore exists between these electrons and the electrons closer to the ligands will have a higher energy than those further away resulting in the d-orbitals splitting in energy. The removal of degeneracy resulting in crystal field splitting increases electron transition possibilities making room for more electron exchange between the split orbitals thus making the coordination system more chemically active. Such ease results to many ligand exchanges and informs antimicrobial activity of coordinating complexes. 2.2.3  Silver coordinating complexes Silver forms complexes with S, Se, P, and As. It forms very stable complexes [Ag(S2O3)]− and [Ag(S2O3)2]3− with sulfur [52]. Other important sulfur ligands for Ag+ are thiolate anions which give oligomers, thioureas, and thioethers [52]. Silver (I) also binds to peptides and proteins with preference for—SR and imidazole nitrogen functionalities. Four-, five-, and sixcoordination is found in Ag+ complexes with sulfur macrocycles [52]. 2.2.4  Applications of silver complexes as antimicrobial agents Apart from silver salts, many complexes of silver are used clinically. The antimicrobial activity and other desirable properties can be changed by varying the number and type of ligands coordinating the silver ion. Key factors used in the design of silver-based antimicrobial complexes are [52]: • • • •

The type of atoms bound to the Ag+ ion (O > N and S ≫P) The ease of ligand replacement and control of Ag+ ion release Chemical and photo-stability Cost

In a complex biological system such as wound fluid, the maximum level of available free Ag+ is approximately 1 µg/mL, above this level, Ag+ complexes with anions such as chlorides to form soluble and inactive salts resulting to loss of efficacy [53]. The rate of metal ion release and the relative stability of any silver complexes can be controlled using variations in the ligand architecture and the overall structural motif of the compound. There is ample evidence that silver ions prevent bacterial cell wall-activity by binding to specific membrane proteins and thereby obstructing the activity of enzymes [51]. The mechanism of antimicrobial activity in aqueous silver ions has been proposed to take place in three ways (1) interference with electron transport, (2) binding to DNA, and (3) interaction with the cell membrane. The antimicrobial activities depend on the type of coordinating donor atoms. The coordinating donor atoms to the silver (I) center and the ease of ligand replacement are the critical factors



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leading to a broad spectrum of antimicrobial activities, and the primary targets for the inhibition of bacteria and yeast by the silver (I) complexes [54]. Ag+ treated bacteria exhibit a characteristic initial stimulation in respiration before cell death occurs. Ag+ has been reported of uncoupling the respiratory chain from oxidative phosphorylation, collapse the proton motion force across the cytoplasmic membrane or interact with thiol groups of membrane-bound enzymes and proteins [55–57]. The site most vulnerable to attack by biocidal species is the cytoplasmic membrane, where many essential proteins reside, including enzymes of the respiratory chain and vital transport channels [55]. The action of Ag+ has been proposed to render the cytoplasmic membrane permeable to protons and to collapse the proton gradient, to compensate for the loss of the proton concentration gradient, the cell respires faster to expel more protons through the respiratory chain. The loss of proton motive force means that the respiration no longer depends on adenosine triphosphate (ATP) synthesis and its rate is no longer determined by the availability of ATP, as it is during normal respiration. When there is an inhibition on the respiratory chain, there is a prevention of the efficient pumping of protons across the membrane and the maintenance of the proton gradient. Such uncontrolled respiration results in the production of superoxide and hydroxyl radicals which are extremely dangerous to the cell. Binding of Ag+ to low-potential enzymes of the bacterial respiratory chain and the resulting production of large quantities of reactive oxygen species (ROS), because of the inefficient passage of electrons at the terminal oxidase, explains why silver (I) ion is toxic to bacteria in such low concentrations. Three possible mechanisms for inhibition by aqueous silver (I) ion have been proposed (1) interference with electron transport, (2) binding to DNA, and (3) interaction with the cell membrane [58]. Even though silver has a low affinity for oxygen donor ligands, it still forms complexes with it. Silver complexes of oxygen donor ligands such as [Ag(hino)]2 (where hino = 4-isopropyltopolone) and water-soluble silver (I) complexes of R (- +)- and (S)-(-)-2pyrollidone2-carboxylic acid display wide range of antimicrobial activity against bacteria, yeasts, and moulds [58,59]. The biological action of the silver (I) oxygen complex comes from the weaker bonding property of the Ag–O bond. In a biological system, the ease of ligand replacement of the silver (I) complexes would result in further replacement with biological ligands. The Ag–O bonding complexes can readily undergo ligand replacement with O–, N–, or S– donor atoms. The antimicrobial activities of silver (I)–oxygen bonded complexes are due to silver (I) ion interacting with biological ligands such as protein, enzymes, and cell membrane [59]. The coordination ligands of the silver (I) complexes play the role of carrier for the silver (I) ion to the biological system. The magnitude of the antimicrobial properties of silver complexes is related to the ease with which they participate in ligand exchange reactions [59]. It has been speculated that the weak Ag–O and Ag–N bond strengths might play an important role in exhibiting a broader spectrum of antimicrobial and antifungal activities and the potential target sites for inhibition of bacteria and yeast growth by silver complexes might be the sulfur containing residues of proteins [59].

2.3  Properties of silver oxides The oxygen states of silver are classified into four (1) physisorbed oxygen, (2) molecular chemisorbed (3) atomic oxygen (dissociatively adsorbed), and (4) oxygen incorporated in the





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subsurface layers or dissolved in the bulk [60,61]. It is peculiarly noted that O2/Ag systems have low sticking coefficient of oxygen. The system therefore, requires the use of high oxygen pressure treatment of silver samples to achieve a high oxygen concentration on the surface. Silver oxide exists in different crystal systems such as cubic, monoclinic, and tetragonal [62]. The silver oxide can be prepared using rf and dc sputtering, thermal evaporation, chemical synthesis, and pulse laser deposition [24,63]. The mixed silver oxide films exhibit p-type semi-conductivity and operates based on Sanderson’s theory of partial ionic charge that the bond between oxygen and silver contains both ionic and covalent components. Silver oxides have different stoichiometry such as Ag2O, Ag3O4, AgO, Ag4O3, Ag2O3, Ag4O4, and Ag3O [46,63]. The formation of oxides depends on growth conditions and reaction kinematics. When deposition parameters were varied, Raju and coworkers [63] observed that the silver oxide compounds formed were polycrystalline. The deposition was done at plasma pressures between 9 and 50 Pa. It was observed in another research that when the oxygen pressure in the growth chamber was increased, the initial hexagonal Ag2O transformed to monoclinic AgO. The grain size in the film increased from 59 to 200 nm, the surface roughness of the film increased from 9 to 42 nm, the resistivity of the films increased from 1 to 4 × 104 Ωm, the surface work function of the films increased from 5.47 to 5.61 eV and the optical band gap of the AgO decreased from 1.01 to 0.93 eV [62]. The deposition was done using the pulsed laser deposition technique. This study shows that single phase AgO thin films required for plasmonic devices can be prepared at room temperature by pulsed laser deposition technique with an oxygen flow rate of 20 Pa [62]. Irrespective of deposition conditions and annealing, silver oxide remains unstable and decompose into metallic silver and oxygen either wholly or partially [46]. The surface morphology and nucleation kinetics of silver oxide is determined by the kinetic energy of the particles, that is, silver, oxygen atoms, or the silver oxide molecules reaching the substrate [63]. Oxidation of silver takes place in the plasma in the gaseous phase because of the small mean free path of silver atoms; consequently, many collisions with oxygen atoms result in the silver oxide molecules reaching the substrate [63]. If there is insufficient oxygen in the chamber due to low oxygen pressure, metallic silver will be deposited on the substrate [63,64]. Analysis reveals that the thickness (1–1.5 µm), the optical band gap (1 eV) and the work function all vary with oxygen pressure in the deposition chamber [62]. In another report, silver was evaporated in the presence of electron cyclotron resonance oxygen plasma and used for the deposition of Ag2−xO films with a range of stoichiometry unto a sapphire substrate [64]. The Ag2−xO films have conductivity of the order of 3 × 10−3 Ω−1cm−1 [64]. Another group of researchers [65] have reported that the rate of evaporation controls p-type or n-type conduction in silver oxides during reactive electron beam evaporation and oxygen plays a dominant role in the conduction process. It was observed that none stoichiometric silver oxide thin films are n-type, while for the oxygen flow rate of 0.54, 1.09, and 1.43 sccm, p-type silver oxide conduction was observed [65]. The most thermodynamically stable oxide of silver is Ag2O. The Ag2O phase possesses a simple cubic structure with a lattice parameter of 0.4728 nm at room temperature. AgO, on the other hand, crystallizes with a monoclinic structure containing both Ag+ and Ag3+. The different stoichiometry of silver gives rise to a wide range of band gap (between 1.2 and 3.4 eV). Rivers and coworkers [64] reported in their research that an Ag2-xO mixed film containing Ag2O phase with (111) and (002) orientation and AgO phase with (111) and (002) orientation [65]. In another research report, Fortin and coworkers [66] reported production of Ag2O



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having a gap of 1.2 eV. In their research, Pierson and coworkers [67] used rf-reactive magnetron sputtering to grow Ag2O films with 2.23 eV band gap. Another research conducted by Varkey and coworkers [68] produced Ag2O using chemical bath deposition with a band gap of 2.25 eV by subsequent air annealing. The formation of different oxides of silver is controlled by the manipulation of deposition conditions.

2.4  Bioinorganic chemistry of silver-based antimicrobials Bioinorganic chemistry is a field of study that investigates the role of the structure, function, mechanism, and dynamics of biologically relevant metals in biological systems. Metals play roles such as transport, speciation, and eventual mineralization in inorganic material [69]. Both deficiencies and excesses of metals in biological systems cause pathological changes. The storage of transition metals and the synthesis of the transporter molecules are carried out by specialized cells. The metal is in ionic form, but the oxidation state can vary depending on biological needs. Bioinorganic chemistry considers the significant roles of metals in biological systems as electron carriers, centers for binding and activating substrates as well as agents for transferring atoms and groups [69]. Many processes in biological systems such as respiration, metabolism, nitrogen fixation, photosynthesis, development, nerve transmission, muscle contraction signal transduction, and protection against toxic and mutagenic agents all require metal ions to function [69,70]. Metal ions such as Na, Mg, K, Ca, V, Fe, Ni, Cu, Zn, Mo, and W as well as Cr, Co, and Cd are used for the performance of vital body activities [70]. The naturally occurring metals are required for the performance of functions such as photosynthesis, hemoglobin formation, etc. Respiration is a function uniquely performed by metalloproteins. Breathing involves taking in dioxygen from the atmosphere and using it in the process of respiration through its interaction with the atoms of iron in hemoglobin and myoglobin [71]. Messages are sent to the nervous system by using changes in the electrical flux caused by movement of potassium and sodium ions across cell membranes [71]. Phosphorus present in DNA double helix, a polymer of nucleotides also plays a vital role in adenosine triphosphate (ATP) which is central to cell energetics. Iron, copper, and silver can readily change their oxidation states and are very useful in electron transfer systems and do produce oxidized substrates which participate in a variety of metabolic cycles. Iron and copper are involved in the carriage and storage of dioxygen [69,71]. Iron is present in hemoglobin, myoglobin, and hemerythrin, while copper is present in hemocyanin. There is iron storage (ferrinin) and carrier proteins (transferrin), lactoferrin and siderophores and a copper-transfer protein (ceruloplasmin). Copper serves as a superacid center in metalloenzyme promoting the hydrolysis or cleavage of a variety of chemical bonds, such as carboxypeptidase, carbonic anhydrase, and alcohol dehydrogenase (converts alcohol to acetaldehyde). Zinc acting as the bimetallic enzyme superoxide dismutase breaks down toxic superoxide to give dioxygen and peroxide from which the peroxide is converted to dioxygen and water by the iron-containing metalloenzyme catalase. Manganese is involved in mitochondrial superoxide dismutase, inorganic phosphatase, glycosyl transferase, and photosynthesis system II. Cobalt is present in vitamin B12 and its coenzyme. Nickel functions in metalloenzymes such as urease and several hydrogenases. Molybdenum and vanadium are found in nitrogenases in clusters containing iron and sulfur. Vanadium is also active in haloperoxidases [69].





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The replacement of any of these metal ions directly affects the functionality of biological systems capable of compromising the survival of the organism such as bacteria. Metal ions such as silver ions are deliberately released to replace some of these biologically important metals with an intention to stop essential cell function such as respiration, cell energetics, etc. which the cell cannot survive and is therefore at the heart of antimicrobial activity. Information gathered from molecular structures in bioinorganic chemistry indicates that nature has carefully selected and located certain metals to perform well defined functions in biological systems in living matter. The main factor used to design antimicrobial materials is to disrupt or interrupt the functionality of the selected metals by removing or blocking them from performing the vital biological functions they are meant to perform. These molecules containing the metals are in the cell membrane, proteins the DNA/RNA, and other biochemical components of living cells.

2.5  Bacteria cell structure Antimicrobial attacks target molecules located at designated areas in the bacteria cell such as the cell membrane and inner biochemical components. A bacterium cell also known as a prokaryotic cell is composed of five components namely; the nucleoid (DNA), ribosomes, cell membrane, cell wall, and a surface layer. The components are arranged in three structural regions—appendages hosts flagella and pili, the cell envelope contains a cell wall, plasma membrane a capsule, and the cytoplasmic region has chromosomes (DNA) and ribosomes. The flagella are made of proteins and are used for movement of the bacteria. The pili also are made of proteins and are used by bacteria for attachment to surfaces and protection. The capsule is usually made of polysaccharides and occasionally polypeptides. The cell wall in Gram-positive bacteria is made of 90% peptidoglycan complexed with teichoic acids and is used in preventing osmotic lysis of the cell protoplast. The cell wall of Gram-negative bacteria is composed of 10% peptidoglycan surrounded by phospholipid and protein-lipopolysaccharide and prevents osmotic lysis, confers rigidity and shape to the cell [72]. The outer membrane serves as a permeability barrier, and the associated lipopolysaccharide and proteins perform various functions. The plasma membrane is composed of phospholipids and proteins and provides a permeability barrier, generates energy for the cell and houses many enzyme systems. The ribosomes are sites for translation and protein synthesis and are made of RNA and proteins. The chromosomes contain the genetic material of the cell; the DNA. The plasmid is extrachromosomal genetic material and is also composed of the DNA. Bacteria biochemical components relevant for antimicrobial activity are: • Proteins [73] which are built from amino acids are contained in the flagella, pili, cell wall, cytoplasmic membrane, ribosomes, and cytoplasm. • Polysaccharides are formed by the combination of sugars and are found in the capsules, inclusions, and cell walls. • Phospholipids [74] are made from fatty acids and are in the membranes. • Nucleic acids are of two types, DNA [75] and RNA [76]. The DNA is located both in the nucleoid as chromosomes as well as in plasmids. The rRNA is in the ribosomes while the mRNA and the tRNA are situated in the cytoplasm.



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2.5.1  Outer cell layer composition of microorganisms The outer cell layer of microorganisms is essential for antimicrobial attack because it is the first point of contact with the antimicrobial surface or entity. The composition of cell layers of microbes differs from each other. The Gram-positive cocci (Staphylococci) has a cell wall composed predominantly of peptidoglycan, while the Gram-negative bacteria (E. coli, P. aeruginosa) have their outer membrane made mainly of phospholipids and lipopolysaccharides [77]. The Mycobacterium tuberculosis has a cell wall made of mycolate of arabinogalactan. Bacteria spores (Bacillus spp.) have their outer spore coats made of alkali-resistant S–S bonds, while their inner layers are composed of alkali-soluble acidic polypeptides, and the cortex is composed of peptidoglycan [78]. Algae (green Chlorophyta, brown Phaeophyta, and red Rhodophyta) have cell walls made of cellulose and other polysaccharides [79]. Each of these structures has molecules performing functions vital for cell survival which when tempered with compromises the functionality of such molecules in the biological system. The target of antimicrobial attack is to remove/replace metals performing cell survival functions depriving the cell of such functions and therefore strangulating the cell to death. If several molecules at different locations such as the cell membrane, DNA/RNA, and other inner biochemical components of the cell are attacked simultaneously and in multiples, the bacteria cells are unable to organize resistance and are rapidly killed or have their viability reduced. 2.5.2  DNA structure The DNA is a critical component of the cell because it instructs and coordinates all cell activities. The integrity of DNA is an essential part of cell survival because the molecule carries genetic information and guides vital biological processes such as transcription and replication of living cells [80]. The interaction between small ligands and DNA involves either nonspecific binding through electrostatic interactions with the negatively charged sugar-phosphate backbones, intercalation of the ligand’s planer aromatic rings between two adjacent base pairs or major or minor groove binding [81]. Most DNA interacting compounds stabilize the double helix, while a few destabilize it leading to various cellular consequences. Genetic information consists primarily of instructions for making proteins, which are the macromolecules responsible for executing cellular functions. Proteins, therefore, serve as building blocks for cellular structures, form the enzymes that catalyze all the cell’s chemical reaction, regulate gene expression, enable cell mobility, and promotes communication with each other [81]. The XRD analysis of DNA shows that it is composed of two strands wound into a helix. A DNA molecule has two long polynucleotide chains consisting of four types of nucleotide subunits. The two strings are held together by hydrogen bonds [79]. The nucleotides are composed of a five-carbon sugar to which one or more phosphate groups are attached and a nitrogen-containing base. The sugar in DNA nucleotides is deoxyribose connected to a single phosphate group, and the base may be either adenine (A), cytosine (C), guanine (G), or thymine (T) [16]. The bases are not randomly paired, but A always pairs with T and G with C [79]. A disruption or manipulation of the DNA of a cell such as that of a microbe can either be of advantage or disadvantage to the germ. It can be of benefit if it causes the germ to adjust and defend itself against external invasion. For example, mutations can make bacteria resistant to 



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antibiotics designed to kill them. The manipulation is disadvantageous if it compromises the security of the microbe. Antimicrobial surfaces and agents are developed with a deliberate intention of creating a non-conducive environment for microbe’s survival or virtually killing them. Disorganization or confusion is sponsored in the coordinating activities of the DNA so that it is unable to organize resistance and is subsequently not able to survive in the environment. The process of manipulating the DNA with an aim to compromise its security and finally killing it is technically referred to as DNA denaturing.

2.6  Mechanisms of antimicrobial activity of silver For centuries and as early as 1000 BC, silver has been known to have bactericidal or antimicrobial properties [19]. A detailed report of the historical review of the early use of silver to treat various conditions is contained in the literature [2,9]. The silver ion is highly reactive and readily combines to negatively charged proteins RNA, DNA, as well as chloride ions [19]. This property is at the heart of its antimicrobial mechanism. The mechanism of the antimicrobial behavior of silver is observed to be in three interactions namely: 1. the inhibition of transport functions in the cell wall which affects respiration, 2. interruption of cell metabolism by changing enzyme structure, and 3. inhibition of cell division in its interaction with DNA. This multiple attack of silver ions on microbes is very effective and unique and reduces the possibility of resistance development by bacteria [18]. The high affinity of silver toward sulfur and phosphorus is a crucial element of the antimicrobial efficacy of silver and its inhibitory effect on bacteria. There is an abundance of protein containing sulfur on the bacteria cell membrane, and when silver reacts with the sulfur containing amino acids inside or outside the cell membrane, it affects the bacterial cell viability [14]. Silver doped materials are chemically durable and release silver ions for a long time. They are therefore a favorable candidate for antimicrobial applications. 2.6.1  Evidence of silver ion attack on Gram-positive and Gram-negative bacteria cell membrane A mechanistic study of the inhibition of silver ions against two different bacteria species (S. aureus and E. coli) was reported by Feng and coworkers [82]. The morphological changes in the cell walls after treatment of the bacteria with silver ions were studied using transmission electron microscope and X-ray microanalysis. Significant morphological changes were observed in the E. coli after exposure to silver ions. An electron light region was seen in the center of the cell that contained a tightly condensed substance twisted together. A significant gap was also found between the cytoplasmic membrane and the cell wall. Also noticed was the presence of some electron-dense granules around the cell wall, and an X-ray microanalysis on the pellets indicated the presence of silver and sulfur presupposing that the silver ions after entering the bacteria cell might have combined with the cell components containing sulfur. The presence of the condensed substance was also observed in the electron-light region of the S. aureus. The cytoplasmic membrane shrank and detached from the cell wall.



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A significant amount of phosphorus was also found in the condensed area of the S. aureus. The electron-dense granules observed in S. aureus and the electron-light region was darker than E. coli cells. The S. aureus bacteria have a much stronger defence system compared to E. coli. The S. aureus is a Gram-positive bacterium with a thicker peptidoglycan cell wall than the E. coli, and there is a visible nuclear region in the center of the cells where DNA molecules are distributed randomly. The thicker cell wall protects the cell from penetration by silver ions in the cytoplasm [32]. These reports indicate that when silver ions are released and interact with bacteria, morphological changes as well as the release of metals such as sulfur and phosphorus which are very vital for bacteria survival. 2.6.2  Evidence of silver ion attack on DNA (DNA denaturing) The silver ions enter the bacteria cells by penetrating through the cell walls and consequently turn the DNA into condensed form and reacts with the thiol group in proteins resulting in cell death. The silver ions also interfere with cell replication process. In a research reported by Martinez-Castanon and coworkers [83], the mechanism of action of silver ions is linked to its interaction with thiol group compounds found in the respiratory enzymes of bacteria cells. Silver ions attack E. coli by inhibiting the uptake of phosphate and releasing phosphate mannitol, succinate, proline, and glutamine from E. coli cells. The effects of the silver ions on bacteria are indicated by the structural and morphological changes. Replication ability is active when DNA molecules are relaxed but is lost when DNA cells are condensed, and this happens when silver ions penetrate the bacteria cell wall [83]. This penetration subsequently results in the death of the bacteria cell. The double helix of a DNA is held together by two significant molecular interactions, the hydrogen bond between the complementary bases and the stacking of the hydrophobic nitrogen bases down the center of the helix. The electrostatic repulsion between charged sugar-phosphate backbones of the DNA, on the other hand, opposes the stabilization interactions. DNA denaturing involves the breaking of the hydrogen bonds holding the double helix together so that the strands are separated from each other [81]. DNA denaturing is achieved by raising the pH value through the addition of bases like NaOH until the H+ shared between N-base electronegative centers is striped from the hydrogen bond [63]. A competition for the formation of hydrogen bonds by chemical compounds containing functional groups such as urea and formaldehyde with electronegative centers of the N-bases in which the denaturants are favored instead of the complementary bases also denatures DNA. Denaturation of the DNA also occurs when the DNA covalent electronegative centers of the N-bases are modified by aldehydes and glycoxals by blocking the formation of hydrogen bonds between complementary bases. 2.6.3  Attack on proteins Protein groups are also inactivated when heavy metals react with proteins by attaching to the thiol groups. Silver ions react with sulfur containing protein leading to the inhibition of enzyme functions [32,58,84–86]. Silver nanoparticles ions get attached to sulfur containing proteins of bacteria cell membrane leading to greater permeability of membrane subsequently causing the death of bacteria [46,87]. The oxides of other metals have been reported to exhibit antimicrobial properties. Such oxides as copper oxides, zinc oxides, and titanium oxides are said to have antimicrobial properties [88]. The oxides of copper, a transition metal in the same group as silver exhibit antimicrobial behavior and this have been well established. It is justifiable to expect the oxides of silver 



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to exhibit antimicrobial activity. A few years ago, some oxides of silver had been reported to exhibit antimicrobial activity [89]. The oxides of silver are reportedly volatile readily dissociating to organic silver, silver ion, and oxygen [17].

2.7  Photocatalysis and antimicrobial activity on the surfaces and generation of reactive oxygen species A photocatalyst is a material that accelerates a photoreaction when a photon of light energy in the visible or ultraviolet spectrum of the electromagnetic radiation is absorbed. The excitation leading to photocatalysis can initially occur in the adsorbate molecule resulting in catalyzed photoreaction, or it may be initiated in the catalyst substrate and the catalyst photoexcited transfers an electron or energy to the ground state in which case it is named sensitized photoreaction. Chemical reactions leading to deactivation through electron or energy transfer occurs in heterogeneous photocatalysis. The electron transfer takes place by one-electron reaction during which an electron jumps from an occupied orbital of a donor reactant to the empty orbital of the acceptor reactant. The initial excitation can be either in the donor molecule or the acceptor molecule. The energy transfer occurs when there is an electron transfer or a dipole–dipole resonant coupling. According to perturbation theory, the probability of an electronic transition is proportional to the square of the amplitude of the radiation field and the square of the transition dipole moment and is expressed as [90]: 2

(12.2) P = Eo2 µif where µif is the transition dipole moment of an electron excitation process from an initial to a final state and Eo is the radiation field amplitude. The radiation field can be manipulated by varying the intensity of light, but the dipole moment cannot, because it is an intrinsic property determined by the molecular structure. The transition dipole moment is a product of the electronic transition moment, the electron spin wave function overlap, and the nuclear wave function overlap and is expressed as [91]: (12.3) µif = φ f|µ|φi = ϕ f|µ|ϕ i + ξ f|ξi + ℵf|ℵi where F is the molecular wave function, ϕ is an electronic spatial wave function, ξ is the electronic spin wave function and ℵ is the nuclear wave function. A photocatalyst engages in the transformation of the reactant partners after absorption of a photon. When excited, a photocatalyst repeatedly interacts with the reactant partners producing reaction intermediates and is itself not consumed but regenerated after the cycle of interaction [92]. Every catalyst has a lifespan which is affected or determined by several factors. Photocatalysis is categorized into two homogeneous and heterogeneous photocatalysis. In homogeneous photocatalysis, the reactants and the photocatalyst are in the same phase, while in heterogeneous photocatalysis, the reactants and the photocatalyst are in different stages and the reaction takes place at the interphase. Ozone and the Photo-Fenton system are the most common homogeneous photocatalysis, and the reactive species functional in them are the hydroxyl radicals (HO•). These reactions represent the generation of hydroxyl radical by ozone: O 3 + hν → O 2 + O (12.4) 

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(12.5) O + H 2 O → HO• + HO• (12.6) O + H2O → H2O2 H 2 O 2 + hν → HO• + HO• (12.7) The equations give the production of hydroxyl radicals by the Photo-Fenton system: (12.8) Fe 2+ + H 2 O 2 → HO• + Fe 3+ + OH – Fe 3+ + H 2 O 2 → Fe 2+ + HO•2 + H + (12.9) (12.10) Fe 2+ + HO• → Fe 3+ + OH – H 2 O 2 + hν → HO• + HO• (12.11) Fe 3+ + H 2 O + hν → Fe 2+ + HO• + H + (12.12) Energy can be supplied to electrons in the valence band and used to activate them to move across the energy gap into the conduction band in semiconductors. When a semiconductor photocatalyst is activated by its absorption of ultraviolet or visible light equal to or higher in energy than its bandgap, an electron initially in the valence band absorbs the energy and is promoted to the conduction band leaving a hole in the valence band. Each absorbed photon of energy produces an electron-hole pair and the electrons and holes are engaged in oxidation and reduction reactions on the surface of the semiconductor. The charge carriers either recombine with the bulk of the material or migrate to the particle surface where they can recombine or be trapped at a defect site as substrate bound O− radicals. The electron-hole pair initiates oxidation and reduction reactions on the adsorbed substrates so that in aqueous solutions the holes are scavenged by surface hydroxyl groups to generate strong oxidizing hydroxyl radicals which promote oxidation in line with the equations [93]. The quantum yield of a photocatalyst is the number of specific primary products such as radicals, photo-excited molecules ions released for each photon absorbed. The combined number of reactant molecules resulting from the absorption of a photon is the overall quantum yield of the photocatalytic process and can be as high as 104. Quantum yield (Φ) is dependent on the incident radiation. The quantum yield in a photocatalytic system is the number of molecules destroyed divided by the number of absorbed photons in the system and is expressed as [94]:  Amount of product formed  Φ= 100 (12.13)  Amount of photons absorbed  The photocatalyst should be stable, that is, to be able to go through several cycles before deactivation. The deactivation of a catalyst occurs due the presence of impurities, recombination of electron-hole pairs, thermal deterioration, volatility, and hydrolysis of active





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components [94]. A stable photocatalyst must have a large surface area available for reaction. A large surface area is assured in thin films with porous structures such as those found in columnar growth mode in SEM imaging. Reactions in a heterogeneous photocatalyst such as a thin film semiconductor take place at the interfaces (solid-liquid, solid-gaseous) which provides sites for the absorption of both reactants and photon energy. The primary reactions in photocatalysis are oxidation and reduction involving the electrons and holes produced by the photo-excited semiconductor. The electrons and holes generated either migrate to the surface of the particle or recombine and cancel out. The electron in the conduction band can be transferred to an electron acceptor if the acceptor has a relatively positive electrochemical reduction potential compared to the potential of the conduction band edge. The hole in the valence band accepts an electron from a donor species with a less positive electrochemical reduction potential than the valence band edge potential. The acceptor species are therefore reduced, while the donor species are oxidized, and both reactions are driven by the potential difference generated due to the absorption of electromagnetic radiation. The primary concern in photocatalytic disinfection is the generation of ROS which is achieved via redox reactions. In the redox reaction, oxygen is an electron acceptor, while hydroxyl ion and water are electron donors. The holes in the valence band oxidize water yielding hydroxyl radical if their electrochemical reduction potential is positive enough. Similarly, the conduction band should be negative sufficiently to reduce oxygen to superoxide radical anion, and subsequently, electron transfers yield hydroxyl and peroxide radicals. Generally, the photocatalytic mechanism in the presence of water and oxygen results in the generation of ROS, with a capacity to inactivate microorganisms as well as degrade organic chemical contaminants among many other applications. The reactions leading to the formation of radicals used for antimicrobial activity of photocatalysts are summarized by the equations [95]: (12.14) Semiconductor + hν = e − + h + h+ + H 2 O → HO• + H − (12.15) O 2 + e − → O•2− (12.16) O•2− + H + → HO•2 (12.17) (12.18) HO•2 + HO•2 → H 2 O 2 + O 2 (12.19) O•2 + HO•2 → O 2 + HO −2 (12.20) HO −2 + H + → H 2 O 2 (12.21) H 2 O 2 + hν → 2HO•



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(12.22) H 2 O 2 + O•2 → HO• + OH − + O 2 H 2 O 2 + e − → HO• + OH − (12.23) The transition metal oxides exhibit heterogeneous photocatalysis, and the reactions have a wide variety of applications which include mild oxidation, total oxidation, dehydrogenation, hydrogen transfer, metal deposition, water detoxification, and gaseous pollutant removal. The best semiconductor materials suitable for photocatalysis should have a reasonable bandgap preferably in the visible range of the electromagnetic spectrum, an active carrier transportation to prevent carrier recombination, should structurally be crystalline to ensure good carrier transport and reasonable bandgap, have a large surface area to provide a high number of active sites and be stable so it can be used severally before losing its lifespan. 2.7.1 Visible light activated photocatalysts Over the years, titanium dioxide (TiO2) nanoparticle thin films have been studied and established as active photocatalytic antimicrobial material and are activated using the ultraviolet-visible range of the electromagnetic spectrum [96]. Titanium dioxide, when exposed to a radiation of frequency corresponding to its bandgap, is an active antimicrobial. The need to generate ultraviolet radiation to activate titanium dioxide as an antimicrobial material is both expensive and limiting in applications, hence the need to develop a biomaterial that can be activated using visible radiation. The adsorption of molecular oxygen on the Ti (III) sites reduces the oxygen to a superoxide radical anion, while the positive charge carrier oxidizes the surface hydroxyl groups or the surface-bound water to surface-bound hydroxyl radicals. The trapping of the charge carriers by water and oxygen molecules serve to suppress the electron-hole combination, thereby increasing the competing ability of the light-induced redox processes. The hydroxyl radicals are responsible for the oxidation observed in photocatalytic degradation in TiO2. Illuminated TiO2 photocatalysts decompose organic compounds by oxidation with holes generated in the valence band and with hydroxyl radicals (produced by the oxidation of water). This photocatalytic oxidation causes damage to microorganisms, which also consist of organic compounds [97]. The reaction can be propagated if the light source has a rich source of UV component (250–300 nm) [98,99]. Hydrogen peroxide adsorbs this wavelength and its interaction with the photo-generated electrons yields hydroxyl radicals. The role of hydrogen peroxide is that of generating the reactive species. The effect of pH is examined to determine the contribution of the reductive photocatalytic pathway; oxygen is reduced to superoxide anion which is less reactive but produces more highly toxic species [96]. If the silver oxide is engineered to have a bandgap of a value lower than that of TiO2, it can equally be photocatalytic within the visible—UV spectrum. Bandgap values reported in the literature for silver oxide range from 1.2 to 3.4 eV. The primary target in the design of photocatalytic materials is to activate then using the visible spectrum of the electromagnetic radiation. This is achieved by developing a narrow bandgap or modification of the bandgap of the existing materials. There is improved performance of photocatalytic materials when activated in the visible spectrum because there are





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a higher number of photoelectrons produced and the material is activated under ambient lighting conditions. Research aimed at developing visible light activated photocatalysis was initiated by using binary materials such as WO3Fe2O3, which can produce reactive oxygen under visible light irradiation [1] and can disinfect [100]. Many of these materials, however, exhibit photo-corrosion when irradiated with visible light [101,102]. The range of bandgap for these materials is 1.8–2.5 eV and includes materials such as Fe2O3, SiC, and Cu2O. Materials with wider bandgaps such as WO3 and MoO6 which have bandgaps in the range 2.6–2.8 eV are more stable and show an improvement in visible light activation [103]. The search for a monolithic, more efficient, and durable visible light activated photocatalyst is still on, and design and combinational chemistry are the main techniques engaged by researchers. Transition metal elements which have electrons in the d orbitals and which absorb strongly in the visible spectrum due to d-d transitions are currently used in designing visible light activated photocatalysts. Materials such as CuFeO2 [104], LaFeO3 [105], BaCr2O4 [106], NiCo2O4 [107], and CuMn2O4 [108] are reported to be photocatalysts designed using transition metals. Another group of photocatalysts have been designed using vanadium because vanadium has a bandgap of less than 1 eV and when nanostructured forms a bcc phase with an optical bandgap about 2.7 eV and is a photocatalyst of very high quantum efficiency [109]. It can also be combined with other elements to design other photocatalysts such as InVO4 [109] and BiVO4. It is further reported that better results are obtained when silver co-catalyst is used on vanadium. Photocatalyst stability has been used as a design factor to develop another set of photocatalysts. The occupation of the d orbitals determines the stability of the materials with d0 and d10 transition elements remaining stable during the reaction and exhibiting no photocorrosion [98]. Another class of photocatalysts been investigated are the two metal-layered oxides with perovskite structures such as SrTiO3 [77], CeCo0.05TiO0.95O3.97 [78], and AgNbO3 [110]. The morphology and bandgap of deposited thin films can be tuned by manipulating deposition conditions such as forward deposition power, the temperature during deposition, deposition rate, and oxygen flow rate [111]. This information can be used to produce thin films with bandgaps within a given range. If the bandgap of silver oxide thin films is tuned to be in the visible range of the electromagnetic spectrum, visible light can be used to activate it as a photocatalyst. The different oxides of silver give rise to a wide range of band gap (between 1.2 and 3.4 eV) [49,64,66]. The rf magnetron sputter decomposition of the oxides of silver makes silver and oxygen ion radicals available for interaction with other materials such as bacteria brought in contact with it. The interaction informs the antimicrobial behavior of the oxides of silver. Electronhole pairs are produced and used to sponsor the production of superoxides and hydroxyl radicals as well as hydrogen peroxide which attack the cell walls, protein, and DNA of microbes denaturing them and thus compromising the security of the bacteria and subsequently killing them [67,68]. Irrespective of deposition conditions and annealing, silver oxide remains unstable and decompose into metallic silver and oxygen either wholly or partially [6] and are abundantly available on the surface. Silver oxide has different stoichiometry and can be deposited using different methods. Ag2O and Ag4O4 are reported to exhibit antimicrobial behavior [82,112].



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12.  Visible light activated antimicrobial silver oxide thin films

2.7.2  Nanostructured modification for photocatalytic activation in the visible spectrum Dye [113] and Plasmon [114] sensitization of photocatalytic materials with a deliberate intention of activating them using visible spectrum has proved to improve the photocatalytic efficiency of materials. The ROS generation in these materials is done through oxygen reduction reactions brought about by electrons in the conduction band of the semiconductor. Surface modification via the introduction of a second semiconductor to the existing photocatalyst promotes charge carrier separation [115], charge carrier sink [116] as well as the promotion of specific reactions [117] in addition to sensitization. Surface sensitisers are used in nanostructure design to modify the bandgap with the aim of increasing photoactivity [118]. It is important to note that strongly polarizing species can affect the lattice structure of a material resulting in a rise in the valence band [119]. The most commonly reported approach aimed at visible light activation of the photocatalyst is that of doping a known photocatalyst such as TiO2. Photocatalysts can be improved to be activated in the visible spectrum by doping with a metal. Examples of metals used include Ag [120], Ce [121], Co [122], Pb [123], and Pt [124]. The materials are prepared using sol-gel [120], microwave [125], or combustion [126] and of these, which are produced through sol-gel prove to be more effective in photocatalytic disinfection. Metal ion dopants have been extensively studied but reports indicate about the same levels of claims of enhancement and reduction in activity of the doped photocatalysts [127]. Photocatalytic disinfection has been investigated on TiO2 materials doped with different elements. Common elements used include Cu, S, C, N, B, the halogens [104,128,129]. Generally, codopants are reported to have higher rates than single regimes [130]. Second generation photocatalysts such as WO3 also offer a good improvement on visible light activation when doped and can be used for the development of new photocatalytic disinfectant materials [131]. Photocatalytic coatings are used for self-disinfection by reducing the risk of infection transmission through environmental surfaces as well as decontamination and disinfection of medical devices. The photoexcitation of some semiconductors produces ROS that are capable of killing or inactivating bacteria and other pathogenic microorganisms. Bacteria were reported to be inactivated by a photocatalyst using TiO2 by Matsunaga and coworkers in 1985 [132]. Since then, several types of research have been conducted establishing the effectiveness of photocatalytic materials against bacteria [133,134], spores [135], protozoa [136], algae [137], viruses [138], and fungi [139]. 2.7.3 The use of photocatalysts as antimicrobials The interaction of biological structures with ROS within microorganisms have been intensively investigated to unravel the mechanism of bacteria inactivation and it is expected that complete understanding of the mechanism will pave the way for researchers to optimize the design of materials and reactors to improve the rate and efficacy of photocatalytic disinfectants [140]. A photocatalytic semiconductor when excited produces a range of ROS at the interface between the particle and the liquid. The hydroxyl group (HO•) is the primary species responsible for the inactivation of microbes. Also, the superoxide radical (O2•−), the hydroperoxyl radical (HO2•), and hydrogen peroxide (H2O2) contribute in the process [141]. The attack of





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ROS on microbes is not restricted to one site or individual pathway, unlike antibiotics which target a biological process within the lifecycle of the microorganism. The ROS generated from photocatalysis can, therefore, be used against a wide range of pathogens with an almost impossibility of developing resistance [142,143]. It is required that the ROS be produced in concentrations higher than that which the pathogens can protect themselves to achieve complete inactivation thereby preventing pathogen recovery and re-growth. Primary contact with microorganism with ROS takes place outside the organism, toward the sensitive metabolic processes as well as in the genetic materials within the organism. The composition of the outer layers of the organism determines the extent of the effect of the ROS on the microbes. The thick protein, carbohydrate, and lipid structures surrounding protozoa and bacteria spores offer more resilience to ROS attack compared to viruses, fungi, and bacteria [144]. The cell cytoplasm is bound by cell membrane which houses a phospholipid bilayer containing cross membrane protein structures responsible for the regulation of chemical transmissions into and out of the cytoplasm. The maintenance of the integrity of this structure is of much importance for bacterial survival and viability. Bacteria survival is threatened when the integrity of this structure is compromised. Bacteria are killed or inactivated during prolonged exposure to ROS in a sequence that is initiated by attack and damage of the cell wall, leading to the compromise of the cytoplasmic membrane and finally a direct attack on the intercellular components. Several microscopy studies have been conducted to study the effects of exposure of microbes to ROS. Wu and coworkers [145] conducted a study that revealed the formation of pores within the cell wall and cell membrane structure of bacteria when exposed to ROS, while Kiwi and coworkers reported the degradation in the peptidoglycan [146], a degradation of pyrin was reported by Kiwi and coworkers [146] within the cell membrane. Several reports indicate that lipid peroxidation takes place within phospholipid membranes on exposure to superoxides [146–150]. Matsunaga and coworkers [124] and Rincon and coworkers [151] confirmed the detection of intercellular and genetic components in the exterior, while direct damage to DNA was reported by Pigeot coworkers [152]. The inactivation of the respiratory track chemistry due to membrane damage was reported by Kiwi and coworkers [146] which can introduce fluid loss as well as allowing ion permeability [150]. A detailed systematic study using levels of genetic and protein biomarkers on TiO2 photocatalysis was conducted by Kubacka and coworkers [37] relating hydroxyl radical-mediated lipid peroxidation of the outer cell wall components. The researchers observed radical cell wall modifications as the primary factor responsible for the high biocidal performance of TiO2 based photocatalytic material. In another research report, Malato and coworkers [144] observed that the irradiation of intercellular chromophores in the presence of oxygen denatures DNA and introduce modifications in nucleic acids. The hydroxyl radical produced in Fenton reaction is reported to have damaged DNA during bacteria recovery [150]. Antibiotic resistance has been reported when a sub-lethal treatment with the photocatalytic material is administered on pathogenic samples [152], and the recommendation from several researchers is that complete inactivation of organisms to prevent bacterial re-growth as well as specific analysis establishing or confirming treatment should be carried out [151]. Photocatalytic coatings have been used to investigate the inactivation of clinically relevant pathogens by Dunlop and coworkers [154]. The researchers developed a method to assess the efficacy of photocatalytic surfaces and determine pathogenic viability as a function of



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12.  Visible light activated antimicrobial silver oxide thin films

treatment time as well as an assessment of the surface for viable surface bound organisms after disinfection and finally measuring the re-growth potential of activated microorganisms. The research was conducted using MRSA, E. coli, P. aeruginosa, C. difficile spores, and extended-spectrum Beta-Lactamase. The result obtained indicate 99.9% reduction in viability for bacteria cells within 80 minutes of photocatalytic treatment using UVA radiation with complete surface inactivation and bacteria re-growth after treatment did not occur. About the same percentage inactivation was observed with C. difficile spores. Three factors influence the efficacy of immobilized photocatalyst in inactivating bacteria. These are: • The penetration ability of radicals into the microbial cells. • The ease with which light and oxygen access the surface of the photocatalyst. • The distance between the surface and pathogens. A team of researchers Dunnill and coworkers [153] used white light to illuminate a material on which sulfur and nitrogen were doped into TiO2 using atomic pressure chemical vapor deposition at different percentage compositions to test the efficacy of bacteria inactivation. The absorption edge for the sulfur and nitrogen incorporated TiO2 shifted to 3.0 eV and 2.9 eV, respectively from the normal anatase TiO2 bandgap of 3.2 eV. The samples were pre- and post-irradiation with visible light for 24 hours before and after E. coli was brought into contact with the surfaces. The control was set up using TiO2 illuminated with visible light and only glass substrates. The results obtained indicated 99.9% kill on the nitrogen-doped at 0.13% to TiO2 compared to the 99.5% observed on the sulfur doped at 0.1% to TiO2 surface. No killing was observed on the un-doped TiO2 and the glass substrates indicating that visible light is unable to activate TiO2 as a photocatalyst. Dunnill and coworkers [1] again investigated the ability of nanoparticulate silver loaded Titania thin films of killing bacteria under hospital lighting conditions. The films were produced using the sol-gel technique with anatase TiO2 on glass slides. The silver was introduced via UV photo-reduction of AgNO3 onto the anatase TiO2. The samples were then annealed at 5000C. The material has an optical absorbance bandgap of 2.8 eV. The purpose of the research was to find out if the coatings are active under visible light and subsequently combine the killing effect of silver ions and photocatalytic effect to achieve faster killing of the microbes under hospital lighting environment. Silver ions attack bacteria and as they move from antimicrobial materials into microorganisms, the cells get damaged. Epidemic Methicillin-resistant S. aureus (EMRSA-16) and E. coli were used to test the antimicrobial effect of the coatings. The inactivation of the microbes was significantly observed with the Ag-TiO2 films under white light radiation compared to anatase TiO2 film. Up to 99.996% of E. coli decreased in viability on exposure to Ag-TiO2 under white light for 6 hours compared to exposure to TiO2 under same conditions. A similar decrease in bacteria population was observed in the Ag-TiO2 for E. coli in the dark indicating that the inactivation was due to Ag ions in the dark. 2.7.4  Nanoparticle photocatalyst semiconductors Nanoparticle semiconductors have properties suitable for exploitation in the design of photocatalysts. They have high surface areas which provide more active sites for catalysis thus making room for faster reactions. The morphology of the nanoparticle can easily be tuned using deposition conditions. The bandgap increases as a material moves from its bulk





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to nanoparticulate form. The presence of surface states on nanoparticulate semiconductors is also advantageous. The positions of the valance and conduction band edges can be manipulated using the particle size variation controlled by deposition conditions. Titanium dioxide (TiO2) is the most researched and utilized nanoparticle semiconductor used for photocatalytic applications and accounts for 36.2% of total usage, while other metal oxide semiconductors account for 10.9% [96]. TiO2 is most widely used because it is commercially available, useful and cheap, has chemical and photochemical stability even under harsh conditions and is non-toxic, exists in different allotropic forms with high photoactivity, can be coated as a thin film, can be readily prepared in the laboratory etc. [155]. The bandgap of TiO2 is between 3.2 and 3.4 eV and absorbs in the ultraviolet spectrum of the electromagnetic radiation and as such can be activated as a photocatalyst using ultraviolet radiation. Metal sulfide semiconductors are known to absorb in the visible region of the electromagnetic spectrum but undergo photo-anodic corrosion [156]. The main issue associated with the use of TiO2 as a photocatalyst is that it is activated using ultraviolet radiation which is only 3%–5% of the total solar radiation reaching the earth compared to visible radiation’s 43% solar energy reaching the earth. Many modifications have been employed to improve on the performance of TiO2 as a photocatalyst. These modifications include the use of impurity doping [157–159], deposition of metals [160–162], and dye sensitization [160,161] as well as deposition of metal oxides [162]. TiO2 based mixed oxides such as TiO2/In2O3, TiO2/SiO2, TiO2/ZrO2, Pt/TiO2, Rh/TiO2, Ru/ TiO2 are also very good photocatalysts [98]. When a metal such as silver is introduced into a photocatalyst, it both alters the structure and mode of action of the photocatalytic activity as well as serves as an electron trap for the photo-activated electrons in the conduction band. This alteration separates the electrons from the holes located in the valence band thereby reducing the tendency of electron-hole recombination. The electrons can then be transferred to molecular oxygen to produce superoxide and subsequently other ROS. The introduction of silver in the TiO2 subject to quantum size effect may introduce surface plasmon resonance (SPR) resulting in increased visible light absorption by the photocatalyst. The collective oscillations of the electrons in SPR serve as an antenna attracting more visible radiation for subsequent absorption by the photocatalyst. Other attempts made to produce visible light activated photocatalysts were made using non-metals such as carbon and nitrogen [163]. Silver is reported to be antimicrobial and has been used over the years in the fight against pathogenic bacteria. Silver is also reported to be less cytotoxic compared to copper and its compounds. Silver has an interesting chemistry that can be tailored toward developing coatings with peculiar properties for specific applications at the nanoscale such as antimicrobial activity. The microstructure and the bandgap of thin films can be tuned using deposition parameters such as oxygen flow rate, forward deposition power, and substrate temperature during or after deposition. Photocatalysis in semiconductors produces electron-hole pairs and initiate redox reactions producing ROS and other radicals that launch simultaneous and multiple attacks on bacteria resulting in effective contact killing of the microbes. Silver oxides have been deposited and used as antimicrobial using chemical reactions, ion beam evaporation, and pulsed laser deposition. In this report, rf magnetron sputtering has been engaged to tune the microstructure and bandgap of silver oxides to produce visible light activated monolithic photocatalytic thin films with excellent antimicrobial activity.



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12.  Visible light activated antimicrobial silver oxide thin films

3  Materials and methods 3.1  RF sputtering and silver oxide thin film coating Two parameters-oxygen flow rates and forward rf power were varied to tune the morphology and bandgap of the thin films using a cryo-pumped vacuum chamber rf magnetron sputtering unit to coat the films. Several advantages attract thin film producers to reactive magnetron sputtering. Such advantages include in-situ cleaning before deposition which is achieved by reversing the electrode potential. High-quality films are deposited at low temperatures such as room temperature. It is a cheap and versatile technique for deposition of alternate layers of different materials with easy control of refractive index and thickness. The deposition technique ensures better adhesion of films to the substrate. The high kinetic energy of the sputtered atoms causes a good re-distribution on the substrate leading to high uniformity, density, and interfacial roughness of the deposited films. The film growth process can easily be controlled by optimizing a few selected deposition parameters such as base vacuum pressure, the sputtering gas pressure during deposition, sputter power, substrate and target temperature, and reactive gas pressure. The interplay of the parameters determines the quality of the deposited thin films regarding its microstructure such as surface roughness, adhesion, the presence of impurities, and density of deposited films. Reproducible film deposition control is also easily achieved through timing for same deposition parameters and at the same rate contrary to what obtains in evaporation techniques [164–168]. These advantages informed the adoption of this technique for thin film deposition. 3.1.1 Substrates The substrates used for the deposition of the thin films were clear glass microscope slides 1–1.2 mm thick and 25.4 mm by 2 mm produced by CiMED, and double polished silicon wafers of (100) orientation, n-type dopant and of resistivity 20 kHz are used resulting in ultrasonication. This is carried out normally in a sonicator. Sonication is a chemical and physical process emanating from acoustic cavitation process resulting in the formation, growth, and implosive collapse of vesicles [7]. Sonication was used in this study to break apart large liposomes into smaller ones. They can be converted from several micrometers to less than 100 nm, which is in the nanoscale range. Sonication overcomes intermolecular forces between the particles and it disperses the liposomes uniformly in the solvent resulting in the formation of an emulsion. The particle diameter is a key parameter in pulmonary delivery to determine whether the particles reach the alveoli [8]. Generally, for spherical particles, the smaller the diameter the further down the particle reaches in the pulmonary gas exchange system. Particle size should be less than 1–2 µm to reach the deep lungs with particles in the nanosize range being preferable. This has been determined using a twin-impinger, which is widely used as a model artificial lung.





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However, a compromise needs to be considered carefully. The larger sized liposomes can entrap a higher amount of anti-asthma drug while finding it more difficult to reach it to the deep lungs due to physical limitations. However, smaller liposomes in the nanosize range will entrap lower amount of antiasthma drug but can readily be targeted to the deep lungs where it is needed. Hence, it is essential to determine and study the size effects on the delivery of the vesicles. The size of the empty and drug-filled SPC liposomes, was analyzed using Malvern mastersizer and zetasizer. The sizes of liposomes formed depend on the preparation method used in addition to the formulations used. As mentioned, previously, probe sonication was used to break up large liposomes into smaller nanosized vesicles. The sonication process was repeated several times until the desired size was achieved at intervals of 1 min to prevent overheating of liposomes, which could disrupt the liposome structure. The effect of sonication on the size of liposomes particles without the incorporation of the anti asthma drug is shown in Fig. 14.1. Clearly the figure shows that sonication deceases the diameter of liposomes even after a few minutes. A size reduction of ×70 is readily achieved after 5 min. There is little change evident between 5 and 6.5 min. When liposome formulations were made without sonication the diameter was about 7.2 µm with a standard deviation (δ) of ± 0.34. Fig. 14.1 shows sonication to be an extremely effective technique for reducing the size of the liposomes. When these liposomes were sonicated for 6.5 min and the average diameter decreased to 110.67 nm with δ ± 2.12. Sonication involves the application of ultrasound usually at frequencies >20 kHz to the solution of liposomes. The shear forces generated break up the liposomes into tiny nanoparticles; in the case of liposomes, to nanosized particles that have important applications in drug delivery systems, particularly for pulmonary delivery. A large drop occurred in size between 0 to 5 min therefore this time domain was investigated in more detail due to a major effect on size reduction. The effect of sonication after every minute interval on the liposome size was investigated. The results are shown in Fig. 14.2. In the time range from 1 to 5 min a linear decrease is observed in the size of liposomes was observed. The trend line shows a negative correlation between the average size and

FIGURE 14.1  Effect of sonication on the size of liposomes without anti-asthma drug BDP (n = 3, p > 0.05, ±SD).

FIGURE 14.2  Effect of sonication time on the average size of empty SPC liposomes.



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14.  Characterization of cochleate nanoparticles for delivery of the anti-asthma drug beclomethasone dipropionate

sonication time as expected since sonication is well known for breaking down larger liposomes into smaller liposomes. The size decreased from 220 nm after 1 min to 120 nm after 5 min. The rate of decrease in size was calculated to be 25 nm/min. Liposomes are spherical particles whose internal volume is related to its radius. There is an obvious relationship between the amounts of drug that can be trapped theoretically within the volume of the liposomes. As the volume of a sphere increases the amount of entrapped is expected to increase. Hence, for liposomes the internal compartment contains the aqueous phase. In this space hydrophilic drugs can be encapsulated and then delivered to the site where they are required. The liposome walls protect the drug from the environment. Therefore, a linear relationship between the amount of hydrophilic drug and volume is expected. Fig. 14.3 is an idealized representation of liposomes, which enable calculations to be made, and relationships to be predicted. In practice liposomes have a bilayer on the outside instead of a single layer. They also contain multiple walls in multilamellar type of liposomes. It is expected that as the number of walls is increased the rate of drug release will decrease. Therefore, for a lower release rate multilamellar liposomes are preferred over unilammellar liposomes. Conversely, if faster drug release is required unilamellar liposomes are preferable. Let’s consider the sphere as representing a liposome vesicle. The volume of a sphere is related to radius is given by V=

4 π r3 3

The size ranges from approximately 7000 nm before sonication to about 100 nm after sonication and the volume varies with the radius as shown in Fig. 14.4. The sphere shown represents large unilamellar vesicles and hydrophilic drugs are trapped with the central fluid whose volume is related to the radius. In practice there is a size range, which is amenable to pulmonary drug delivery. The size of the particles determines the amount of entrapped drug reaching the lungs [9]. Liposomal vesicles about approximately 2 µm will not reach the deep lungs even if they are carrying a higher volume of entrapped drug whereas smaller vesicles can reach the deep lungs but carry a lower

FIGURE 14.3  Sphere with r representing the

FIGURE 14.4  Theoretical variation of the volume

radius.

with the sphere radius.





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volume of entrapped drug. Smaller amount of targeted anti-asthma drug have a better therapeutic effect requiring less frequent dosing. The volume increases exponentially with radius as shown in Fig. 14.4. This theory applies well to hydrophilic drugs, which are entrapped into the aqueous core in the middle. Many drugs are hydrophobic and get entrapped in between the hydrocarbon tails in between the phospholipid bilayer with the phosphate heads being exposed to the aqueous phase due to their hydrophobicity. Hence, the important region of the concentric biosphere is the region in between two concentric spheres for hydrophobic drug such as BDP. The volume of space (Vs) between two concentric spheres is given by the equation Vs = 4/3 π r23 − 4/3 π r13 This equation can be rearranged as follows:

(

Vs = 4/3 π r23 − r13

)

This simplifies to Vs = 4/3 π ∆r 3 where ∆r 3 = r23 − r13 Therefore, the larger the sphere the greater the entrapped volume between the two spheres. Another factor that needs to be considered is the hydrocarbon chain length i.e., the hydrophobic group. Generally, the larger the chain length the greater the potential to entrap hydrophobic drug in between the layers. The length of the hydrocarbon determines the difference the value of r23 − r13 . The carbon chain length determines is the most important factor determining the value of ∆r. This project focuses on the application for the delivery of anti asthma drug. Therefore, it is also important to understand the effect of sonication on the drug-filled liposomes. A formulation of liposomes was also made using the drug beclomethasone dipropionate (BDP) and the effect of sonication on the size studied (Fig. 14.5).

FIGURE 14.5  Effect of probe sonication on the size of DBP drug-filled SPC liposomes (n = 3, p