Degradation Rate of Bioresorbable Materials: Prediction and Evaluation

Degradation rate of bioresorbable materials

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Degradation rate of bioresorbable materials Prediction and evaluation Edited by Fraser Buchanan

Woodhead Publishing and Maney Publishing on behalf of The Institute of Materials, Minerals & Mining CRC Press Boca Raton Boston New York Washington, DC

Cambridge England

Woodhead Publishing Limited and Maney Publishing Limited on behalf of The Institute of Materials, Minerals & Mining Woodhead Publishing Limited, Abington Hall, Granta Park, Great Abington, Cambridge CB21 6AH, England www.woodheadpublishing.com Published in North America by CRC Press LLC, 6000 Broken Sound Parkway, NW, Suite 300, Boca Raton, FL 33487, USA First published 2008, Woodhead Publishing Limited and CRC Press LLC © 2008, Woodhead Publishing Limited The authors have asserted their moral rights. This book contains information obtained from authentic and highly regarded sources. Reprinted material is quoted with permission, and sources are indicated. Reasonable efforts have been made to publish reliable data and information, but the authors and the publishers cannot assume responsibility for the validity of all materials. Neither the authors nor the publishers, nor anyone else associated with this publication, shall be liable for any loss, damage or liability directly or indirectly caused or alleged to be caused by this book. Neither this book nor any part may be reproduced or transmitted in any form or by any means, electronic or mechanical, including photocopying, microfilming and recording, or by any information storage or retrieval system, without permission in writing from Woodhead Publishing Limited. The consent of Woodhead Publishing Limited does not extend to copying for general distribution, for promotion, for creating new works, or for resale. Specific permission must be obtained in writing from Woodhead Publishing Limited for such copying. Trademark notice: Product or corporate names may be trademarks or registered trademarks, and are used only for identification and explanation, without intent to infringe. British Library Cataloguing in Publication Data A catalogue record for this book is available from the British Library. Library of Congress Cataloging in Publication Data A catalog record for this book is available from the Library of Congress. Woodhead Publishing ISBN 978-1-84569-329-9 (book) Woodhead Publishing ISBN 978-1-84569-503-3 (e-book) CRC Press ISBN 978-1-4200-7971-5 CRC Press order number: WP7971 The publishers’ policy is to use permanent paper from mills that operate a sustainable forestry policy, and which has been manufactured from pulp which is processed using acid-free and elementary chlorine-free practices. Furthermore, the publishers ensure that the text paper and cover board used have met acceptable environmental accreditation standards. Typeset by SNP Best-set Typesetter Ltd., Hong Kong Printed by TJ International Limited, Padstow, Cornwall, England

Contents

Contributor contact details Preface

x xiii

Part I: Introduction

1

1

3

1.1 1.2 1.3 1.4 1.5 1.6 1.7 1.8 1.9 1.10 1.11 1.12 1.13 2

2.1 2.2 2.3 2.4 2.5

An overview of bioresorbable materials K. J. L. Burg and D. E. Orr, Clemson University, USA Introduction Degradation mechanisms Resorbable ceramics Resorption process Intended medical application guides the design of an absorbable implant Understanding the in vivo environment Naturally-derived materials Synthesized polymers Fabrication of absorbable materials Sterilization of absorbable implants Commentary Sources for further information and advice References The biological environment for bioresorbable materials D. Lickorish, N. Zebardast and J. E. Davies, University of Toronto, Canada Introduction to a hostile environment Blood Plasma protein cascades Fibrin formation Biomaterial interactions

3 4 5 6 7 8 10 11 13 16 16 17 17

19

19 22 28 31 33 v

vi

Contents

2.6 2.7

Host response to injury Practical demonstration of acute inflammation: The triple response Chronic inflammation Conclusion and future trends References

2.8 2.9 2.10

34 36 37 37 38

Part II: Degradation mechanisms

41

3

Synthetic bioresorbable polymers R. E. Cameron and A. Kamvari-Moghaddam, University of Cambridge, UK Introduction Bioresorbable polymers Degradation of aliphatic polyesters Factors affecting aliphatic polymer degradation Processing and devices Conclusions Sources of further information and advice References

43

Natural bioresorbable polymers W. Paul and C. P. Sharma, Sree Chitra Tirunal Institute for Medical Sciences and Technology, India Introduction Chitin and chitosan Alginates Cellulose Conclusion Acknowledgments References

67

Bioresorbable ceramics M. Bohner, Dr Robert Mathys Foundation, Switzerland Introduction Solubility Kinetics In vivo transformation Other bioresorbable ceramics Modelling resorption Future trends

95

3.1 3.2 3.3 3.4 3.5 3.6 3.7 3.8 4

4.1 4.2 4.3 4.4 4.5 4.6 4.7 5 5.1 5.2 5.3 5.4 5.5 5.6 5.7

43 44 48 54 60 60 61 61

67 68 75 83 88 88 88

95 99 101 102 104 107 108

Contents 5.8 5.9

Conclusion References

Part III: Bioresorption test methods 6

6.1 6.2 6.3 6.4 6.5 7

7.1 7.2 7.3 7.4 7.5 8

8.1 8.2 8.3 8.4 8.5 8.6 8.7 9

9.1 9.2

In vitro physicochemical test methods to evaluate bioresorbability S. Li, University Montpellier I, France Introduction Protocol for in vitro degradation studies In vitro physicochemical test methods Conclusion References In vitro biological test methods to evaluate bioresorbability G. Mabilleau and A. Sabokbar, University of Oxford, UK Introduction Methods of degradation of biomaterials Methods of assessing resorbability in vitro Characterization of the resorbability in vitro: Microscopic analysis of the surface References In vivo test methods to evaluate bioresorbability S. A. Clarke and G. R. Jordan, Queen’s University Belfast, Northern Ireland Introduction In vivo models Outcome measures Histomorphometric measurements Imaging Summary References Modelling of the degradation processes for bioresorbable polymers D. Farrar, Smith & Nephew Research Centre, UK Introduction Overview of degradation processes for bioresorbable polymers

vii 109 109

115

117 117 120 121 140 140

145 145 146 152 154 159 161

161 162 170 172 175 176 177

183 183 184

viii

Contents

9.3 9.4 9.5 9.6 9.7 9.8

Modelling of key processes Modelling of surface erosion Temperature effects Future trends Conclusion References

Part IV: Factors influencing bioresorption 10

10.1 10.2 10.3 10.4 10.5 10.6 10.7 10.8 11

11.1 11.2 11.3 11.4 11.5 11.6 11.7 11.8

Influence of processing, sterilisation and storage on bioresorbability F. Buchanan and D. Leonard, Queen’s University Belfast, Northern Ireland Introduction Processing techniques Processing-related degradation Sterilisation Maximising shelf-life: Packaging and storage Additives for reducing degradation Conclusion References Influence of porous structure on bioresorbability: Tissue engineering scaffolds P. Tomlins, National Physical Laboratory, UK Introduction Materials Processing Characterisation of tissue scaffolds Methods for monitoring the degradation of polymeric tissue scaffolds Conclusion Acknowledgement References

Part V: Clinical application 12

Influence of clinical application on bioresorbability: Host response

186 199 201 202 203 203

207

209

209 210 212 220 226 229 231 231

234 234 235 238 244 247 257 257 257

265

267

Contents

12.1 12.2 12.3 12.4 12.5 12.6 12.7 12.8 12.9 12.10 12.11 13

13.1 13.2 13.3 13.4 13.5 13.6 13.7 14

14.1 14.2 14.3 14.4 14.5 14.6

J. C. Y. Chan, K. Burugapalli, J. L. Kelly and A. S. Pandit, National University of Ireland, Galway, Republic of Ireland Introduction Host response cascade Host factors influencing biodegradation Influence of site of implantation on biodegradation Influence of species and repeated implantation Adverse outcomes of biodegradable polymers Mechanisms of in vivo biodegradation Material factors influencing biodegradation Biomaterial design parameters Conclusion References

ix

267 268 271 283 285 285 286 290 298 302 302

Scaffold and implant design: Considerations relating to characterization of biodegradability and bioresorbability 319 D. W. Hutmacher, Queensland University of Technology, Australia and C. X. F. Lam, National University of Singapore, Singapore Introduction 319 Biodegradation and bioresorption 320 Hydrolytic degradation of polycaprolactone 323 Hydrolytic degradation of medical polycaprolactone (mPCL) versus research polycaprolactone (PCL) 325 In vivo degradation of polycaprolactone-based scaffolds 343 Conclusions 352 References 353 Drug release from bioresorbable materials M. Westwood and D. S. Jones, Queen’s University of Belfast, Northern Ireland Introduction Examples of biodegradable pharmaceutical polymers Mechanisms of drug release from biodegradable polymers Drug delivery applications of biodegradable polymers Conclusions References

357

Index

393

357 359 373 377 385 385

Contributor contact details

(* = main contact)

Editor F. Buchanan Queen’s University Belfast Ashby Building Stranmillis Road Belfast BT28 3B Northern Ireland UK E-mail: [email protected]

Chapter 1 K. J. L. Burg* and D. E. Orr Hunter Endowed Chair and Professor of Bioengineering Department of Bioengineering Clemson University 501 Rhodes Engineering Research Center Clemson, SC 29634 USA E-mail: [email protected]

Chapter 2 D. Lickorish* and J. E. Davies Institute of Biomaterials and Biomedical Engineering University of Toronto Ontario Canada E-mail: [email protected]; [email protected] x

N. Zebardast and J. E. Davies Faculty of Dentistry University of Toronto Ontario Canada E-mail: nazlee.zebardast@ utoronto.ca

Chapter 3 R. E. Cameron* and A. Kamvari-Moghaddam Department of Materials Science and Metallurgy University of Cambridge Pembroke Street Cambridge, CB2 3QZ UK E-mail: [email protected]

Chapter 4 W. Paul and C. P. Sharma* Division of Biosurface Technology, Biomedical Technology Wing Sree Chitra Tirunal Institute for Medical Sciences & Technology Poojappura, Thiruvananthapuram 695012 Kerala, India E-mail: [email protected]

Contributor contact details

xi

Chapter 5

Chapter 8

M. Bohner Dr Robert Mathys Foundation Bischmattstrasse 12 CH-2544 Bettlach Switzerland E-mail: [email protected]

S. A. Clarke* and G. R. Jordan Dept of Surgery and Perioperative Care Queen’s University Belfast Whitla Medical Building 97 Lisburn Road Belfast BT9 7BL Northern Ireland UK E-mail: [email protected]

Chapter 6 S. Li Max Mousseron Institute on Biomolecules, UMR CNRS 5247 Faculty of Pharmacy University Montpellier I 34093 Montpellier France E-mail: [email protected]

Chapter 7 G. Mabilleau* and A. Sabokbar Nuffield Department of Orthopaedic Surgery University of Oxford Nuffield Orthopaedic Centre Headington Oxford OX3 7LD UK E-mail: guillaume.mabilleau@ndos. ox.ac.uk; afsie.sabokbar@ndos. ox.ac.uk

Chapter 9 D. Farrar Smith & Nephew Research Centre York Science Park Heslington York YO10 5DF UK E-mail: [email protected]

Chapter 10 F. Buchanan* and D. Leonard Queen’s University Belfast Ashby Building Stranmillis Road Belfast BT28 3B Northern Ireland UK E-mail: [email protected]

xii

Contributor contact details

Chapter 11

Chapter 13

P. Tomlins Innovation and Industry Division National Physical Laboratory Hampton Road Teddington TW11 OLW UK E-mail: [email protected]

D. W. Hutmacher* Institute of Health and Biomedical Innovation, Queensland University of Technology 60 Musk Avenue Kelvin Grove QLD 4059 Australia E-mail: dietmar.hutmacher@qut. edu.au

Chapter 12 J. C. Y. Chan,* K. Burugapalli J. L. Kelly and A. S. Pandit National Centre for Biomedical Engineering Science Orbsen Building National University of Ireland, Galway University Road Galway Republic of Ireland E-mail: [email protected] J. C. Y. Chan and J. L. Kelly Department of Plastic, Reconstructive and Hand Surgery University Hospital Galway Newcastle Road Galway Republic of Ireland

C. X. F. Lam Division of Bioengineering Faculty of Engineering National University of Singapore 9 Engineering Drive 1 Singapore 117576 E-mail: [email protected]

Chapter 14 M. Westwood and D. S. Jones* School of Pharmacy Queen’s University of Belfast Medical Biology Centre 97 Lisburn Road Belfast BT9 7BL UK E-mail: [email protected]

Preface

Bioresorbable materials, whether they be polymer, ceramic or composite based systems, are firmly established as the material of choice for a wide range of biomedical applications. For example, in tissue engineering a bioresorbable scaffold is required to provide initial strength and form to guide regenerating tissue. The mechanism of bioresorption differs depending on the class of material used. For example, synthetic polymers generally bioresorb by hydrolytic degradation, whereas bioresorption of certain natural polymers is controlled by enzymic action. For ceramic materials, such as calcium phosphate, the process is considered to be one of dissolution. In all cases however, the bioresorption rate is likely to be influenced by the local environment, including cellular activity. The tissue response to bioresorbable materials also needs to be considered, with the release rate of degradation products (or incorporated bioactive additives) influencing tissue regeneration. The bioresorption rate is one of the most important considerations when designing a bioresorbable device and therefore accurate methods of predicting bioresorption of this category of material are required. Long-term bioresorption needs to be predictable in terms of both changes in mechanical properties and biological response to degradation products. This presents unique challenges to the biomaterials scientist, where materials may remain in the body for weeks, months or even years. By combining the expertise of an international multidisciplinary team of authoritative contributors within this single publication, the main aspects of prediction and evaluation of bioresorbable biomaterials are covered. Therefore the book should provide a unique and valuable text for biomaterials scientists whether they be from an engineering, biology or other scientific background. Fraser Buchanan

xiii

1 An overview of bioresorbable materials K. J. L. BURG* and D. E. ORR, Clemson University, USA

Abstract: This chapter focuses on absorbable materials, beginning with a description of the cellular involvement in material degradation. Degradation terms are defined, to allow a complete understanding of the process by which absorbable materials are reduced and removed from the body. Multiple chemical and biological mechanisms that are at work in the body during material bioresorption are described. Additionally, the design criteria and characteristics for absorbable materials are outlined. The particularly appealing characteristic of an absorbable material is the elimination of a second surgery for implant removal. The lack of a follow-up surgery translates into reduced cost, pain, and recovery time for the patient; the chapter concludes by highlighting the exciting potential of absorbable and degradable materials in biomedicine. Key words: absorbable, biomaterials, bioresorbable, ceramics, degradable, polymers.

1.1

Introduction

Absorbable materials can be considered a specialized, biological subset within the larger classification of degradable materials. Multiple chemical and biological mechanisms are at work in the body during material bioresorption. Existing definitions of degradation terms can be conflicting and confusing, as authors have directed the various definitions around different issues, including level of material elimination, degradation process, degradation geometry, and implant locale. Bioresorbable is often perceived as synonymous with absorbable and is generally applied to those materials where the implant or its degradation products are removed through cellular activity in a biological environment.1 Biodegradation is a broader term used when a biological agent, such as an enzyme or microbe, is a dominant component in the degradation process and where a degraded material may simply be remodeled in the human body; similarly, biodegradation can refer to, for example, the process by which starch packing peanuts are degraded in a landfill. Bioerosion is used to describe, without regard to the specific mechanism involved, the process by which a water-insoluble polymer is converted into water-soluble products. Most medical researchers reserve the term degradation for the transformative step that changes the 3

4

Degradation rate of bioresorbable materials

material physicochemical properties. Not all degradable materials are absorbable; that is, the degradation process may not render the material soluble (and therefore will not permit its elimination from the body); multiple degradation processes can often occur under physiologic conditions. Ultimately, the descriptive terms describe various aspects of the transformation and/or disappearance of a material or implant after introduction into the body. This discussion will focus on absorbable materials, beginning with a description of the cellular involvement in material degradation. Periodic inclusion of degradation terms will be necessary to provide a complete understanding of the process by which absorbable materials are reduced and removed from the body. Additionally, the design criteria and characteristics for absorbable materials will be outlined. The particularly appealing characteristic of an absorbable material is the elimination of a second surgery for implant removal. The lack of a follow-up surgery translates into reduced cost, pain, and recovery time for the patient.

1.2

Degradation mechanisms

In the case of polymers, the underlying mechanism of material degradation is cleavage of the chemical bonds of the polymer backbone or side groups. Hydrolysis is the basic mode of degradation, as water reacts with the covalent bonds in the polymer to separate the polymer into smaller functional groups. Polymers most readily hydrolyzed are those containing carbonyl functional groups and bonded with heterochain elements of oxygen, nitrogen, and sulfur. Esters, amides, and carbonates are all included in this category. Weaker intermolecular bonds increase the rate at which hydrolysis can segment the molecular chain into oligomeric and monomeric species. Several additional factors may influence the degradation rate. Depending on the chemical characteristics of the polymer, hydrolysis can be expedited or catalyzed through the involvement of acids, bases, salts, and enzymes. Monomer generation and collection within the implant bulk can result in a localized pH shift that increases the degradation rate in the surrounding polymer. The relative hydrophilicity or hydrophobicity of a polymeric implant will govern water uptake into the physical structure of the implant; i.e. a hydrophilic polymer will initially attract much more water than a hydrophobic polymer. Hydration of the bulk material will allow more direct interaction with the chemical bonds, resulting in an increased rate of degradation. Hydrophilicity, as well as other polymeric characteristics, can be controlled through copolymerization. Two or more monomeric species may be combined to target desired properties of the degradable material, such as crystallinity or glass transition temperature. Accordingly, copolymers can

An overview of bioresorbable materials

5

Table 1.1 Degradable polymer parameters and the effect on degradation rate Parameter

Degradation rate

↑ ↑ ↑ ↑ ↓ ↑ ↑

↓ ↓ ↑ ↑ ↑ ↑

Crystallinity Molecular weight Temperature pH Surface area Enzyme activity

be developed with increased or decreased hydrolytic stability in relation to other polymer formulations. Increased crystalline regions, i.e. areas of closely aligned molecular chains, inhibit the introduction of water into the bulk of a polymer and thereby decrease the degradation rate. Assuming chains of similar length, amorphous polymer regions are less densely packed and are therefore more susceptible to degradation than crystalline regions. Table 1.1 outlines the typical relationship between specific polymer properties and the resultant degradation rate. An implant undergoes a series of degradation stages, beginning with the hydration of the polymer and hydrolysis of secondary and primary chemical bonds. Cleavage of the covalent bonds decreases the molecular weight as the larger polymer molecules are broken down into smaller oligomeric and monomeric forms. An increasing quantity of shorter oligomeric and monomeric segments allows increased movement of molecular chains in the bulk material. Shorter chains act as plasticizers, allowing more freedom for longer chains to slide past one another, ultimately leading to a reduction in the mechanical properties of an implant. As degradation progresses, chain fragmentation continues and results in mass loss as the lower molecular weight species are solubilized and exit into the surrounding medium. A gelatinized mass may be formed as the implant loses physical integrity and mass density. For truly absorbable implants, the process continues until the implant is completely dissolved and no biomaterial remnants remain at the site. Further physiological activity is then involved in the complete removal of degradation by-products from the body. Chemical chains and metabolites that are small enough can be eliminated through the lungs during normal respiration or via the kidneys in urine and feces.

1.3

Resorbable ceramics

Ceramics have been used for many years in orthopaedic applications in conjunction with traditional artificial joint implants. Bioresorbable ceramic

6

Degradation rate of bioresorbable materials

implants offer potential advantages as alternative bone void filler and in other such applications. Bioresorbable ceramics will provide stable and insoluble substrates for bone repair, until acted upon by osteoclasts, the bioresorbing cells of bone tissue.2 In normal bone, osteoclasts work in concert with osteoblasts to continually turn over and remodel the mineralized and organic matrix of bones in accordance to acting loads. Osteoclasts continually take up the calcium phosphate bone mineral while osteoblasts lay down new bone matrix relative to the stresses experienced by the skeleton. Bioresorbable ceramics can provide good initial support to a damaged or deficient bone region until osteoclasts naturally resorb the implant material through normal bone remodeling.

1.4

Resorption process

Cells use endocytosis to ingest macromolecules and particulate substances, including degrading polymeric material. The cell plasma membrane closes around the particulate and eventually pinches off to form a vesicle within the cell cytoplasm. Phagocytosis is used to describe the endocytic process of ingesting larger particulate material (>250 nm). Two classes of specialized phagocytic cells exist in the body: macrophages and neutrophils. The primary function of macrophages and neutrophils is to ingest invading microorganisms as well as consume damaged cells and cell debris. Foreign material from a degradable implant can be identified with adherent antibodies that trigger a phagocytic response. A phagocytic vesicle eventually fuses with lysosomes to activate particulate degradation through a variety of oxidative chemicals and hydrolytic enzymes that accelerate the breakdown of the material. Figure 1.1 demonstrates the process of phagocytosis. Macrophages can merge to form multinucleated foreign body giant cells (FBGCs) and osteoclasts (in bone) for purposes of engulfing and breaking down larger material fragments. While the bioresorption process can occur intracellularly, cells can also dispense enzymatic proteins extracellularly to biodegrade a degradable implant. However, enzymatic transfer and its effect on degradable substrates is more effective via phagocytosis. Enzymes are relatively large in relation to the spacing commonly found amongst polymer chains. Therefore, some level of hydrolytic activity must first take place to develop larger pores and cracks that increase the accessible surface area of the implant. These voids in the bulk structure then provide openings for the enzymes to enter and begin catalytic degradation of the material. Researchers have therefore studied the effect of modulating surface roughness to induce increased enzymatic activity in the degradation process.

An overview of bioresorbable materials (a)

(b)

(c)

(d)

(e)

(f)

7

1.1 Cellular phagocytosis of a bioresorbable material fragment. (a) Phagocyte comes in contact with material fragment; (b) material is engulfed by phagocyte membrane; (c) intracellular vesicle is formed; (d) lysosomes and vesicle fuse and release enzymes; (e) enzyme catalyzed hydrolytic action further degrades the material; (f) residual bodies and by-products are released from the cell via exocytosis.

1.5

Intended medical application guides the design of an absorbable implant

Absorbable and/or degradable materials may be used in a variety of applications, including temporary scaffolds, temporary barriers, drug delivery systems, and multifunctional devices. An absorbable scaffold is used to provide mechanical support that gradually transfers load to the healing, natural tissue. An absorbable suture provides an excellent example of a scaffold, as it initially seals a wound but absorbs over time while the skin heals. Hernia repair can often result in scar tissue between the abdominal wall and vital internal organs. A degradable barrier is well suited to prevent such unwanted attachment of tissues until healing is complete. Absorbable drug delivery devices provide opportunities for controlled release in the immediate vicinity of the targeted tissue, without the need to remove the carrier implant in a follow-up operation. Some implants may require

8

Degradation rate of bioresorbable materials

the incorporation of a device with two or more functions. For example, an absorbable orthopaedic implant might need to provide initial mechanical support while eluting beneficial drugs or growth factors into the surrounding bone tissue. Multifunctional degradable implants can be extremely diverse, depending on their intended functions. Several characteristics are important in developing an absorbable implant. Regardless of the intended application, the implant must have adequate initial strength and a documented, controlled rate of strength loss. Two types of erosion occur during the degradation of an implant. The first, surface erosion, causes the implant to decrease in size as the surface is removed – not unlike a melting ice cube. The second type of erosion is bulk degradation, where the degradation occurs throughout the entire implant, generally resulting in a surface–center differential; in the case of absorbable polyester materials, the degradation of the central regions causes pooling of acidic products, which can in turn cause pH-induced autocatalytic degradation. The danger of bulk erosion is that eventually, when the biomaterial surface is breached, a large pool of acidic products may be released in mass and may cause local tissue necrosis before being cleared from the area. Surface erosion is a more predictable process in comparison to bulk erosion. The process of bulk erosion and the resulting shift in implant strength is much more difficult to model.3 Depending on the application, a degradable implant will be designed according to the appropriate erosion mechanism that best suits the intended implant site. Figure 1.2 demonstrates the physical breakdown of degradable materials via surface and bulk erosion. The chemical composition of the implant should provide an endpoint of complete dissolution. In line with the removal of the implant, the degradation by-products and the implant itself must have a low order of toxicity to prevent damage to the surrounding cells and tissue. The degradable material must be malleable in order to form desired shapes and sizes and must accommodate sterilization following fabrication.4 The degradation rate of degradable materials can vary dramatically depending on the chemical composition, geometry, molecular weight, and crystallinity. Therefore, it is important to understand the inherent degradation rate of any materials used and the subsequent loss of mechanical strength over time, with attention to changes in physicochemical and/or mechanical characteristics due to processing, sterilization, and implantation effects. Deviation from standard medical implant fabrication procedures will substantially increase the cost of an absorbable implant.

1.6

Understanding the in vivo environment

An absorbable implant must be carefully designed for the intended in vivo environment. The degradation rate of a particular implant can be drastically

An overview of bioresorbable materials

9

(a)

(b)

1.2 Degradation mechanisms: (a) surface erosion and (b) bulk erosion.

different when placed in one tissue type versus another, due to biochemical and biomechanical differences in the local environment. Enzymatic activity can vary within the human body; implant location therefore determines the degradation rate of an enzymatic sensitive material. Toxicity is also a major issue in implant design. Both the bulk material and the degradation by-products that leach into the surrounding tissue must be relatively inert. Monomeric species and other degradants from degrading polymers must be examined through cytotoxicity testing, using cell and tissue types for which the application is intended. Toxic effects can lead to an inflammatory or immune response that may induce a shift in the local pH. The physiologically induced shift in pH can increase the rate of degradation, leading to a premature reduction in mechanical strength. One of the necessary aspects of a temporary scaffold or barrier is to remain intact long enough to allow proper tissue healing and growth. Degradation products can remain

10

Degradation rate of bioresorbable materials

in the body indefinitely if they are too large and/or incapable of being further metabolized. Identification of these non-degradable by-products is crucial to avoid future complications from remnant material. Degradation products found to be too large for removal through normal physiological channels may still be small enough to travel into regions of the body and cause harm. Typical physiological reaction to a foreign material includes the generation of a collagen-based fibrous capsule around the implant. Containment of degradable implants by a capsule can lead to a detrimental build-up of leachable by-products that may destabilize the implant’s structural integrity through a related shift in pH. Additionally, potential tissue damage may occur as by-products break free of the fibrous layer and infiltrate the surrounding tissue. A fibrous capsule can also be devastating to the function of a drug delivery device and its intended steady release of medicinal agents. Thorough evaluation of material properties must be conducted to prevent problems that may disrupt the natural tissue healing process.

1.7

Naturally-derived materials

Historically, naturally-derived (from animal tissue and plant fibers) degradable materials have proven effective as sutures for wound closure. Ever increasing applications such as wound dressings and temporary barriers have been addressed using natural polymers such as collagens, gelatins, polysaccharides and poly(β-hydroxy acids). As the most abundant protein found in mammalian tissue, Type I collagen exhibits a wide array of favorable characteristics. Collagen has good tensile strength and favorable cellular attachment, which meets the needs of a variety of medical applications, ranging from skin dressings to potential ligament and tendon repair. Gelatin is a protein-based material derived from the degradation of collagen. Gelatin is particularly well suited for low-strength applications requiring foam or films, such as hemostasis agents or void fillings. Both collagen and gelatin stimulate minimal antigenic and inflammatory responses. Chemical cross-linking can be used to vary the degradation rate and increase the physical strength of a collagen or gelatin implant. Antigenicity can also be reduced through the use of enzymatic and chemical treatments.5 One drawback is the high production cost associated with preparing these natural materials for medical applications. Polysaccharides are abundant in nature in the form of chitosan, dextran, and cellulose. Water-soluble chitosan has shown promise in wound dressings and dialysis membranes. Certain modified dextrans are structurally similar to heparin and may demonstrate similar anti-thrombogenic properties. Meanwhile, oxidized regenerated cellulose is being considered for several applications, including membranes, drug delivery, sutures, and temporary

An overview of bioresorbable materials

11

barriers. The oxidation process converts the hydroxyl groups into carboxylic acid groups, rendering the material susceptible to hydrolysis and enzymatic catalysis. Poly(β-hydroxybutyrate) is derived from microorganisms and is an intracellular storage polymer with reserves of carbon and energy. Highly crystalline and brittle in nature, poly(β-hydroxybutyrate) has been shown to degrade into D-3-hydroxybutyric acid through the work of soil bacteria. The fact that D-3-hydroxybutyric acid is a normal constituent of human blood is a positive factor in the case for using poly(β-hydroxybutyrate) in implant development. However, researchers have shown that poly (β-hydroxybutyrate) implants do not, to a significant extent, degrade or undergo loss of mechanical strength in vivo.6 Alginate is another naturally-derived polymer thought to hold promise as a cell carrier or gelling agent. Derived from algae, the material demonstrates many physical properties conducive to medical applications. Alginate gels are cross-linked with calcium or other 2+ ions, resulting in a wide range of sizes and shapes. Naturally-occurring, enzymatically degraded macromolecule polymers require specific enzymes to be present in the body in order to degrade the implant material into chemical chains small enough for filtration and excretion from the body. The necessary enzymes exist in the human body for polysaccharides and a variety of proteins. However, the human body does not contain alginase, the enzyme responsible for breaking down the alginate macromolecule. Therefore, medical applications with alginate may be limited to cases of special interest and unique circumstances. Natural polymers provide a large source from which to harvest such materials, but inherent difficulties reside due to batch-to-batch variability.

1.8

Synthesized polymers

A means of producing a wider range of more repeatable material properties in unlimited quantities is obtained through the production of synthetic degradable polymers. Synthetic degradable materials were first introduced for medical use in the late 1960s, as polyglycolide (PG) sutures entered the market. PG is highly crystalline yet very hydrophilic in nature, which contributes to the material’s relatively rapid degradation properties. Polylactide (PL) was introduced a short time later and provided a more hydrophobic material with a slower degradation rate. Two particular forms of PL, ‘L-PL’ and ‘D,L-PL’, were found to be useful in medical applications. L-PL is a stereoregular polymer while D,L-PL is the optically inactive racemic form. The increase in mechanical strength and toughness found in L-PL makes it better suited for sutures, orthopaedic implants and other applications that require more physical strength. The amorphous nature of D,L-PL offer potential advantages for drug delivery devices where consistent degradation

12

Degradation rate of bioresorbable materials

properties are important in providing homogenous dispersion of a drug over time. Polymers prepared directly from lactic acid or glycolic acid, by step growth polycondensation, are termed polylactic acid and polyglycolic acid, respectively. Polycondensation occurs at high temperature and results in low molecular weight species with impurities. Biomedical devices are generally prepared through ring-opening polymerization of lactide or glycolide, resulting in a purer, much higher molecular weight product. PL and PG, both aliphatic polyesters, have received considerable investigation over the years, leading to the incorporation of each, or their copolymers, in a variety of applications. Prior regulatory approval for devices comprised of these polymers contributes to the wide use of PL and PG. The Food and Drug Administration (FDA) approval encompasses those polymers derived from glycolide and lactide, provided synthesis was achieved via ring opening polymerization of the cyclic dimers, lactide diastereomers and glycolide. The expense involved in verifying the efficacy and safety of new materials in a given medical application can be quite inhibitory. Therefore, companies often turn to clinically familiar materials such as PL and PG to develop new applications.7 One characteristic of aliphatic polyesters is the acidic nature of the degradation products, prior to further processing into carbon dioxide via cellular metabolism. The acidic by-products tend to contribute to bulk erosion, as the pH shifts in internally-degrading regions of the implant. Studies have shown that fluid flow can slow the degradation rate of porous polymers by removing those acidic moieties that would otherwise contribute to the catalyzed hydrolysis of the implant.8 Another commonly-used degradable polymer is polydioxanone, which was commercially introduced in 1981. Polydioxanone provides an alternative to PG and PL for unloaded orthopaedic applications, e.g. those requiring sutures and staples. An etherester homo-polymer of p-dioxanone polydioxanone has a longer degradation time in comparison to PG. As with PG and PL, the prior approval of polydioxanone for medical use has led to its development for additional applications. Poly(ε-caprolactone) is a semi-crystalline material with high solubility and a low melting temperature. The degradation of poly(ε-caprolactone) is slower than that of PL (assuming relatively similar molecular weights) and lends to longer duration drug delivery of over one year. Toxicology evidence demonstrates nontoxic outcomes and has lead to further investigations to incorporate poly(ε-caprolactone) into additional medical devices in the future. Additional synthetic polymers of interest include polyanhydrides, poly(amino acids) and poly(ortho esters). Each formulation provides characteristics that may be of future interest in the development of degradable implants. Polyanhydrides demonstrate varying degrees of hydrolytic instability, depending on their aliphatic or aromatic formulation.

An overview of bioresorbable materials

13

Aliphatic–aromatic polyanhydrides were developed to generate a polymeric family with tunable levels of degradation. Limitations in the use of polyanhydrides include reactivity with drugs containing free amino groups and nucleophilic functionalities. Despite these limitations, polyanhydrides are degraded by surface erosion and were therefore initially envisioned for use in drug delivery, where a predictable release profile is required. The common chemical feature of poly(amides) is the presence of a hydrolyzable amide connecting group. Bulk erosion is the common method of degradation due to high water absorption. However, enzymatic and ioncatalyzed hydrolysis can lead to surface erosion. Poly(amino acids) present unique possibilities, where synthetic peptide chains of natural and possibly non-natural amino acids may be used to mimic extracellular proteins normally found in the body.9 A large number of polymers and copolymers can be generated through the modification of amino acid sequences within the molecular chains. Amino acids make up the bulk of the molecule, resulting in a low level of toxicity. Side chains can also be used as drug attachment sites and points of attachment for cross-linking agents. Existing problems with amino acid processing include handling difficulties (and therefore high expense), high reactivity, and extreme sensitivity to moisture. Poly(amino acids) are highly water-insoluble, requiring enzymatic hydrolysis of the amide bond for degradation. Accordingly, poly(amino acid) implant degradation rate will vary for each patient (as well as for each implantation location), making drug release characteristics difficult to predict. Finally, poly(orthoesters) demonstrate promising characteristics as controlled drug delivery agents. Surface erosion is their common form of physical degradation, allowing the continual release of embedded drugs. Additional research has found that the degradation rate can be controlled by modifying the level of acidic and basic molecules incorporated into the polymer structure, as the orthoester link is more stable in base than in acid. Each of the aforementioned polymers offers distinct characteristics, and promises degradable device innovation in the further development of degradable polymers. Table 1.2 presents the characteristics and chemical formulae for select synthetic degradable polymer materials.

1.9

Fabrication of absorbable materials

The fabrication of degradable and absorbable implants can often be accomplished with common processes used for non-degradable polymers. However, special practices must be adapted to ensure consistent and appropriate degradation properties for each batch of implants. Extrusion, injection and compression molding, and solvent spinning or casting are all relevant options when working with most degradable materials. A primary concern

Table 1.2 Degradable polymer characteristics and properties Polymer

Formula

Poly-L-lactide

Glass transition temperature (°C)

CH3 O O

Poly-D,L-lactide

CH

C

CH

C

Polyglycolide CH2

C

Poly(β-hydroxybutyrate)

O

CH2

O C

1200–3000

50–53

–

1900–2400

35–45

210–230

6000–7000

–

1500

171

2500

−10–0

O (CH2)2

160–180

n

Polydioxanone O

55–65

n

O O

Young’s modulus (MPa)

n

CH3 O O

Melting temperature (°C)

C

n

1 CH2

CH

O

n

CH3

Poly(ε-caprolactone)

−62

O C

(CH2)3

O

n

59–64

400

Polyanhydrides

O R

C

n/a

O O

C

C

O

R O

Poly(bisphenol A iminocarbonate)

NH O

CH3

Poly(amino acids)

H

H

O

N

C

C

R

n

55

–

820

69

–

2150

n

CH3 C

45

n

Poly(ortho esters) O

49

C O

n

16

Degradation rate of bioresorbable materials

is the possible reduction in molecular weight as bonds hydrolyze in response to thermal processing and exposure to oxygen. Vacuum drying and nitrogen purging has been shown to reduce the extent of degradation during processing by lowering the moisture content and the hydrolytic activity. Processing temperatures must be maintained at appropriate levels as overheating can actually lead to depolymerization and monomer formation. Excessive monomers can directly affect degradation and mechanical properties by catalyzing bond hydrolysis or by simply acting as plasticizing agents. As with conventional materials, thermal properties become even more important when developing processing parameters for degradable materials. Shelf-life of degradable polymers is extended by packaging implants in double, opaque bags, sealed using either inert gas or vacuum. Desiccators maintained under vacuum and sub-ambient temperature provide favorable storage conditions for the prevention of moisture infiltration into the implant environment.

1.10

Sterilization of absorbable implants

Sterilization of degradable implants provides yet another hurdle in the fabrication process. Steam autoclaving is not an option due to the inherent heat and moisture levels that will prematurely degrade the polymer. Gamma-radiation is acceptable only for the sterilization of select polymers as it can cleave polymer chains and cause a reduction in molecular weight. While currently used to sterilize some commercially available devices, ethylene oxide (EtO) gas is highly toxic and requires long periods of degassing and monitoring to ensure that no residual gas remains in the implant; protective gear is necessary to protect EtO operators.10 Given that degradable implants are ultimately destined for in vivo applications and given environmental considerations with respect to disposal of spent chemicals, there is little surprise that companies are investigating less toxic and more environmentally friendly means for sterilization. There appears to be no truly inert form of sterilization for degradable polymers to date. However, recent advances in methods such as gas plasma and supercritical carbon dioxide processing are developing to new levels and may improve operational throughput and outcome. As time progresses, manufacturing improvements will continue to stimulate increased efficiencies that will drive down costs associated with producing degradable implants for a variety of medical applications.

1.11

Commentary

The field of tissue engineering is broad, as research has come to encompass every major organ and tissue in the body. Degradable materials play

An overview of bioresorbable materials

17

a vital role in the application of tissue engineering principles. The fundamental advantage of degradable implants is the prevention of a second surgery for implant removal. Engineered tissues, manufactured using degradable biomaterials, may one day provide wide-spread application in organ replacement, three-dimensional void fillers, alternative tissue grafts, and cell-scaffold constructs. Approximately 450 000 orthopaedic autografts are performed each year; more than 84 000 individuals wait while approximately only 25 000 receive an organ transplant each year. The American Cancer Society estimates that 216 000 women are diagnosed with breast cancer annually; and, nearly 900 000 fracture cases require hospitalization per year. These statistics highlight the need for alternative therapies using tissue engineering for both treatment and reconstruction. ‘Smart materials’ must encompass the mechanical and biochemical properties of the tissue which they are intended to replace. A thorough understanding of material degradation rate and characteristics must exist in order to tune material replacement with tissue regeneration. The incorporation of drugs, growth factors, and surface proteins affects the properties of degradable materials. Proper kinetic alignment of the two parallel, dynamic systems, i.e. the degrading implant and the developing tissue, will prevent mechanical failure as well as stress shielding at the operative site. Extending the tissue engineering knowledge base of degradable and absorbable materials is critical to designing successful next-generation implants responsible for ever increasing physiological functions.

1.12

Sources for further information and advice

shonaike go and advani sg, Advanced Polymeric Materials. Structure Property Relationships, Boca Raton, CRC Press, 2003. von recum af, Handbook of Biomaterials Evaluation. Scientific, Technical, and Clinical Testing of Implant Materials, Philadelphia, Taylor & Francis, 1999. shalaby sw and burg kjl, Absorbable and Biodegradable Polymers, Boca Raton, CRC Press, 2004. schutte e, kaplan ds and picciolo gl, Tissue Engineered Medical Products, West Conshohocken, ASTM International, 2004.

1.13

References

1 kohn j and langer r, ‘Bioresorbable and bioerodible materials’, in Ratner BD, Hoffman A, Schoen F, and Lemons J, Biomaterials Science: An Introduction to Materials in Medicine, San Diego, Academic Press, 64–73, 1996. 2 langstaff s, sayer m, smith tjn and pugh sm, ‘Resorbable bioceramics based on stabilized calcium phosphates. Part I: Rational design, sample preparation and material characterization’, Biomater, 1999 20 1727–41.

18

Degradation rate of bioresorbable materials

3 gopferich a, ‘Mechanisms of polymer degradation and erosion’. Biomater, 1996 17(2) 103–14. 4 kronenthal rl, ‘Biodegradable polymers in medicine and surgery’, in Kronenthal RL, Aser Z, and Martin E, Polymers in Medicine and Surgery, New York, Plenum Press, 119–137, 1974. 5 yannas iv, ‘Natural materials’, in Ratner BD, Hoffman A, Schoen F, and Lemons J, Biomaterials Science: An Introduction to Materials in Medicine, San Diego, Academic Press, 84–94, 1996. 6 vert m, li sm, spenlehauer g and guerin p, ‘Bioresorbability and biocompatibility of aliphatic polyesters’, J Mater Sci-Mater Med, 1992 3(6) 432–46. 7 pulapura s and kohn j, ‘Trends in the development of bioresorbable polymers for medical applications’, J Biomater Appl, 1992 6(3) 216–50. 8 agrawal cm, mckinney js, lanctot d and athanasiou ka, ‘Effects of fluid flow on the in vitro degradation kinetics of biodegradable scaffolds for tissue engineering’, Biomater, 2000 21(23) 2443–52. 9 chaikof el, matthew h, kohn j, mikos ag, prestwich gd and yip cm, ‘Biomaterials and scaffolds in reparative medicine’, Ann NY Acad Sci, 2002 961 96–105. 10 middleton jc and tipton aj, ‘Synthetic biodegradable polymers as orthopedic devices’, Biomater, 2000 21(23) 2335–46.

2 The biological environment for bioresorbable materials D. LICKORISH, N. ZEBARDAST and J. E. DAVIES, University of Toronto, Canada

Abstract: Implantation of biomaterials triggers a number of interrelated physiological processes aimed at protecting the body from the foreign material. Depending on the specific material and application, this response may be undesirable. In order to generate the most appropriate tissue response for a given application, the biomedical engineer should be cognizant of the many ways in which the biological environment will interact with their material. An understanding of the processes described in this chapter will facilitate appropriate biomaterial design and implementation. Key words: inflammation, complement, neutrophil, macrophage, coagulation, extracellular fluid.

2.1

Introduction to a hostile environment

Despite maintaining a generally constant temperature and stable pH, the biological environment presents a remarkably hostile environment to foreign materials. The combined effect of cells and tissue fluid provides a chemically active and highly corrosive environment. In addition, it is a physically oppressive system with cells whose sole task is to remove or isolate intruders from the body. Given the numerous biological, mechanical and chemical insults to which we have been exposed (and to which we continue to be exposed), it is not surprising that a highly successful and integrated system to respond to these insults has evolved. The ability of the body to deal with a range of foreign materials whether they be a virus, a microbe, a recalcitrant splinter or a polymeric implant is remarkable. That this environment is largely controlled speaks to the fidelity of the system in coordinating appropriate responses to injury. It is the purpose of this chapter to introduce and describe the components that comprise the biological environment and how these components interact, with particular reference to the implanted resorbable polymer. In general, these components are introduced and described individually for ease of illustration. However, it is important to keep in mind that they do not necessarily act independently, but normally occur concurrently in numerous ways and with other systems. 19

20

Degradation rate of bioresorbable materials

2.1.1 Water and electrolytes The human body is approximately 60–70% water. In an average human male weighing 70 kg there will be approximately 42 L of body fluid divided into two main compartments; intracellular and extracellular (Guyton and Hall, 2006). The intracellular fluid (ICF) accounts for about 28 L with the remaining 14 L being extracellular. The extracellular fluid (ECF) can be further divided into plasma (3 L) and interstitial fluid (11 L). Interstitial fluid represents the fluid bathing the cells. In a normal individual the ionic concentrations of both ICF and ECF are very similar; however, their ionic composition is not (Table 2.1). ECF contains predominantly sodium and chlorine ions whilst ICF contains potassium and phosphate ions. Differences in ionic concentrations of ECF and ICF will drive the flow of water to balance these compartments. The sodium concentration in ECF is the primary driver for water shifts between compartments and is tightly controlled via the kidneys. Extracellular fluid compartment – Interfacial reactions The high dielectric ECF environment inevitably plays a pivotal role in determining the early biological responses to an implanted material, although relatively little is still known of the biological relevance of these complex interfacial interactions. In addition to water, many ionic species of both positive and negative charge are immediately available (nanoseconds) to engage in interfacial ion exchange events which will govern the spatial properties and strengths of charged double layers in the fluid phase. The relaxation of metastable defects present within, and at the surfaces of, inorganic materials, such as wide band gap calcium phosphates, may also contribute to the solid/liquid interface as we have discussed some time ago (Davies, 1985). However, in addition to these ionic species, a variable mixture of amino acids, sugars, fatty acids, and a library of proteins including enzymes and hormones will also be present. These macromolecules will adsorb to the implant surface within milliseconds, but each will display dif-

Table 2.1 Ionic concentration of ECF and ICF (Adapted from Guyton and Hall, 2006) Ion

[ECF] (mmol/L)

[ICF] (mmol/L)

Ca2+ Na2+ K+ Cl− HCO3− HPO4−

1.2–1.3 139–142 4.0–4.2 108 24–28.3 2

0 14 140 4 10 11

The biological environment

21

ferent adsorption kinetics to establish a dynamically changing interfacial composition (see discussion of the Vroman effect later). Proteins may be polar, charged molecules or hydrophobic, and thus the surface composition of the implanted material will have a profound effect on the adsorption dynamics of these differing species. Figure 2.1 is a stylized cartoon which illustrates the main components of the interface created at the surface of any implanted material. As we shall see later, it is the interaction of the proteinaceous macromolecules with the material that plays the driving role in cellular interfacial interactions with the implant surface, and more is known about these events. Nevertheless, it is salutary to keep in mind that the chemistry of this environment is highly complex.

A

m

m

B

m

m surface

C bulk

2.1 Diagram to illustrate the components of the early cell/biomaterial interface. The cell (above) is bathed in an extracellular fluid of complex composition that comprises both various inorganic ions (geometrical shapes) and organic molecules (irregular shapes). The surface of all materials, and particularly inorganic materials, can be radically different from the bulk composition. These deviations at the surface are shown to be contributing to the establishment of various charge double layers which will influence both ion exchange and protein adsorption (see text). The cell membrane itself is also negatively charged and geometrically complex. (From Davies, 1985.)

22

Degradation rate of bioresorbable materials

Blood and extracellular pH pH is defined as the negative log of the concentration of the hydrogen ion [H+] (Moffett et al., 1993) and many biological processes take place within a restricted range of pH. Blood has a pH of 7.4, which is slightly alkaline. Transient changes in levels of H+ or OH− ions can be dealt with by a very effective bicarbonate buffering system that attempts to prevent systemic pH from straying too far from 7.4. The pH of the intracellular space is approximately 7.0. However, there are times when a localized drop in extracellular pH in the tissue spaces can be an effective tool for the activation of cells during healing. An example of this is the pH at the site of injury. Wounds involve necrotic cells and the accumulation of acidic waste products lowers the local pH. This can provide a chemotactic stimulus to the ingress of leukocytes and macrophages that respond to injury. Of note for the polymer engineer is the acidic degradation products of the poly alphahydroxyacids. These commonly used, resorbable polymeric materials have been implicated in local drops in pH (Agrawal and Athanasiou, 1997), thereby activating macrophages and which may result in foreign body giant cell formation. Such an outcome may be advantageous for the rapid degradation of polymeric sutures but be more problematic for polymeric screws or fracture fixation plates in orthopedic applications (Bostman and Pihlajamaki, 2000).

2.1.2 Body temperature As with pH, core body temperature is controlled to within a very tight range around 37°C, as this is the temperature at which most enzymatic activity is optimal. The most obvious fluctuations in temperature that affects us all is a transient increase due to fever (pyrhexia), in which the body attempts to create conditions favourable for the generation of host defence cells whilst providing a temperature inhospitable to pyrogens.

2.2

Blood

2.2.1 Blood cells and plasma In a 70 kg male the volume of blood is approximately 5 L. Of this, 3 L is plasma whilst the remaining 2 L is cellular (Guyton and Hall, 2006). Plasma is a straw coloured solution of water, salts and proteins and is the liquid medium in which a variety of cells, termed the ‘formed elements’ are carried. By weight, plasma is 92% water, 7% protein and the remaining 1% is organic molecues, ions, vitamins, minerals and dissolved gases, such as carbon dioxide and oxygen (Silverthorn, 2006).

The biological environment

23

2.2.2 Formed elements of the blood The cells circulating in blood are broadly divided into two groups: red blood cells (erythrocytes) and white blood cells (leukocytes). Platelets (thrombocytes) are not considered to be cells as they are simply membrane-bound fragments of megakaryocytes. One milliliter of blood contains approximately 5 × 109 erythrocytes, 8–11 × 106 leukocytes and 300 × 106 platelets.

2.2.3 Erythrocytes Erythrocytes are by far the most numerous cell in blood, accounting for over 90% of the circulating blood cells (Tefferi, 2001). Males have approximately 5–6 × 106 red blood cells (RBC) per microliter (μL) of blood whilst females have slightly less, approximately 4–5 × 106 cells/μL. RBCs are derived from erythroblasts in the bone marrow and during maturation in the marrow they lose their nucleus. RBCs are approximately 7 μm in diameter and have a biconcave disc shape. Once in the bloodstream the principal physiological function of RBCs is to bind and distribute oxygen to the tissues through the globular protein, haemoglobin.

2.2.4 Leukocytes White blood cells develop in the bone marrow and can be divided into cells containing granules (granulocytes) and those that do not (agranulocytes). The cytoplasmic granules contain a wide variety of enzymes and antimicrobial proteins that leukocytes utilize to destroy invading pathogens or communicate with other cells involved in host defence. In addition, when stimulated by chemical mediators of tissue damage, leukocytes respond by secreting an array of signaling molecules, chiefly the interleukin family, which acts as a homing signal for other inflammatory cells.

2.2.5 Granulocytes Granulocytes are further subdivided based on the tendency of the cytoplasmic granules to take up either acidic, basic or neutral stains in histological preparations. Hence their respective names, acidophil, basophil and neutrophil. The acidophil is more commonly referred to as the eosinophil given its tendency to take up eosin, a very common acidic dye. Eosinophils Eosinophils constitute approximately 1–4% of the circulating leukocytes (Tefferi, 2001). They measure approximately 9 μm in diameter and are phagocytic cells. The primary function of the eosinophil is in host defence

24

Degradation rate of bioresorbable materials

against infection and parasites which is mediated through the secretion of the protease, cathepsin (Bainton, 1999). Eosinophils also produce major basic protein which plays a role in the release of histamine from mast cells. Thus eosinophils also play a role in mediating the allergic response. Neutrophils Neutrophils are the most common leukocyte, accounting for approximately 60% of circulating leukocytes. They remain in the blood stream, unactivated, for between 5–10 hours. However, when activated, in the case of the inflammatory response, they will exit the vasculature and remain resident in the injured tissue for several hours. The neutrophil is a phagocytic cell and is the first leukocyte to arrive at the site of injury from the bloodstream. The neutrophil’s main role is to remove bacterial and cellular debris by the production of reactive oxygen species such as hypochlorite and superoxide (Bainton, 1999). Whilst the neutrophil is responsible for the oxide mediated destruction of bacteria, this can result in significant damage to healthy tissue due to its non specific nature. Pus is largely composed of dead neutrophils, tissue debris and fluid – its characteristic yellow/green colour is attributed to the enzyme, neutrophil myeloperoxidase. Basophils The basophil is the least common of the circulating leukocytes, accounting for between 0.5% of the leukocyte count (Min and Paul, 2008). The secretory granules in the basophil contain histamine and proteolytic enzymes. Mast cells Mast cells and basophils are closely related. Not surprisingly, the mast cell is functionally similar to the basophil but, unlike the basophil which is restricted to the blood stream, the mast cell is resident in a wide variety of connective tissues. Due to their wide tissue distribution and close proximity to blood vessels, the degranulation of the histamine from mast cells is perhaps one of the first events leading to vascular dilation in acute inflammation. The mast cell is perhaps better known for its role in mediating allergic reactions due to its interaction with the immunoglobulin IgE (Min and Paul, 2008); widespread degranulation of histamine from mast cells can result in life-threatening anaphylaxis.

2.2.6 Agranulocytes Agranulocytes can be divided into lymphocytes (20–40% of leukocytes) and monocytes (5–10% of leukocytes). The lymphocytes are further

The biological environment

25

subdivided into B lymphocytes, T lymphocytes and dendritic cells. Lymphocytes are effector cells of the immune system, generating and reacting to antibodies derived from contact with antigens. Monocytes are phagocytic cells that ingest foreign material. B lymphocytes B lymphocytes are the cells responsible for antibody production and form part of the humoral immune system. When challenged with a new antigen, such as a virus, B cells will synthesize immunoglobulins (IgE, IgG, IgM, IgA, IgD), proteins that will specifically recognize the antigen when it next presents, enabling its binding to T lymphocytes and rapid removal from the body. T lymphocytes T lymphocytes, which develop in the thymus, form part of the cell-mediated immune system. T cells are responsible for the direct lysis of antibodybound cells. T lymphocytes are distinguished from B lymphocytes by the presence of the T cell receptor on the cell surface. There is a vast array of T cell subsets, each with a specific role in mediating the immune response.

2.2.7 Dendritic cells Dendritic cells (so named because their multiple branched cell processes resemble a tree) function to present antigen to immune cells. They are generally located close to the surface of tissues that communicate with the external environment; for example, nose, skin, gut. Dendritic cells recognize pathogens through Toll-like receptors on their cell surface (Hochrein and O’Keeffe, 2008). This interaction with immune cells has made dendritic cells a particularly interesting target cell to mediate delivery to lymphocytes of antigens encapsulated in biodegradable polymers (Gerelchuluun et al., 2007).

2.2.8 Monocytes/Macrophages The terms monocyte and macrophage describe essentially the same cell, the difference being in their tissue distribution. Monocytes are restricted to the bloodstream and function to remove cellular and other debris. However, like neutrophils, circulating monocytes once activated will leave the blood stream via a process of extravasation and are then termed macrophages. Monocytes/macrophages are much longer-lived than neutrophils, with macrophages present for several days during acute inflammation and longer in cases of chronic inflammation. While transient increases in the macrophage

26

Degradation rate of bioresorbable materials

population can be brought about through monocyte extravasation during inflammation, macrophages are also located within the connective tissue of specific organs, e.g. lung, liver, spleen, and are thus termed ‘tissue resident macrophages’. It is these tissue resident macrophages that provide a first line of defense for the body when the tissue is damaged. The macrophage, literally ‘big eater’, is the cell responsible for engulfing and removing cellular and foreign debris and is arguably the most important cell in mediating the medium- to long-term reaction to implanted biomaterials. Monocytes and macrophages are activated by a wide range of signaling molecules. These include bacterial products such as lipopolysaccharide (LPS), complement and inflammatory chemical mediators synthesized by neutrophils (Schmidt and Kao, 2007). Macrophages will ingest pathogens, dead cells, etc. by invagination of the plasma membrane. These food vacuoles are then fused with lysosomes (bags of proteolytic lysozymes) to degrade the material. Macrophages can ingest only a limited amount of material before rupture of full lysosomes causes digestion of the cell itself. Macrophages also interact with lymphocytes by presenting digested antigen to B lymphocytes for the generation of antibodies. Foreign body giant cells Foreign body giant cells (FBGC) are the result of macrophage fusion and are a histological hallmark of chronic inflammation. In the fibrous tissue surrounding a foreign body, numerous macrophages will fuse to form the FBGC in order to more effectively neutralize and degrade foreign materials. FBGCs secrete proteases and superoxides and can have a deleterious effect on implanted medical devices. FBGCs dwarf those cells in the surrounding tissue, measuring 100–200 μm in diameter, and can have upwards of 20 nuclei (Fig. 2.2). They are generally linked with chronic inflammatory responses. FBGCs, along with macrophages, will often remain present at the interface between host and implanted biomaterial for the lifetime of that material. As these cells play a pivotal long-term role in the reaction to implanted materials, the reader is directed to excellent reviews of FBGC formation and function (Anderson et al., 2007; Collier et al., 2004; Hu et al., 2001).

2.2.9 Platelets Platelets are cell membrane fragments, measuring 1–4 μm in diameter, that are released from megakaryocytes in the bone marrow (Scott et al., 2004a). Each milliliter of blood contains approximately 200 000–400 000 platelets. Platelets are not true cells as they lack a nucleus; however, they have many functional characteristics of whole cells. Their cytoplasm contains

The biological environment

27

2.2 Foreign body glant cells (FBGC, arrows) engulfing polylactide-co-glycolide (PLGA) particles. Field width = 300 μm.

contractile proteins such as actin and myosin; they contain endoplasmic reticulum and Golgi apparatus capable of synthesizing enzymes, large calcium stores, mitochondria, and stores of various growth factors and coagulation factors. The cell membrane of platelets is coated with glycoproteins, preventing their adherence to the intact endothelium but causing adherence to injured areas of the vessel wall. In addition, the platelet membrane contains phospholipids, which participate in the clotting process and selectins which mediate leukocyte exit from the vasculature.

2.2.10 Ionic and protein content of the blood Aside from the formed elements, plasma also contains numerous ions and proteins. Ions dissolved in plasma play roles in maintaining osmotic pressure and pH and in regulating cell membrane permeability. A wide variety of proteins are carried in the blood, including the albumins, immunoglobulins, fibrinogen, blood clotting factors and complement. The principal plasma proteins, accounting for 60% of circulating protein, are the albumins, which function as transport proteins for a variety of hormones and other signaling molecules. Immunoglobulins, or antibodies, are proteins that function to recognize a wide variety of foreign antigens and direct the host immune response. There are five isotypes of immunoglobulins, termed IgG, IgA, IgD, IgE and IgM. Proteins of the clotting and complement cascades will be discussed later. Irrespective of their specific function, plasma proteins as a

28

Degradation rate of bioresorbable materials

whole serve to balance the flow of water between blood and the extracellular compartment. Specifically, plasma proteins raise the oncotic pressure of the blood which tends to draw water into the venous end of the capillary circulation, balancing the water lost at the arterial end of the capillary due to hydrostatic pressure.

2.3

Plasma protein cascades

Two families of plasma proteins are of particular significance in generating and coordinating host response to injury. These are the coagulation cascade and the complement cascade. Whilst they are described separately they occur simultaneously and are intimately linked with the inflammatory process (Levi and van der Poll, 2005; So et al., 2003).

2.3.1 Haemostasis and coagulation Hemostasis is a vital host defence mechanism which prevents blood loss from damaged vessels. Hemostasis is achieved by several mechanisms: (i) (ii) (iii) (iv)

Vasoconstriction Formation of a platelet plug Formation of a blood clot as a result of blood coagulation Eventual growth of fibrous tissue into clot

We will examine each of these mechanisms in detail, with particular attention given to the blood clotting mechanism. Vasoconstriction Injury of the vessel wall results in contraction of the smooth muscle layer of the wall. This is a result of local myogenic spasm as well as the release of vasoconstrictor substance from platelets, namely, thromboxane A. Platelet plug When platelets come in contact with injured blood vessel walls, especially exposed collagen fibres or foreign surfaces (i.e. biomaterials), they become activated. This means they begin to swell and pseudopods extend from their surface. Contractile proteins cause release of granules containing multiple factors and they become sticky and adhere to the exposed subendothelial surface. In addition, platelets release ADP and thromboxame A2 that both act on nearby platelets, causing them to adhere to other activated platelets thereby amplifying the initial signal. The number of platelets increases at

The biological environment

29

the site of injury resulting in the formation of a platelet plug. This loose plug is strengthened as fibrin threads form during the process of coagulation. The platelet plug itself is sufficient for closing of small injuries and minute vessel ruptures. These types of injuries occur regularly in a normal, healthy individual. Coagulation Whilst platelet plug confers primary hemostasis, a family of proteins found in plasma are activated to yield secondary hemostasis, or coagulation. More than 50 substances found in the host blood and tissue are thought to play a role in coagulation. Procoagulatants are substances that promote coagulation while anticoagulants are substances that inhibit the process. It is the balance between these two substances that determines whether the blood will clot. In an uninjured vessel, anticoagulants normally predominate preventing the clotting of blood as it circulates in the host. However, when a vessel is ruptured, procoagulants, normally present in precursor forms (zymogens), become activated, overriding anticoagulants and resulting in blood clot formation. In certain pathologies, blood clots may form in the lumen of intact vessels (a process known as thrombosis). Portions of the thrombus may break loose forming emboli, which may lodge in arteries supplying the heart or brain, causing myocardial infarction and stroke, respectively. The currently accepted model of blood coagulation involves a series of zymogen activation reactions (Lwaleed et al., 2007). At each stage, zymogen activation into a protease is accomplished by cleavage of one or more peptide bonds. In the initiating stage, a complex cascade of reactions occurs, involving over a dozen coagulation factors. The end result of this is the formation of a complex of activated substance called prothrombin activator. Prothrombin activator then catalyzes the conversion of prothrombin to thrombin, which then converts fibrinogen to fibrin fibers. It is the network of fibrin fibers that results in the entrapment of platelets, blood cells and plasma to form the blood clot. We will now discuss these processes in detail. Initiation of coagulation Coagulation is initiated by trauma to the vessel walls, blood, or adjacent tissue, contact of blood with damaged enthothelial cells or exposed collagen/tissue products underneath the vessel walls. Two pathways to the final formation of prothrombin activator have been identified: intrinsic and extrinsic. It is important to realize that in reality these two pathways

30

Degradation rate of bioresorbable materials

constantly interact with one another. In both pathways, a number of plasma proteins known as blood-clotting factors, initially inactive, participate in a cascade of activation reactions. These factors are designated by roman numerals. To indicate activation, a lowercase letter ‘a’ is added after the roman numeral. These two biochemical pathways meet in the common pathway at the conversion of Factor X to Factor Xa. Extrinsic pathway The extrinsic pathway is set in motion as a result of trauma to the vascular wall and surrounding tissues. Traumatized tissue releases tissue factor comprising phospholipids and lipoproteins. Tissue factor complexes with blood coagulation Factor VII and in the presence of calcium ions, enzymatically converts Factor X to Factor Xa. Activated Factor X then combines with Factor V and tissue phospholipids that are part of tissue factor or with additional phospholipids released from platelets to form the prothrombin activator complex. This complex in the presence of calcium ions splits prothrombin into thrombin. Feedback loops Initially Factor V is inactive. Once thrombin begins to form, it acts proteolytically on Factor V, activating it. Activated Factor V accelerates the protease activity of Factor X, resulting in increased splitting of prothrombin to thrombin. This positive feedback effect greatly accelerates the clotting process. The extrinsic pathway is therefore explosive in nature and, once initiated, its speed is limited only by the amount of tissue factor released, and the quantity of Factors X, VII and V present in the blood. Intrinsic pathway The intrinsic factor begins due to trauma to the blood itself or exposure of blood to collagen or an artificial surface (for example, some polymeric biomaterials) in the vessel wall. The intrinsic pathway proceeds much slower than the extrinsic pathway. This alters Factor XII and platelets in the blood. When Factor XII is disturbed, its molecular configuration changes, converting it into activated Factor XII. In addition, blood trauma results in platelet damage/adherence to collagen or the artificial surface and the release of platelet phospholipids containing platelet factor 3. Activation of Factor XII results in the activation of Factor XI. This activation step requires the presence of high-molecular-weight (HMW) kininogen and is accelerated by prekallikrein. Activated Factor XI acts enzymatically on Factor IX, activating it which in turn acts in concert with Factor VIII, platelet factor 3 and

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31

platelet phospholipids to activate Factor X. The remaining steps in this pathway are the same as those of the extrinsic pathway. This overlap between the two pathways is known as the common pathway.

2.4

Fibrin formation

Thrombin is an enzyme with proteolytic activity. It acts on fibrinogen present in the blood to cleave four peptides from each molecule, forming one fibrin monomer. Fibrin monomers have the capability to polymerize automatically with other monomers to form fibrin fibers. Initially, the fibrin monomers are held together via weak hydrogen bonds. However, thrombin also activates fibrin-stabilizing factor which is present in plasma and also released from entrapped platelets. This factor acts as an enzyme that covalently binds and crosslinks fibrin monomers. This results in increased strength of the fibrin meshwork.

2.4.1 Clot retraction Once the clot a formed, with the help of platelets, it slowly begins to retract, resulting in the compression of the fibrin meshwork. Contraction is activated and accelerated by thrombin and calcium ions, which are released, from platelet stores. In addition, as the clot retracts, the edges of the damaged vessel are pulled together, which contributes to hemostasis.

2.4.2 Fibrous tissue formation and dissolution of blood clot Once a blood clot is formed it can either become invaded by fibroblasts resulting in growth of connective tissue into the clot or it can dissolve. If the clot has formed due to the rupture of a vessel, it will become invaded by fibroblasts and fibrous tissue forms resulting in the permanent repair of the injury. However, if the clot has been formed due to leakage of excess blood into surrounding tissue, fibrinolytic enzymes are released that function to dissolve the clot. One such enzyme is plasmin. Plasma contains plasminogen that becomes activated to plasmin, a proteolytic enzyme. Blood clot formation results in the entrapment of a large amount of plasminogen along with other plasma proteins. Traumatized tissue and vascular endothelium slowly release tissue plasminogen activator. This enzyme, over the course of a few days, slowly activates plasminogen to plasmin which results in the lysis of fibrin clots. Plasmin achieves this by the digestion of fibrin fibers and other coagulation proteins including fibrinogen, Factor V, Factor VIII, prothrombin and Factor XII.

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Degradation rate of bioresorbable materials

2.4.3 Complement cascade Blood contains a family of proteins that help identify foreign materials. They are so named as this group of proteins, now standing at over 30 members, was initially assumed to provide a ‘complementary’ role to antibodies in the context of host immunological reactions (Walport, 2001). The role of complement is more important and complex than its name suggests and the pathways by which complement is activated/controlled is of particular relevance to the biomedical engineer. Complement refers to a series of nine proteins, termed C1–C9, that circulate in the blood as inactive zymogens. When activated, these proteins then catalyse an enzymatic cascade, resulting in the generation of protein fragments that can enable the host to target invading pathogens. These fragments are designated with a lowercase letter, for example C3 is converted into C3a and C3b fragments. Complement proteins can promote inflammation, facilitate phagocytosis and even direct lysis of cells. Given the powerful effect of activated complement proteins, their activation is tightly controlled via inhibitors. Studies have shown that complement proteins are covalently bound to proteins at the interface of biomaterials rather than being directly bound to the material surface (Andersson et al., 2005). Nilsson et al. (2007) describe in great detail the molecular pathway for complement biomaterial interactions. There are two pathways of activation of complement, though both will arrive at the production of the important C3b fragment. Classical pathway The classical pathway mediates the complement response to specific antigen–antibody complexes. Circulating C1 will bind to antibody fragments once the antibody has bound antigen. This interaction initiates an enzymatic cascade which results ultimately in the generation of a C3b fragment. Macrophages and neutrophils have cell surface receptors for C3b and so will use this interaction as a way of opsonization (from the Latin opsonere, ‘to obtain food’) and this is the way in which the complement system can mark substances for elimination by these cells. In addition, the continuing enzymatic cascade of C5b, C6, C7, C8 and C9 results in the formation of the membrane attack complex (MAC) which, as its name suggests, creates a channel in the plasma membrane of the target cell causing lysis. Alternative pathway The alternative pathway does not require antibody interactions and is initiated by hydrolysis of the C3 protein into C3a and C3b fragments. This spontaneous hydrolysis occurs naturally but is tightly controlled. Formation of the MAC proceeds as with the classical pathway.

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2.5

33

Biomaterial interactions

An understanding of blood–biomaterial interactions is important in interpreting the biological response to an implant. This is because, following surgical insult and subsequent implantation of a biomaterial, blood is the first tissue of the body to come into contact with the introduced material and it is these initial interactions between blood and implant that profoundly affect the subsequent biological reaction to an implanted material.

2.5.1 Plasma protein interactions with implanted materials Leo Vroman is credited almost 30 years ago with the observation that fibrinogen preferentially adsorbs to tantalum from blood plasma (Vroman et al., 1980). As such, the effect of adsorbed proteins on surfaces is referred to as the ‘Vroman effect’ (Noh and Vogler, 2007). The interaction of proteins with an implant surface is complex and the reader is referred to an excellent review on this subject (Slack and Horbett, 1995). It is important to appreciate that once implanted, a biomaterial is immediately covered in blood and plasma proteins. These proteins include albumins, fibrinogen and complement. The type and extent of protein adsorption is variable and the chemistry of the surface along with parameters such as protein molecular weight dictate the protein interaction. Furthermore, different proteins have different affinities and early adsorbed proteins may be displaced by later arriving proteins that have a higher affinity (‘later’ in this context meaning a temporal scale range that can be measured in microseconds). Why is protein adsorption so important in the context of implanted materials and their interactions with the body? Simply because host cells, such as macrophages and fibroblasts, do not recognize materials per se but interact with proteins that preferentially adsorb to those materials. Thus an understanding of the types of proteins that will adsorb to a particular material is helpful not only in determining the surface interactions but in tailoring surfaces where specific protein interactions are required. A good example of this is the popular strategy of modifying biological interactions by pre-coating materials with proteins involved in cellular attachment, for example fibronectin and vitronectin (Karakecili and Gumusderelioglu, 2007; Anderson et al., 1999). Another may be to use resorbable polymeric materials as delivery vehicles for a range of growth factors embedded within the polymer matrix. In trying to control the biological reaction through surface functionalization techniques one must be mindful that what may well work in a highly controlled in vitro system may not be directly transferable to the complex, dynamic in vivo environment.

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Degradation rate of bioresorbable materials

2.6

Host response to injury

We have described many of the individual components and processes that comprise the biological environment but it is when we encounter injury to our tissues that we see the complex ways in which these processes interact.

2.6.1 Inflammation The term ‘inflammation’ stems from the Latin for ‘fire’ and relates to the heat and redness commonly associated with the vascular response to tissue damage (Scott et al., 2004b). Once a tissue is injured, a myriad of cells, plasma proteins and signaling molecules combine in a coordinated response that is designed to remove the injurious agent and/or render it inert. Understanding that inflammation is a natural process that has been fine-tuned as we have evolved is important, as there is no way to fully engineer a solution to prevent this response. Nor should there be need to. Even the simplest, sterile surgical incision will generate an inflammatory response. Nevertheless, research efforts continue to target key signaling molecules with a view to dampening inflammation (Henson, 2005; Han and Ulevitch, 2005). The inflammatory response is often referred to in the literature as ‘innate immunity’ to be distinguished from ‘adapted or acquired immunity’ which describes the conferring of lasting protection through generation of antibodies. However, it is clear that there are many interactions between the two systems which Litman et al. (2005) suggest blur this distinction. For example, B-lymphocytes have been described as playing a key role in regulating host susceptibility to secondary infection following acute inflammation (Rajakariar et al., 2008).

2.6.2 Mediators of inflammation Communication between cells must occur in order to coordinate the host response to injury. These molecules are either stored inside the cell, as is the case with the secretory granules of the leukocyte and mast cell, ready to be released on contact with noxious stimuli, or synthesized and released in response to this stimuli. In general, the effects of these molecules can be described by whether they affect the cell from which they were synthesized (autocrine) or cells in the local tissue area (paracrine). The secreted signaling molecules are loosely termed cytokines and can be described as soluble factors that bind to specific receptors on the target cell to elicit a specific response. The most prominent pre-formed molecules are the histamines, bradykinins and prostaglandins. Histamine receptors are found on endothelial and

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35

smooth muscle cells; thus histamine plays a role in vasodilation and capillary permeability. Bradykinin is a peptide found in plasma that, like histamine, is a powerful vasodilator and increases vascular permeability by a direct affect on endothelial cells. Prostaglandins are a family of lipid-derived molecules with varied effects including vasodilation and vasoconstriction. Aspirin inhibits the formation of prostaglandin and thomboxane, hence its use as an anti-inflammatory and blood thinner. Cytokines secreted by cells in the context of an inflammatory response are too numerous to list. However, there are important groups of cell signaling molecules that can be identified. The interleukins mediate signaling between leukocytes. There are approximately 18 named interleukin molecules, designated IL-1, IL-2, etc. The principal function of the IL-1 family is the direct effect they have on neutrophils and monocytes in the early stages of an acute inflammatory response. Other important signaling molecules are the interferon (IFN) and tumour necrosis factor (TNF) families. Importantly, all of these groups provide a means of communication with the immune system and have contradictory pro-inflammatory and anti-inflammatory functions. Thus the specific effect of each cytokine is highly dependent on the microenvironment. For excellent reviews of biological mediators of inflammation, specifically IL, IFN and TNF function, the reader is directed to Rankin, 2004; Dinarello, 2007; Kovarik et al., 2008; and Chen and Goeddel, 2001.

2.6.3 Acute inflammation Acute inflammation is defined as the response of a vascularized tissue to injury or infection. The four cardinal signs of inflammation, described in the first century AD by Celsus, are redness (rubor), swelling (tumor), heat (calor) and pain (dolor) (Attur et al., 2002). Virchow later added loss of function (functio laesa) in the last century, given the observation that inflamed tissue often does not function correctly. Regardless of the specific cause, injury to a vascularized tissue will generate an immediate, acute inflammatory response that is localized to the injured area and is of a relatively short duration, in the order of a few days. This process involves two stages – changes in vascular flow and permeability followed by leukocyte migration from the bloodstream.

2.6.4 Alterations in vascular flow The first response of the local vasculature to an injury is a very brief vasoconstricton followed by release of histamines and bradykinins from mast cells in the damaged tissue, resulting in local arteriolar vasodilation. This increases the blood flow to the area and the capillary bed is soon engorged with blood. Alterations in hydrostatic pressure at first force fluid out of the

36

Degradation rate of bioresorbable materials

vasculature into the interstitial space. The increased extracellular fluid, or transudate, functions initially to dilute the potential toxin. The lymphatic system, which functions to drain excess interstitial fluid, cannot drain the transudate particularly rapidly, leading to edema, or swelling of the tissue.

2.6.5 Increased vascular permeability Ordinarily, the walls of the vasculature that are lined with endothelial cells remain impermeable to both cells and proteins, allowing only exchange of ions and fluid. In response to soluble factors such as bradykinin, histamine and leukotrienes generated during inflammation, endothelial cells of the post capillary venules are stimulated to reorganize their cytoskeleton. What results is a contraction of the endothelial cells sufficient for protein components of plasma, such as fibrinogen and complement, to leave the circulation. This exudate generates an increase in osmotic pressure that draws more fluid from the capillaries into the tissue.

2.6.6 Leukocyte extravasation Following implantation of a biomaterial, the surgical incision will itself have created sufficient local hemorrhage for the implant and wound area to be covered in blood. However, white blood cells, such as neutrophils and monocytes, from undamaged vessels adjacent to the wound must be able to leave the circulation in order to reach the site of injury during the course of healing. This process of leukocyte extravasation is made possible by the increase in vascular permeability. A series of selectin protein binding events, between membrane bound proteins of both leukocytes and platelets with proteins on the endothelial cell membrane, facilitate this egress from the vasculature (Zarbock and Ley, 2008). Selectins are surface proteins expressed by platelets (P-selectin), endothelial cells (E-Selectin) and leukocytes (L-selectin). As more fluid leaves the vasculature there is a local increase in blood viscosity, slowing the flow of blood allowing the leukocytes to escape the flow and come into contact with the endothelial cell wall, a process called margination. Marginated leukocytes then roll along the endothelial cell wall and are slowed by progressive selectin binding. Eventually the leukocyte will stop rolling and will squeeze between the retracted endothelial cell walls into the tissue space.

2.7

Practical demonstration of acute inflammation: The triple response

A simple method of demonstrating the hallmarks of the acute vascular response to injury is to simply scratch your skin firmly with your thumbnail.

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37

Traditionally, the triple response refers to the generation of three cardinal signs of inflammation; redness, heat and swelling. However, sensory nerve endings in your skin will register the pain of this action (dolor). For a brief moment, usually a few seconds, the area of the scratch will turn white. This is due to vasoconstriction, contraction of the smooth muscle walls of nearby arterioles which transiently restricts blood flow to the site. This constriction is due to release of vasoactive factors, such as bradykinin and histamine, from cells in the dermis damaged in the scratch. This blanching of the wounded area is an initial response to limit blood loss in case of a significant wound. Following the blanching of the scratched site, the smooth muscle walls of the arterioles relax and blood flow is restored to the area. The increased flow of blood, designed to bring the circulating neutrophils to the area of damage, results in erythema, or reddening of the tissue due to the large numbers of erythrocytes in the congested capillary bed (rubor). This also results in a local warming of the inflamed tissue (calor). Over the next ten minutes or so, extravasation of plasma will occur, resulting in the slight swelling (tumor) around the scratched area. As the damage to your skin is minimal, the response will resolve rapidly without invoking Virchow’s loss of function. This demonstration may appear simplistic but as we have seen, there are a profound number of cellular signaling events that orchestrate this apparently simple process.

2.8

Chronic inflammation

Acute inflammation, in most cases, resolves with no adverse sequelae within a few days. However, in cases where the initiating stimulus remains, such as that encountered in prolonged exposure to chemical, physical or biological agents, acute inflammation will progress to chronic inflammation. There is no precise definition of what constitutes the transition to chronicity, though histologically, chronic inflammation is characterized by a frustrated cycle of tissue destruction and repair with repeated activation and recruitment of the macrophage and FBGC. This damage may well occur through a recapitulation of the acute inflammatory phase events (Kushner, 1998) underscoring the notion that acute and chronic are simplistic descriptors. Nevertheless, chronic inflammation usually results in destruction of the parenchyma and the stroma with a fibrous scar repair of the damaged area.

2.9

Conclusion and future trends

When taken in its entirety, the biological environment is an incredibly complex system. In the last 50 years, researchers have made great strides in unraveling the components of this system and the myriad of interactions in

38

Degradation rate of bioresorbable materials

both normal and pathological states. This deconvolution of physiological processes will continue at an ever increasing pace, providing insight into the molecular mechanisms by which the body functions and specifically the way in which the body deals with introduced materials. Increasing knowledge of the molecular mechanisms that play a role in inflammation provides possible therapeutic strategies to promote or reduce inflammatory responses specifically within the context of implanted resorbable polymeric materials.

2.10

References

agrawal cm and athanasiou ka (1997) Technique to control pH in vicinity of biodegrading PLA-PGA implants. J Biomed Mater Res 38:105–14. anderson jm, defife k, mcnally a, collier t and jenney c (1999) Monocyte, macrophage and foreign body giant cell interactions with molecularly engineered surfaces. J Mater Sci Mater Med 10(10/11):579–88. anderson jm, rodriguez a and chang dt (2008) Foreign body reaction to biomaterials; Seminars in Immunology, 20(2):86–100. andersson j, ekdahl kn, lambris jd and nilsson b (2005) Binding of C3 fragments on top of adsorbed plasma proteins during complement activation on a model biomaterial surface. Biomaterials 26(13):1477–85. attur mg, dave mn, tsunoyama k, akamatsu m, kobori m, miki j, abramson sb, katoh m and amin ar (2002). Osteoarthritis or osteoarthrosis: The definition of inflammation becomes a semantic issue in the genomic era of molecular medicine. Osteoarthritis Cartilage 10(1):1–4. bainton df (1999) Developmental biology of neutrophils and eosinophils. In Inflammation: Basic principles and clinical correlates, 3rd ed., Gallin, Goldstein and Snyderman (eds), publ. Lipincott, Williams and Wilkins, Philadelphia 13–34. bostman o and pihlajamaki h (2000) Clinical biocompatibility of biodegradable orthopaedic implants for internal fixation: A review. Biomaterials 21:2615–21. chen g and goeddel dv (2001) TNF-R1 signaling: A beautiful pathway. Science 31;296(5573):1634–5. collier to, anderson jm, brodbeck wg, barber t and healy ke (2004) Inhibition of macrophage development and foreign body giant cell formation by hydrophilic interpenetrating polymer network. J Biomed Mater Res 69(4):644–50. davies je (1985) Exoemission for Biomaterials Research. Japanese Journal of Applied Physics 24 Suppl. 24–4:43–47. dinarello ca (2007) Historical insights into cytokines. Eur J Immunol 37: S34–45. gerelchuluun t, lee yh, lee yr, im sa, song s, park js, han k, kim k and lee ck (2007) Dendritic cells process antigens encapsulated in a biodegradable polymer, poly(D,L-lactide-co-glycolide), via an alternate class I MHC processing pathway. Arch Pharm Res 30(11):1440–6. guyton, ac and hall je (2006) Textbook of Medical Physiology, 11th ed, Elsevier Saunder, Pennsylvania, US, 2006.

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han j and ulevitch rj (2005) Limiting inflammatory responses during activation of innate immunity. Nature Immunology 6:1198–205. henson pm (2005) Dampening inflammation. Nature Immunology 6:1179–81. hochrein h and o’keeffe m (2008) Dendritic cell subsets and toll-like receptors. Handb Exp Pharmacol 2008;(183):153–79. hu wj, eaton jw, ugarova tp and tang l (2001) Molecular basis of biomaterialmediated foreign body reactions. Blood 98:1231–8. karakecili ag and gumus¸derelioglu m (2008) Physico-chemical and thermodynamic aspects of fibroblastic attachment on RGDS-modified chitosan membranes. Colloids Surf B Biointerfaces 61(2):216–23. kovarik p, sauer i and schaljo b (2008) Molecular mechanisms of the antiinflammatory functions of interferons. Immunobiology 212(9–10), 895–901. kushner i (1998) Semantics, Inflammation, Cytokines and Common Sense; Cytokine and Growth Factor Reviews 9(3–4):191–6. levi m and van der poll t (2005) Two-way interactions between inflammation and coagulation. Trends in Cardiovascular Medicine 15:254–9. litman gw, cannon jp and dishaw lj (2005) Reconstructing immune phylogeny: New perspectives. Nat Rev Immunol 5(11):866–79. lwaleed ba, cooper aj, voegeli d and getliffe k (2007) Tissue Factor: A Critical Role in Inflammation and Cancer; Biological Research Nursing 9(2): 97–107. min b and paul we (2008) Basophils and type 2 immunity. Curr Opin Hematol 15(1):59–63. moffett d, moffett s and schauf c (1993) In Human Physiology, 2nd ed. Mosby, St. Louis. nilsson b, ekdahl kn, mollnes te and lambris jd (2007) The role of complement in biomaterial-induced inflammation. Molecular Immunology 44:82–94. noh h and vogler ea(2007) Volumetric interpretation of protein adsorption: Competition from mixtures and the Vroman effect. Biomaterials 28:405–22. rajakariar r, lawrence t, bystrom j, hilliard m, colville-nash p, bellingan g, fitzgerald d, yaqoob mm and gilroy dw (2008) Novel biphasic role for lymphocytes revealed during resolving inflammation; Blood 111(8):4184–92. rankin ja (2004) Biological Mediators of Acute Inflammation; AACN Clinical Issues 15(1):3–17. schmidt dr and kao wj (2007) The interrelated role of fibronectin and interleukin-1 in biomaterial-modulated macrophage function. Biomaterials 2007 Jan;28(3): 371–82. scott a, khan km, roberts cr, cook jl and duronio v (2004a) What do we mean by the term ‘inflammation’? A contemporary basic science update for sports medicine; Br J Sports Med 38:372–80. scott a, khan km, cook jl and duronio v (2004b) What is ‘inflammation’? Are we ready to move beyond Celsus? Br J Sports Med 2004 Jun;38(3):248–9. silverthorn du (2006) Human Physiology: An Integrated Approach; 4th edition; Pearson Education Benjamin Cummings, San Francisco CA. slack sm and horbett ta (1995) The Vroman effect: A critical review. In: Horbett TA, Brash JL, editors. Proteins at interfaces. II. Fundamentals and applications. Washington: American Chemical Society; 112–28.

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so ak, varisco pa, kemkes-matthes b, herkenne-morard c, chobaz-peclat v and gerster jc (2003) Arthritis is linked to local and systemic activation of coagulation and fibrinolysis pathways. Journal of Thrombosis and Haemostasis, 1:2510–15. tefferi a (2001) Primary Hematology; Humana Press, Towtowa NJ. vroman l, adams al, fischer gc and munoz pc (1980) Interaction of high molecular weight kininogen, factor XII, and fibrinogen in plasma at interfaces. Blood 55(1):156–9. walport mj (2001) Complement. First of two parts. N Engl J Med 5;344(14): 1058–66. zarbock a and ley k (2008) Mechanisms and consequences of neutrophil interaction with the endothelium. Am J Pathol 172(1):1–7.

3 Synthetic bioresorbable polymers R. E. CAMERON and A. KAM VARI-MOGHADDAM, University of Cambridge, UK

Abstract: Synthetic bioresorbable polymers find wide medical application as temporary mechanical supports such as sutures, as tissue engineering scaffolds, and as mediators of release rate for the controlled release of drugs. This chapter outlines the main families of synthetic bioresorbable polymers, discusses the physical and chemical mechanisms by which they degrade and outlines the factors which can affect their rates of degradation. Key words: poly(alpha-hydroxy acids), degradation rate, polyglycolide, polylactide, degradation.

3.1

Introduction

Biomaterials are materials of natural or man-made origin that are used to direct, supplement, or replace the functions of tissues of the human body.1 The use of biomaterials dates back to ancient civilizations; artificial eyes, ears, teeth and noses were found on Egyptian mummies, and waxes, glues and tissue were used by the Chinese and Indians in reconstructing missing or defective parts of the body.2 Advances made in synthetic materials, sterilisation methods and surgical techniques have allowed progress to be made in the way biomaterials are used today. Biomaterials in the form of implants (such as sutures, screws, bone plates, joint replacements, heart valves, intraocular lenses and dental implants) and medical devices (such as pacemakers and biosensors) are regularly used to replace and restore damaged or dysfunctional tissues and organs within the body, thus helping to improve quality of life. Polymer chemical structure can be tailored to control degradation behaviour, making them, under physiological conditions, bioinert or bioresorbable over a defined period. Polymer degradation is generally denoted by a deterioration in the functionality of the polymeric material caused by a change in its physical and/or chemical properties.3 This chapter aims to present the different degradation mechanisms for synthetic bioresorbable polymers, highlighting factors influencing their degradation. 43

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Degradation rate of bioresorbable materials

3.2

Bioresorbable polymers

Synthetic biodegradable polymers were first used commercially as medical implants in the late 1960s with the introduction of the biodegradable suture material Dexon®.4 A steady progression in biodegradable polymer development has lead to a growth in their experimental and clinical use in the fields of orthopaedics and traumatology as fracture fixation devices, in the pharmaceutical industry as drug delivery devices and, most recently, as scaffolds for tissue engineering.2,5–7 The terms biodegradable, bioabsorbable, and bioresorbable are often used interchangeably to represent natural or synthetic polymers that degrade over time. However, the terms can be more precisely defined.8 Biodegradation can be thought of as the disintegration, erosion, dissolution, breakdown and/or chain scission of a polymer into metabolisable or excretable fragments in the human body, in animal models, or in ex vivo or in vitro test media, which represent, mimic or approximate the body environment. Williams9 emphasized the biological agent (enzyme or microbe) as being the dominant component in the degradation process. Bioresorbable polymers however, can be defined as those classes of polymers whose degradation products ‘resorb’ in the body, i.e. can be metabolised and enter the general metabolic pathways.8,9 These include poly(α-hydroxy acids) such as polyglycolide, polylactide and their copolymers. In contrast, bioabsorbable polymers are those which dissolve (or disperse) in biofluids, and are eliminated from the body without chain scission, as measured by molecular weight. Poly(vinyl alcohol) (PVAlc) and poly(ethylene glycol) (PEG) are examples of bioabsorbable polymers.8,9 A list of common synthetic bioresorbable polymers is given in Table 3.1.8 What follows is a short overview of the properties and degradation mechanisms for the most widely used of the polymers listed in Table 3.1, with particular attention paid to the poly(α-hydroxy acids).

3.2.1 Poly(α-hydroxy acids) Poly(α-hydroxy acids) are bioresorbable polyesters derived from α-hydroxy acids, HO-CHR-COOH, especially glycolic (GA, RKH) and lactic (LA, RKCH3) acids.10 Poly(lactide) (PLA) and poly(glycolide) (PGA) Poly(glycolic acid) is the simplest linear polyester and exists only in one form. Lactic acid on the other hand is a chiral compound, and thus exists under two enantiomeric forms, giving rise to different polymers, namely poly(L-Lactic acid) (PLLA), and poly(D-Lactic acid) (PDLA). The racemic mixture is signified by DL and is optically inactive (P-dl-LA).

Synthetic bioresorbable polymers

45

Table 3.1 List of common synthetic bioresorbable polymers (Arshady, 20038) Poly(α-hydroxy acids)

Polyglycolide Polylactides Polycaprolactone Malates

Other polyesters

Poly[(benzyl malate)-malate]s Poly-p-dioxanone Polyesters of diacids and diols

Other polycondensates

Polyanhydrides Polyorthoesters Polycarbonates Polyaminoacids Poly(amino-esters) Poly(amido-carbonate)s Polyphosphazenes Polyethers

Vinylic polymers

Poly(alkyl cyanoacrylate)s Poly(vinyl alcohol) Polyvinylpyrrolidone

The polymers can be made via two routes: direct polycondensation of lactic and glycolic acids or ring-opening polymerisation of the cyclic lactide and glycolide dimers. The nomenclature for polymers prepared by different routes is full of contradiction in literature, but polymers prepared from lactic acid by polycondensation are strictly referred to by the acid, as in poly(lactic acid), and those prepared by ring opening polymerisation by the dimer, as in polylactide (PLA).11 Direct polycondensation is the cheaper of the two routes, but the polymers produced have lower molecular weights and are more polydisperse than those produced by ring-opening polymerisation. High molecular weight polylactide and polyglycolide can be obtained by ring opening polymerisation of cyclic diesters using inorganic metal salts such as tin,12 aluminium13 and zinc11,14 as catalysts in the process. The temperature, time, concentration of catalyst and concentration of the chainlength determining agent control the molecular weight of the final polymer. PGA synthesis is possible through simple polycondensation of glycolic acid with antimony trioxide,15,16 but the resulting polymer has low molecular weight, and optimum properties are not obtained. Figure 3.1 shows the syntheses and chemical formulae for the poly(α-hydroxy acids). LA-GA polymers are biocompatibile and bioresorbable, and wide ranges of physical, thermal, mechanical and biological properties can be covered by varying the chemical and configurational structures in the polyester chain.2,11,15,17–19

46

Degradation rate of bioresorbable materials O HRC

O

H

n

O CO C O

H

C C

O

CO

C

CHR R

O

R

n

(R = H, glycolide ; R = CH3, lactide) 3.1 Chemical formula of polylactide and polyglycolide.

Table 3.2 Summary of homopolymer properties for poly (α-hydroxy acids) (adapted from Van de Velde & Kiekens 20025 and Nakamura et al., 198935) Polymer Melting temp (°C) Glass transition Modulus temp (°C) (GPa)

Degradation time (months)

PGA PLA PLLA P-DL-LA PCL

6 to 12 Several years >24 12 to 16 Several years

220–230 150–162 170–200 Amorphous 59–64

35–45 45–60 55–65 50–60 −60

6.0–7.0 0.35–3.5 2.7–4.14 1–3.45 0.2–0.4

The homopolymer of L-lactide (PLLA) exhibits high tensile strength (50–70 MPa) and low elongation (∼4%), and consequently has a high modulus (3 GPa)4,6,11 that makes it suitable for load bearing applications such as orthopaedic fixation and sutures. PLLA is partially crystallisable with a melting temperature of 175–178°C and a glass transition temperature of 60–65°C.6,20 It has the longest degradation time of the poly(α-hydroxy acids) requiring between 2–3 years to be completely absorbed.21,22 However, extensive in vitro degradation studies10,23–25 have shown that copolymerising PLLA with dl-Lactide and glycolide results in considerably shorter degradation times. Poly(D,L-lactide), P-dl-LA, is an amorphous polymer exhibiting a random distribution of both isomeric forms of lactic acid, and accordingly is unable to arrange into an organised crystalline structure. This material has lower tensile strength, higher elongation, and a much more rapid degradation time (see Table 3.2), making it attractive as a drug delivery system.10,26 Polyglycolide (PGA) is partially crystallisable, with a high melting point (220–230°C) and a glass transition temperature of 34–40°C.4,27 It has a tensile strength, in the order of 57 MPa,6 and a tensile modulus ranging from 6–7 GPa.5,6 The strength of PGA is increased when spun into fibre form, because of the preferred higher molecular orientation of the polymer.28

Synthetic bioresorbable polymers

47

It is not soluble in most organic solvents, except for extremely fluorinated organics such as hexafluoroisopropanol.29 Copolymers of PLA and PGA tend to be less crystalline and less hydrophobic than the homopolymer and subsequently have shorter degradation times.30 Poly(lactide-co-glycolide) (PLGA) has been used extensively for drug delivery31,32 and device manufacture.4,30 Copolymers of glycolide with trimethylene carbonate (TMC), called polyglyconate, have been used as sutures (Maxon ®) and as screws.4 Polycaprolactone Polycaprolactone (PCL) was first synthesised in the 1930s by ring opening polymerisation of ε-caprolactone. PCL is partially crystalline, with a melting point of 59–64°C and a glass transition temperature of −60°C.4 It has a tensile strength of 16 MPa and tensile modulus of 0.4 GPa.6 PCL is highly hydrophobic and thus has longer degradation times than PLA (2–5 years) which makes it suitable for applications where long degradation times are required.7,8 Due to its low melting temperature, PCL is easily processed by conventional melting techniques and can be filled with stiffer materials (particles or fibres) for better mechanical properties. PCL scaffolds have been used for tissue engineering of bone and cartilage.7 Table 3.2 summarises some of the properties of the poly(α-hydroxy acids) mentioned thus far.

3.2.2 Other polyesters Of the common members of this group of polymers, poly-p-dioxanone and polyesters of diacids and diols will be covered briefly. Poly-p-dioxanone Polydioxanone (PDS) is a colourless, partially crystalline, poly(ester-ether), which is produced by ring opening polymerisation of p-dioxanone monomers. The ring opening polymerisation requires heat and an organometallic catalyst such as zirconium acetylacetone or zinc-L-lactate.8 PDS is crystallisable and has a glass transition temperature in the range −10 to 0°C.4 PDS is also the tradename for the degradable suture made from poly-p-dioxanone. Polyesters of diacids and diols Poly(propylene fumarate) (PPF) is an unsaturated linear polyester which degrades via bulk hydrolysis. The products of its degradation are propylene

48

Degradation rate of bioresorbable materials

glycol, poly(acrylic acid–fumaric acid), and fumaric acid which occur naturally as part of the Krebs cycle.8

3.2.3 Other polycondensates Polyanhydrides hydrolyse to their constituent dicarboxylic acids in aqueous media. They have been used and investigated as an important biomaterial for short-term release of drugs for more than two decades.33 They are produced by melt polycondensation of diacids and have glass transition temperatures in the range of 50–100°C.34 Due to their highly hydrophilic nature, polyanhydrides are highly susceptible to hydrolysis and degrade by surface erosion. Poly(sebacic acid) (PSA) is the most prominent example, and degrades relatively quickly within 54 hours in saline. Other polymers within the polycondensate group are also susceptible to hydrolysis and degrade predominantly via surface erosion. Polyorthoesters, polycarbonates, polyphosphazenes have been utilised on their own and as copolymers for drug delivery.4,8

3.3

Degradation of aliphatic polyesters

In vitro and in vivo degradation of aliphatic polyesters have been investigated by many authors.10,23–25,35 However, data have been generally collected from experiments carried out with compounds of different origins, for devices of different natures, shapes and sizes, and at different implantation sites including different tissues and animals, making comparisons between studies difficult. Nonetheless, in vitro studies have been shown to be a good model for simulating polymer degradation in in vivo conditions.36–38 The most widely adopted in vitro model is that of degradation in pH 6.9–7.4 phosphate buffer solution at 37°C. Once implanted in the body, mechanical bioresorbable devices should ideally maintain mechanical properties, allowing gradual transfer of load to the surrounding tissue before being fully degraded, absorbed and excreted by the body, leaving no trace. The micromechanisms of deformation are affected by changes within the miscrostructure with degradation.39 Factors such as the size of the crystals within the polymer, pore size and pore volume within the implant, and the overall shape and size of the implant affect the mode of degradation within polymers which can then be used to control the release profile in drug delivery applications.40–44

3.3.1 The chemistry of reaction mechanisms Simple chemical hydrolysis of the hydrolytically unstable backbone is the prevailing mechanism for polyester degradation, which depends on the pH

Synthetic bioresorbable polymers +OH

O H+ C

R1

R1

OR2

OH

C+

R1

OR2

OH OH

slow

C

R1

OH OH

R1

OH

C HO+R2

HO+R2

HOR2

O+H2

C

OR2

R2OH R1

OH

OH2

C+

R1

OH H+

C

C OR2

O R1

OH H2O

49

3.2 Acid catalysed ester hydrolysis mechanism (based on Sykes, 198645).

O–

O OH– R1

C OR2

R1

C

O

O OH

R1

OR2

C

OH

+ O(R2)–

R1

C

O–

+ HOR2

3.3 Base catalysed ester hydrolysis mechanism (based on Sykes, 198645).

of the solution and may be catalysed by an acid or a base. The acid-catalysed reaction mechanism is given in Fig. 3.2.45 The reaction is reversible and is displaced towards hydrolysis by an excess of H2O. This mechanism is characterised by acyl-oxygen cleavage, and the removal of the alcohol is in fact the slow stage of the reaction. The base-catalysed ester hydrolysis reaction mechanism is outlined in Fig. 3.3.45 This reaction has an irreversible acyl-oxygen cleavage mechanism, and the slow stage is the attack of the ester by the OH−.

3.3.2 Bulk and surface degradation Hydrolytic degradation is commonly categorised under bulk and surface erosion, in which bulk erosion may be further split into homogeneous and heterogeneous degradation. The mechanisms may be identified by their effects on the sample diameter during degradation. In surface erosion the rate at which water penetrates the device is slower than the rate of conversion of the polymer into watersoluble materials.46 Surface erosion results in the device thinning over time while maintaining its bulk integrity. Hydrophobic polymers in which the

50

Degradation rate of bioresorbable materials

chemical bonds are highly susceptible to hydrolysis undergo surface erosion (e.g. polyanhydrides and polyorthoesters). By contrast, in bulk erosion, degradation takes place throughout the whole of the sample and the rate of ingress of water into the sample is greater than the rate of degradation. Sample dimensions remain constant until a critical stage of degradation where there is dramatic mass loss. Homogeneous bulk erosion describes a gradual drop in molecular weight throughout the sample until a critical value is reached at which the material becomes soluble and erosion occurs. Heterogeneous erosion describes a situation where the central region of the polymer becomes more degraded than the surface over time. The PLA/GA polymer system degrades through a mechanism of bulk heterogeneous degradation. The system can be subdivided into three classes of material, depending on the material’s crystallisability. In this section the notation PLAXGAY, where X refers to the percentage of L-lactide, Y refers to the percentage of glycolide units and consequently (1-X) refers to the percentage of D-lactide units will be used.47 (a) Intrinsically amorphous PLAXGAY that cannot crystallise with degradation (b) Intrinsically amorphous PLAXGAY that crystallises with degradation, OR partially crystallisable PLAXGAY that has been quenched to the amorphous state (c) Semicrystalline PLAXGAY In the first category, (a) the sample is considered to be completely amorphous and in a homogenous state. Water uptake is the first event that occurs upon implantation. The diffusion of water into the sample causes hydrolytic cleavage of the ester bonds throughout the sample. Initially, surface degradation exceeds the degradation in the centre of the sample as there is more water present at the surface. As hydrolysis proceeds, the molecular weight of the sample falls, and degradation products (monomers and oligomers) form. The degradation products at the surface are rapidly dissolved in the surrounding fluid and removed from the bulk polymer. The inability of large polymeric degradation products to diffuse away from the bulk device results in a local acidic environment in the interior of the implant. The increased acidic environment catalyses further degradation resulting in accelerated hydrolysis of the ester linkages in the interior, a process known as autocatalysis.48 Meanwhile, the surface continues to degrade at its original rate, resulting in a surface–centre differentiation. Over time, the degradation at the centre continues and the inner material can become a viscous liquid of oligomers contained by an outer shell also known as a ‘skin’ layer. The hollow skin layer eventually degrades away at a reduced rate. Li et al.10,23,24 demonstrated that both in vivo and in vitro, large PLA/GA devices degrade heterogeneously, with the degradation being faster in the

Synthetic bioresorbable polymers

H2O

51

Soluble oligomers

H2O

t

H 2O

0

H2O

ions?

5

Time in weeks

Water absorbtion + Degradation without loss of matter Beginning of the loss of matter

10

15

Central degradation + Formation of the outer membrane Formation of hollow structure

3.4 Schematic representation of the diffusion–reaction phenomena that occur during the hydrolytic degradation of implants, according to the model of heterogeneous degradation (modified from Grizzi et al., 199549).

internal part than at the surface where a layer of less degraded material is formed. Studies carried out by Grizzi et al.49 confirmed the formation of an outer skin layer as demonstrated in Fig. 3.4. The thickness of the skin, t, depends on many factors such as the diffusion rates of the various species involved and the rate of ester bond cleavage. Diffusion coefficients of the soluble oligomers in PLA/GA depend on factors such as molar mass, degree of matrix swelling, and macromolecular conformation and rigidity. In the same way, degradation rates depend on the degree of swelling, and on the sequential distribution of chiral and achiral units along the polymer chains. Finally, the release of the soluble carboxyl-terminated oligomers depends on their solubility in the surrounding aqueous medium and thus on factors such as pH, ionic strength, temperature and buffering capacity. The second class of materials considered, (b), are those which are initially amorphous but crystallise with hydration and degradation. As before, once placed in an aqueous medium, water diffuses into the sample from the outside in and initially surface degradation exceeds the degradation in the centre of the sample. During the second step of degradation for this class of material, autocatalysis predominates in the centre of the sample. However,

52

Degradation rate of bioresorbable materials

in this case as molecular weight decreases, factors such as preferential degradation of some units, shorter chain length and plasticisation by water allow crystallisation to occur in the centre of the sample. As hydrolysis continues, the volume of crystalline material in the centre of the sample increases, while the surface degrades at a reduced rate, becoming only slightly crystalline. The centre of the sample reaches stable crystallinity and becomes very resistant to further degradation. Degradation occurs in the amorphous regions, so in the final stages, the centre is often found in the form of a white powder or a porous structure. A dense outer layer remains in which cracking may occur due to shrinkage of the bulk sample on degradation.24 Li et al.24 found this mechanism in initially amorphous PLLA samples, in which hydrolytic chain scission resulted in lower molecular weight and absorbed water acted as a plasticiser, both resulting in increased chain mobility, and the ability of the chain segments to crystallise, which in turn improved the material’s resistance to degradation. Schwach et al.38 also confirmed the ability of the PLAXGAY polymers to crystallise with degradation. The third category of material to be considered is (c) where a PLAXGAY polymer is fundamentally crystallisable and has a measure of crystallinity before degradation. When placed in an aqueous environment, water diffuses into the sample. Hydrolytic cleavage of ester bonds starts immediately where water is present, and so initially surface degradation exceeds degradation in the centre of the sample. At later time points, a slight heterogeneity occurs as the outer surface remains resistant to degradation (since degradation products diffuse into the surrounding medium), while the central material undergoes autocatalysis with preferential degradation in the amorphous regions. Such samples are less likely to show a hollow structure. Li et al. found crystalline PLLA to degrade in this manner.24 It is important to note that in all these instances water diffuses into the amorphous regions first because they are less organised and they allow water to penetrate more easily than the highly ordered, densely packed crystalline regions.25,50,51 Commercially available polylactide and polyglycolide devices and sutures degrade by the bulk erosion mechanisms described above.46 Knowledge of these mechanisms has been used to tailor the degradation rates of biodegradable polymers. For example, replacing carboxylic acid end groups with ester end groups (the end group being determined by the choice of initiator) on poly(DL-lactide-co glycolide) (DLPLGA) polymers slows both water uptake and degradation rate in vitro as reported by Tracy et al.52 The acid end groups add to the hydrophilicity of the polymer and catalyse degradation. A study conducted by Middleton et al.53 in which the ester end groups in DLPLGA rods were replaced with poly(ethylene glycol) (mPEG) demonstrated enhanced water uptake without accelerated

Synthetic bioresorbable polymers

53

degradation. This was because the presence of ethylene glycol units enhanced polymer hydrophilicity without lowering the pH of the local environment. By increasing water uptake it may also have allowed the acidic degradation products to diffuse away from the interior of the rod more easily. The incorporation of acidic and basic groups within the polymer matrix allows control of the local pH of the system during degradation.54–59 In such cases, the control of the autocatalytic effect dominates over the effect of changing the reaction product concentration, and the effect of end-group neutralisation is to lower the reaction rate. Renouf-Glauser et al.60 observed an increase in the degradation rate of PLLA with the addition of lauric acid – an acid found in soaps, cosmetics and food additives. Lauric acid was found to catalyse the hydrolytic degradation, reducing the time to loss of tensile strength. The addition of inorganic and ceramic fillers (such as CaCO3, HA, TCP) to bioresorbable polymers has been widely reported in literature.61–64 Both the basicity and the solubility of the inorganic phase play a role in degradation of the polymers. The presence of such fillers is generally thought to retard polymer degradation by neutralising the acidic degradation products during polymer hydrolysis. The autocatalytic effect associated with the accumulation of acidic products in the internal part of the polymer device is delayed if not hindered by the presence of ceramic fillers.63

3.3.3 Reaction erosion fronts Work carried out by Hurrell and Cameron,41,43,65 and Milroy et al.66 extends the theories of Li et al., by paying particular attention to polyglycolide degradation. This series of work suggest that in PGA reaction, erosion fronts contribute to a four-stage degradation process (Fig. 3.5). In stage I, small quantities of water diffuse rapidly into the sample, reaching an equilibrium level (less than 1 wt%). Little further water is absorbed in Stage II but the hydrolysis causes the molecular weight of the polymer to decrease. Insertion secondary crystallisation, in which new crystals are inserted between existing crystals, takes place in this stage.41,43 This is facilitated by both the presence of water molecules and the decrease in molecular weight, which causes the polymer chains to become more mobile. At the beginning of Stage III, a critical molecular weight is reached where oligomers begin to diffuse from the surface of the sample. This creates space into which water molecules diffuse. A cooperative diffusion of water molecules in and oligomers out gives rise to sharp reaction–erosion fronts close to the surface in a mechanism comparable to that for the creation of the surface layer as proposed by Li et al.10 and Grizzi et al.49 for PLA/GA polymers. Behind the fronts, the polymer is hydrated and porous, whilst ahead of the fronts, oligomers have yet to diffuse out. During Stage III the fronts move linearly towards the

54

Degradation rate of bioresorbable materials I

II

Mw

III

IV

3.5 A schematic diagram illustrating the hypothesised water distribution in polymer samples during stages I–IV (modified from Hurrell et al., 200343).

centre of the sample. The co-operativity of the mechanism has analogies to Case II diffusion. The front movement is a particular feature of PGA which seems to differ from other members of the PLA/GA family in which the division between layers is more static with time. Presumably the differences arise because of the particular balance of kinetics between diffusion and reaction rates within the system. Stage IV begins when the fronts meet at the centre of the sample. After this point, the degradation becomes more homogeneous throughout the now highly-porous sample. The linear movement of fronts leads to the potential for zero order drug release, since active agents embedded in the sample are released quickly from the porous regions behind the fronts and the kinetics of release are determined by front movement.40,43,66

3.4

Factors affecting aliphatic polymer degradation

The degradation rate of bioresorbable polymers depends on their intrinsic properties such as reactivity, hydrophilicity, molecular weight, degree of crystallinity, and glass transition temperature. However, other external factors such as the degradation media, sterilization and sample size also play a role in the degree of degradation.

3.4.1 Polymer composition Many literature sources can be found in which the interplay between aliphatic polyester composition and degradation is considered.19,23,30,67,68

Synthetic bioresorbable polymers

55

T½ Months

6

4

2

0

0 100

PLA PGA Copolymer ratio

100 0

3.6 Half-life of PLA and PGA homopolymers and copolymers (from Middleton and Tipton4).

Particular attention is given to PLA, PGA and their copolymers. Figure 3.6, shows the schematic variation in the degradation half-life of PGLA.4 PLA shows the longest degradation time of the PLA/PGA series, with the degradation time decreasing with increasing amounts of PGA, as random copolymerisation decreases the degree of crystallinity of the polymers and because of the faster hydrolysis rate and greater hydrophilicity displayed by PGA compared with PLA.24 The half-life rises again when the copolymers contain mostly PGA as the polymer becomes more crystallisable with lower levels of copolymerisation. Amorphous regions are more accessible by water molecules than crystals, and this will increase hydrolysis and chain degradation in these regions.

3.4.2 Degradation media Media pH can affect polymer degradation, either directly, by altering the polymer chemistry involved during degradation, or indirectly, by interacting with degradation products and affecting consequent kinetics. Chu55 found there was a greater loss in the mechanical properties of PGA sutures in buffered compared with unbuffered solutions. Similarly, the presence of buffer ions increases the degradation rate of poly(α-hydroxy acid) copolymers as a result of the enhanced removal of degradation products through neutralisation of the end groups.23,54,60 Work carried out by King et al.50 and Hurrell and Cameron54 showed no significant change in PGA degradation rate for increases in ionic strength and for changes to phosphate-salt concentration (from 0.01 m to 0.1 m), but for increase in pH (from 6 to 8) a small but significant increase was observed.

56

Degradation rate of bioresorbable materials

Ginde and Gupta69 also observed alkaline media (pH 9.2 and 10.6) to have a more drastic effect on the degradation and mechanical properties of PGA fibres than acidic or neutral media. Makino et al.70 and Zaikov71 also reported greater degradation in alkaline compared with acidic conditions for homopolymers and copolymers of polylactide and polyglycolide. Their results imply that in alkaline conditions, surface erosion dominates over bulk erosion. Hurrell and Cameron54 concluded that the onset of reaction erosion fronts and the rate at which they move through the polymer depends on the rate at which acidic degradation products diffuse out, creating space for water in the structure. If the solution is made more alkaline or more concentrated, more of the acid is neutralised, driving the reaction forward. This hastens the onset of Stage III, and encourages drug release from loaded drug releasing polymers. Other authors23,54,72,73 also observed that the presence of ions in solution initially decreases water uptake into the polymer compared with the uptake from distilled water, but, as degradation products are released from the polymer the presence of ions increases the erosion rate.

3.4.3 In vitro and in vivo correlation There is some contradiction in the literature concerning the extent of the effect of enzymes on polylactide and polyglycolide degradation. For example, Vert et al.26 reported no difference in degradation between PLA50 (poly(DL-lactide) and PLA100 (100% L-lactide) in buffer and esterases solutions. However, Mason et al.74 reported a degradation rate between two and six times higher for PLA50 in plasma, which contains enzymes, compared with simple buffers. Despite these results, many authors have found in vitro and in vivo degradation rates for a range of PGLA copolymers to be the same.4,23,24

3.4.4 Crystallinity and polymer morphology Crystallinity has a significant effect on the degradation rate of aliphatic polyesters because it determines how easily water molecules can access the ester linkages to cause chain cleavage. Aliphatic polyesters can be random copolymers which are amorphous and do not crystallise during degradation; homopolymers and blocky copolymers can be semicrystalline or amorphous depending on processing; and polymers in between are amorphous in the undegraded state but become more crystalline as they degrade. Crystallisable polymers can be given different morphologies by quenching or annealing before degradation.15

Synthetic bioresorbable polymers

57

Li et al.10,23 examined morphological changes of amorphous polymers during degradation. In vitro investigations of PLA50 (poly(DL-lactic acid) and PLA37.5GA25 (75% DL-lactide and 25% glycolide) demonstrated that the polymers started and remained amorphous throughout degradation. However, PLA75GA25 (75% L-lactide and 25% glycolide) samples were crystalline after 7-weeks degradation. H-NMR studies (proton based Nuclear Magnetic Resonance Spectroscopy) suggested that PLA and PLA/ PGA copolymers all degrade via the same mechanism but that the GA units on PLA/GA copolymers constitute vulnerable points on the macromolecular chains, so degradation occurs preferentially on the GA bonds. This phenomenon was observed only up to 12 weeks degradation, after which point the lactide–glycolide ratio remained constant. Li et al. concluded that the glycolide removal caused crystallisation of the remaining lactide regions. The remaining glycolide was then incorporated into the new crystals protecting glycolide from hydrolysis after 12 weeks. Gel Permeation Chromatography (GPC) and Size Exclusion Chromatography (SEC) tests were carried out on 100% L-lactide (PLA100), 75% DL-lactide and 25% glycolide (PLA37.5GA25) and 50% DL-lactide and 50% glycolide (PLA25GA50) samples by Hakkarainen et al.25 Results exhibited development of a multimodal distribution in molecular weight during degradation for the semicrystalline PLA100 samples, whereas the amorphous copolymers degraded smoothly with no additional peaks in the molecular weight distribution. This was attributed to the difference in degradation rates between the amorphous and crystalline micro-domains in the semicrystalline polymer. Li et al.10,23 observed similar patterns for semicrystalline PLA. Gilding and Reed15 also observed partial crystallisation of PGA and PLA/PGA copolymers after 3 days degradation. It is clear that an increase in crystallinity during degradation need not be an indication of preferential amorphous material removal. Degrading PGA can show significant crystallinity increase in the early stages of degradation. Hurrell and Cameron41 reasoned that removal of amorphous material from the polymer structure should result in a significant loss of mass, which was not observed. Further analysis using Small Angle X-ray Diffraction (SAXS) deduced that insertion of secondary crystals within the amorphous region was the cause of the increase in crystallinity and decrease in long period. A slowing down of insertion crystallisation with time caused the subsequent levelling off of crystallinity and the rise in long period.

3.4.5 Molecular weight The molecular weight and molecular weight distribution of polymers affects the volume fraction of chain ends and hence the free volume of the polymer, and consequently, determines the chain mobility and crystallisation. It also

58

Degradation rate of bioresorbable materials

determines the number of acidic end groups that can participate in hydrolysis and their possible catalytic effect. Overall, the molecular weight affects both the chemical and physical properties of a polymer.44 Zhu et al.75 and Park76 both report faster degradation of PGA, PLA and their copolymers with decreasing molecular weight. This is attributed to an increase in the free volume and number of acidic reactive end groups. Processing techniques will also affect the molecular weight of polymer samples. The thermal treatments experienced in processing cause thermal degradation, which in turn results in a lowering of the molecular weight. Hurrell and Cameron65 concluded that for thermal degradation to occur, the thermal treatments in general need to be quite severe. They observed no molecular weight changes between PGA samples with no prior treatment and those treated at 160°C for 5 minutes, and at 120°C for 2 hours.

3.4.6 Sample size Some literature addresses the possible interplay between sample size and degradation of aliphatic polyesters. Sample size is commonly interchanged with sample geometry. Production of different shaped devices introduces different processing routes, which results in further changes to polymer degradation. Grizzi et al.49 and Ginde and Gupta’s69 work suggested that the degradation rate of PLA and PGA devices was very much dependent on size, and the greater the thickness of the device, the faster the degradation. In a separate experiment, Visscher et al.77 investigated the effect of size on the release of drugs in microparticles made of a copolymer of 25% Dlactide, 25% L-lactide and 50% glycolide. The microparticles ranged in size from 45–177 μm, and in accordance with the surface layer hypothesis the larger microparticles degraded first. Tormala et al.37 degraded SR-PGA rods with diameters ranging from 1.5–4.5 mm, which were prepared by sintering bundles of PGA sutures (Dexon) together. In these experiments, in contrast to those reported above, tensile measurements revealed that the larger rods degrade more slowly. This was attributed to the smaller surface area to volume ratio, which resulted in slower diffusion of water into the samples. Another factor that would have affected the degradation is a change in internal porosity of the rods with sample size. Hurrell et al.43 found that during PGA degradation the progression of reaction-erosion fronts through the samples was linear and at the same rate irrespective of sample size. Stage I of degradation is diffusion-controlled and is therefore strongly affected by the thickness of the sample. However, since this diffusion occurred quickly in comparison with the time scale of the degradation it did not have a strong effect. Once water had diffused

Synthetic bioresorbable polymers

59

through the bulk of the sample by the beginning of Stage II, hydrolysis occurred at the same rate regardless of sample size, until the critical molecular weight was reached. The reaction-erosion fronts (whose appearance marks the onset of Stage III) started at the same time in samples of different size. They concluded that it was only the onset of Stage IV of degradation (explained in Section 3.3.3), when the reaction erosion fronts meet in the centre of the sample, that showed size dependence, with the onset time increasing for thicker samples.

3.4.7 Sterilisation and packaging It is necessary to sterilise all medical implants after fabrication and before their surgical placement to reduce the risk of infection and associated complications. The most common sterilisation techniques utilise heat, steam, radiation, or a combination of these methods. PLA and PGA polymers, in addition to being susceptible to damage by moisture and radiation, are heat sensitive. Thus, selection of the correct sterilisation technique is of crucial importance to their physical and mechanical performance in vivo. Table 3.3 gives an overview of the sterilisation techniques commonly used and their advantages and disadvantages.19 γ-radiation sterilisation is known to cause chain scission in PLA and PGA polymers. At doses of 2.5 Mrad, Co60 γ-radiation causes deterioration of Dexon and Vicryl sutures.78 Other studies have reported decreases in tensile strength of PLA and PGA polymers and an increased rate of degradation upon γ-radiation79 as a result of a decrease in molecular weight. Table 3.3 Standard sterilisation techniques and their applicability to PLA-PGA (Athanasiou et al., 199619) Sterilisation technique

Advantages

Disadvantages

Steam sterilisation (high steam pressure, 120–135°C)

No toxic residues

Dry heat sterilisation (160–190°C)

No toxic residues

Radiation (Ionising or γ)

High penetration, low chemical reactivity, quick effect Low temperature range

Deformation/degradation due to water attack, limited usage for PLA-PGA. Melting and softening of polymers, not usable for PLA-PGA. Instability and deterioration, crosslinking/breakage of polymer bonds. Lengthy process due to degassing, residues are toxic.

Gas sterilisation (ethylene oxide)

60

Degradation rate of bioresorbable materials

The best method of sterilisation for polymers sensitive to heat and moisture, such as PLA and PGA, is chemical sterilisation by gases such as ethylene oxide (EO). However, this leaves residues in harmful quantities on the surface and within the polymers.19 Hence, the polymeric implants must then be subjected to degassing or aeration so that residual EO concentrations are reduced to acceptable levels.46 A further point to note is that polymers and their composites are fully compatible with modern diagnostic methods such as computer tomography (CT) and magnetic resonance imaging (MRI) as they are non-magnetic. Their radio transparency can be adjusted by adding contrast medium to the polymer.2 Peltoniemi et al.51 reported no change to the mechanical properties of biodegradable implants as a result of postoperative radiation doses.

3.5

Processing and devices

There are many medical and surgical devices of various shapes and sizes made of aliphatic polyesters. These devices are made by different processing routes. In general, large-scale devices such as sutures [e.g. Dexon (100%PGA), Vicryl (copolymer of glycolide in combination with L-lactide), Monocryl (copolymer of ε-caprolactone) or Maxon (copolymer of trimethylene carbonate)] and macroscopic implants used for bone fixation can be manufactured by solvent or melt spinning processes.11 The fibre forms can then be drawn under different conditions in order to orient the polymer chains. Fibres prepared by solvent-spinning usually have higher mechanical properties because of the thermal degradation during melt spinning. Micro and nano-particles used for oral administration of drugs are solvent cast using water or organic solvents. The stirring rate and temperature under which the particles are processed greatly affects their drug release rate and degradation properties. Porous biodegradable polymer scaffolds have been used as matrices for bone regeneration in tissue engineering. Current approaches for their fabrication include fibre bonding, solvent casting and particulate leaching, melt modelling, phase separation, emulsion freeze drying, gas foaming, and combinations of these.80

3.6

Conclusions

Bioresorbable polymers are put to extensive use as medical materials because of their diverse biodegradability, good mechanical properties and biocompatibility. The ability to tailor their chemical structures to control their degradation behaviour and rate is a great advantage when it comes to designing implants with suitable mechanical and degradational properties for their intended use.

Synthetic bioresorbable polymers

61

Poly(α-hydroxy acids) are the predominant group of polymers regarded as bioresorbable. Simple chemical hydrolysis of the ester backbone is the main chemical mechanism for their degradation whereas physical degradation is governed by a combination of homogeneous and heterogeneous degradation and the formation of reaction erosion fronts within the polymer systems. The lactide/glycolide bioresorbable polymers are thermoplastics which can be processed by many methods, including fibre spinning, extrusion, and injection moulding, which means they can be fabricated into a variety of wound closure items (e.g. sutures), implantable devices (e.g. bone plates, bone screws), and drug delivery systems, which include microspheres, fibres, films, rods and others. With the advancement of research, synthesis routes and implant manufacturing techniques, there is no doubt that bioresorbable polymers (PLA and PGA) will be used more and more as biomaterials to replace and/or augment damaged tissues within the body. The future for such polymers is bright and currently there is much attention paid to the suitability of using these bioresorbable polymers in solving difficult orthopaedic problems such as osteotomy fixation, articular cartilage and meniscal repairs,20 ligament and tendon reconstructions, and substitutes of autologous bone fills. Furthermore, a biodegradable implant system that provides a steady, controlled release of drugs or bioactive factors can be used as a delivery vehicle for substances that can enhance repair processes in the musculoskeletal system.

3.7

Sources of further information and advice

Encyclopedic Handbook of Biomaterials and Bioengineering: Part A: Materials – eds.: Wise, D. L., Trantolo, D. J., Altobelli, D. E., Gresser, J. D. & Yaszemski, M. J. (Marcel Dekker, New York, USA, 1995). The PBM Series: Biodegradable Polymers: Volumes 1 & 2. ed: Arshady, R. (Citus Books, London, U.K., 2003). The Biomedical Engineering Handbook – ed: Bronzino, J. D. (CRC, USA, 2nd edition, 1999). The Williams Dictionary of Biomaterials – ed: Williams, D. F. (Liverpool University Press, Liverpool, UK, 1999).

3.8

References

1 black, j. Biological Performance of Materials: Fundamentals of Biocompatibility (Marcel Dekker, New York, 1998). 2 ramakrishna, s., mayer, j., wintermantel, e. and leong, k. w., Biomedical applications of polymer-composite materials: A review. Composite Science and Technology 61, 1189–1224 (2001).

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4 Natural bioresorbable polymers W. PAUL and C. P. SHARMA, Sree Chitra Tirunal Institute for Medical Sciences and Technology, India

Abstract: Bioresorption is a desirable property for materials and devices used in tissue engineering and drug delivery. Various natural and synthetic bioresorbable polymers are utilized for biomedical and drug delivery applications. Most natural polymers, such as cellulose, chitin, chitosan, alginates, hyaluron, dextran and pullulan, are bioresorbable. Bioresorption takes place by chemical degradation mediated by water, enzymes, or microorganisms. Bioresorbable polymers are increasingly being used to replace metallic orthopaedic devices where only temporary fixation of tissue is required. It is possible that bioresorbable natural polymer materials will reduce the need for synthetic polymers and thereby produce a positive effect both environmentally and economically. The present chapter provides a brief outline of work that is under way in the area of natural bioresorbable polymer research and development. Key words: natural polymers, bioresorption, toxicity, biomedical applications, pharmaceutical applications, chitosan, cellulose, alginates.

4.1

Introduction

In many medical devices, biodegradation or bioresorption is an undesired phenomenon whereas items such as sutures, bone plates, interference screws, tissue fasteners, staples and other tissue and fracture fixation devices should completely resorb at a rate allowing the repair device to maintain sufficient integrity only while the soft tissue or bone heals. Similarly for devices used in tissue engineering and drug delivery, bioresorption is a desirable property. Various natural and synthetic bioresorbable polymers are available for these uses. Natural polymers such as cellulose, natural rubber, silk and other proteins and starch were known even in the early nineteenth century. Most natural polymers are bioresorbable, e.g. cellulose, chitin, chitosan, alginates, hyaluron, dextran and pullulan. When a material is biodegraded, it decomposes (is chemically reduced) into substances excretable by natural biological processes. Bioresorption is the loss of substance by physiological or pathological means. However, in this case, some compounds may remain in the organism without being degraded. Bioresorbable polymers are increasingly being used to replace metallic orthopaedic devices where only temporary 67

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fixation of tissue is required. Applications include fracture fixation and sports medicine (e.g. ligament repair, meniscal repair). Major advantages of bioresorbables over metal implants are that they are replaced by host tissue, do not require a second surgical procedure to remove them, avoid metal-related complications such as stress shielding and corrosion, and allow targeted delivery of bioactives. However, current applications of resorbables are limited by their low load-bearing capacity and concerns over long-term response to the implants. Bioresorption of polymers takes place by chemical degradation mediated by water, enzymes, or microorganisms. The mechanisms of chemical degradation are usually mediated by cleavage of crosslinks between chains, cleavage of side chains, cleavage of polymer backbone or combinations of the above. In recent years, there has been a marked increase in interest in bioresorbable materials for use in orthopaedic and drug delivery applications. Many researchers are devoting their time to modifying traditional materials to make them more user-friendly, and into designing novel polymer composites out of naturally-occurring bioresorbable materials. A number of biological materials may be incorporated into bioresorbable polymer materials, the most common being starch and fiber extracted from various types of plants (chitin, chitosan, alginate and cellulose). It is possible that bioresorbable natural polymer materials will reduce the need for synthetic polymers, thereby producing a positive effect both environmentally and economically. An attempt has been made in this chapter to provide a brief outline of the work that is under way in the area of natural bioresorbable polymer research and development.

4.2

Chitin and chitosan

Chitin is a major structural nitrogenous polysaccharide (Fig. 4.1) found in invertebrate animals and lower plants. It is a dense, tough and semitransparent horny substance. It is noticeably present in the outer skeletons of arthropods in particular; for example, in the epidermis of crustaceans such as crabs and prawns. In plants, chitin is present in hyphae or spores of molds. It was first described by Braconnot in 1811 (Braconnot, 1811), as a distinct substance among those identified in plants. It occurs naturally in three polymorphic forms (Muzzarelli, 1977), namely α-chitin, β-chitin and γchitin. It is one of the most abundant polysaccharides found in nature, second only to cellulose. Chitin is being produced naturally worldwide in the order of 100 billion tonnes per year. It has been found to have applications towards wound healing, as films, sponges, mats and fibers. Chitin has exhibited a significant accelerating effect on the wound-healing process. Even coatings of chitin on silk and catgut show almost similar woundhealing activities to all-chitin fibers. Chitin-coated surgical gauzes also

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4.1 Structure of chitin, chitosan and cellulose.

demonstrate significant activity compared to uncoated gauze. Chitin is made up of acetylglucosamine groups, where an acetyl group is attached to individual glucosamine molecules. When these acetyl groups are removed by the process called deacetylation, chitin is converted to chitosan (Fig. 4.1). This process releases amino groups which imparts polycationic characteristics to the polymer. Chitosan (pronounced kite-o-san), the principal derivative of chitin, was discovered by Rouget in 1859 (Rouget, 1859). He found that, when chitin was boiled in a concentrated potassium hydroxide solution, a product was obtained that dissolved in dilute iodine and acids, unlike chitin that only stained brown. However, this chitin derivative was formally named ‘Chitosan’ only in 1894 by Hoppe-Seyler (Hoppe-Seyler, 1894). Chitosan [α (1∼4) 2-amino 2-deoxy β-D glucan], the deacetylated form of chitin, is a mucopolysaccharide having structural characteristics similar to glycosaminoglycans with a chemical formula (C6H11O4N)n. It is marketed under a variety of forms with different molecular weights and degrees of deacetylation, or as chitosan base or salt. It is produced in different parts of the world as a solution and in flake, powder, bead and fiber forms. High-purity chitosan

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(biomedical grades) is also available for biomedical applications including pharmaceutics and drug delivery. Chitosan is synthesized by deacetylation of chitin. Briefly, to prepare chitin, crab and shrimp shells are decalcified in dilute hydrochloric acid and deproteinated in a dilute sodium hydroxide solution. The chitin is then deacetylated to become chitosan in a concentrated sodium hydroxide solution at boiling point. The biomedical grade/ purified chitosan is prepared by repeating the deacetylation process. Pharmaceutical grade chitosan is deacetylated between 90 and 95% and the food grade between 75 and 80%. Chitin, chitosan and their derivatives have received much attention from scientists in different parts of the world. Due to chitin’s natural abundance and versatility, many investigations have focused on its properties and various applications. Although studies on chitin and chitosan were initiated in the early nineteenth century, most of the reports available today on their medical and pharmaceutical applications have been obtained only during the last couple of decades. Chitosan has been investigated in depth or applied to various industrial and medical applications (Chandy, 1990; Savage, 1995). Medical and pharmaceutical applications include drug delivery, wound-healing ointments and dressings, artificial skin, hemostatic agents, enzyme immobilization, dialysis membranes, contact lenses or eye bandages, orthopaedics, surgical sutures and dentistry. It has been reported in 2005 that the world market for chitin and chitosan flourished in the year 2004 with an annual market of US $2 billion (Frost and Sullivan Market Report, 2005). The major share, accounting for more than US $1.4 billion (about 75%), was used for health-care products. Given all these different applications, chitosan – a biodegradable material – is emerging as an important biomedical material of the future. Chitosan is used in pharmaceutical preparations due to its availability, good biocompatibility and other desirable characteristics. It is inexpensive and digestible, which makes it a promising vehicle for the development of drug delivery systems. The use of chitosan in the development of drug delivery preparations is based on experience with chitosan intragastric tablets and studies of chitosan-coated drug delivery systems (Paul, 2006b). Drugs dispersed in chitosan were found to be released at a constant rate, thus highlighting its potential as a sustained-release matrix. The dissolution of several poorly-soluble drugs was increased by kneading low molecular weight chitosan with the drugs. The significant increase in dissolution rates could be due to improved wetability, crystallinity or crystal size and shape. Chitosan also exhibited significant mucoadhesive properties (Lehr, 1992) which were, however, less pronounced than those of polycarbophil. The positive attributes of excellent biocompatibility and biodegradability, ecological safety and low toxicity, combined with versatile biological activities such as antimicrobial activity and low immunogenicity, and the

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possibility of preparing a variety of chemically or enzymatically modified products and processes for specific end uses, have provided ample opportunities for immense development (Jayakumar, 2006, 2007; Rinaudo, 2006; Kurita, 2006; Muzzarelli, 2005; Yi, 2005). It has become of great interest not only as an under-utilized resource but also as a new functional biomaterial of high potential in various fields (Kumar, 2004; Kurita, 1995; Hirano, 1999). In addition to these properties, the biodegradable nature of chitosan makes it a suitable candidate for drug delivery applications. The latter include the nasal, mucosal, ocular, oral, parenteral and transdermal delivery of drugs and the delivery of pesticides, nematocides and insecticides in agriculture.

4.2.1 Physicochemical properties Chitin is a type of polysaccharide composed of an amino sugar and consisting of a straight chain molecule by binding D-glucosamines in which each amino sugar is acetylated by β-1,4 linkage (Fig. 4.1). Among the three forms, α-chitin is the most abundant, crystalline, tightly compacted and stable form in which the chains are arranged in an antiparallel (one up, one down) fashion (Muzzarelli, 1977). In β-chitin the chains are arranged in parallel whereas in γ-chitin there are two ‘up’ chains to every ‘down’ chain. α-Chitin is found where extra hardness proves essential whereas β- and γ-chitin provide toughness, flexibility and mobility together with physiological functions such as control of electrolytes and polyanionic characteristics. In both α- and β-chitins, the chains are organized in sheets where they are tightly held by a number of intra-sheet hydrogen bonds (Minke, 1978; Blackwell, 1969; Ogawa, 2004; Yui, 1994, 2007). This tight network, dominated by the rather strong CH—O—NH hydrogen bonds, maintains the chains at a distance of about 0.47 nm (Blackwell, 1969). Such a feature is not found in the structure of β-chitin, which is therefore more susceptible than α-chitin to intracrystalline swelling (Minke, 1978). The current model for the crystalline structure of α-chitin indicates that the inter-sheet hydrogen bonds are distributed in two sets, with half occupancy in each set (Blackwell, 1969). These aspects make evident the insolubility and intractability of chitin. In chitin, the degree of acetylation is typically 90%, indicating the presence of some amino groups. As some amount of deacetylation might take place during extraction, chitin may also contain a small portion of amino groups (5–15%) (Campana-Filho, 2007; Dong, 2002). So, the degree of N-acetylation, i.e. the ratio of 2-acetamido-2-deoxy-D-glucopyranose to 2-amino-2-deoxy-D-glucopyranose structural units has a striking effect on chitin solubility and solution properties (Rinaudo, 2006). Chitosan is derived from chitin by removing and refining the acetyls through a process called deacetylation. In chitosan, the typical degree of acetylation is less than 35%. It is, thus, a copolymer composed of glucosamine and

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N-acetylglucosamine. The physical properties of chitosan depend on the molecular weight (from approximately 10 000 to 1 million Dalton) and degree of deacetylation (in the range of 50–95%), sequence of the amino and the acetamido groups and the purity of the product (Rinaudo, 2006). The crustacean shells (crabs, etc.) which are waste products (now byproducts) of the food industry are commercially employed for production of chitin and chitosan. It is believed that at least 1011 tons of chitin are synthesized and degraded, but only 150 000 tons of chitin is made available for commercial use (Van Luyen, 1996). Chitosan, the long unbranched polysaccharide, is similar to that of a cellulose derivative (Roberts, 1992) where amino groups have replaced the hydroxyl groups at the C2 position. The structure of chitin and chitosan is as depicted in Fig. 4.1. Their similarity with cellulose is obvious from the figure. Its reactive hydroxyl and amino groups can be modified chemically for various applications (Chandy, 1990). Chitosan molecules have a strong positive charge at pH of <6.5 that attracts negatively-charged molecules and ionically bond to them. They also form gels with polycations at low pH and have antacid and anti-ulcer activities that may prevent or reduce drug irritation in the stomach (Paul, 2006b) on oral administration.

4.2.2 Toxicity of chitosan Chitosan has been used by the water purification industry for more than 20 years. When spread over the surface of water it absorbs the impurities (grease, oil, heavy metals, etc.), thus making the pollutants easy to remove. The very properties that made chitosan so popular in waste purification have also made it a powerful fat-absorber in the human digestive system. Thus Chitosol capsules and Chitosan Plus+ tablets have been proposed as a dietary supplement in the case of obesity (each gram of chitosan can safely bind between 3 and 12 g of fat). The strong positive charge of chitosan causes it to bind negatively-charged substances such as fats, lipids and bile. The US Food and Drug Administration (FDA) have approved chitosan as a food additive in animal feed (McCurdy, 1992). Chitosan-fed animals exhibited significantly low blood cholesterol levels and significant inhibition of atherogenesis (Ormrod, 1998). Dietary chitosan also reduces the incidence of early preneoplastic markers of colon carcinogenicity (Torzsas, 1996). The oral LD 50 of chitosan in mice was reported to be more than 16 g/kg (Arai, 1968). Oral chitosan-coated dialdehyde cellulose has also been reported to decrease elevated urinary protein and blood urea nitrogen to normal levels in nine days following administration to experimental rats (Yagi, 1998). Chitosan as a food additive in normal and neonatal streptozotocin-induced diabetic mice significantly reduced blood glucose, cholesterol and triglyceride levels (Miura, 1995). This study suggests that chitosan alone may be useful for the treatment of lean type non-insulin-

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dependent diabetes mellitus. Glutaraldehyde-treated porcine pericardial substitutes exhibited weakest epicardial reaction and adhesion to the heart on subsequent chitosan treatment (Chanda, 1996). Chitosan treatment helped to stop the slow release of residual glutaraldehyde, which thereby reduced toxicity and helped to prevent calcification (Kuribayashi, 1996). Intramuscular injection of chitosan as micro spheres to rats confirmed the tissue compatibility of this compound (Jameela, 1994). Chitosan has antigenotoxic activity (Ohe, 1996). In vivo toxicity tests in rats have shown that chitosan is non-toxic and that the sterilized films are pyrogen-free (Rao, 1997). Thanks to its safe toxicity profile, chitosan is an interesting and promising excipient for the pharmaceutical industry. However, contradictory reports suggesting side effects following oral and parenteral administration to mice (Tanaka, 1997) and implantation in Wistar rats (Szymonowicz, 1997) have also been documented.

4.2.3 Biomedical and pharmaceutical applications Although restoration of tissue continuity after injury is a natural phenomenon, infection, quality of healing, speed of healing, fluid loss, and other complications that prolong the healing time are major problems in wound management. Extensive research has been carried out for the past 35 years on wound coverings. An ideal wound covering should mimic the properties of human skin. The principal reasons for applying a dressing can be summarized as follows: (i) to produce rapid and cosmetically acceptable healing; (ii) to remove or contain odor; (iii) to reduce pain; (iv) to prevent or combat infection; (v) to contain exudate; (vi) to cause minimum distress or disturbance to the patient; (vii) to hide or cover a wound for cosmetic reasons; or a combination of two or more of the above. Wound dressings are generally classified as (i) passive products, (ii) interactive products, and (iii) bioactive products, based on the nature of materials used. Traditional dressings such as gauze and tulle dressings that account for the largest market segment are passive products. Interactive products are comprised of polymeric films and forms, which are mostly transparent, permeable to water vapor and oxygen but impermeable to bacteria. These films are recommended for low exuding wounds. Bioactive dressings deliver substances active in wound healing either by delivery of bioactive compounds or because the dressings are constructed from material having endogenous activity. These materials include proteoglycans, collagen, non-collagenous proteins, alginates, or chitosan. Chitosan is a natural biopolymer known in the wound management field for its hemostatic properties. Furthermore, it possesses other biological activities and affects macrophage function that helps in faster wound healing (Balassa, 1984). It also has an aptitude to stimulate cell proliferation and

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histoarchitectural tissue organization (Muzzarelli, 1989). The biological properties including bacteriostatic and fungistatic properties are particularly useful for wound treatment. Similarly to alginate materials, there is also a number of references on chitosan in wound treatment (Prudden, 1970; Biagini, 1991; Minami, 1992a, b). A variety of chitosan-based skin graft material was tested in guinea pigs and rabbits (Jayasree, 1995). Faster wound healing was observed in these animal trials. This dressing was studied on human subjects with chronic non-healing ulcers (unpublished data). It was observed that the material made the ulcer cleaner and had a beneficial effect in the control of infection. The epithelialization process was faster. Application of this dressing was extremely satisfactory and has great scope in the management of chronic ulcers. The dressing possesses antibacterial and anti-inflammatory properties. Flexible, thin, transparent, novel chitosan–alginate polyelectrolyte complex (PEC) membranes caused an accelerated healing of incision wounds in a rat model compared with conventional gauze dressings. Closure rate and appearance of PEC membrane treated wounds were comparable with Opsite®-treated wounds (Wang, 2002). Application of the photo-crosslinkable chitosan hydrogel on fullthickness skin incisions made on the backs of mice significantly induced wound contraction and accelerated wound closure and healing compared with the untreated controls (Ishihara, 2001, 2002). Healing at split skin graft donor sites was studied with chitosan by dressing half with chitosan and half with a conventional dressing. This showed that chitosan facilitated rapid wound re-epithelialization and the regeneration of nerves within a vascular dermis. Early returns to normal skin color at chitosan-treated areas were demonstrated (Stone, 2000). Treatment with chitin and chitosan demonstrated a substantial decrease in treatment time with minimum scar formation on various animals (Minami, 1999). The biochemistry and histology of chitosan in wound healing have been reviewed by Muzzarelli (1999) and Feofilova (1999). A silver sulfadiazine-incorporated bilayer chitosan wound dressing showed excellent oxygen permeability, controlled water vapor transmission rate, and water-uptake capability. It exhibited excellent antibacterial activity in in vitro culture for 1 week (Mi, 2002). Chitosan has been frequently examined as a wound dressing material (Mi, 2001; Loke, 2000; Kim, 1999); however, a wound-dressing product based on chitosan is yet to be commercialized. Chitosan has been widely studied as a drug delivery matrix (Paul, 2006b). It has been looked at in the form of tablet delivery systems, as microspheres, microparticles, nanoparticles, beads or granules, and gels or films. Chitosan has been studied mainly as a diluent or filler in tablet manufacturing; however, it has also been studied as an adhesive, a disintegrator, and an anti-adherent. Chitosan has been used as an excipient for directly

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compressed tablets for various drugs such as aminopyrine–barbital complex, aspirin, diclofenac sodium, diltiazem, insulin, ketoprofen, nifedipine, prednisolone, propranolol hydrochloride, and theophylline. Chitosan microspheres were tested for the controlled release of antibiotics, antihypertensive agents, anticancer agents, protein and peptide drugs and vaccines, etc. Various cross-linking agents for specific applications can control chitosan microsphere degradation. Dissolution of chitosan in gastric pH could be controlled by cross-linking or by using chitosan derivatives. Microspheres with controlled particle size can be prepared using various techniques and can be modified appropriately for the oral, nasal, and parenteral delivery of drugs. However, most of the studies conducted with chitosan microspheres or granules involved oral delivery systems. Other forms of drug delivery systems based on chitosan are nanoparticles, gels, and films (Paul, 2006b; Cui, 2001). Other applications of chitosan include biodegradable sutures (Hirano, 1998, 2000), ophthalmic application, in dentistry, as a fat absorber (Ylitalo, 2002; Gallaher, 2000), as an anticoagulant (Vongchan, 2002), and as an antitumor agent. With its unique combination of biological, physical, and chemical properties, chitosan is widely used in both the industrial and medical fields. These properties produce a novel, versatile biopolymer that can be tailor-made to suit a specific application with the required modes of function. The development of, and commercial applications for, chitosan have expanded in recent years. There is great potential for exploitation of chitosan in the field of biomedical, pharmaceutical, food and beverage, cosmetic, agriculture, industry, and water treatment, as given in Table 4.1. With a wide range of potential applications in medicine, pharmaceutics, and industry, there is tremendous scope for future research on chitosan and its derivatives.

4.3

Alginates

Alginic acid is an unbranched binary copolymer of L-guluronic acid (G) and D-mannuronic acid (M). It is a polysaccharide found in brown seaweed (Phaeophyceae: Laminaria hyperborea, Macrocystis pyrifera, and Ascophyllum nodosum). It exists in the kelp cell wall as an insoluble mixed salt of calcium, magnesium, sodium, and potassium. The salts of this polysaccharide are generally named alginates. The two monomers (G and M) are linked together in blocks of poly G and poly M or randomly intermingled, depending on its origin, as shown in Fig. 4.2. The proportion as well as the distribution of the two monomers determines, to a large extent, the physicochemical properties of alginate. The first scientific studies on the extraction of alginates from brown seaweed were reported by a British chemist, Stanford, at the end of the nineteenth century. He found that the extracted

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Degradation rate of bioresorbable materials Table 4.1 Applications of chitosan Field

Application

Healthcare

Contact lens/eye bandages Wound-healing ointments and dressings (skin treatments) Orthopaedics Anticholesterol and fat-binding (hypo cholesterolemic agent) Surgical sutures Drug delivery Ophthalmology Dentistry Transportation of cells

Food and beverages

Anticholesterol and fat-binding Food stabilizer Flavors and tastes Food packing Nutritional additives Fruit preservation

Agriculture

Seed treatments (coating) Feed ingredients (animal feed) (controlled release formulations) Nematocides and insecticides

Cosmetics and toiletries (personal Hair treatment Skin care care) Oral care Waste and water treatment (clarification)

Sewage effluents Drinking water Recovering metals Pools and spas Treating food wastes (food processor waste)

Product separation and recovery (bioapplications)

Membrane separations Chromatographic matrix Immobilization of enzymes/cells Recover valuable bioproducts

substance, which he named algin, possessed several interesting properties. These included the ability to thicken solutions, to make gels, and to form films. From these, he proposed several industrial applications. However, large-scale industrial production of alginate was not introduced until 50 years later. Alginates are also synthesized by some bacteria (e.g. Azotobacter and Pseudomonas species). Industrial manufacture of alginate is based on the extraction of the polymer from brown algae. The M/G ratio will vary from one species of brown algae to another, as shown in Table 4.2.

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4.2 Structure of alginates.

Table 4.2 Typical M/G composition and structural sequences of various species of brown algae Seaweed

M/G

%M

%G

%MM

%GG

Laminaria hyperborean (stem) Laminaria hyperborean (leaf) Laminaria digitata Macrocystis pyrifera Lessonia nigrescens Ascophyllum nodosum Laminaria japonica Durvillea antarctica Durvillea potarum

0.45

30

70

18

58

1.22 1.22 1.5 1.5 1.86 1.86 2.45 3.33

55 55 60 60 65 65 71 77

45 45 40 40 35 35 29 23

36 39 40 43 56 48 58 69

26 29 20 23 26 18 16 13

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Degradation rate of bioresorbable materials

4.3.1 Physicochemical properties Alginate is a natural and biodegradable biopolymer (Smidsrød, 1996). Several studies have shown that cells immobilized in alginate maintain good morphology and metabolism during long-term culture. There are, however, numerous studies using uncharacterized commercial alginates to immobilize cells that have resulted in inconsistent outcomes. This is mainly related to different levels of fibroblastic overgrowth when used as an implantation device. It is known that alginates may contain small amounts of pyrogens, polyphenols, proteins, and complex carbohydrates. The presence of polyphenols might possibly harm the immobilized cells, and the presence of pyrogens, proteins, and complex carbohydrates may induce immunological reactions by the host. It has also been demonstrated that perfectly spherical and smooth alginate droplets may be formed only by using highly purified alginates. To avoid such problems, it is therefore important to use alginates of high quality. To obtain repeatable results in all respects, the chosen alginates should also be well characterized with respect to all critical parameters (impurities, M/G content, molecular weight, etc.) The primary functions of the alginates are as thermally stable cold-setting gelling agents in the presence of calcium ions, gelling at far lower concentrations than gelatin. Such gels can be heat-treated without melting, although they may eventually degrade. Gelling depends on the ion binding (Mg2+ << Ca2+ < Sr2+ < Ba2+), with the control of the dication addition being important for the production of homogeneous gels (e.g. by ionic diffusion or controlled acidification of CaCO3). High G content produces strong brittle gels with good heat stability (except if present in low molecular weight molecules) but prone to water weepage (syneresis) on freeze–thaw, whereas high M content produces weaker more-elastic gels with good freeze–thaw behavior. However, at low or very high Ca2+ concentrations, high M alginates produce the stronger gels. As long as the average chain lengths are not particularly short, the gelling properties correlate with average G block length (optimum block size ∼12) and not necessarily with the M/G ratio, which may be primarily a result of alternating MGMG chains. The future prospects are excellent since recombinant epimerases with different specificities may be used to produce novel ‘designer’ alginates. Alginate’s solubility and water-holding capacity depend on pH (precipitating below about pH 3.5), ionic strength (low ionic strength increasing the extended nature of the chains), and the nature of the ions present. Generally, alginates show high water absorption and may be used as low-viscosity emulsifiers and shearthinning thickeners. They can be used to stabilize phase separation in lowfat fat substitutes, e.g. as alginate/caseinate blends in starch three-phase systems. Alginate is used in pet food chunks, onion rings, stuffed olives, and pie fillings.

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4.3.2 Biomedical applications Wound dressings based on alginic material are well known in the art and in literature, as well as from the commercial point of view, in wound management. With calcium alginate being a natural hemostat, alginate-based dressings are indicated for bleeding wounds. The gel-forming property of alginate helps in removing the dressing without much trauma and reduces the pain experienced by the patient during dressing changes (Winter, 1962). It provides a moist environment that leads to rapid granulation and re-epithelialization. Alginate is highly absorbent, gelforming, hemostatic, and biocompatible. This makes alginate a good candidate for burns and wound management. Cytotoxic effects on proliferation of fibroblasts and epidermal cells in culture (Ikarashi, 1992; Rosdy, 1990; Sank, 1989; Fry, 1986; Suzuki, 1995) and residual debris with inflammation in healing wounds (Suzuki, 1995; Matthew, 1995; Lansdown, 1994) have been reported recently with commercial alginate dressings (Table 4.3). However, alginates are well known for their hemostatic properties. In a study (Doyle, 1996), calcium alginate increased the proliferation of fibroblasts but decreased the proliferation of microvascular endothelial cell and keratinocytes. It decreased fibroblast motility but had no effect on keratinocyte motility. There was no significant effect of calcium alginate on the formation of Table 4.3 Alginate-based wound dressings commercially Product

Manufacturer

AlgiDERM AlgiSite

Bard Smith & Nephew, Inc., Wound Management Division Johnson and Johnson Carrington Kendall Kendall Sherwood-Davis & Geck B. Braun Gentell Hyperion Medical, Inc.

Algosteril CarraSorb H CURASORB CURASORB zinc Dermacea FyBron Gentell Hyperion Advanced Alginate dressing KALTOSTAT KALGINATE Maxorb PolyMem Restore SORBSAN SeaSorb Tegagen HG Tegagen HI

ConvaTec DeRoyal Medline Ferris Mfg. Hollister Dow Hickam Coloplast Sween Corp. 3M Health Care 3M Health Care

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capillary-like structures by human dermal fibroblasts and microvascular endothelial cells. These results suggest that the effects of calcium alginate on cell proliferation and migration may have been mediated by released calcium ions, and the alginate tested in this study may improve some cellular aspects of normal wound healing, but not others. In a controlled clinical trial, patients who received calcium alginate were compared with patients who received paraffin gauze. The donor sites were assessed at 10 days postharvesting to determine if they were completely healed (100%) or not. A significant number of patients dressed with calcium alginate was completely healed at day 10 compared with few of the members of the paraffin gauze group. Thus calcium alginate dressings provide a significant improvement in healing split skin graft donor sites (O’Donoghue, 1997). In another study with burn patients, calcium alginate significantly reduced the pain severity and was favored by the nursing personnel because of its ease of care. However, there was no clinical advantage as a dressing for skin graft donor sites (Bettinger, 1995). The combined use of calcium sodium alginate and a bio-occlusive membrane dressing in the management of split-thickness skin graft donor sites eliminated the pain and the problem of seroma formation and leakage, seen routinely with the use of a bio-occlusive dressing alone (Disa, 2001). A study with various alginate wound dressings indicated that some alginate-containing dressings have the potential to activate macrophages within the chronic wound bed and generate a pro-inflammatory signal which may initiate a resolving inflammation characteristic of healing wounds (Thomas, 2000). It has been recommended that calcium alginate be used as the dressing of choice for split skin graft donor sites (O’Donoghue, 2000; Lyall, 2000). As an alternative to standard alginate dressing, new alginate dressings have also been trialled and claimed to have superior properties (Suzuki, 1999; Williams, 1999; Steenfos, 1998). Alginate wound dressing materials have been commercially utilized and have been reviewed widely in the literature (Seymour, 1997; Sedlarik, 1992; Piacquadio, 1992). The use of alginate as an immobilizing agent in most applications resides in its ability to form heat-stable strong gels, which can develop and set at room temperatures (Dulieu, 1999; Fraser, 1997). It is the alginate gel formation with calcium ions that has been of interest in most applications. However, alginate forms gels with most divalent and multivalent cations. Monovalent cations and Mg2+ ions do not induce gelation, while ions such as Ba2+ and Sr2+ will produce stronger alginate gels than Ca2+. The gel strength will depend upon the guluronic content and also on the average number of G-units in the G-blocks. Gelling of alginate occurs when divalent cations take part in the interchain binding between G-blocks, giving rise to a threedimensional network in the form of a gel. The binding zone between the G-blocks is often described by the so-called ‘egg-box model’ as shown in Fig. 4.3. During gelling, the concentration of divalent cations has a large

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Ca2+

more Ca2+

4.3 Egg-box model for alginate gel formation.

impact on the gel network and gel homogeneity. When the gelling takes place in the presence of excess calcium, a modified egg-box model has been suggested, involving more than two alginate chains in the gelling zone. This may have an impact on the porosity of the gel. If no shrinking of the gel occurs, there may be more space in between the chains, leading to an increased porosity. Because alginates may form strong complexes with polycations, such as chitosan or polypeptides, or synthetic polymers such as polyethylenimine, they may be used to stabilize the gel. When used as coating materials, such complexes may also reduce the porosity. Alginate gels have been found to be stable in a range of organic solvents and are therefore, in contrast to other hydrogels, potentially useful in applications involving entrapment of enzymes in nonaqueous systems. Chitosan and poly (L-lysine) coatings on alginate beads were almost totally inert to enzymatic hydrolysis by lysozyme, chitosanase and trypsin, chymotrypsin, or proteinase (Quong, 1999b). Alginic acid has been used to encapsulate living cells and this is perhaps the most widely used technique. The usage of alginate gel capsules is mainly because of the simple procedure employed and the mild environment for living cells. The living cells are encapsulated by dropping a suspension of live cells in aqueous alginate solution into a mild gelling bath. The automated procedure for this preparation, with uniform bead size and under sterile conditions, is now available. Polycations such as poly-L-lysine or chitosan can coat onto alginate beads by ionic interaction. This will give strength and reduce porosity of the beads. Chitosan coating also makes the beads bioadhesive. It has been shown that pancreatic islet

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Degradation rate of bioresorbable materials

cells encapsulated in polylysine alginate can function for more than two years (O’Shea, 1986). Applications involving cell immobilization in alginate as a device, or artificial organs, are insulin regulation in diabetes with islet encapsulation, treatment of hemophilia B, growth factor for treatment of dwarfism, treatment of kidney failure (urea removal), treatment of liver failure, interferon production, treatment of Parkinson’s disease, treatment of hypocalcemia, treatment of myocardial infarction, treatment of chronic pain, and study of tumor angiogenesis (Skja˚k-Bræk, 1996; Onsøyen, 1996). The polyionic complex based on alginate was used to encapsulate porcine islet and was studied for its ability to induce an antibody response in the recipient or to act as an adjuvant to antibody responses against antigens leaked from the capsule (Kulseng, 1999). Although it has previously been suggested that alginate rich in mannuronic acid (high M) is more immunogenic than alginate rich in guluronic acid (high G), this encapsulation did not protect against the generation of antibodies against islet-like cell clusters. However, somatic gene therapy using non-autologous recombinant cells immunologically protected with alginate microcapsules has been successfully used to treat rodent genetic diseases (Ross, 1999). Encapsulation of drugs in alginate matrix is a mild process, unlike other presently available devices where harsh solvents or extreme conditions such as high temperatures are involved, which have the potential of destroying the encapsulated drug. Calcium alginate has been extensively studied and employed in a number of pharmaceutical applications (Edwards-Levy, 1999; Fundueanu, 1999), especially in controlled drug release from microspheres, nanospheres, and polymeric films (Aynie, 1999; Imai, 2000; Acarturk, 1999; Takka, 1999). Alginate matrix has been studied for encapsulating a variety of drugs such as theophylline and indomethacin (Liu, 1999). In this study, it has been shown that alginate enhanced the absorption of drug across Caco-2 cell monolayers by about three times. Wheatley (1991) encapsulated protein drugs (myoglobin) in sodium alginate solution and sprayed into buffered calcium chloride solution to form cross-linked microcapsules. The drug-loaded capsules were coated with a final layer of poly (L-lysine). The biodegradable microspheres based on sodium alginate were used to encapsulate plasmid DNA, and it expressed LacZ in the intestine, spleen, and liver (Aggarwal, 1999). Alginate gels produced by an external and internal gelation technique gave a yield of double-stranded DNA of over 97% and 80%, respectively, regardless of the composition of the alginate. Homogeneous gels formed by internal gelation absorbed half as much DNAse as compared with the heterogeneous gels formed by external gelation. It has been shown that high alginate concentration, high G content, and homogeneous gels (internal gelation) result in the lowest bead shrinkage and alginate leakage. These characteristics appear best suited for stabilizing DNA during GI transit (Quong, 1998). Many studies have been conducted on calcium alginate for DNA encapsulation (Quong, 1999a).

Natural bioresorbable polymers

4.4

83

Cellulose

Cellulose is a complex carbohydrate (polysaccharide) consisting of 1000– 3000 (or more) glucose units in a linear chain structure (Fig. 4.1) that can pack into fibers of great tensile strength. Mammals (including humans) cannot digest cellulose, but bacteria in the rumens of cattle and other ruminants, and protozoans in the gut of termites produce enzymes that can break it down. Soil fungi can also break down cellulose. Its most important uses are in wood, paper, and fiber products, as an ethanol and methanol source, and specialized applications. Cellulose derivatives are used in plastics, photographic films, rayon fibers, cellophane, coatings, explosives (e.g. nitrocellulose), and foods (e.g. the stabilizer and thickener carboxymethylcellulose). Cellulose is the most abundant polysaccharide found in nature. It is distributed in its native form in trees, shrubs, and grass. In these materials, cellulose serves as a structural material within the plant cell wall. Commercially, cellulose is produced mainly from cotton and wood, which are composed of 90–99% and 40–50% cellulose, respectively. Cellulose is isolated from wood and cotton by a chemical pulping process (Hon, 1996). Cellulose is a structural polysaccharide derived from β-glucose which condenses through (1–4)-glycosidic bonds, with the formula (C6H10O5)n. Cellulose was discovered and isolated in the mid-nineteenth century by the French chemist Anselme Payen and, as of the year 2006, the estimated annual production was 1.5 × 109 tons (Kim, 2006). Cellulose is a straight chain polymer: Unlike starch, no coiling occurs, and the molecule adopts an extended rod-like conformation. In microfibrils, the multiple hydroxyl groups on the glucose residues hydrogen bond with each other, holding the chains firmly together and contributing to their high tensile strength. This strength is important in cell walls, where they are meshed into a carbohydrate matrix, conferring rigidity to plant cells. In contrast to starch, cellulose is also much more crystalline. Whereas starch undergoes a crystalline to amorphous transition at 60–70°C in water (as in cooking), cellulose requires 320°C and 25 MPa to become amorphous in water (Deguchi, 2006). Its basic monomeric unit is D-glucose, which links successively through a glycosidic linkage in the β-configuration between carbon 1 and carbon 4 of adjacent units to form long chain 1,4-glucans. It possesses one primary and two hydroxyl groups per unit of glucose. These hydroxyl groups can undergo addition, substitution, and oxidation reactions. The structure of cellulose is as shown in Fig. 4.1. Cellulose and its derivatives (methyl, ethyl, amino ethyl, hydroxy ethyl, hydroxy propyl, carboxy methyl, acetate, phthalate, nitrate, etc.) have been widely used for applications such as the dialysis membranes in artificial kidneys, membranes in plasmapheresis, and as drug delivery matrices for controlled release.

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Cellophane was the first semipermeable membrane used for hemodialysis back in 1944 by Prof. Willem J. Kolff. Hemodialysis is a procedure where the excess of water and uremic toxins are removed from the patient’s blood by dialysis for the treatment of chronic renal diseases. This is achieved by the diffusion of metabolites and waste products from blood through semipermeable membranes. Although cellulose membranes are comparatively less hemocompatible, such membranes and their derivatives are still being widely used as dialysis membranes because of their high permeability and mechanical stability. The hemocompatibility of cellulose membranes is improved by surface modification via various techniques, or by utilizing synthetically modified cellulose (Paul, 2006a). They are used in large scale for hemodialysis applications, as reported in the literature. Cellulose acetate membranes are used for plasmapheresis. It is a therapeutic treatment in which the plasma is separated from the whole blood through extracorporeal circulation, where the plasma contains pathogenic substances of high molecular weight. The plasma is then substituted with fresh plasma or protein solution. Here the pore size of the membranes is much larger than that for hemodialysis. The cellulose acetate membrane-based plasma filters are used clinically for the therapy of renal, hepatic, and immune diseases and for autoimmune diseases whose athogenic substances are of antibody type. Microcrystalline cellulose is a highly-crystalline particulate cellulose consisting primarily of crystallite aggregates obtained by removing amorphous (fibrous cellulose) regions of a purified cellulose source material by hydrolytic degradation, typically with a strong mineral acid such as hydrogen chloride. The acid hydrolysis process produces microcrystalline cellulose of predominantly coarse particulate aggregates, typically having a mean size range of about 15 to 40 microns. Microcrystalline cellulose is made up of a chain of about 250 glucose molecules in the form of a microcrystal. Microcrystalline cellulose is FDA-approved as generallyregarded-as-safe (GRAS) and the spheres have been studied as a solid embolic material for the permanent occlusion of blood vessels (Kai, 2000). With its high biocompatibility, and with few adverse effects, the material seems to be suitable for intravascular applications. Cellulose derivatives have been studied as drug delivery matrices. Ethyl cellulose has been studied as matrix tablets for theophylline, caffeine, and dyphylline, allowing the drugs to be released by a diffusion mechanism (Neau, 1999). Microcrystalline cellulose, sodium carboxymethylcellulose, hydroxypropylmethylcellulose, hydroxyethylcellulose, hydroxypropylcellulose, and cellulose acetophthalate have been studied for the sustained release of acetaminophen, heparin, steroid anti-inflammatory drugs, diclofenac sodium, ciprofloxacin hydrochloride, theophylline, ketoprofen, etc. (Chandy, 2002; Pillay, 2002; Ceschel, 2001; Saha, 2001; El-Kattan, 2000).

Natural bioresorbable polymers

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4.4.1 Cellulose application in hemodialysis Cellulose, a powerful and versatile polymer commonly used in dialysis, can be easily chemically modified to yield different products such as ethers, esters, and carbamates. Regenerated cellulose membranes are prepared by the cuprammonium process (i.e. regeneration of refined cotton linters of high purity and molecular weight from cuprammonium solution) (Cheung, 1997). These membranes are uniform in their pore distribution and are hydrophilic. Cuprophan is a popular regenerated cellulose membrane introduced three decades back and is excellent for its low molecular weight clearance. Cellulosic membranes are typically more hydrophilic than synthetic membranes. This is an advantage because the adsorption of proteins decreases as the hydrophilicity increases. In the presence of water, Cuprophan membranes form hydrogels. The solute diffusion occurs through the highly water-swollen amorphous regions. The primary advantages of Cuprophan membranes are their low cost, high tenacity in the wet state, ability to form thin membranes, high mechanical strength and chemical stability, and effective diffusive transport properties of small solutes (e.g. urea and sodium chloride) (Kamide, 1994; Galletti, 2000). Some of the drawbacks of Cuprophan membranes are their limited transport of middle molecules; the presence of hydroxyl groups, which initiate complement activation and transient leucopenia during the first hour of contact with blood; and the asymmetry of the membrane. The symmetry of a membrane affects the permeability of the membrane. Asymmetry allows high diffusion, thus asymmetrical membranes are better for passing small solutes. Cellulose acetate membranes are made from cellulose diacetate polymers and have the advantage of greater permeability to water and larger solutes than the Cuprophan membranes. Cellulose triacetate (Hemophan) is also used in hemodialysis. Hemophan is a diethyl aminoethyl ether of cellulose, with unsurpassed phosphate clearance and excellent hemocompatibility. This membrane is symmetrical (i.e. membrane porosity is similar throughout the entire thickness of the membrane). In a study with six different membranes (polyacrylonitrile, Hemophan (from two manufacturers), polysulfone, cuprammonium, and cellulose acetate), the cellulose acetate membrane induced a significant drop in mean platelet count, and a dramatical increase of 6-keto prostaglandin F1a was noted. Other membranes did not significantly influence prostanoid levels and platelet count (Verbeelen, 1991). All cellulosic membranes induce platelet activation during hemodialysis, although there are quantitative differences among them. Although Cuprophan induced the highest degree of platelet activation, Hemophan was the only cellulosic membrane that showed a degree of platelet activation similar to the biocompatible membrane polysulfone (Cases, 1997).

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4.4 Chemical structure of synthetically modified cellulose.

SMC is another cellulose-based low-flux membrane where aromatic benzyl groups have replaced some hydroxyl groups by ether bonds (Bowry, 1998). The chemical structure of SMC is shown in Fig. 4.4. The molecular mosaic of hydrophobic (benzyl) and hydrophilic (hydroxyl and cellulose) regions contributes to an optimum hydrophobic–hydrophilic balance to achieve a unique surface structure for hemocompatibility and innovative spacer yarn for optimal performance (Hoenich, 1997). To improve the blood compatibility of cellulosic membranes, the surface hydroxyl groups have been replaced by grafting polyethylene glycol (PEG) and compared with cellulose diacetate and polymethylmethacrylate. It has been suggested that derivatizing cellulose is a useful approach to improve the blood compatibility of the cellulose polymer. However, there has been great variability in the blood compatibility profile of derivatized cellulosic membranes (diacetate and PEG grafted), which seems to be because of the different type of structural modification, rather than the degree of hydroxyl group replacement (Sirolli, 2000). Excebrane (Terumo, Japan), a covalently bonded vitamin E-modified cellulose membrane is an efficient, blood-compatible membrane with no deleterious effects on β2-microglobulins or lipids (MacGinley, 2001; Dhondt, 2000; Sasaki, 2000). Through a reaction mechanism where the surface treatment is done during the fiber spinning process by dissolving the modifier (a synthetic copolymer) into the core solution, a layer of several hundred angstroms is formed only on the inner surface of the hollow fiber. The cellulose surface is bound to a fluororesin polymer and to an oleyl alcohol chain with complement activation–inhibition and platelet aggregation–inhibition effects, respectively. The vitamin E, which has an antioxidant function is bonded to oleyl alcohol with hydrophobic bonding. It has been hypothesized that oxidative DNA damage is a consequence of uremia. In a comparative study with cellulose (complement-activating membrane) and synthetic membranes such as polysulfone and polymethylmethacrylate (non-complement-activating), it has been concluded that

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Table 4.4 Cellulose and modified cellulose derivative membranes currently used commercially in hemodialyzers Membrane

Material

Cellulose Cellulose acetate Cellulose diacetate Cellulose triacetate Cuprammonium rayon Cuprophan Hemophan Saponified cellulose ester Synthetically modified cellulose (SMC) Excebrane Modified cellulose

Regenerated cellulose Esterified with 2.0 acetate Esterified with 2.5 acetate Esterified with 3.0 acetate Regenerated cellulose Regenerated cellulose Etherified with diethylaminoethyl (DEAE) Saponified cellulose acetate Benzyl group substitution of the hydroxyl groups Regenerated cellulose coated with vitamin E Cuprammonium rayon coated with PEG

complement-activating membranes (cellulose) further augment the oxidative DNA damage (Tarng, 2000a). However, with the vitamin E-modified, regenerated cellulose (CL-E) membrane, oxidative DNA damage was significantly less compared with cellulosic membrane, and this was similar to synthetic membranes (Tarng, 2000b). Some cellulose and cellulose-derived membranes currently available commercially for dialysis are shown in Table 4.4.

4.4.2 Degradation of cellulose (cellulolysis) Cellulolysis is the process relating to or causing the hydrolysis of cellulose (e.g. by cellulolytic bacteria, fungi or enzymes). Mammals do not have the ability to break down cellulose directly. Cellulose is not digestible by humans and is often referred to as ‘dietary fiber’ or ‘roughage’, acting as a hydrophilic bulking agent for feces. Typically, this ability is possessed only by certain bacteria (which have specific enzymes) such as Cellulomonas, and which are often the flora on the gut walls of ruminants such as cows and sheep, or by fungi, which in nature are responsible for cycling of nutrients. Some animals, particularly ruminants and termites, can digest cellulose with the help of symbiotic micro-organisms. The enzymes utilized to cleave the glycosidic linkage in cellulose are glycoside hydrolases including endoacting cellulases and exo-acting glucosidases. Such enzymes are usually secreted as part of multienzyme complexes that may include dockerins and cellulose binding modules, referred to in some cases as cellulosomes. Many cellulolytic bacteria, fungi or enzymes break down cellulose into shorter linked chains known as cellodextrins (Young, 1986).

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4.5

Conclusion

Just as important as the way in which a material is formed is the way in which it is degraded. A general statement regarding the breakdown of polymer materials is that it may occur by microbial action, photodegradation, or chemical degradation. By scientific definition, biodegradable refers to a biological mediated degradation process such as enzymatic and/or cellular processes. Bioresorption refers to a chemically mediated degradation process such as hydrolysis where the degradation products are then incorporated into normal metabolic pathways. Bioasbsorbable also refers to a chemically mediated degradation, but the degradation products are generally excreted through one of the body’s organ systems. All three methods are classified under biodegradation, as the end products are stable and found in nature. This article has reviewed the general biomedical applications of polysaccharides with special emphasis on cellulosics, chitin, chitosan, and alginates. Considering the toxicity profile and their auspicious properties – such as favorable interaction with cells and other biological microcomponents including DNA and its biodegradable properties – these polysaccharides appear to be very attractive toward many biomedical applications, from cell encapsulation to drug delivery systems and wound-care materials. Although many scientific studies have taken place in the last few decades, these materials may still be the most attractive materials of the 21st century. More studies are demonstrating and reinforcing the wider applications of these materials in years to come. Obviously, much more information remains unknown as far as the science of bioresorbable natural polymers are concerned with a wider scope of biomedical applications.

4.6

Acknowledgments

We are grateful to the Prof. K. Mohandas, Director, and Dr. G.S. Bhuvaneshwar, Head BMT Wing of Sree Chitra Tirunal Institute for Medical Sciences & Technology for providing facilities for the completion of this review. We are thankful to the laboratory staff and library staff for their assistance.

4.7

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5 Bioresorbable ceramics M. BOHNER, Dr Robert Mathys Foundation, Switzerland

Abstract: Ceramics are being increasingly used as biomaterials, particularly for bone defect filling. However, despite numerous studies in the field, there is still a poor understanding of the biological, chemical and physical aspects involved in ceramic bioresorption. To improve this situation, a comprehensive review of the topic is presented. Even though the main emphasis is set on calcium phosphates, other ceramics such as calcium carbonate and sulfate are also considered. The chapter is divided into eight sections. After a general introduction, the three main parameters involved in bioresorption, namely solubility, kinetics, and in vivo conversion, are discussed extensively. Examples, taken mostly from calcium phosphate materials, are provided to illustrate the presented concepts. These sections are followed by a section devoted to bioresorbable ceramics other than pure calcium phosphates. Then, bioresorption models are briefly discussed. Finally, future trends are presented and a general conclusion is provided. Key words: calcium phosphate, silicate, carbonate, sulfate, bone graft, solubility, resorption, degradation.

5.1

Introduction

In the last three decades, the number of articles devoted to bioresorbable ceramics has increased dramatically. This evolution is mainly driven by the rapid increase of the clinical use of ceramics, particularly in the field of bone defect filling. However, new discoveries have also contributed to it. For example, the discovery of calcium phosphate (CaP) cements1,2 has eased the handling and application of CaP bone substitutes as well as the synthesis of low-temperature CaPs such as dicalcium phosphate dihydrate (DCPD; CaHPO4·2H2O), dicalcium phosphate (DCP; CaHPO4), octocalcium phosphate (OCP; Ca8H2(PO4)6·5H2O), or the wide range of precipitated hydroxyapatites (PHA; see Table 5.1).3,4 These CaPs are particularly relevant for in vivo use because low-temperature CaPs can all be found in vivo contrary to traditionally implanted CaPs such as β-tricalcium phosphate (β-TCP; β-Ca3(PO4)2) and sintered hydroxyapatite (SHA; Table 5.1).4,5 Moreover, recent studies have shown the in vivo potential of DCPD,6–8 DCP,9,10 OCP,10,11 amorphous calcium phosphate (= ACP; Table 5.1),10 and PHA.12,13 Parallel to this evolution, various studies have demonstrated the importance of trace 95

Table 5.1 Main calcium phosphate compounds: the first six compounds precipitate at room temperature in aqueous systems; the last six compounds are obtained by thermal decomposition or thermal synthesis Name

Formula

Ca/P (molar)

Mineral

Symbol

In vivo behavior

Monocalcium phosphate monohydrate Dicalcium phosphate Dicalcium phosphate dihydrate Octacalcium phosphate Precipitated hydroxyapatite* Amorphous calcium phosphate

Ca(H2PO4)2 ⋅ H2O

0.50

–

MCPM

Biosoluble

CaHPO4 CaHPO4⋅2H2O Ca8H2(PO4)6⋅5H2O Ca10-x(HPO4)x(PO4)6-x(OH)2-x Ca3(PO4)2⋅nH2O where n = 3 − 4.5; 15–20% H2O Ca(H2PO4)2 α-Ca3(PO4)2 β-Ca3(PO4)2 Ca10(PO4)6(OH)2 Ca10(PO4)6O Ca4(PO4)2O

1.00 1.00 1.33 1.33–1.67 1.50

Monetite Brushite – – –

DCP DCPD OCP PHA ACP

Biosoluble Biosoluble Bioresorbable Bioresorbable Biosoluble

0.50 1.50 1.50 1.67 1.67 2.00

– – – Hydroxyapatite – Hilgenstockite

MCP α-TCP β-TCP SHA OXA TetCP

Biosoluble Biosoluble Bioresorbable Bioresorbable Bioresorbable Biosoluble

Monocalcium phosphate α-Tricalcium phosphate β-Tricalcium phosphate Sintered hydroxyapatite Oxyapatite Tetracalcium phosphate

* Tricalcium phosphate (TCP) Ca9(HPO4)(PO4)5(OH) is a sub-category of PHA.

Bioresorbable ceramics

97

amounts of ionized metals such as magnesium,14 silicon,15 strontium16 and zinc,17,18 on the regulation of mineralization. As a result, attempts have been made to incorporate these ions into CaP to modify their biological properties.19–28 In vivo studies performed with bioresorbable ceramics have shown that large differences exist between their in vivo properties.29,30 For example, calcium sulfate dihydrate (CSD; CaSO4·2H2O; Table 5.2) disappears within weeks from the implantation defect,31 whereas SHA would probably require centuries.32 Alternatively, some CaPs are converted to a less soluble phase, e.g. apatite, after in vivo implantation, hence strongly reducing their removal rate.10,33–35 These differences are partly related to geometrical aspects,36 and partly to chemical aspects.37,38 Numerous attempts have been made to determine the effect of geometry on bone ingrowth and bone substitute removal rate, either with an experimental39–44 or a theoretical approach.36 The effect of chemistry on the in vivo behavior of bioresorbable ceramics has also been subject to many studies, with either an in vitro,45–50 an in vivo30,51–57 or a theoretical approach.58–62 Despite this large number of publications, there is presently no broad understanding of the differences in in vivo behavior of various bioresorbable ceramics. This fact is due partly to a poor understanding of the kinetics and thermodynamics aspects, as well as to the absence of a clear terminology to describe the in vivo behavior of bone substitutes: the terms ‘degradation’, ‘resorption’, ‘erosion’, ‘absorption’, with or without the prefix ‘bio’ are all used,63,64 very often as alternatives in the same article.38,65 There are three important aspects when dealing with the in vivo behavior of ceramic bone substitutes. The first aspect is related to the solubility of the compound in the body: if the compound is soluble in certain physiological conditions, dissolution and hence removal can occur. The second aspect is related to the speed at which the particular ceramic is removed from the body, i.e. to its dissolution kinetics. Experiments show that a compound does not necessarily dissolve even though it is soluble.66 The third aspect is related to the tendency to convert into another compound via a dissolution–precipitation reaction. Taking the example of calcium phosphates, apatites are the most stable phases in physiological conditions. Therefore, all non-apatitic CaPs compounds implanted in bone tend to transform into an apatite. These three aspects, i.e. solubility, dissolution, and conversion, are strongly related but are still independent. A given bioresorbable ceramic might be soluble in the body, e.g. α-tricalcium phosphate (α-TCP; α-Ca3(PO4)2), but might require years to be removed due to its in vivo conversion, e.g. into an apatite. The topic of CaP bioresorption has been addressed extensively by Driessens et al. two decades ago,58–62 but considering the large evolution of the field, the more extended use of all bioresorbable ceramics, and still a lack

Table 5.2 Non exhaustive list of doped CaP ceramics and non-CaP ceramics proposed for in vivo use Name

Formula

Ca/P (molar)

Mineral

Symbol

In vivo behavior

Carbonated apatite Calcium carbonate53,54,95 Calcium carbonate93,95 Calcium carbonate Calcium magnesium silicate114,115 Calcium metaphosphate98,100–102 Calcium Potassium phosphate119 β-calcium pyrophosphate103–106 Calcium silicate109 Calcium silicate111 Calcium silicate – phosphate composite110, 122 Calcium sodium phosphate62,79 Calcium sodium potassium phosphate45 Calcium sulfate hemihydrate55 Calcium sulfate dihydrate57 Silicon-substituted tricalcium phosphate20,21 Silicon-substituted hydroxyapatite22

Ca8.8(HPO4)0.7(PO4)4.5(CO3)0.7(OH)1.3 CaCO3 CaCO3 CaCO3 Ca7MgSi4O16

1.96 – – – –

Dahlite Calcite Aragonite Vaterite Bredigite

CA – – – –

Bioresorbable Bioresorbable Bioresorbable Biosoluble ?

Ca(PO3)2

0.50

–

–

CaKPO4

1.00

–

–

Biosoluble to bioresorbable* Biosoluble

β-Ca2P2O7

1.00

–

CPP

Bioresorbable

CaSiO3 CaSiO3 CaSiO3 − Ca3(PO4)2

– – –

Wollastonite Pseudowollastonite –

– – –

Biosoluble Biosoluble Bioresorbable

CaNaPO4

1.00

Rhenanite

–

Biosoluble

Ca2NaK(PO4)2

1.00

–

–

Biosoluble

CaSO4 ⋅ ½H2O

–

Plaster of Paris

CSH

Biosoluble

CaSO4 ⋅ 2H2O

–

Gypsum

CSD

Biosoluble

Ca3.08(P0.92Si0.08O4)2 or Ca3(P0.9Si0.1O3.95)2

>3

–

–

Bioresorbable

Ca10(PO4)6−x(SiO4)x(OH)2−x

1.73 (x = 0.19)

–

–

Bioresorbable

*Depends on the degree of crystallinity

Bioresorbable ceramics

99

of understanding in the scientific community of the mechanisms involved in ceramic bioresorption, there is a need to review the present knowledge on the topic. This is the goal of this chapter, which is structured as follows: firstly, it discusses factors affecting the bioresorption of CaPs (listed in Table 5.1) and secondly, it looks at the behavior of other ceramics such as calcium carbonates, silicates, sulfates, as well as doped CaPs (listed in Table 5.2). Finally, the topic is summarized in a general conclusion.

5.2

Solubility

According to Driessens et al.,60 ‘the solubility-controlling phase in living bone is OCP and the presence of this phase regulates the CaP concentrations of the body by simple physicochemical equibilibrium’. As a result, all CaPs that are more soluble than OCP in physiological conditions, such as DCP, DCPD, α-TCP, tetracalcium phosphate (TetCP; Ca4(PO4)2O), monocalcium phosphate (MCP; Ca(H2PO4)2) and monocalcium phosphate monohydrate (MCPM; Ca(H2PO4)2·H2O) are expected to dissolve spontaneously, whereas CaPs that are less soluble than OCP in physiological conditions such as β-TCP, SHA, PHA can be dissolved only with an active involvement of cells. Cells such as osteoclasts and macrophages can indeed reduce the local pH value67 and as some ceramics such as CaPs or calcium carbonates are much more soluble in low pH values (Fig. 5.1),68 dissolution and hence 1E+01 1E+00

[Ca]Tot

1E-01 1E-02 1E-03 1E-04 1E-05 1E-06 3

4

5

6

7

8

9

10

11

pH 5.1 Solubility isotherms of calcium-based bone substitutes at 25ºC in water. At pH 7.4, the solubility curves are from top to bottom: MCPM, CSD, α-TCP, TetCP, DCPD, DCP, OCP, β-CPP (flat curve), CDHA (parallel to β-TCP), β-TCP, and SHA. The dashed line represents the concentration of calcium ions in human serum/blood (Turitto and Flack, 1998116).

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resorption can occur (see the paper of Koerten and van der Meulen for more details on the cellular events involved in bioresorption52). As the solubility of the ceramic is a key parameter in defining its mechanism of resorption, it is important to emphasize the difference between ceramics that are more soluble than OCP and ceramics that are less soluble than OCP. Ceramics that are more soluble than OCP are called hereafter ‘biosoluble’ and those that are less soluble than OCP are described by the term ‘bioresorbable’. In a first approximation, the removal rate of the CaP from the body can be assumed to be directly related to the CaP solubility. In that case, the removal rate should occur in the order: MCPM > TetCP > αTCP > DCPD > DCP > OCP > PA ≈ β-TCP > SHA In the literature, there is plenty of data showing that β-TCP is more bioresorbable than SHA.11,69 There is also evidence that the in vivo removal rate of DCPD is larger than that of β-TCP.33 Other data show that the in vivo removal rate of OCP is larger than that of β-TCP30 and HA.70 So far, two important factors affecting solubility have been neglected: the stoichiometry and the size of the crystals. These two parameters have indeed a large effect on solubility as described in the next few lines. Regarding stochiometry, Liu et al.71 observed a large variation of the solubility of PHA when the Ca/P molar ratio was reduced from 1.67 to 1.50. Regarding crystals size, most precipitated CaPs such as PHA are micro or nanosized. When crystals reach nanometric dimensions, the crystal surface energy is a nonnegligible fraction of the global crystal energy. As a result, the crystal is thermodynamically less stable, or in other words more soluble. This concept is expressed by the equation of Kelvin:72 S ⎞ 2Mγ = ⎝ So ⎟⎠ ρr

(m + n + o) RT ln ⎛⎜

[5.1]

where R is gas constant (8.3144 J·mol−1·K−1), T is the temperature (in K), S and So are the solubilities of the spherical and flat particles of a compound MmXnYo (for HA, i.e. Ca5(PO4)3OH, m = 5, n = 3 and o = 1), M and ρ are the molecular weight and the density of the compound (e.g. 0.502 kg/mole and 3100 kg/m3 for HA), γ is the interfacial energy between the liquid and the solid (two very distinct values have been reported for hydroxyapatite: 47 mJ/m2 73 and 105 mJ/m2 66), and r is the particle radius. An exemplar calculation using the smallest surface energy (47 mJ/m2) shows that a monodispersed SHA particle population with a radius of 68.6 nm (specific surface area (SSA) = 14.1 m2/g) has a 1% higher solubility, whereas a radius of 7.16 nm (SSA = 135 m2/g) present a 10% increase of solubility

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(compared to a flat surface). This effect explains why sintered hydroxyapatite, where crystals are in the microsize range, is generally quoted to be non-bioresorbable, whereas apatite CaP cements, where crystals are in the nanosize range (typically 10–20 nm), are generally quoted to be bioresorbable. Other CaPs, such as DCPD, have generally larger crystal sizes than apatite crystals (in the micrometer range), but the Kelvin equation has been considered to play an important role in calcium phosphate solubility and hence dissolution rate.66 So, to summarize, it appears that the bioresorption rate of implanted CaP materials depends on solubility, which itself depends on chemical (= composition) and physical (= crystal size) aspects. But, as previously mentioned, kinetics aspects also play a very important role, explained in the next section.

5.3

Kinetics

In vitro, numerous studies exist demonstrating the effect of kinetics on crystallization and dissolution.4,74 For example, thermodynamic data (solubility) predict that DCP and hydroxyapatite should be the only two phases potentially precipitating in an aqueous solution: DCP in acidic conditions (pH lower than 4.0–4.5) and hydroxyapatite at other pH values.68 Experimentally, DCPD and OCP are the first phases precipitating in acidic conditions (pH < 7).4 In terms of dissolution kinetics, it is well known that the dissolution rate of CaP materials decrease over-proportionally with the dissolution time as mentioned by Tang et al.66 The latter authors suggested that this effect was related to the nucleation kinetics of dislocations. Recently, Brunner et al.75 observed that α-TCP nanopowders obtained by calcining ACP could be very poorly reactive in certain conditions, even though α-TCP is the main component of most CaP cements. Clearly, kinetics effects play a role in CaPs dissolution but, so far, there is no report in the literature describing a relation between effects affecting the dissolution rate and bioresorption. One kinetic effect influencing resorption is related to the implant geometry: the implant size, shape and porosity have proved to have a very large effect and thus have received much attention, as reviewed recently.76 Specifically, the implant size and shape influence the total bioresorption time.36 Large interconnected pores with a diameter and interconnection size larger than 50–100 μm allow blood vessel and cell invasion and, as a result, faster bioresorption. Furthermore, microporosity is believed to play a large role in bioresorption: Klein et al.77 concluded one study by mentioning that ‘microporosity plays a more important role in the biodegradation rate of β-whitlockite (= β-TCP) ceramics than difference in the macroporosity degree’.

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Degradation rate of bioresorbable materials

5.4

In vivo transformation

Biosoluble CaPs, which are by definition soluble in vivo,60,62 can be removed from the body by simple dissolution and ionic transport. However, the ions released by the dissolution of these compounds can interact with body fluids and form new compounds that are less soluble.74 Based on this observation and similar observations with bioresorbable CaPs, Driessens62 concluded that ‘the solubility behavior of CaPs is, in general, not determined by their solubility product, but rather by transformation of the surface layer (in the case of ceramics of hydroxyapatite, β-TCP and probably also α-TCP) or even complete transformation (in the case of rhenanite – see Table 5.2). Hence the solubility product is not a good measure for the in vivo behavior of CaP implants’. Hydroxyapatite is the most stable CaP compound at physiological pH (Fig. 5.1). Therefore, all CaPs other than hydroxyapatite are thermodynamically unstable in such conditions and should eventually convert to hydroxyapatite or an apatitic compound (as hydroxyapatite can easily incorporate foreign ions such as carbonates, sulfates, Mg, Na, etc . . .). This type of reaction is, for example, seen in early mineralization: despite some controversies about the type of mineral precursor of bone (OCP, DCPD, or ACP depending on the author), all authors agree that the bone mineral precursor will eventually convert to an apatite.59,74 Looking at in vivo data, transformation into an apatite has been reported for TetCP, α-TCP, OCP, DCP and DCPD but not for β-TCP.10,33–35 More specifically, α-TCP readily converts to CDHA when placed in an aqueous solution.78 Suzuki et al.10 reported also that DCP, OCP and amorphous calcium phosphate (ACP) converted to an apatitic phase when these compounds were implanted into the subperiosteal area of the calvaria of BALB/c mice. In a recent study, Bohner et al.33 showed that the center of brushite CaP cement was transformed after 6 to 8 weeks of implantation into an apatite compound. Simultaneously to this transformation, there was a strong decrease in the removal rate, suggesting that the transformation provoked a decrease of the removal rate. In another study, brushite cements were found to rapidly transform into an apatite compound, hence leading to a very slow resorption rate.35 A similar behavior was found by Constantz et al.34 Therefore, and in summary, it appears that all phases that are biosoluble convert, at least partially, into an apatite upon implantation. Interestingly, the extent and location of the transformation is a function of the material and perhaps also of the location and the volume. As mentioned, Bohner et al.33 showed that one particular DCPD cement formulation started converting into PHA in its core after 6–8 weeks of implantation. In another DCPD cement formulation, the whole cement remnants were observed to be converted to carbonated apatite after

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24 weeks’ implantation.35,56 The latter authors postulated that different observations made in different in vivo models were related to a change of cement volume. Another formulation tested by Constantz et al.34 also converted to apatite, but the location of the conversion reaction was not mentioned. Interestingly, the rate of conversion was a function of the implant location. Driessens et al.79 observed a full transformation of rhenanite into apatite within 6 weeks of implantation. As carbonation of the implant proceeded centripetally from the outside to the inside, results suggest that transformation proceeded in a similar way. Presently, there is no understanding of the factors determining the location of the conversion reaction, but clearly, not only solubility but also kinetics and diffusion aspects must play an important role. To illustrate this comment, Suzuki et al.10 observed that DCP conversion to an apatite proceeded more slowly than that of OCP, even though DCP is more soluble than OCP. The conversion of CaPs into thermodynamically more stable compounds such as PHA may provoke biocompatibility problems, particularly for large implant volumes, as suggested by a study of Flautre et al. 56 Assuming that the end product of the conversion reaction is calcium-deficient hydroxyapatite (CDHA; Ca9(HPO4)(PO4)5OH), the following reactions can be written: 9CaHPO4⋅2H2O → Ca9(HPO4)(PO4)5OH + 3H3PO4 + 17H2O [5.2] 9CaHPO4+H2O → Ca9(HPO4)(PO4)5OH + 3H3PO4

[5.3]

9Ca8H2(PO4)6⋅5H2O → 8Ca9(HPO4)(PO4)5OH + 6H3PO4 + 37H2O

[5.4]

3α-Ca3(PO4)2 + H2O → Ca9(HPO4)(PO4)5OH

[5.5]

3β-Ca3(PO4)2 + H2O → Ca9(HPO4)(PO4)5OH

[5.6]

9Ca4(PO4)2O + 6H3PO4 → 4Ca9(HPO4)(PO4)5OH + 5H2O

[5.7]

From these reactions, it appears that TetCP is expected to basify the medium during conversion (Eq. 7), that α-TCP and β-TCP are expected to be converted without pH change (Eqs 5.5–5.6), whereas DCPD, DCP and OCP are expected to release phosphoric acid and hence decrease the pH (Eqs 5.2–5.4). As a decrease of pH increases DCPD, DCP and OCP solubility (Fig. 5.1), DCPD, DCP and OCP conversion to an apatite is expected to be autocatalytic. The latter transformations should proceed as long as the equilibrium pH between DCPD, DCP or OCP and the precipitated apatite has not been reached (close to pH 6 between DCPD and CDHA; Fig. 5.1). Similar considerations can be made for TetCP since the solubility of PHA increases at high pH value. In reality, the situation is different, as body fluids can buffer the release of acids or bases. Moreover, there is to our knowledge

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Degradation rate of bioresorbable materials

no experimental evidence showing that apatites precipitate below pH 6.5–7 in physiological conditions (1 Atm, 37°C), probably because the interfacial energy between apatite and an aqueous solution is very high,66 hence hindering apatite nucleation. Therefore, the conversion of TetCP, DCPD, DCP and OCP is expected to proceed at, or close to, physiological pH. But as mentioned previously, kinetic effects are expected to play a very large role. This explains why Grover et al.80 proposed to use a highly reactive TetCP powder for antimicrobial purposes. The present review of the literature shows that Driessen’s statement62 that ‘the solubility behavior of calcium phosphates is, in general, not determined by their solubility product’ appears to be correct for DCPD, DCP, OCP, and α-TCP. Interestingly, there are probably limitations to this comment as DCPD, DCP, OCP, and α-TCP in vivo transformation can probably be modulated. Rousseau et al.81,82 have indeed shown in vitro that the presence of a soluble magnesium salt in DCPD cement can postpone its conversion to OCP, which is often a precursor of apatite formation. This soluble magnesium salt is, in fact, a release system for magnesium ions, which is a strong inhibitor for OCP and apatite precipitation.66,83 Similarly, Grover et al.80 observed in vitro that a DCPD cement based on pyrophosphoric acid was not hydrolyzed to an apatite, unlike a DCPD cement based on phosphoric acid. Furthermore, Constantz et al.34 have shown that in vivo phase transformation depends on implant location, or, in other words, on factors related to kinetics (flow of body fluids).

5.5

Other bioresorbable ceramics

Beside CaPs, there are other bioresorbable ceramics of interest for biomedical applications. These include calcium carbonate, metaphosphate, pyrophosphate, silicate and sulfate, as well as CaPs doped with small amounts of impurities, such as Si or carbonates. In this section, the biological behaviour of these compounds is briefly reviewed.

5.5.1 Doped calcium phosphates Schematically, two types of approaches have been used to dope CaPs: a chemical and a biological approach. In the chemical approach, ions have been added to modify the solubility and hence the bioresorption rate of the compound. One example is the incorporation of carbonate ions into apatites to increase the solubility and hence the bioresorption rate.12 Another example is the use of fluoride in toothpastes to reduce the enamel solubility and hence reduce the risk of caries. In the biological approach, compounds such as Mg,14 Si,15 Sr16 and Zn17,18 have been added to bioresorbable ceramics to modify their biological activity.19–22,24–28,84 It would appear logical

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to believe that the latter strategy has to be applied to a compound that is fairly soluble to be potent. But even when poorly-soluble compounds such as sintered HA are doped, doping appears to modify the biological behaviour.22,84

5.5.2 Calcium carbonates The use of calcium carbonate has been triggered by the discovery that corals could be used with some success as bone substitute.85,86 Coral consists of aragonite, one of the three calcium carbonate polymorphs, viz. calcite, aragonite, and vaterite. The solubility of polymorphs is in the order calcite < aragonite < vaterite.87,88 Vaterite generally transforms readily to calcite when placed in water. The in vivo behavior of the various polymorphs was generally found to be similar: (i) all polymorphs were found to be bioresorbable;53,54,85,86,89–93 (ii) direct bonding was seen between calcite and bone;53,54,94 (iii) bone formation in calcium carbonate derived from marine corals (aragonite) was comparable to that of hydroxyapatite;90 (iv) coral crystallized in the aragonite or calcite form presented a similar degree of specific cytocompatibility.95 The main difference between aragonite and calcite seemed to be the interfacial reaction: various authors detected the formation of a calcium phosphate layer on aragonite,89,90 whereas no such layer could be found on calcite,53,54 hence suggesting that aragonite was biosoluble contrary to calcite. Interestingly, Combes et al.96,97 used the polymorphicity of calcium carbonate to design calcium carbonate hydraulic cements.

5.5.3 Calcium metaphosphate Calcium metaphosphate (CMP; Ca(PO3)2) has been proposed as scaffolding material98–101 or as reinforcing material for bone substitutes.93 In vivo results suggest that bioresorption is moderate: Grynpas et al.101 mentioned that ‘the degradation is rapid initially (within the first six weeks) and slows down thereafter’, whereas Baksh et al.100 observed that after two weeks of implantation ‘the morphology of the individual surface grains of the CMP scaffolds employed in vivo changed to a more rounded form’. In fact, the bioresorption rate can be controlled by playing with the degree of crystallinity.102

5.5.4 Calcium pyrophosphate Calcium pyrophosphate (CPP; Ca2P2O7) has been proposed as a bone substitute.103–107 Interestingly, in vivo results suggest that CPP could have a positive effect against osteoporosis, probably due to its chemical similarity to bisphosphonates.104 In terms of resorption, there is little relevant in vivo

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Degradation rate of bioresorbable materials

data. Kitsugi et al.107 compared the in vivo behavior of HA, β-TCP, CPP and TetCP and observed signs of bioresorption only for β-TCP and TetCP, suggesting that CPP is less bioresorbable than β-TCP. Sun et al.106 observed very little resorption of CPP but some resorption of Na-doped CPP. To correlate in vivo with in vitro data, Sun et al.106 measured the dissolution rate of CPP, Na-doped CPP, β-TCP and HA in various media at various pH values. Results were consistent with each other and showed the following order of dissolution rate: β-TCP > doped CPP > undoped CPP > HA.106

5.5.5 Calcium silicate There are various forms of calcium silicate such as wollastonite (CaSiO3), pseudowollastonite (CaSiO3), larnite (β-Ca2SiO4) and calcium olivine (γCa2SiO4).87 Initially, wollastonite was studied as part of a glass ceramic containing hydroxyapatite and fluoroapatite crystals.108 But authors started looking at the synthesis and behavior of pure wollastonite,109 and wollastonite – β-TCP composites.110 Moreover, new silicates were investigated such as pseudowollastonite,111 hydraulic cements based on tricalcium silicate (Ca3SiO5),112 Si-rich CaP nanocomposite,113 and bredigite (Ca7MgSi4O16).114,115 In terms of resorption, Ni et al.109 cultured wollastonite with cells and observed that the calcium concentration in the culture medium was close to 5 mM, i.e. larger than 2.4 mM, the concentration present in serum.116 These results suggest that wollastonite is biosoluble. The same should be true for pseudowollastonite since it is more soluble than wollastonite.87 Regarding Si-rich CaP nanocomposite, the fact that the compound can quickly dissolve in serum suggests that it is also biosoluble.117

5.5.6 Calcium sulfates Plaster of Paris (= calcium sulfate hemihydrate = CSH; CaSO4·½H2O) has a long history as bone substitute as its first use was already reported in 1892.55 The hydration reaction of CSH leads to the formation of gypsum or calcium sulfate dihydrate (CSD; CaSO4·2H2O): 2CaSO4⋅½H2O + 3H2O → 2CaSO4⋅2H2O

[5.8]

Whereas CSH has a solubility close to 50 mm, that of CSD is close to 15 mm.118 For comparison, the calcium concentration in serum is close to 2.4 mM.116 There are two types of CSH: α-CSH and β-CSH. These two forms possess a different crystallographic structure and, as a result, a slightly different solubility (α-CSH is slightly less soluble than β-CSH87). Despite various differences between α-CSH, β-CSH and CSD, all calcium sulfates are biosoluble, i.e. freely soluble in physiological conditions. As a result, implant removal is expected to occur chemically rather than biologically

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through cell mediation. But as seen with the more soluble CaPs, all calcium sulfates can potentially be transformed after implantation into a calcium phosphate due to the presence of phosphate ions in physiological solutions and the release of calcium ions during calcium sulfate dissolution. In vivo, gypsum has been observed to disappear very fast, much faster than brushite cements.31 To decrease the removal rate and improve the biological result, calcium sulfates are generally applied as dense pellets57 or composites with HA granules have been proposed.48

5.5.7 Rhenanite Rhenanite (CaNaPO4)62,79 and two related compounds, i.e. (CaKPO4)119 and Ca2NaK(PO4)2,45 have been proposed as bone substitutes. These compounds have a high solubility and hence are biosoluble. This feature resulted, after in vivo implantation, in a full transformation of rhenanite into apatite79 and in poor compatibility with CaKPO4.119

5.6

Modelling resorption

The previous sections of this chapter show that the in vivo behavior of CaP materials can be fairly well understood with the help of solubility data. However, is it possible to predict the in vivo behavior based on solubility data? In other words, is it possible to predict how much faster DCP would be removed from the implantation site compared to β-TCP (keeping the geometry constant)? The traditional approach to predict the in vivo behavior of CaP bone substitutes is to make dissolution experiments, and use the latter results for the prediction. For example, Chow et al.46 determined the dissolution rate of CaP compounds in simulated serum at pH 3 using a dual constant composition titration system. These authors found out that DCPD dissolved three times faster than an apatitic cement. Ducheyne et al.47 measured the dissolution rate of a large range of CaPs in TRIS buffer at pH 7.3. The dissolution rate (in terms of calcium concentration) was found to increase in the order: HA < CDHA < OXA < β-TCP < α-TCP < TetCP. Fulmer et al.48 and Berger et al.45 also used TRIS buffer to compare various apatitic cements and CaP compounds, but at pH 7.4. LeGeros49 used a 0.1 M potassium acetate solution at pH 6.0. Finally, Koerten and van der Meulen52 placed their ceramics in a sodium acetate buffer at pH 4.0, 4.8 and 5.6. As can be seen from these studies, it is not clear what pH value and what solution should be chosen. The physiological pH is close to 7.4, but lower pH values are found within cells responsible for the removal of most calcium phosphates, such as osteoclasts and macrophages. For example, a study of Silver et al.67

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reports a pH value around or below 3.0 for osteoclasts and around 3.6–3.7 for macrophages. Therefore, it would make sense to simulate the in vivo behavior by measuring the dissolution kinetics at a pH in the range of 3 to 7.4. But which pH value? Chow et al.46 used a pH of 3.0 and observed that DCPD was dissolved only 3 times faster than SHA, which appears to be too small a difference compared to the in vivo reality.29,120 In fact, it is likely that neither pH 7.4 nor pH 3.0 are adequate pH values to model the in vivo behavior of CaP. The removal of a bioresorbable CaP occurs via a dynamic process where there is: (i) a decrease of the pH value from physiological range to an acidic pH, e.g. pH 3–4,67 at which the CaP becomes thermodynamically unstable (= soluble), and (ii) dissolution of the ceramic with release of neutral or basic ions. For example, dissolution of DCP and DCPD releases Ca2+ and HPO4− ions. Therefore, if osteoclasts do not produce enough acid to lower the pH values to dissolve the CaP, CaP dissolution and hence resorption does not occur. So, it appears that the simulation of the resorption process requires a knowledge of the rate of acid production of osteoclasts (as a function of pH value) which is unfortunately not the case. If this acid production rate was known anyway, resorption rate could probably be modelled by measuring the dissolution rate of a CaP when placed in an acidic environment where acid is produced according to this acid production rate function. Practically, a setup similar to that of Chow et al.46 could be taken, but the initial pH for the measurement should be pH 7.4, and the pH should not be kept constant, but should be decreased by adding continuously some acid. The rate of acid addition should be a function of the solution pH.

5.7

Future trends

The present document has reflected many trends in the field of bioresorbable ceramics. However, three major trends can be seen. First, a more adequate balance between resorption and bone formation is sought. Whereas the main focus of the 1970s and 1980s was set on hydroxyapatite, a shift has occurred in the past ten years towards the use of more bioresorbable materials that provide a faster turnover from a bone defect to a fully mature and functional bone. As a result, much effort has been placed on bioresorbable materials such as β-TCP, OCP, DCP and DCPD, and on newer materials or revisited materials such as, for example, calcium silicate or calcium carbonate. The second trend is to dope the materials to modify the biological response of these materials upon implantation, in particular in terms of bioresorption. Finally, the third trend is to optimize the structure of bioresorbable materials to better control their bioresorption rate. Such an optimization has been eased by the discovery/development of rapid prototyping techniques.121

Bioresorbable ceramics

5.8

109

Conclusion

The in vivo fate of ceramics is very complex and depends on (i) biological factors such as body fluid flow or local pH value (produced by, e.g. osteoclasts), (ii) chemical factors such as ceramic solubility and dissolution kinetics, and (iii) physical factors such as crystal size and geometry. Schematically, ceramics can be removed either via direct dissolution if the compound is soluble in an extracellular environment such as serum, or via cell-mediated dissolution if the compound is soluble in an intracellular environment, or after a combination of in vivo dissolution, precipitation in a new poorlysoluble phase, and bioresorption. The first and last mechanisms are strongly controlled by kinetics, whereas the second mechanism is controlled by cellular activities. For applications where the mechanical integrity of bone should be restored as soon as possible, an ideal bone substitute should be able to bind tightly to the bone at the edges of the defect, to restore the mechanical properties of the defect, and eventually be replaced by new bone to make sure that fatigue failure of the bone substitute does not occur. Presently, there is no such material available. Therefore, the strategy that appears most adequate is to use a bone substitute that resorbs as fast as bone grows. To prevent the formation of a large gap between bone and the bone substitute, the bone substitute should be either bioresorbable, i.e. resorbed by cell processes, or close to the bioresorbable domain (serum almost saturated towards the ceramic). In that respect, ceramics such as DCPD, DCP, OCP, β-TCP, precipitated apatites, calcite and aragonite appear to be the most interesting because serum is slightly under or over saturated towards these compounds, and the degradation products are present in the physiological environment. Among these materials, β-TCP is the most predominantly used ceramic, but recently proposed techniques that ease the synthesis of DCP, DCPD, and OCP macroporous granules and blocks could change this fact in the future.

5.9 1 2 3 4 5 6 7 8

References r. z. legeros, a. chohayeb and a. shulman, J Dental Res, 1982, 1482, 343. w. e. brown and l. c. chow, J Dental Res, 1983, 62, 672. m. bohner, Injury, 2000, 31 Suppl 4, 37–47. j. c. elliott, Structure and Chemistry of the Apatites and other Calcium Orthophosphates, Elsevier, Amsterdam, 1994. r. z. legeros, Adv Dent Res, 1988, 2, 164–180. e. munting, a. mirtchi and j. lemaître, J Mater Sci Mater Med, 1993, 4, 337–344. k. ohura, m. bohner, p. hardouin, j. lemaître, g. pasquier and b. flautre, J Biomed Mater Res, 1996, 30, 193–200. f. m. tamimi, j. torres, i. tresguerres, c. clemente, e. lopez-cabarcos and l. j. blanco, J Clin Periodontol, 2006, 33, 922–928.

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6 In vitro physicochemical test methods to evaluate bioresorbability S. LI, University Montpellier I, France

Abstract: Degradable aliphatic polyesters deriving from lactides, glycolide and ε-caprolactone present great interest for applications in the fields of surgery, tissue repair and regeneration, and sustained drug delivery. The hydrolytic degradation of aliphatic polyesters is a complex process involving many physico-chemical phenomena, namely water uptake, ester bond cleavage, neutralization of carboxyl endgroups at the surface, autocatalysis inside, crystallization, and diffusion and solubilization of soluble oligomers. These phenomena can be monitored by using various in vitro physicochemical test methods including visual examination, environmental scanning electron microscope (ESEM), proton nuclear magnetic resonance (1H NMR), size-exclusion chromatography (SEC), X-ray diffraction (XRD), differential scanning calorimetry (DSC), electron spray injection mass spectrometry (ESIMS), capillary zone electrophoresis (CZE), weight loss, and mechanical properties. Factors such as polymer morphology, chemical composition and configurational structures, MW and MW distribution, size and porosity, and chemically reactive additives, can strongly affect the degradation of polymers. This chapter aims to present the state-of-theart on the basis of recent advances in this domain. Key words: polylactide, polyglycolide, polycaprolactone, hydrolysis, degradation, autocatalysis.

6.1

Introduction

Synthetic polymers, or plastics, have been widely used in many domains of our society due to their excellent properties such as good strength, toughness, lightness and barrier properties. During past decades, surgeons and pharmacists have been trying to utilize polymers as biomaterials for both permanent and temporary therapeutic applications (Hoffman, 1984; Dunn; 1995; Li, 1999a). Permanent applications require biostable polymeric materials that are resistant to degradation in the body. In contrast, temporary applications necessitate materials for a limited period of time only, i.e. the healing time. In this regard, degradable polymers became of great interest in surgery, in pharmacology as well as in tissue engineering. The 117

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first degradable synthetic polymer was poly(glycolic acid) or polyglycolide (PGA), which was synthesized in 1954 (Charles, 1954). This polymer was first discarded because of its poor thermal and hydrolytic stabilities, which precluded any applications. However, it was realized later that one could take advantage of the hydrolytic sensitivity of PGA to make polymeric devices that can degrade in a human body, which led to the first bioresorbable suture material made of a synthetic polymer (Schmitt, 1967; Frazza, 1971). It is worth noting that terminology is one of the sources of confusion in the field. Nowadays, people tend to use the word ‘degradable’ as a general term and reserve ‘biodegradable’ for polymers which are biogically degraded by enzymes introduced in vitro or generated by surrounding living cells. A polymer able to degrade and to have its degradation by-products assimilated by or eliminated from a living system is then designated as ‘bioresorbable’ (Li, 2002). Degradable and biodegradable polymers generally contain labile linkages, namely ester, orthoester, anhydride, carbonate, amide, urea or urethane, along the chain backbone (Albertsson, 1995; Mochizuki, 1996). Aliphatic polyesters are particularly interesting because of their outstanding biocompatibility and variable physical, chemical and biological properties (Holland, 1986; Vert, 1992; Li, 1999b). The main members of the aliphatic polyester family are listed in Table 6.1. In this family, PGA, poly(lactic acid) (PLA), and poly(ε-caprolactone) (PCL) are the most attractive and the most investigated. A number of products have reached the stage of clinical use, such as Dexon®, Vicryl®, Maxon® and Monocryl® sutures, Lactomer® and Absolok® clips and staples, Biofix® and Phusiline® plates and screws, as well as Decapeptyl®, Lupron Depot®, Zoladex®, Adriamycin®, Capronor®drug delivery devices (Dunn, 1995). High molecular weight (MW) PLA, PGA, PCL and various copolymers are obtained by ring opening polymerization of corresponding cyclic monomers, i.e., lactides, glycolide and ε-caprolactone, respectively (Reed, 1979; Spinu, 1996). The direct polycondensation of hydroxy acids such as lactic acids and glycolic acid leads to low MW oligomers only (Fukuzaki, 1989; Wang, 1997), although post-solid polycondensation allows one to obtain high MW polymers (Moon, 2001). In the case of LA-containing polymer chains, chirality of LA units provides a worthwhile means of adjusting bioresorption rates as well as physical and mechanical characteristics (Miller, 1977; Vert, 1984; Nakamura, 1989). The use of PGA homopolymer is limited to suture material because of its high crystallinity and absence of practical solvents. In contrast, PLA, PCL and various copolymers have been widely used to make implants, microparticles, nanoparticles and scaffolds in tissue engineering. For the sake of simplicity, PLA stereocopolymers and PLAGA copolymers are identified in this chapter by acronyms PLAX or PLAXGAY where

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Table 6.1 Aliphatic polyesters Polymer and acronym

Structure

Poly(glycolic acid) (PGA)

-[-O-CH2-CO-]nH

Poly(lactic acid) (PLA)

-[-O-*C-CO-]nCH3

Poly(ε-caprolactone) (PCL)

-[-O-(CH2)5-CO-]n-

Poly(para-dioxanone)

-[-O-(CH2)2-O-CH2-CO-]nH

Poly(hydroxybutyrate) (PHB)

-[-O-*C-CH2-CO-]nCH3 H

Poly(hydroxyvalerate) (PHV)

-[-O-*C-CH2-CO-]nCH2-CH3 H

Poly(β-malic acid) (PMLA)

-[-O-*C-CH2-CO-]n-

Poly(alkylene dicarboxylate)

COOH -[-O-(CH2)X-O-CO-(CH2)Y-CO-]n-

* = asymmetric carbon.

X is the percentage of L-LA units in the monomer feed, Y is that of GA units, the rest being the percentage of D-LA units. Similarly, copolymers of lactides and ε-caprolactone can be denoted as PLAXCLZ where X and Z represent the percentages of L-LA and CL units, respectively. This nomenclature presents the advantage of reflecting clearly the chemical and configurational compositions of the polymers, the average polymer chain composition being generally close to that of the feed (Chabot, 1983). The in vivo and in vitro degradations of aliphatic polyesters have been extensively investigated during the past two decades. It was established that degradation is catalyzed by carboxyl endgroups formed by chain cleavage and that amorphous regions are preferentially degraded. Nevertheless, there existed discrepancies or controversial data in the literature, especially concerning the degradation rates, degradation location (at the surface or in the bulk) and the involvement of enzymes in vivo. Our group is one of the pioneers to undertake long-term systematic investigations by using various

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analytical techniques. Important advances have been accomplished in the understanding of the hydrolytic degradation characteristics, in particular, faster internal degradation and degradation-induced morphological and compositional changes (Li, 1999a, b). In this chapter, we will describe in detail the in vitro physicochemical test methods to evaluate bioresorbability of PLA, PGA and PCL homo- and copolymers. Discussion is largely based on recent advances in the domain.

6.2

Protocol for in vitro degradation studies

Various in vitro degradation conditions have been used to investigate the hydrolytic degradation behaviors of aliphatic polyesters. Polymers can be processed in the form of plates, cylinders, films, fibers, microparticles or nanoparticles, prepared by different processing methods. Degradation is generally realized at 37°C in pH = 7.4 phosphate-buffered saline (PBS). Sodium azide (0.02%) is added in the solution to prevent bacterial growth. In our long-term systematic investigations, the polymers are obtained by ring-opening polymerization of cyclic monomers using zinc powder as initiator. This initiator is selected because it is more acceptable than tin derivatives from the viewpoint of biocompatibility, zinc ions being normally present in living organisms. Polymer powder or granules are compression molded to yield circular plates of 75 mm diameter and of 2 mm thickness. Sets of 15 × 10 × 2 mm parallelepiped specimens are obtained by machining from the plates. The specimens are placed in a degradation medium and allowed to degrade at 37°C. After predetermined degradation time intervals, the specimens are withdrawn, and washed with distilled water. After wiping, the specimens are weighed and examined visually and by ESEM. They are then vacuum dried at room temperature up to constant weight before being subjected to analyses by various techniques, namely visual examination, environmental scanning electron microscope (ESEM), proton nuclear magnetic resonance (1H NMR), size-exclusion chromatography (SEC), X-ray diffraction (XRD), differential scanning calorimetry (DSC), electron spray injection mass spectrometry (ESI-MS), capillary zone electrophoresis (CZE), gravimetry, mechanical property tests, etc. Enzymatic degradation of polymers is usually realized in enzymecontaining buffer solutions whose pH should be in the optimal enzymatic activity range. A pH = 8.6 Tris/HCl buffer (0.05 m) is used in the case of PLA-degrading Proteinase K, whereas a pH = 7.0 phosphate buffer (0.05 m) is used for PCL-degrading Pseudomonas lipase. The buffer/enzyme system is changed periodically to restore the original level of enzymatic activity. Similar techniques are used to monitor the enzymatic degradation as in the case of hydrolytic degradation. Controls are realized in the buffer solutions without enzyme.

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6.3

121

In vitro physicochemical test methods

The hydrolytic degradation of aliphatic polyesters is a complex process involving several phenomena, namely water absorption, ester bond cleavage, neutralization of carboxyl endgroups at the surface, autocatalysis inside, and diffusion and solubilization of soluble oligomers. These phenomena can be monitored by using various in vitro physicochemical test methods including visual examination, environmental scanning electron microscope (ESEM), proton nuclear magnetic resonance (1H NMR), size-exclusion chromatography (SEC), X-ray diffraction (XRD), differential scanning calorimetry (DSC), electron spray injection mass spectrometry (ESI-MS), capillary zone electrophoresis (CZE), and weight loss.

6.3.1 Visual examination Degradation of aliphatic polyesters is generally regarded as a bulk phenomenon (Kenley, 1987; St Pierre, 1987; Helder, 1990; Lewis, 1990). Faster internal degradation has been detected during hydrolytic degradation of various large-size PLAGA polymers (Li, 1990a; Gogolewski, 1993; Schliephake, 1993; Pistner, 1994). The heterogeneous degradation was assigned to diffusion-reaction phenomena as summarized in the following. When an initially homogeneous polymer is placed in an aqueous medium, water penetrates the matrix leading to hydrolytic cleavage of ester bonds (Ali, 1993; Schmitt, 1994). Each ester bond cleavage yields a carboxyl endgroup which, according to autocatalysis, accelerates the hydrolytic reaction of the remaining ester bonds (Pitt, 1981a, b, 1992; Huffman, 1985). However, no autocatalysis occurs at the surface because the carboxyl endgroups are neutralized by the surrouding buffered medium and lose their catalytic capacity. Moreover, when soluble oligomeric species are generated, those which are located close to the surface can escape from the matrix before total degradation, while those located inside can hardly diffuse out of the matrix. Therefore, autocatalysis occurs only in the internal part of the matrix, thus leading to a surface/interior differentiation. As degradation proceeds, more and more carboxyl endgroups are formed inside to accelerate the internal degradation and thus enhance the surface/interior differentiation. Bimodal MW distributions are observed due to the presence of two populations of macromolecules degrading at different rates. Finally, hollow structures are formed when the internal material is totally transformed to soluble oligmers and dissolves in the aqueous medium. This is the case with amorphous polymers such as PLA50, PLA62.5, PLA75, and PLA37.5GA25 (Li, 1990a, b, 1994). Figure 6.1 shows that, after 5 weeks’ degradation in 0.13 m pH = 7.4 phosphate buffer at 37°C, the cross-section

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Degradation rate of bioresorbable materials (a)

(b)

6.1 (a) Cross-section of a PLA50 specimen after 5 weeks’ degradation in 0.13 M pH = 7.4 phosphate buffer at 37°C; (b) Hollow structure of a PLA50 specimen after 14 weeks’ degradation in 0.13 M pH = 7.4 phosphate buffer at 37°C (from Li, 1990a).

of a PLA50 specimen exhibits a very pronounced surface/interior differentiation. After 14 weeks, the specimen is totally hollowed out, only a hollow structure remaining. In contrast, in the case of crystallizable polyesters such as PLA87.5, PLA96, PLA100, PLA75GA25 and PLA85GA15, no hollow structures are obtained due to the crystallization of internal degradation by-products. Nevertheless, surface–interior differentiation was also observed with faster internal degradation (Li, 1990b, c, 1994 Vert, 1991, 1992; Therin, 1992). For example, after as long as 18 and 40 weeks in 0.13 m pH = 7.4 phosphate buffer at 37°C, initially amorphous PLA96 exhibited a very pronounced

In vitro physicochemical test methods

123

(a)

(b)

6.2 Cross-section of a PLA96 specimen after 18 (a) and 40 (b) weeks’ degradation in 0.13 M pH = 7.4 phosphate buffer at 37°C (from Vert, 1991).

surface/interior differentiation (Vert, 1991), the interior being composed of highly crystalline material (Fig. 6.2).

6.3.2 Environmental scanning electron microscope (ESEM) ESEM is a technique of choice to monitor the changes of surface morphology of degrading materials because it provides original images, without necessitating high vacuum or gold coating, which often results in artifacts. It is particularly interesting in the case of enzymatic degradation, which is a surface phenomenon. In fact, enzymes are generally proteins of high

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Degradation rate of bioresorbable materials

molecular weights. As a consequence, enzymes are not able to penetrate a solid polymeric material in contrast to small molecules such as water, but can eventually adsorb on the polymer surface. Enzymatic degradation consists of two steps: first, enzyme adsorbs on the surface of a polymer matrix through its binding domain; second, ester bonds are cleaved due to the effect of the calalytic domain of the enzyme. PLA is degradable in the presence of proteinase K, a protease secreted by the fungus Tritirachium album Limber which has a strong activity to hydrolyze proteins (Williams, 1981). However, proteinase K can degrade only L-LA units as opposed to D-LA ones, amorphous zones as opposed to crystalline ones (Reeve, 1994; MacDonald, 1996; Li, 1999c, 2000; Tsuji, 2001a, b). Figure 6.3 shows the surface morphology changes of amorphous PLA25 film during enzymatic degradation (Li, 2001). PLA25 exhibits initially a smooth surface with the presence of some stripes resulting from molding (Fig. 6.3a). After 168 hours’ enzymatic degradation, the film appears strongly eroded and a porous structure is obtained (Fig. 6.3b). PCL is degradable mainly in the presence of lipase-type enzymes such as Rhizopus delemer lipase (Fukuzaki, 1990), Rhizopus arrhizus lipase (Mochizuki, 1995), and Pseudomonas lipase (Gan, 1999; Liu, 2000). Highly crystalline PCL can totally degrade in 4 days (Gan, 1999), in contrast to hydrolytic degradation which lasts several years. Figure 6.4 pesents the ESEM micrographs of solution-cast PCL film (Liu, 2000). Well defined spherulites of 50–100 μm are initially observed, with boundaries clearly distinguished (Fig. 6.4a). Dramatic changes are observed after degradation in the presence of Pseudomonas lipase. The surface appears strongly eroded to leave sponge-like and fibrillar structures (Fig. 6.4b), showing that both amorphous and crystalline zones can be degraded. The boundaries between these zones become better defined because they are mainly composed of crystallite defects or amorphous material which can be preferentially degraded. The presence of both sponge-like and fibrillar structures might reflect different degrees of degradation, the former reflecting a more advanced degradation. In the case of PCL/PLA100 blends, the selective degradation of PCL or PLA100 components revealed their inner morphology (Liu, 2000). After degradation by proteinase K, the surface of PLA100/PCL(25/75) showed numerous pores of several micrometers (Fig. 6.5a). Spherulitic structures are distinguishable. In the presence of Pseudomonas lipase, the surface presents a microsphere-like pattern (Fig. 6.5b). In fact, PLA100 spherical microdomains are regularly dispersed within the PCL continuous matrix. Proteinase K selectively degraded the PLA100 microdomains, leaving a porous structure. In contrast, Pseudomonas lipase selectively degraded the PCL continuous matrix, leaving PLA100 microdomains in the form of microspheres.

In vitro physicochemical test methods

125

(a)

(b)

6.3 ESEM micrographs of the surface of compression molded PLA25 film after (a) 0 and (b) 168 hours’ enzymatic degradation in the presence of proteinase K (from Li, 2001).

6.3.3 Size-exclusion chromatography (SEC) SEC allows one to monitor the weight-average molecular weight (Mw), number-average molecular weight (Mn) and MW distribution changes of polymers during degradation. In fact, chain cleavage leads to MW decrease, while faster internal degradation, degradation-induced crystallization and preferential degradation of amorphous domains result in MW distribution changes (Li, 1999b).

126

Degradation rate of bioresorbable materials (a)

(b)

6.4 ESEM micrographs of the surface of a solution cast PCL film after (a) 0 and (b) 72 hours’ degradation in the presence of Pseudomonas lipase (from Liu, 2000).

Figure 6.6 presents the SEC chromatograms of intrinsically amorphous PLA37.5GA25 after 0, 17, and 35 days’ degradation in distilled water at 37°C (Li, 1990b). The MW distribution is initially monomodal and narrow, with Mw = 52 000 and Mw/Mn = 1.6 (Fig. 6.6a). After 17 days in water, the MW of the interior (Mw = 4500) becomes much lower than that of the surface (Mw = 18 400) due to faster internal degradation (Fig. 6.6b). Finally, after 35 days, the Mw of the interior further decreases to 1600, whereas the surface exhibits a bimodal MW distribution with Mw = 6000 due to the presence of two populations of macromolecules resulting from faster internal degradation (Fig. 6.6c).

In vitro physicochemical test methods

127

(a)

(b)

6.5 ESEM micrographs of the surface of a PLA100/PCL(25/75) blend film after 72 hours’ degradation in the presence of (a) Proteinase K and (b) Pseudomonas lipase (from Liu, 2000).

The situation is different in the case of intrinsically semi-crystalline polymers (Li, 1990c; Vert, 1991). Figure 6.7 shows the SEC chromatograms of initially amorphous PLA96 screws after 0, 20, and 40 weeks’ degradation in 0.13 m pH = 7.4 phosphate buffer at 37°C. The Mw distribution is initially monomodal and narrow with Mw = 79 000 and Mw/Mn = 1.6 (Fig. 6.7a). After 20 weeks, the Mw of the interior (Mw = 16 100, Mw/Mn = 2.3) becomes much lower than that of the surface (Mw = 31 000, Mw/Mn = 2.7) due to faster internal degradation, and the Mw distribution becomes larger (Fig. 6.7b). Moreover, a shoulder is detected in the low Mw range for the surface. After 40 weeks, a bimodal Mw distribution is obtained

128

Degradation rate of bioresorbable materials (a)

0 days

20

25 106

30 105

40

35 104

103

Ve/ml

Masse

(b)

17 days

20

25 106

30 104

105

40

35 103

Ve/ml Masse

(c)

35 days

20

25 106

30 105

40

35 104

103

Ve/ml Masse

6.6 SEC chromatograms of PLA37.5GA25 after (a) 0, (b) 17, and (c) 35 days’ degradation in distilled water at 37°C (from Li, 1990b).

In vitro physicochemical test methods

129

(a)

0 weeks

25

20 106

30 104

105

40

35 103

Ve/ml Masse

(b)

20 weeks

20

25 106

30 104

105

40

35 103

Ve/ml Masse

(c)

40 weeks

20

25 106

30 105

40

35 104

103

Ve/ml Masse

6.7 SEC chromatograms of initially amorphous PLA96 after (a) 0, (b) 20, and (c) 40 weeks’ degradation in 0.13 M pH = 7.4 phosphate buffer at 37°C.

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Degradation rate of bioresorbable materials

for the surface since the shoulder becomes a small peak (Mw = 17 900, Mw/Mn = 2.5). In contrast, the interior exhibits a narrow Mw distribution with Mw = 4000 and Mw/Mn = 1.3 (Fig. 6.7c). The appearance of a narrow peak in the low Mw range is assigned to the crystallization of degradation by-products whose Tg is lower than that of longer chains. With the plasticizing effect of absorbed water, which further decreases the Tg, these short chains are mobile enough to crystallize under the degradation conditions (37°C). The resulting crystallites are more resistant to further degradation. As a consequence, a very narrow peak corresponding to low Mw crystalline zones is detected in SEC chromatograms (Fig. 6.7c). Mw distribution changes are also observed for initially semi-crystalline polymers (Li, 1990c, 1996a). Figure 6.8 shows the Mw changes of PCL with in vitro degradation. The Mw of PCL is initially 58 700 with Mw/Mn = 1.7. After 63 weeks, Mw decreases to 19 500 and Mw/Mn increases to 2.3. Shoulders are detected in the low Mw range. After 133 weeks, the Mw further decreases to 9200. Moreover, the Mw distribution becomes trimodal, the Mw of the three peaks being 2600, 5200 and 8800, respectively. Finally after 200 weeks, the Mw slightly decreases to 7000, but the three peaks remain at the same positions (Figure 6.8). The appearance of trimodal Mw distribution is assigned to the selective degradation of amorphous domains of semi-crystalline polymers (Fischer, 1973). It is now well known that degradation of semicrystalline polyesters in aqueous media proceeds in two stages (Chu, 1981; Carter, 1984; Fredericks, 1984). The first stage consists in water diffusion into the amorphous regions with random hydrolytic scission of ester bonds. The second stage starts when most of the amorphous regions are degraded. The hydrolytic attack then progresses from the edge towards the center of the crystalline domains. This phenomenon was first observed by Fischer et al., who obtained a trimodal Mw distribution after chemical degradation of PLA92.5 single crystals, the three peaks corresponding respectively to onefold, twofold and threefold the crystalline lamellae thickness (Fischer, 1973). Bimodal Mw distributions are also detected at the later stages of in vitro degradation of semi-crystalline PLA100, the two peaks corresponding to onefold and twofold thicknesses of crystalline lamellae, respectively (Li, 1990c). Therefore, multimodal Mw distributions can be observed by SEC during degradation of aliphatic polyesters, resulting from faster internal degradation for amorphous polymers (Fig. 6.6), from crystallization of degradation by-products for crystallizable but made amorphous polymers (Fig. 6.7), or from the selective degradation of amorphous domains for semi-crystalline polymers (Fig. 6.8).

In vitro physicochemical test methods

0 weeks

Mw = 58700

63 weeks

Mw = 19500

131

Mw = 9200

133 weeks 5200

2600

8800

Mw = 7000

200 weeks

10

12

14

16

18

Elution time (min) 6.8 SEC chromatograms of PCL after 0, 63, 133 and 200 weeks’ degradation in 0.13 M pH = 7.4 phosphate buffer at 37°C (from Li, 1999a).

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Degradation rate of bioresorbable materials

6.3.4 X-ray diffraction (XRD) Polymer morphology, i.e. amorphousness or semicrystallinity, plays a critical role in the degradation process because of the selective degradation of amorphous domains. XRD is usually used to monitor crystalline structure and crystallinity changes of polymers during degradation. Figure 6.9 presents the XRD spectra of PLA100 during degradation in pH = 7.4 phosphate buffer at 37°C. PLA100 is initially amorphous with only an amorphous halo (Li, 1990c). Small diffraction peaks are detected at theta = 8.4° and 9.6° after 18 weeks, indicating the appearance of crystallinity. After 50 and 110 weeks, the two peaks become much more intense, the degree of crystallinity being 21% and 50%, respectively. The crystallization is assigned to low Mw degradation by-products only because high Mw chains cannot crystallize at 37°C, which is much lower than the Tg (56°C), as mentioned above. Morphological changes are also observed during degradation of intrinsically amorphous polymers. PLAX is intrinsically amorphous if X is in the 10 to 90 range (Li, 1999b). For PLAXGAY copolymers, the polymer is intrinsically amorphous when Y is in the 10 to 70 range (L-LA/GA) or in the 0 to 70 range (DL-LA/GA) (Reed, 1979). However, one must be cautious with these divisions because of the variable extents of transesterification reactions, which tend to randomize the comonomer distribution along the chains. PLA50 is an intrinsically amorphous polymer which cannot crystallize even after long annealing. However, a crystalline oligomeric stereocomplex composed of L-LA and D-LA segments is detected at the later stages of PLA50 degradation (Li, 1994, 1996b, 1999d), as shown in Fig. 6.10. In fact,

0 weeks

Intensity

18 weeks

50 weeks

110 weeks

2

4

6

8

10

12

14

16

18

20

Theta (degrees)

6.9 X-ray diffraction spectra of initially amorphous PLA100 after 0, 18, 50 and 110 weeks’ degradation in pH = 7.4 phosphate buffer at 37°C (from Li, 1990c).

In vitro physicochemical test methods

133

38 weeks 0 weeks

2

4

6

8

10

12

14

16

18

20

Theta (degrees) 6.10 X-ray diffraction spectra of PLA50 after 0 and 38 weeks’ degradation in 0.13 M pH = 7.4 phosphate buffer at 37°C (from Li, 1996b).

stereoregular segments are initially present within the PLA50 matrix, which is known to have a predominantly isotactic structure resulting from the pair-addition mechanism of DL-lactide polymerization (Chabot, 1983). Once released by degradation, these L-LA and D-LA segments are susceptible to crystallize, leading to formation of an oligomeric stereocomplex.

6.3.5 Differential scanning calorimetry (DSC) Morphological changes of polymers during degradation can also be monitored by using DSC in comparison with XRD. Two heating scans are generally realized on the polymers: after the first heating at 10°C/min, the sample is quenched and a second run is immediately performed. The first run allows the thermal history to be erased in order to better determine the thermal properties of the polymers. PLA50 is initially amorphous with only a small peak around 42.5°C reflecting physical ageing of polymer chains, as shown in Fig. 6.11. A glass transition (Tg) is observed at 46.4°C on the second heating. After 98 days’ degradation, hollow structures are formed due to faster internal degradation. A small melting peak (Tm) is detected at 98.5°C, indicating the appearance of crystallinity. Tm disappears on the second heating, whereas Tg decreases to 22.9°C due to MW decrease. After 126 days, Tm increases to 115.7°C, indicating the perfection of crystallites. Tg further decreases to 20.0°C on the second heating. Moreover, a crystallization peak is detected at 83.8°C, followed by a melting peak at 107.8°C (Li, 1999d). The appearance of crystallinity is assigned to the formation of a crystalline oligomeric stereocomplex composed of L-LA and D-LA segments, as mentioned above.

134

Degradation rate of bioresorbable materials 0.2

First heating

0 days

0 days 0.0

Second heating

46.4

42.5 98 days 98.5 126 days

–0.6

DSC heat flow (W/g)

DSC heat flow (W/g)

98 days –0.2

–0.6

22.9

–0.4 126 days

8.3.8

20.0 –0.2

115.7 –1.0 10

30 90 50 70 110 130 150 Temperature (°C)

107.8 –0.8 10

30 50 70 90 110 130 150 Temperature (°C)

6.11 DSC thermograms of PLA50 after 0, 98 and 126 days’ degradation in 0.13 M pH = 3.7 phosphate buffer at 37°C (from Li, 1999d).

6.3.6 Proton nuclear magnetic resonance (1H NMR) 1

H NMR is generally used to follow chemical composition changes of various copolymers during degradation. Figure 6.12 presents the variation of lactyl content (mol%) of PLA75GA25 and PLA85GA15 (Vert, 1991; Li, 1999b). PLA75GA25 exhibits a faster increase at the early stages than PLA85GA15 due to faster degradation. Beyond, the LA content tends to a limit in both cases. This preferential degradation of GA units is assigned to the higher hydrophilicity of GA units as compared to LA ones. In consequence, the L-LA-enriched segments crystallize under the in vitro degradation conditions to form crystalline zones of relatively low MW. Crystallization prevents, in return, the preferential degradation of GA units because of their inclusion in crystalline zones. 1 H NMR also allows changes of chain microstructure of copolymers to be followed (Dobrzynski, 2005; Li, 2005). Figure 6.13 shows the spectra of PGA43CL57 after 0, 2, 10 and 26 weeks’ degradation. The various signals are assigned to different sequences as shown on the figure. It is of interest to note the presence of CGC (7) and CGGGC (3) sequences with odd numbers of glycolyl units, indicating the occurrence of transesterification reactions. An intensity decrease is observed for all glycolyl signals during degradation, except for Signals 3 and 7. The relative intensity of GC also decreases as compared to CC, in agreement with the preferential cleavage of G-C bonds.

In vitro physicochemical test methods

135

100

LA mol percentage (%)

PLA85GA15 90 PLA75GA25 80

70

60 0

10

20

30

40

50

Degradation time (weeks) 6.12 Variation of lactyl (LA) mole percentage of PLA75GA25 and PLA85GA15 during degradation in 0.13 M pH = 7.4 phosphate buffer at 37°C (from Li, 1990b; Vert, 1991).

6.3.7 Electron spray injection mass spectrometry (ESI–MS) ESI–MS analysis can be used to follow the appearance of low Mw degradation by-products in polymer bulk, giving insights into the chain microstructure of copolymers (Li, 2005). Figure 6.14 shows the ESI–MS spectrum of PGA82CL18 after 10 weeks’ degradation. Various GmCntype oligomers are detected, where m ranges from 2 to 4, and n from 8 to 15, respectively. On the other hand, Gm homo-oligomers with m ranging from 7 to 16 are observed, while Cn oligomers are not present (Fig. 6.14).

6.3.8 Capillary zone electrophoresis (CZE) CZE is used to monitor the release of water soluble species in the buffer solution during degradation, which is complementary to ESI–MS (Li, 2005). Figure 6.15 shows the buffer solution containg PGA65CL35 and PGA45CL55 copolymers after 2 and 26 weeks’ degradation. Assignments of the various signals are given on the figure. After 2 weeks, an intense peak corresponding to glycolate monomer (G) is detected in the PGA65CL35-containg buffer solution. Other peaks with lower intensity are also observed, corresponding to C, CG, CGG, CC, CCG and CCC sequences. It should be noted that the peaks of CGG could also result from GGC or GCG sequences,

136

Degradation rate of bioresorbable materials 1 – GGGGGG 2 – CGGGG + GGGGC 3 – CGGGC 4 - -GGGGC 5 – CGGGG + CGGGC 6 – CGGC 7 – CGC 8 – GC 9 – CC

1

2 0

3 4

5 6

9

7

8

1

2 2

9

56 3 4

7

8 9

1

2

3

10

26 5.0

2 3 4.8

56 4

5 4 6

8

7

9

8

7

4.6

4.4

4.2

4.0

ppm

3.8

6.13 1H NMR spectra of PGA43CL57 after 0, 2, 10 and 26 weeks’ degradation in 0.13 M pH = 7.4 phosphate buffer at 37°C (from Li, 2005).

and that of CCG from GCC or CGC ones. Peaks A and G correspond to phosphate and cholate, respectively. Cholate is generally used as an internal reference. The intensity of peaks increases after 26 weeks’ degradation, except CGG which is easily degraded to yield CG and G. The small peak beside CG might be assigned to GC which is supposed to exhibit a slightly different mobility as compared to CG. Similar patterns are found for PGA45CL55, although the intensity of the peaks is lower due to slower degradation.

In vitro physicochemical test methods S#: 111 RT: 3.58 AV: 1 NL: 7.68E5 T: - c Full ms [370.00 - 2000.00]

C9G3

1218.2

100

80

1104.1

C10G2 C11G3 C11G4 C12G3 1274.1 1446.2 1504.1 1560.1 C G 12 4 C13G4 1618.0

Relative abundance

1732.0

C9G2

70

C13G3

G12

G G10 11 655.6 597.9 G G8 9

50

C14G3 1788.0

C14G4 C15G4

G15

713.8

60

30

C10G3 1332.1

90

40

137

G13 771.4

1846.0

1960.0

C15G3

887.7

1902.1

G14 C8G3

829.4

1046.8

481.8 539.7

G16 989.4

G7 423.7

20 10 0 400

600

800

1000

1200 m/z

1400

1600

1800

2000

6.14 ESI–MS spectrum of PGA82CL18 after 10 weeks’ degradation in 0.13 M pH = 7.4 phosphate buffer at 37°C (from Li, 2005).

6.3.9 Weight loss Weight loss reflects the release of soluble oligomers from the bulk during degradation. Figure 6.16 presents the weight loss profiles of PLA37.5GA25 during degradation in 0.13 m pH = 7.4 phosphate buffer or in distilled water at 37°C (Li, 1990b). No weight loss is detected during the first 21 days. Between 21 and 24 days, a dramatic weight loss up of to 95% is obtained in the buffer, indicating that the internal material is totally dissolved. Beyond 24 days, the residues degrade slowly with a weight loss of 98% after 35 days. Only a tiny weight loss of 0.3% is found in distilled water after 24 days. Weight loss then increases to reach 45% after 35 days. The difference between the weight loss profiles in the two media is assigned to the fact that dissolution of carboxyl-terminated oligomers is better facilitated in the pH = 7.4 buffer than in distilled water.

6.3.10 Mechanical tests Mechanical properties are key parameters of polymeric biomaterials which are supposed to fulfill their functions during a predetermined period of time. In the literature, most mechanical tests are carried out on highly crystalline PGA or PLAGA fibers in which alternative amorphous and

138

Degradation rate of bioresorbable materials

Cop2-2 C CG CGG CCG B CC CCC

G

A

Cop2-26 CG

A

CGG CCG B CC CCC

C G

Cop5-2 G

C

CG

B

A Cop5-26

C A

–3.8

CG

CC CCG B CCC

G

–3.4

–3.0

–2.6

–2.2

–1.8

–1.4

–1.0

Effective electrophoretic mobility (x E4 cm2/ V. s) 6.15 CZE diagrams of the buffer solutions containing PGA/PCL copolymers (Cop2-PGA65CL35 and Cop5-PGA45CL55) after 2 and 26 weeks’ degradation: A – phosphate; B – cholate (from Li, 2005).

crystalline regions arrange in the direction of the fiber axis (Chu, 1981; Deng, 2005). Two types of tie-chain segments (interfibrillar and intrafibrillar) are present in the amorphous regions of the microfibrillar structure. Their major function is to tie crystalline regions together and to transmit tensile load. The scission of these tie-chain segments leads to rapid loss of tensile properties (Chu, 1981). Figure 6.17 presents the storage modulus changes of initially amorphous PLA100a and semi-crystalline PLA100c during degradation (Li, 1990c). PLA100c initially exhibits higher modulus than PLA100a due to the presence of crystalline domains (4.2 GPa versus 3.7 GPa). The modulus remains unchanged during the first 3 weeks for PLA100c. Thereafter, an almost constant decrease is observed. In contrast, PLA100a keeps its storage modulus unchanged for

In vitro physicochemical test methods

Weight loss (%)

90

139

Buffer

70 50 Water

30 10 –10

0

5

10

15

20

25

30

35

40

Degradation time (days) 6.16 Weight loss profiles of PLA37.5GA25 during degradation in 0.13 M pH = 7.4 phosphate buffer or in distilled water at 37°C (from Li, 1990b).

Storage modulus (GPa)

5

4

3

2 PLA100c

1 PLA100a 0 0

10

20 30 40 Degradation time (weeks)

50

60

6.17 Storage modulus changes of initially amorphous PLA100a and semicrystalline PLA100c during degradation in 0.13 M pH = 7.4 phosphate buffer (from Li, 1990c).

12 weeks. The rapid loss of mechanical properties observed for PLA100c is assigned to the scission of tie-chain segments in the amorphous region, as in the case of fibers. PLA100a exhibits a uniform structure, and the mechanical properties seem less sensitive to chain cleavage.

6.3.11

Accelerated degradation tests

The in vitro degradation rate of aliphatic polyesters depends on many factors, including polymer composition, MW, morphology and size. The half-life of PLA100 and PCL is estimated to be more than two years

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Degradation rate of bioresorbable materials

(Li, 1996a, 1999b). Efforts have been devoted to develop models of accelerated degradation tests at elevated temperatures or in acidic and alkaline media (Lyu, 2007; von Burkersroda, 2002). Lyu et al. established a time– temperature equivalent relationship to predict polymer degradation based on accelerated tests. However, the proposed master did not take into account many phenomena, in particular, the most important ones of faster internal degradation and degradation-induced crystallization (Lyu, 2007). Von Burkersroda et al. developed a theoretical model to predict the degradation mechanisms and, to support this model, the authors tried to change the degradation mechanism from bulk degradation to surface erosion by acid and base catalysis. Similar bulk degradation behaviors of PLA and PLAGA were observed at pH 2 and at pH 7.4. In contrast, at pH > 13, surface erosion was found to be the predominant phenomenon. Therefore, in vitro and in vivo degradation of aliphatic polyesters is a very complex process, which can hardly be simulated by using accelerated degradation tests.

6.4

Conclusion

Degradable aliphatic polyesters, in particular, those deriving from lactides, glycolide and ε-caprolactone, have been most widely investigated for applications in the field of surgery, tissue repair and regeneration, as well as sustained drug delivery. Factors that can affect the degradation of polymers have been gradually identified. Degradation can be monitored by using vatrious analytical techniques. This paper has tried to present the state-ofthe-art on the basis of recent advances in this domain. The hydrolytic degradation of aliphatic polyesters is a complex process involving many physico-chemical phenomena, namely water uptake, ester bond cleavage, neutralization of carboxyl endgroups at the surface, autocatalysis inside, crystallization, and diffusion and solubilization of soluble oligomers. These phenomena should be monitored by using a large number of physicochemical test methods if one wants to succeed in a real control of degradation of the polymeric devices because of the interdependence of corresponding phenomena.

6.5

References

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carter b k, wilkes g l (1984), ‘Some morphological investigations on an absorbable copolyester biomaterial based on glycolic and lactic acid’. In Shalaby S. W. , Hoffman A. S. , Ratner B. D. and Horbert T. A. eds., Polymers as Biomaterials, Plenum Press, New York, 67–92. chabot f, vert m, chapelle s, granger p (1983), ‘Configurational structures of lactic acid stereocopolymers as determined by C—H n.m.r’, Polymer, 24, 53–60. charles e l, buffalo n y (1954), Preparation of high molecular weight polyhydroxyacetic ester. U.S. Pat. 2668162 (Feb. 2, 1954) (to E. I. DuPont de Nemours). chu c c (1981), ‘Hydrolytic degradation of poly(glycolic acid): Tensile strength and crystallinity study’, J Appl Polym Sci, 26, 1727–1734. deng m, zhou j, chen g, burkley d, xu y, jamiolkowski d, barbolt t (2005), ‘Effect of load and temperature on in vitro degradation of poly(glycolide-co-L-lactide) multifilament braids’, Biomaterials, 26, 4327–4336. dobrzynski p, li s, kasperczyk j, bero m, gasc g, vert m (2005), ‘Structure property relationships of copolymers obtained by ring-opening polymerization of glycolide and ε-caprolactone. Part 1. Synthesis and characterization’, Biomacromolecules, 6, 483–488. dunn r l (1995), ‘Clinical Applications and Update on the Poly(α-hydroxy acids)’. In Hollinger J O, ed., Biomedical Applications of Synthetic Biodegradable Polymers, Boca Raton, CRC Press, 17–31. fischer e w, sterzel h j, wegner g (1973), ‘Investigation of the structure of solution grown crystals of lactide copolymers by means of chemical reactions’, Kolloid-Z. u. Z. Polymere, 251, 980–990. frazza e j, schmitt e e (1971), ‘A new absorbable suture’, J Biomed Mater Res Symposium 1, 43–58. fredericks r j, melveger a j, dolegiewtz l j (1984), ‘Morphological and structural changes in a copolymer of glycolide and lactide occurring as a result of hydrolysis’, J Polym Sci: Polym Phys Ed, 22, 57–66. fukuzaki h, yoshida m, asano m, kumakura m (1989), ‘Synthesis of copoly(D,Llactic acid) with relatively low molecular weight and in vitro degradation’, Eur Polym J, 25, 1019–1026. fukuzaki h, yoshida m, asano m, kumakura m, mashimo t, yuasa h, imai k, yamanaka h (1990), ‘Synthesis of low molecular weight copoly(L-lactiic acid/εcaprolactone) by direct copolycondensation in the absence of catalysts, and enzymatic degradation of the polymers’, Polymer, 31, 2006–2014. gan z, yu d, zhong z, liang q, jing x (1999), ‘Enzymatic degradations of poly(εcaprolactone)/poly(DL-lactide) blends in phosphate buffer solution’, Polymer, 40, 2859–2863. gogolewski s, jovanovic m, perren s m, dillon j g, hughes m k (1993), ‘Tissue response and in vivo degradation of selected polyhydroxyacids: Polylactides (PLA), poly(3-hydroxybutyrate) (PHB), and poly(3-hydroxybutyrate-co-3hydroxyvalerate) (PHB/HV)’, J Biomed Mater Res, 27, 1135–1148. helder j, dijkstra p j, feijen j (1990), ‘In vitro degradation of glycine/DL-lactic acid copolymers’, J Biomed Mater Res, 24, 1005–1020. hoffman a s (1984), ‘Synthetic Polymeric Biomaterials’. In Gebelein C G, Polymeric Materials and Artificial Organs. ACS Symposium Series, No. 256, pp. 13–29. holland s j, tighe b j, gould p l (1986), ‘Polymers for biodegradable medical devices. 1. The potential of polyesters as controlled macromolecular release systems’, J Control Rel, 4, 155–180.

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huffman k r, casey d j (1985), ‘Effects of carboxylic end groups on hydrolysis of polyglycolic acid’, J Polym Sci: Polym Chem Ed, 23, 1939–1954. kenley r a, lee m o, mahoney t r, sanders l m (1987), ‘Poly(lactide-co-glycolide) decomposition kinetics in vivo and in vitro’, Macromolecules, 20, 2398–2403. lewis d h (1990), ‘Controlled release of bioactive agents from lactide/glycolide polymers’, Drugs Pharm Sci (Biodegrad Polym Drug Delivery Syst), 45, 1–41. li s, garreau h, vert m (1990a), ‘Structure–property relationships in the case of the degradation of massive poly(α-hydroxy acids) in aqueous media. Part 1: Poly(DL-lactic acid)’, J Mater Sci: Mater Med, 1, 123–130. li s, garreau h, vert m (1990b), ‘Structure–property relationships in the case of the degradation of massive poly(α-hydroxy acids) in aqueous media. Part 2: Degradation of lactide/glycolide copolymers: PLA37.5GA25 and PLA75GA25. J Mater Sci: Mater Med, 1, 131–139. li s, garreau h, vert m (1990c), ‘Structure–property relationships in the case of the degradation of massive poly(α-hydroxy acids) in aqueous media. Part 3: Influence of the morphology of poly(L-lactic acid). J Mater Sci: Mater Med, 1, 198–206. li s, vert m (1994), ‘Morphological changes resulting from the hydrolytic degradation of stereocopolymers derived from L- and DL-lactides’, Macromolecules, 27, 3107–3110. li s, espartero j l, foch p, vert m (1996a), ‘Structural characterization and hydrolytic degradation of a Zn metal initiated copolymer of L-lactide and ε-caprolactone’, J Biomater Sci Polym Ed, 8, 165–187. li s, girod-holland s, vert m (1996b), ‘Hydrolytic degradation of poly(DL-lactic acid) in the presence of caffeine base’, J Control Rel, 40, 41–53. li s, vert m (1999a), ‘Biodegradable polymers: Polyesters’, in Mathiowitz E, Encyclopedia of Controlled Drug Delivery, New York, John Wiley & Sons, 71–93. li s (1999b), ‘Hydrolytic degradation characteristics of aliphatic polyesters derived from lactic and glycolic acids’, J Biomed Mater Res: Appl Biomater, 48, 142–153. li s, mccarthy s p (1999c), ‘Influence of crystallinity and stereochemistry on the enzymatic degradation of poly(lactide)s’, Macromolecules, 32, 4454–4456. li s, mccarthy s p (1999d), ‘Further investigations on the hydrolytic degradation of poly(DL-lactide)’, Biomaterials, 20, 35–44. li s, tenon m, garreau h, braud c, vert m (2000), ‘Enzymatic degradation of stereocopolymers derived from L-, DL- and meso-lactides’, Polym Degr Stabl, 67, 85–90. li s, girard a, garreau h, vert m (2001), ‘Enzymatic degradation of PLA stereocopolymers with predominant D-lactyl contents’, Polym Degr Stabl, 71, 61–67. li s, vert m (2002), ‘Biodegradation of aliphatic polyesters’, in Scott G, Degradable Polymers: Principles and Applications, 2nd edition, Dordrecht (NL), Kluwer Academic Publishers, 71–131. li s, dobrzynski p, kasperczyk j, bero m, vert m (2005), ‘Structure property relationships of copolymers obtained by ring-opening polymerization of glycolide and ε-caprolactone. Part 2. Influence of composition and chain microstructure on the hydrolytic degradation’, Biomacromolecules, 6, 489–497. liu l j, li s, garreau h, vert m (2000), ‘Selective enzymatic degradations of poly (L-lactide) and poly(ε-caprolactone) blend films’, Biomacromolecules, 1, 350– 359.

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lyu s p, schley j, loy b, lind d, hobot c, sparer r, untereker d (2007), ‘Kinetics and time – temperature equivalence of polymer degradation’, Biomacromolecules, 8, 2301–2310. macdonald r t, mccarthy s p, gross r a (1996), ‘Enzymatic degradation of poly(lactide): Effects of chain stereochemistry and material crystallinity’, Macromolecules, 29, 7356–7361. miller r a, brady j m, cutright d e (1977), ‘Degradation rates of oral resorbable implants (polylactates and polyglycolates): Rate modification with changes in PLA/PGA copolymer ratios’, J Biomed Mater Res, 11, 711–719. mochizuki m, hirano m, kanmuri y, kudo k and tokiwa y (1995), ‘Hydrolysis of polycaprolactone fibers by lipase: Effects of draw ratio on enzymatic degradation’, J. Appl. Polym. Sci., 55, 289–296. mochizuki m, hirami m (1996), ‘Structural effects on the biodegradation of aliphatic polyesters’, Polym Adv Technol, 8, 203–209. moon s i, lee c w, taniguchi t, miyamoto m, kimura y (2001), ‘Melt/solid polycondensation of L-lactic acid: An alternative route to poly(L-lactic acid) with high molecular weight’, Polymer, 42, 5059–5062. nakamura t, hitomi s, watanabe s, shimizu y, jamshidi k, hyon s h, ikada y (1989), ‘Bioabsorption of polylactides with different molecular properties’, J Biomed Mater Res, 23, 1115–1130. pistner h, stallforth h, gutwald r, müling j, reuther j, michel v (1994), ‘Poly (L-lactide): Along-term degradation study in vivo, II. Physico-mechanical behavior of implants’, Biomaterials, 15, 439–450. pitt c g, chaslow f i, hibionada y m, klimas d m, schindler a (1981a), ‘Aliphatic polyesters. I. The degradation of poly(ε-caprolactone) in vivo’, J Appl Polym Sci, 26, 3779–3787. pitt c g, gratzl m m, himmel g l, surles j, schindler a (1981b), ‘Aliphatic polyesters. 2. The degradation of poly(DL-lactide), poly(ε-caprolactone) and their copolymers in vivo’, Biomaterials, 2, 215–220. pitt c g (1992), ‘Non-microbial Degradation of Polyesters: Mechanisms and Modifications’. In Vert M, Feijen J, Albertsson A C, Scott G, Chiellini E, Biodegradable Polymers and Plastics, London, Royal Society of Chemistry, 7–19. reed a m, gilding d k (1979), ‘Biodegradable polymers for use in surgery – poly(glycolic)/poly(lactic acid) homo- and copolymers: 2. In vitro degradation’, Polymer, 20, 1459–1464. reeve m s, mccarthy s p, downey m j, gross r a (1994), ‘Polylactide stereochemistry: Effects on enzymatic degradability’, Macromolecules, 27, 825–831. schliephake h, klosa d, rahlff m (1993), ‘Determination of the 3-D morphology of degradable biopolymer implants undergoing in vivo resorption’, J Biomed Mater Res, 27, 991–998. schmitt e e, polistina r a (1967), Surgical sutures, U.S. Pat. 3297033 (Jan. 10, 1967) (to Am. Cyanamid Co.). schmitt e a, flanagan d r, linhardt r j (1994), ‘Importance of distinct water environments in the hydrolysis of poly(DL-lactide-co-glycolide)’, Macromolecules, 27, 743–748. spinu m, jackson c, keating m y, gardner k h (1996), ‘Material design in poly(lactic acid) systems : Block copolymers, star homo- and copolymers, and stereocomplexes’, J Macromol Sci: Pure Appl Chem, A33, 1497–1530.

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st pierre t, chiellini e (1987), ‘Biodegradability of synthetic polymers for medical and pharmaceutical applications: Part 2 – Backbone hydrolysis’, J Bioact Comp Polym, 2, 4–30. thérin m, christel p, li s, garreau h, vert m (1992), ‘In vivo degradation of massive poly(α-hydroxy acids): Validation of in vitro findings’, Biomaterials, 13, 594–600. tsuji h, miyauchi s (2001a), ‘Poly(L-lactide): VI: Effects of crystallinity on enzymatic hydrolysis of poly(L-latide) without free amorphous region’, Polym Degr Stabl, 71, 415–424. tsuji h, miyauchi s (2001b), ‘Poly(L-lactide): 7. Enzymatic hydrolysis of free and restricted amorphous regions in poly(L-lactide) films with different crystallinities and a fixed crystalline thickness’, Polymer, 42, 4463–4467. vert m, christel p, chabot f, leray j (1984), ‘Bioresorbable Polymers for Bone Surgery’. In Hastings G W, Ducheyne P, Macromolecular Biomaterials, Boca Raton, CRC Press, 119–142. vert m, li s, garreau h (1991), ‘More about the degradation of LA/GA-derived matrices in aqueous media’, J Control Rel, 16, 15–26. vert m, li s, spenlehauer g, guerin p (1992), ‘Bioresorbability and biocompatibility of aliphatic polyesters’, J Mater Sci: Mater Med 3, 432–446. von burkersroda f, schell l, gopferich a (2002), ‘Why degradable polymers undergo surface erosion or bulk erosion’, Biomaterials, 23, 4221–4231. wang n, wu x s, lujan-upton h, donahue e, siddiqui a (1997), ‘Synthesis and Characterization of Lactic/glycolic Acid Oligomers’, Proceedings of the American Chemical Society, 76, 373–374. williams d f (1981), ‘Enzymic hydrolysis of polylactic acid’, Eng Med, 10, 5–7. zhang x, wyss u p, pichora d, goosen m f a (1994), ‘An investigation of poly(lactic acid) degradation’, J Bioact Compat Polym, 9, 80–100.

7 In vitro biological test methods to evaluate bioresorbability G. MABILLEAU and A. SABOKBAR, University of Oxford, UK

Abstract: Biomaterials are commonly employed in medicine, odontology and biotechnology. The degradation of a material will depend on the chemical structure of the material employed such as polymer, metal and/ or ceramic. In this chapter the degradation mechanisms of biomaterials are dealt with, using in vitro methods to assess the characterization of their resorption. This chapter addresses specifically biodegradation, bioresorption, bioerosion and bioabsorption. Key words: biodegradation, bioresorption, bioerosion, bioabsorption, mechanism.

7.1

Introduction

Nowadays, biomaterials are commonly employed in medicine, odontology and biotechnology. A biomaterial must fulfil criteria such as: (i) biocompatibility; (ii) biomechanics adapted for the host tissue and (iii) degradability or un-degradability according to its future in the body. The degradation of a material will depend on its chemical structure (e.g. polymer, metal, ceramic); a polymer will be degraded by hydrolytic or oxidative mechanisms, whereas metals will be degraded mainly by corrosion. After implantation in the body, the material is recognized as a foreign body by the cells, mainly immune cells, whose function is to eliminate it. This starts by the adsorption of specific protein on the surface of the material which, depending on the surface material characteristics, serve as substrates for the membrane receptors located on the outer layer of the cytoplasmic membrane (mainly integrins). In the literature, four terms are employed to describe the degradation of polymers:1 (i)

Biodegradation. A biomaterial is described as biodegradable when it is degraded into macromolecules, but the macromolecules stay in the body and can migrate. For instance, the macromolecules of Ultra High Molecular Weight Polyethylene (UHMWPE) generated from joint prostheses. 145

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(ii)

Bioresorption. A material is resorbable if it can be entirely bulk degraded in vivo. It is eliminated from the body via low molecularweight molecules. The major factor for resorbability is the time required for the total elimination of the material; a bioresorbable polymer such as polylactic acid will be eliminated comparatively quickly. Surprisingly, metals could be classified as bioresorbable because corrosion could lead to the total elimination of the material, but the time required is much longer than for other resorbable polymers. (iii) Bioerosion. This term refers to materials that first are degraded on the surface and then resorbed in vivo. (iv) Bioabsorption. In this case the material could dissolve in the body without modification of their molecular weight, a process which is in contrast to bioresorption. In this chapter we will deal with the in vitro (ultimately translated into in vivo circumstances) modes of degradation of biomaterials; with in vitro methods to assess their resorption; and finally with the characterization of resorption.

7.2

Methods of degradation of biomaterials

In the in vitro assay, or after implantation, materials are exposed to saltcontaining solutions and biomechanical stress, which can lead to the generation of particles and wear debris and activation of the immune system (Fig. 7.1). Some materials can be inert in a bulk shape but biologically active when they are in a particulate form (e.g. UHMWPE).

Choice of the material

Polymer

Hydrolytic degradation

Oxidative degradation

Reaction with H2O2

Reaction with free radicals

Ceramic

Dissolution

7.1 Degradation paths for polymers and ceramics.

Cell-mediated resorption

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147

7.2.1 Degradation of polymers There are two different modes of polymer degradation: (i) hydrolytic degradation and (ii) oxidative degradation, both of which lead to the generation of debris and then to the bioresorption of the materials. Hydrolytic degradation Hydrolytic degradation is the cleavage of functional groups by water. This process can be catalyzed by acids, bases or enzymes. In this kind of degradation, the cleavage rate is directly proportional to the initiator rate.2 The susceptibility of a polymer to hydrolytical degradation is the result of its chemical structure (presence of hydrolyzable groups) and of the surrounding biological environment. Many hydrolyzable polymers contain CKO groups linked to another element (O, N or S). In this family, there are esters, amides, thioesters, urethanes, carbonates, imides, anhydrides (Fig. 7.2). Other polymers such as ethers, nitriles, polyphosphonates, sulfonates, sulfonamides and cyanoacrylates can be hydrolyzable under certain conditions3 (Fig. 7.3). The hydrolytic rate increases with the number of hydrolyzable groups, low crystallinity, hydrophilicity, absence of cross-links and the surface area of the material (a porous material will be degraded faster than a non-porous, similarly a rough surface will be degraded sooner than a smooth surface). Inversely, hydrophobic materials, cross-links, high crystallinity and bulk shape decrease the hydrolytic rate. Hydrolysis can be catalyzed by ions. Usually biological and cell culture fluids contain ions such as H+, OH−, Na+, Cl−, HCO3−, H2PO42−, K+, Mg2+, Ca2+ and SO42− (Table 7.1).

Table 7.1 Ionic concentration in blood and extracellular fluid (Ratner et al., 200430) Anion/Cation

Cl− HCO3− HPO42− SO42− H2PO4− Na+ Mg2+ Ca2+ K+

Concentration (mM) Blood

Extracellular fluid

96–111 16–31 1–1.5 0.35–1 2 131–155 0.7–1.9 1.9–3 35–56

112–120 25.3–29.7 102–193 0.4 141–150 1.3 1.4–1.55 3.5–4

148 (1)

Degradation rate of bioresorbable materials O

O H2O

C

X

C (H+, OH–, enzyme)

O C

O O

(2)

O H N

C

Ester

OH + XH

C

Amide

S

Thioester

O

O H2O X

C

X

X'

C

OH + X'H

(H+, OH–, enzyme)

XH + CO2 O O

C

O H N

O

Urethane

C

O

Carbonate

(3) O

O

O

O

H2O C

X

O C

O

Anhydride

C

(H+, OH–, enzyme)

O

O

C

C

C

OH + HX

O H N

C

Imide

7.2 Groups of polymers hydrolyzable containing CKO bonds.

C

In vitro biological test methods to evaluate bioresorbability (1)

H2 C

O

H+

H2 C

H+ or OH– C

C

N

O

H2O

Nitrile (3)

C

NH2

OH

O O

P

O

O + NH4OH

2 H2O

O

H+ or OH– R

H2 C

OH + HO

H2O

Ether (2)

H2 C

149

R'

R

H2O

OH + HO

O

P

OH + HO

O

R"

R"

Phosphonate (4)

O

O H+ or OH– S

S

X

OH + HX

H2O O

O

Sulfonamide or sulfonate (5) CN H2 C

C

CN H2 C

CN

CN OH–

C

H2 C

C

H2 C

OH + HC

H2O CH2

CH2

CH2

CH2

O

O

O

O

R

R

R

R

Polycyanoacrylate 7.3 Other chemical groups susceptible to hydrolysis under certain conditions.

R'

150

Degradation rate of bioresorbable materials

It has been shown that ions such as PO43−, can accelerate the hydrolytic rate of degradation.4 The ion-mediated hydrolysis can be limited only to the surface of the polymer (hydrophobic material) or it can penetrate into the deeper layers (hydrophilic material). Modifications of pH, due to increased inflammatory reactions, can increase the hydrolytic rate in the immediate proximity of the biomaterial. Enzymes are also known to recognize some structural motifs and initiate the degradation of the polymer.5,6 For this reason, synthetic polymers are more resistant to degradation than natural polymers. However, some studies have shown that hydrolases can initiate degradation of synthetic polyesters and polyamides,7–9 but it is still unclear whether certain enzymes can initiate the degradation of JH2CJCH2J bonds. Generally, enzymes exert their activity only at the surface of materials due to their high molecular weight hindering their penetration into the deep layers of the material. Oxidative degradation Oxidation of polymers occurs by a homolytic (formation of free radicals) or heterolytic (one of both atoms keeps the pair of electrons of the chemical bond) reaction. The favourite sites for oxidative degradation are those which allow the withdrawal of an atom or an ion and the stabilization of the corresponding radical or ion (Fig. 7.4). The most suitable sites are carbon substituted by an aliphatic chain, carbon substituted by an aromatic cycle or by an allylic group, ethers, phenols, alcohols, aldehydes and amines. In this kind of degradation, the free radicals generated by the host directly induce the oxidation of the polymer. It is well known now that activated phagocytes (e.g. macrophages, neutrophils) are capable of releasing reactive oxygen species (ROS) and reactive nitrogen species (RNS) in the close proximity of the materials.10 During the first days of implantation, neutrophils, which respond to chemical mediators at the trauma site in the acute or early phase, induce a strong and transitory attack on the material surface by the release of free radicals (ROS and RNS).11,12 Activated macrophages are the second actors of the immune defence and replace the neutrophils in the late phase of the immune response. Activated macrophages are present in the trauma site after two to three days and can persist in the case of infection. The multinucleated giant cells, observed sometimes at the nearest vicinity of the biomaterials, are generated from the fusion of activated macrophages and these polykaryons persist at the site of implantation for months and even years. Neutrophils and activated macrophages can metabolize oxygen to generate the superoxide anion (O2−•) via the NADPH oxidase. This free radical is extremely reactive and can be transformed into the hydroxyl radical (OH•) and initiate the oxidation of the polymer surface. Inside the cells,

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151

7.4 Oxidative attack of polymers (*susceptible point of attack).

specific enzymes are capable of removing these free radicals; superoxide dismutase (SOD), present in the peroxysomes, can convert the O2−• into hydrogen peroxide (H2O2), which itself can be converted into hypochlorite acid (HOCl) by the myeloperoxidase (MPO) inside the lysosomes of the neutrophils. HOCl can oxidize the amine groups of proteins into chloramines that are one of the sources of chloride in the body. HOCl can also oxidize functional amino-groups (amide, urethane) with the possibility of a cleavage of these groups. The macrophages do not possess MPO, so their H2O2 is not converted directly into HOCl. However, neutrophils can release their active MPO into the surrounding tissue which can bind the surface of the biomaterials.3,13 The H2O2 released by the macrophages in the surrounding tissue can be catalyzed into HOCl via the MPO released by the neutrophils.

7.2.2 Degradation of ceramics Ceramics are extremely sensitive to variations in pH, and in acidic conditions, ceramics such as β-TCP, hydroxyapatite are degradable.14 It has been

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Degradation rate of bioresorbable materials

a major concern in the past decade to characterize the cells responsible in vivo for the degradation of the implanted ceramic. Degradation of a bone-implanted ceramic can be mediated by (i) inflammatory multinucleated giant cells, formed by the fusion of macrophage, which degrade the material by phagocytosis with a low-grade resorption, and (ii) osteoclasts, corresponding to physiologic multinucleated cells, involved in the resorption of bone tissue.15 There has been an intense debate whether multinucleated giant cells involved in biomaterial degradation actually are osteoclast as it appeared that they do not possess all the osteoclast features.16,17 To date, there has been accumulating evidence that osteoclasts are capable of resorbing calcium phosphate (CaP) ceramics in vitro and in vivo.18–20 Osteoclasts cultured on CaP ceramic develop typical ultrastructural features of bone osteoclasts, such as a polarized dome shape, a clear zone and a ruffled border.19,21 Modification of the shape and density of CaP crystals under the ruffled border can include an acidic microenvironment. Moreover, osteoclasts can degrade ceramic by simultaneous resorption and phagocytic mechanisms. Many environmental factors are involved in the gradual degradation of CaP ceramic after implantation, including physicochemical processes (dissolution–precipitation) and the effect of various cell types. Heymann et al. (1999) reviewed the cellular mechanisms of CaP ceramic degradation.22 Ceramics can be degraded by phagocytic mechanisms by cells such as fibroblasts, osteoblasts, monocytes/macrophages or by an acidic mechanism generated by a proton pump to reduce the pH of the microenvironment and resorb this synthetic substrate; the main function of an osteoclast. Osteoclasts resorb CaP ceramics in a similar process to that observed with the natural bone involving the following three phagocytic steps: (i) Crystal phagocytosis, (ii) disappearance of the endophagosome envelope membrane and (iii) fragmentation of phagocytosed crystals in the cytoplasm.21 However, the physicochemical properties of the ceramic, mainly its solubility and composition of Ca, influence this degradation process. Ceramics with high rates of dissolution increase the intracellular concentrations of Ca in the osteoclast which, in turn, results in the disorganization of the intracellular actin network present in the osteoclast podosomes and subsequent detachment of the osteoclast from the material surface.

7.3

Methods of assessing resorbability in vitro

In vitro assessment of the biological resorbability is essentially determined by using cell culture techniques. The key points worth considering herein are the method to be employed to culture the cells at the surface of the material and the choice of the cell type.

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153

One of the important parameters to take into account is the choice of a relevant cell type susceptible to being in contact with the biomaterial after implantation. For example, bone prosthesis should be tested with monocyte–macrophage, bone cells or endothelial cells which are likely to be in contact with but certainly not with brain cells. As soon as the appropriate cell type is chosen, the next step is to choose a suitable shape and size of the material for the cell culture and analyses. The choice of the material is a key point in the in vitro assay. The material itself will determine the adhesion or non-adhesion of the cells at its surface. Post-implantation, the surface of the material is adsorbed by proteins circulating in the biological fluids. These proteins serve as a substrate for cell surface receptors, mainly integrins, for the cells to adhere at the surface of the material. However, the nature of the protein layer and the quantity of the protein adsorbed at the surface of the material is dependent on the surface material properties. Protein adsorption is influenced by the roughness of the surface; the rougher the surface, the greater the concentration of proteins deposited. Protein adsorption is also dependent on the chemical composition of the surface and its surface energy (influenced by the free surface energy, the zeta potential and the wettability of the surface), thus influencing the nature of the protein that could adhere on the surface. It is well known that the positively charged surfaces induce cell adhesion and proliferation, whereas negatively charged surfaces exhibit a very low rate of cell adhesion. The surface charge can be modified by the pH of the medium used for the cell culture. The pH of the medium is also an important factor to take into account, as it will influence the charge of the material, and the cell behaviour. For example, mature osteoclasts are more re-sorbing in acidic pH than neutral or basic pH. For these reasons, it is usually necessary to pre-incubate (preferably overnight) the material in the culture media required for the cell culture in order to allow the protein deposition on the biomaterial surface to have maximal adhesion. The shape (e.g. disk, cylinder, beads) and the thickness of the material are also important. The choice of the shape and thickness of the material under investigation depends on the type of study to be conducted. In order to visualize the cells at the surface of the material after some days in culture (albeit by light, fluorescence or scanning electron microscopy), having the material in a disk shape would be the most suitable form; facilitating the rapid fixation of the specimens on a glass slide or metal stub. Furthermore, any changes in the surface roughness (an indicator of any degradation) can be detected on the disk by either contact profilometry, scanning probe microscopy or image analysis with a fractal algorithm. However, if the experimental design requires the detection of the resorption notches, the use

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Degradation rate of bioresorbable materials

of ‘round’ shape materials in the form of cylinders or beads, would be preferable as it facilitates the embedding the semi-thin or ultra-thin sections.

7.4

Characterization of the resorbability in vitro: Microscopic analysis of the surface

The characterization of the resorbability of a material is achieved mainly by microscopic observation of the surface. It is well known that resorbability of the surface can be measured by an increase in the surface roughness. However, when the material contain specific cleavable sites, it is possible to assess the quantity of lateral chain cleaved by biological means, hence facilitating the quantification of the rate of resorption. Microscopic analyses of any surface can be determined at different magnitude scales: microscopic levels by scanning electron microscopy, confocal and fluorescent microscopy, contact and laser profilometry with image analysis using a fractal algorithm; nanoscopic levels by transmitting electron microscopy and scanning probe microscopy.

7.4.1 Electron microscopy There are two types of electron microscopies: scanning electron microscopy (SEM) and transmitting electron microscopy (TEM). With the SEM technique, the surface of the samples is scanned by an electron beam. The energy brought by the electron beam induces the emission of new electrons from the sample surface, mainly secondary electrons, which are collected to a specific detector for secondary electrons and form the image. However, the sample surface can emit other electron types such as backscattered electrons, specific for the chemical elements, auger electrons and Xelectrons; the presence of specific detectors for these electrons is required to obtain the specific image. To assess whether a material is resorbable, it is recommended to remove the cells from the surface. Using SEM, it is possible to scan the sample surface before and after cell culture and compare the surface roughness. As such, if the material has been exposed to degradation by cells, the presence of peaks and valleys on the surface induced by the enzymatic activity of the cells increases the surface roughness. However, it would be rather difficult to quantify the depth of the resorption lacunae using this approach. With transmitting electron microscopy (TEM), the sample is embedded into a specific resins to be cut in semi-thin (1.5–2 μm thickness) and ultrathin (75–100 nm thickness) sections. Semi-thin sections are stained with the usual stains e.g. toluidine blue or Azur II to visualize the presence of the cells and eventually the presence of a notch in the material below the cell. It is worth noting that, usually, semi-thin sections serve only as a location

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155

of the sample, and as soon as the region of interest has been identified, the sample is cut into ultra-thin sections to be observed by TEM. An ultra-thin section is deposited on a specific metallic grid (mainly made of copper) and stained with heavy elements, mainly uranium and lead. Then the grid is introduced into the electron microscope and the electron beam will focus on the ultra-thin section. Because of the thickness of the sample, the electrons emitted from the probe of the microscope can cross the sample and are collected on a specific detector (and only transmitted electrons). The presence of heavy elements is used to impair the electrons from the source from crossing the sample, so actually the image collected on the detector is an 8-bit image in grey levels, the white zone corresponding to the electrons which cross the samples and the black zone to the electrons collected by the heavy elements. On a TEM image it is easy to assess the presence of the notch on the surface of the biomaterial and the presence of the material inside cell vacuolations. On a TEM image, it is possible to assess the depth of the notch. For both SEM and TEM, the sample needs to be dehydrated and some artefacts, due to the dehydration process, can interfere with the ‘real’ size and shape of the surface. Moreover, with the TEM technique, it is worth noting that the shape and the depth of the notch could be slightly modified by the embedding and cutting processes.

7.4.2 Contact microscopy This kind of microscopy divides into profilometry (contact or laser profilometry) and scanning probe microscopy [Atomic Force Microscopy – (AFM), Scanning Tunnelling Microscopy – (STM)]. Profilometry Two different types of profilometry have been documented: contact profilometry where a metallic probe scans the surface, and laser (or optical) profilometry, where a light or laser beam scans the surface without direct interaction with the surface in contrast to the contract profilometry. Contact profilometry is the oldest technique. It is a mechanical technique which uses a thin probe (2.5 μm at its extremity) which scans the sample surface. The diameter of the probe at its extremity is the limiting factor (Fig. 7.5). The movements of the probe are transmitted and amplified by a cantilever to determine the Zo line. Below this line are defined the notches or pits at the surface of the material, above this line are defined the surface asperities. Optical profilometry is a more recent and modern approach and has been developed to increase accuracy. Briefly, a source of light is used to scan the

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Degradation rate of bioresorbable materials

Profilometer probe

Biomaterial

Surface topography determined by contact profilometry 7.5 Surface topography determined by contact profilometry. The size of the probe is the limiting factor. Note that the probe cannot go to the end of the valley and that the surface roughness is therefore underestimated.

sample surface and the light beam diffracted by the surface roughness is collected on a mirror. The image generated is the deviation of the light beam on the mirror. With this technique, it is possible theoretically to assess the roughness as low as a nanometer. However, the limitation of this technique is to have a surface capable of reflecting the light beam and often it is necessary to modify the sample surface to obtain a better reflection. Scanning probe microscopy Atomic force microscopy is now widely used to assess the physical surface properties of materials. The principle of this technique is schematically represented in Fig. 7.6. Briefly, a probe is located at the end of a microlever and scans the surface of the sample. The interactions between the probe and the sample surface induce deflections of the microlever. The deflections of the microlever are detected by an optical system: a laser beam is focused on the extremity of the microlever, just above the probe. The samples are located at the surface of a piezoelectric tube which allows displacement in the x, y and z directions. With the interaction force varying from point to point (depending of the roughness of the sample), it is possible to observe deflections of the microlever. A feedback loop is used to maintain as a

In vitro biological test methods to evaluate bioresorbability

157

Mirror Laser

Microlever

Feedback loop Sample AFM image Piezoelectric tube 7.6 Principle of AFM microscopy. A laser can detect the deflections of the microlever, which reflect it to the mirror. A feedback loop maintains the constant position of the microlever to the sample surface by dilating or retracting the piezoelectric tube. The movements of the piezoelectric tube are used to generate the AFM images.

constant the position of the microlever compared to the sample surface and the image recorded corresponding to the current sent by the feedback loop to the piezoelectric tube to maintain constant the position of the microlever compared to the sample surface. The size of the probe is thinner than the one used for contact profilometry (average diameter 45–200 nm). The other advantage of the AFM technique is that it is possible to operate in a contact mode (which represents a distance of 10–100 angstroms between the end of the probe and the sample surface) or in a tapping mode, when the probe comes to the surface and is deflected from each point, with a specific vibration frequency. With this last mode it is possible also to determine differences in hardness of the surface material. Both profilometry and AFM allow one to calculate the roughness coefficient (Ra), which corresponds to the average height of the peaks and valleys on the surface of the material. It is calculated using the following equation: ⎛ n zn ⎜∑ i =1 Ra = ⎜ ⎜ N ⎜⎝

⎞ ⎟ ⎟ ⎟ ⎟⎠

[7.1]

where ⏐Zn⎪ is the absolute value of the difference in levels between the Z0 line and the peak or the valley for each point and N is the number of points. However, profilometry and AFM are limited by the geometry of the

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sample (examples: orthopaedic screw) and cannot be used when the geometry of the surface is complex. As such, for the past five years, surface roughness of materials has been used intensively to characterize degradation processes.23–25 Image analysis with a fractal algorithm Surface roughness can be evaluated by texture analysis of microscopic images using a fractal algorithm (the ‘skyscraper’ algorithm).26–28 It is necessary to take pictures of the surface by episcopic illumination or by SEM. Briefly, images of the surface need to be converted and coded on 8 bits (i.e. in 256 gray levels, where black = 0 and white = 255). Pixels, which constitute an image A, can be considered as skyscrapers whose height is represented by the gray level. The roof of a skyscraper is a square of side ε. The surface area of the image A(ε) is obtained by measuring the sum of the top surfaces (ε) and sum of the exposed lateral sides of the skyscrapers. The gray levels of adjacent pixels are then averaged in squares of ε: 2, 4, 8, 16 and 32 pixels to produce new images and A(ε) is calculated for each ε according to: A ( ε ) = ∑ ε 2 + ∑ ε [abs [Z ( x, y) − Z ( x + 1, y)] + abs [Z ( x, y) − Z ( x, y + 1)]] [7.2] where Z(x,y) is the height of a skyscraper in the x, y plane and abs is the absolute value. The fractal dimension of the surface (D) is determined by plotting a graph of log A(ε) vs. log ε. The linear regression line is computed, only on the aligned points, by the least-squares method. The fractal dimension was obtained as D = 2 − slope. The more a surface is degraded, the higher is the likelihood the fractal dimension will reach 3. This method has been successfully used to assess the surface roughness of titanium disks on SEM images and the surface roughness of hydrogel disks by episcopic illuminations.26–28

7.4.3 Fluorescent microscopy and biological testing Fluorescent microscopy employs the presence of specific molecules capable of absorbing energy of photons to re-emit the energy at a different wavelength. To increase the flexibility of the approach, increasing numbers of fluorescent probes have been coated on materials to assess their biodegradability. The principle of the microscopic examination is simple; a fluorescent probe (usually conjugated to a fluorophore such as FITC or Rhodamine) is deposited on the surface. When the material surface is degraded, the probe is released in the medium and the degraded surface of the material appears ‘non-fluorescent’. A classical example is a study by Rogers and his team29 whereby the dentine slice was exposed to a fluorophore-conjugated

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bisphosphonate. When the osteoclasts had resorbed the organic matrix, the pits appeared dark, whereas the rest of the dentine surface and the osteoclasts appear fluorescent. Biological and biochemical tests can also be used to evaluate the resorption rate. These techniques are suitable to assess the degradation rate in vivo. The principle of the assay is to collect the supernatant of the cell culture and to quantity the degradation products. For example, the resorption of dentine slice can be followed by a biochemical test, such as enzyme linked immunosorbent assay (ELISA), by measuring the amount of the degradation products of collagen Type I. The increase in the release of the degradation products is directly proportional to the rate of the resorption process.

7.5

References

1 vert m, feijen j, albertson a, scott g, chiellini e. Biodegradable Polymers and Plastics. Cambridge, U.K.: Royal Society of Chemistry, 1992. 2 schnabel w. Polymer Degradation Principles and Practical Applications. New York: Macmillan, 1981:15–7. 3 coury aj. Chemical and biochemical degradation of polymers. In: Ratner BD, Hoffman AS, Schoen FJ, Lemons JE, eds. Biomaterials Science: An Introduction to Materials in Medicine. London: Elsevier, 2004:411–30. 4 zaikov ge. Quantitative aspects of polymer degradation in the living body. JMSRev. Macromol. Chem. Phys. 1985;C25:551–97. 5 kopecek j, ulbrich k. Biodegradation of biomedical polymers. Prog. Polym. Sci. 1983;9:1–58. 6 smith r, oliver c, williams df. The enzymatic degradation of polymers in vitro. J Biomed Mater Res 1987;21–8:991–1003. 7 labow rs, erfle dj, santerre jp. Neutrophil-mediated degradation of segmented polyurethanes. Biomaterials 1995;16–1:51–9. 8 pitt cg. Non microbial degradation of polyesters: Mechanisms and modifications. In: Vert M, Feijen J, Albertson A, Scott G, Chiellini E, eds. Biodegradable Polymers and Plastics. Cambridge: R. Soc. Chem, 1992:1–19. 9 santerre jp, labow rs, duguay dg, erfle d, adams ga. Biodegradation evaluation of polyether and polyester-urethanes with oxidative and hydrolytic enzymes. J Biomed Mater Res 1994;28–10:1187–99. 10 rajagopalan s, meng xp, ramasamy s, harrison dg, galis zs. Reactive oxygen species produced by macrophage-derived foam cells regulate the activity of vascular matrix metalloproteinases in vitro. Implications for atherosclerotic plaque stability. J Clin Invest 1996;98–11:2572–9. 11 northup sj. Strategies for biological testing of biomaterials. J Biomater Appl 1987;2–1:132–48. 12 rosen gm, pou s, ramos cl, cohen ms, britigan be. Free radicals and phagocytic cells. Faseb J 1995;9–2:200–9. 13 locksley rm, wilson cb, klebanoff sj. Role for endogenous and acquired peroxidase in the toxoplasmacidal activity of murine and human mononuclear phagocytes. J Clin Invest 1982;69–5:1099–111.

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14 koerten hk, van der meulen j. Degradation of calcium phosphate ceramics. J Biomed Mater Res 1999;44–1:78–86. 15 suda t, nakamura i, jimi e, takahashi n. Regulation of osteoclast function. J Bone Miner Res 1997;12–6:869–79. 16 blouin s, moreau mf, weiss p, daculsi g, basle mf, chappard d. Evaluation of an injectable bone substitute (betaTCP/hydroxyapatite/hydroxy-propyl-methylcellulose) in severely osteopenic and aged rats. J Biomed Mater Res A 2006;78–3:570–80. 17 dersot jm, colombier ml, lafont j, baroukh b, septier d, saffar jl. Multinucleated giant cells elicited around hydroxyapatite particles implanted in craniotomy defects are not osteoclasts. Anat Rec 1995;242–2:166–76. 18 basle mf, chappard d, grizon f, filmon r, delecrin j, daculsi g, rebel a. Osteoclastic resorption of Ca-P biomaterials implanted in rabbit bone. Calcif Tissue Int 1993;53–5:348–56. 19 heymann d, guicheux j, rousselle av. Ultrastructural evidence in vitro of osteoclast-induced degradation of calcium phosphate ceramic by simultaneous resorption and phagocytosis mechanisms. Histol Histopathol 2001;16–1:37–44. 20 yamada s, heymann d, bouler jm, daculsi g. Osteoclastic resorption of biphasic calcium phosphate ceramic in vitro. J Biomed Mater Res 1997;37–3:346–52. 21 wenisch s, stahl jp, horas u, heiss c, kilian o, trinkaus k, hild a, schnettler r. In vivo mechanisms of hydroxyapatite ceramic degradation by osteoclasts: Fine structural microscopy. J Biomed Mater Res A 2003;67–3:713–8. 22 heymann d, pradal g, benahmed m. Cellular mechanisms of calcium phosphate ceramic degradation. Histol Histopathol 1999;14–3:871–7. 23 dumas a, gaudin-audrain c, mabilleau g, massin p, hubert l, basle mf, chappard d. The influence of processes for the purification of human bone allografts on the matrix surface and cytocompatibility. Biomaterials 2006;27–23: 4204–11. 24 mabilleau g, bourdon s, joly-guillou ml, filmon r, basle mf, chappard d. Influence of fluoride, hydrogen peroxide and lactic acid on the corrosion resistance of commercially pure titanium. Acta Biomater 2006;2–1:121–9. 25 mabilleau g, moreau mf, filmon r, basle mf, chappard d. Biodegradability of poly(2-hydroxyethyl methacrylate) in the presence of the J774.2 macrophage cell line. Biomaterials 2004;25–21:5155–62. 26 chappard d, degasne i, hure g, legrand e, audran m, basle mf. Image analysis measurements of roughness by texture and fractal analysis correlate with contact profilometry. Biomaterials 2003;24–8:1399–407. 27 mabilleau g, basle mf, chappard d. Evaluation of surface roughness of hydrogels by fractal texture analysis during swelling. Langmuir 2006;22–10:4843–5. 28 mabilleau g, stancu ic, honore t, legeay g, cincu c, basle mf, chappard d. Effects of the length of crosslink chain on poly(2-hydroxyethyl methacrylate) (pHEMA) swelling and biomechanical properties. J Biomed Mater Res A 2006;77–1:35–42. 29 coxon fp, thompson k, ebetino fh, rogers mj. Resorbing osteoclasts increase the availability of mineral-bound bisphosphonates to non-resorbing cells. Symposium: Frontiers of Skeletal Biology and What’s New in Bisphosphonate. Davos: Bone, 2006. 30 ratner bd, hoffman as, schoen fj, lemons je. Biomaterials Science: An Introduction to Materials in Medicine. London: Elsevier, 2004.

8 In vivo test methods to evaluate bioresorbability S. A. CLARKE and G. R. JORDAN, Queen’s University Belfast, Northern Ireland

Abstract: Although material characterization and in vitro testing can provide valuable information, it is only possible to fully understand the bioresponse to a material when it is implanted in vivo. This chapter describes animal models that have been previously used to test bioresorbable materials, with practical considerations for choosing the most appropriate. It also describes key outcome measures, such as histology and histomorphometry, that can specifically measure material resorption. Key words: animal models, bone defects, material resorption, histomorphometry.

8.1

Introduction

It is important to adequately characterize a new, or modified, material in terms of its chemical and physical properties before in vivo testing. Furthermore, in vitro tests, either with primary explant cells or cell lines, are screening tools that can provide valuable preliminary results of biocompatibility and bioresorbability of the material. Yet, as much as one would wish to replace the use of live animals in research, it is possible only to fully understand the bioresponse to a material when it is implanted in vivo. Accepting this, however, does not negate our responsibility to the philosophy of the 3 Rs of animal research – reduction, refinement and replacement – and we must ensure that the testing we perform is necessary. Therefore, although this chapter is primarily concerned with measuring material resorption, it will also cover some general principles of good practice for in vivo research. The chapter includes descriptions of models that have been previously used to test bioresorbable materials, with practical considerations for choosing the most appropriate, and considers outcome measures that will specifically measure material resorption. 161

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8.2

In vivo models

8.2.1 Pretest considerations Before committing to in vivo testing, it is important to ensure that the tests are absolutely necessary. If animal testing is being considered, it is assumed that the researchers have obtained the relevant licences and ethical approvals applicable to the host country and, in many cases, this will automatically mean they are committed to the 3Rs of animal research. Therefore it is the responsibility of the project licence holder (or equivalent) to perform the checks necessary to ensure that the work is required. •

Replacement – If it is possible to obtain the information by any other means, then the use of animals is not supported. As mentioned previously, this is not always possible because information gleaned from material characterization and in vitro testing will not provide the whole picture. An example of this in terms of bioresorption, is that in in vitro resorption assays, osteoclasts have permanent access to the material surface but in vivo, a layer of bone can form very quickly on the material surface, blocking osteoclast access.1 However, it is sometimes possible, especially if only a minor modification has been made to an existing material, to find the information within previously published results. More rarely, there may be an established method or predictive model that will provide an alternative to animal testing. Therefore comprehensive searches of (a) scientific databases to ensure the novelty of the work and (b) 3R-specific databases to ensure that there is no alternative method to animal testing are required. Details of 3Rs-related databases can be found on The National Centre for the Replacement, Refinement and Reduction of Animals in Research website (www.nc3rs.org.uk). • Reduction – The number of animals used in each experiment should be the minimum required to provide meaningful results and should involve the least sentient species possible (i.e. species with the least capacity for pain and distress). One of the basic steps to determine the minimum number of animals/samples per group is to perform a power calculation during the experimental design phase, but this can sometimes be difficult. If it is the first time the model has been used, then the magnitude of the response and the inter-group variation may be difficult to predict. It is recommended that these parameters are based either on previous results with the model or on the smallest difference between groups that one would consider to be clinically relevant. In some cases, the number of animals in each experiment can be reduced by implanting more than one test material, e.g. the use of bilateral defects to test experimental and control materials. Ethically, it is acceptable to use this method of reducing animal numbers only as long as the level of distress

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and suffering inflicted on the remaining animals is not significantly increased – it is more correct to use more animals who suffer less pain than fewer animals who suffer more pain. Furthermore, if there is a possibility that the implant will have systemic effects (e.g. when testing drug delivery), obviously, the use of a contralateral limb as control or further test site is not scientifically viable. For further consideration of these and other points to consider when determining animal numbers per group, see Festing et al. The Design of Animal Experiments – Reducing the Use of Animals in Research Through Better Experimental Design.2 There are also international guidelines which state the recommended maximum number and size of multiple defect sites for each species (ISO 10993:6). • Refinement – The procedures inflicted upon animals and their experience during the research process should cause as little distress as possible. This often indicates considerations of husbandry and care, such as group housing for rodents and environmental enrichment wherever possible, but can also mean refinements to experimental protocols. Quite aside from the obvious measures of adequate peri-operative analgesia and anaesthesia, other refinements can be made such as taking radiographs every two weeks instead of weekly, if no significant increase in information is gleaned from the additional monitoring. Refinement of the procedures, however, must be balanced with the need to maximize the information obtained from each set of experiments so that they do not need to be repeated at some future date for further analysis. Each researcher must be responsible for implementing refinements to the specific procedures involved in their experiments but for refinements on procedures applicable to all animal experimentation, such as noninvasive methods of animal identification, monitor the NC3Rs website. Finally, it is necessary to discuss the importance of control materials. As with any scientific study, lack of proper controls renders the results unreliable and invalid, but when performing in vivo research, it also makes the experiments unethical. If a new material is being tested, it should be compared to a known standard; if a treated material is to be tested (e.g. growth factor loaded cement), it should be tested against the untreated material at an absolute minimum. We would also stress the importance of using positive controls and would suggest that any experimental material should be compared to the current best treatment (also recommended by ISO10993 guidelines). For example, when testing synthetic bone substitutes which are designed to improve or shorten healing times, autograft or allograft controls, which are the current clinical treatments, should be used. If the model does not allow this, adding fresh autologous bone marrow aspirate or a known effective growth factor such as bone morphogenetic protein-2 to the

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material may provide a suitable alternative. Finally, when considering bone defect models in particular, it is often the practice to include an empty defect group. This control is necessary if information about spontaneous healing of the defect is required, for example if one is creating a ‘criticalsized’ defect, i.e. one that will not heal itself over time. In our view, however, this control is inappropriately used and often results in the use of a greater number of animals with very little experimental gain. It should be used only to determine the natural progression of healing in order to characterize the model initially in a small number of animals. There is no need to include this group every time the model is used subsequently; it is not an appropriate negative control (the untreated material should be used for comparison), neither is it clinically relevant (a bony defect would never be left unfilled).

8.2.2 Described models There are a host of in vivo models already described in the literature for testing biomaterials. Details of some of these are given in Table 8.1, but this list is not exhaustive. Species Both small and large animals have been employed in this type of research. The smallest species used is generally the rat; muscle and air-pouch models in mice have been used to test the osteoinductive properties of biomaterials but they are too small for any meaningful evaluation of resorbability. The other small animal popular for this type of research is the rabbit. It has an advantage over the rat in that it allows larger defects to be created and therefore larger granules or segments of biomaterial, more closely resembling those used clinically. This is partly the reason why large animals, such as sheep, goat, dog and pig, are also widely used. Implantation sites Obviously the site of implantation will depend to a great extent upon the intended use of the material. Dental materials tend to be tested in mandibular or other periodontal defects3–6 and there are some models specifically designed to test spinal disc repair7 and vertebral fusion.8,9 Initial tests of biocompatibility or bioresorption of a new or modified material, however, tend to be performed in calvarial or long bone defects. This leads to large variation within the in vivo modeling conditions, as each site is a unique combination of bone architecture and loading conditions (Table 8.1). For example, femoral condylar defects test the resorption of a

Table 8.1 Selected features of published in vivo models Species

Site

Bone type

Defect size

Loaded/unloaded

rat rat rat rat rat

mandible calvaria radius femur tibia

cancellous can + cort cortical cortical cancellous

20 mm3 4–8 mm ø 5 mm 4.5–8 mm 27 mm3

loaded unloaded loaded loaded ?

rat rabbit rabbit rabbit rabbit rabbit

tibia calvaria ulna radius femur femur

cortical can + cort cortical cortical cancellous cortical

4 mm 6–8 mm ø 5–20 mm 10–20 mm 3–8 mm ø ×6–10 mm 3–12 mm

rabbit

femur

cancellous

3.5 mm ø ×6 mm

loaded unloaded loaded? loaded? loaded loaded (bicortical) unloaded (unicortical) loaded

rabbit minipig minipig minipig

tibia mandible tibia calvaria

cancellous cancellous cancellous can + cort

10 × 15 × 2 mm 5 cm3 3.5–4.7 cm3

unloaded loaded loaded unloaded

Features

metal implant in situ and open to joint

Reference 6 31, 49–51 52 53–56 17

57 58–62 28 63, 64 65–70 67, 71 osteochondral defect

19 72 4 73 74

Table 8.1 Continued Species

Site

Bone type

Defect size

Loaded/unloaded

pig pig pig goat goat dog

intervertebral disc calvaria orbital iliac crest femur periodontal

– can + cort can + cort cancellous cancellous can + cort

– 10 mm 2 cm × 2 cm 8.7 × 106 cm3 1 mm varied

loaded unloaded unloaded unloaded loaded varied

dog dog dog dog sheep sheep sheep sheep sheep sheep

ulna femur femur tibia calvaria femur metatarsus tibia orbital rib

cortical cortical cancellous cortical can + cort cortical cortical cortical can + cort can + cort

25 mm 21–40 mm 5 × 6 × 10 mm 15 mm ø × 20 mm 20–23 mm ø 5 mm ø ×10 mm 25 mm 3–4 cm 7 cm2 5 mm

loaded loaded loaded loaded unloaded loaded loaded loaded unloaded unloaded

ø = diameter.

Features

number of periodontal sites described

Reference 7 14 75 38, 76 77 3, 5, 78–80

81 82, 83 84 15 29, 85 86 87 88–90 22 11

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material in a wholly cancellous bony site and mid-shaft or diaphyseal defects test materials in a cortical bone/marrow milieu. Both of these sites will generally result in loads being transmitted through the implanted material but the extent of load transfer will depend upon (a) the degree to which the animal weight-bears on the limb post-surgery (the animal may guard the limb if it is in pain), (b) the long bone and animal involved (radial and ulnar defects in rabbits may not be loaded to any significant degree if the animal is in pain or if there is limited movement around the cage. It may be beneficial to allow rabbits regular access to a larger run which will encourage weight bearing and mechanical loading through the material but care must be taken to exercise all animals equally.) and (c) the method of fixation (segmental defects in the femur or tibia must be fixed in some way [Radial and ulnar bicortical defects do not need additional fixation as the unaffected bone acts as a natural splint] and the rigidity of the fixation may limit load transfer). Bone architecture of the implantation site may also affect host response. Cancellous bone has a higher rate of bone turnover than cortical10 which, in theory, should lead to greater resorption of the biomaterial. However, Schopper et al. found that there was slightly more implant resorption in the cortical side of a corticocancellous costal defect in sheep.11 This may be because a slower remodeling rate allows osteoclasts greater access time to the material surface, similar to the effect seen when basic fibroblast growth factor (bFGF) is loaded onto the biomaterial (bFGF delays osteoblast differentiation).12 Walsh et al. found that healing progressed more quickly in the cortical area of a metaphyseal/diaphyseal rabbit tibial defect compared to the cancellous site; however, although they anecdotally reported greater resorption in cortical bone, they reported no statistically significant difference between the sites in this parameter.13 Schopper et al.11 also highlighted the replacement of like with like; therefore, biomaterials implanted in cortical sites become surrounded by bone and, over time, are remodeled into the osteon/haversian architecture. In cancellous sites, in the short term, the histological appearance may be disorganized but in the longer term, trabecular architecture begins to be restored.12,14 Lamellar bone is laid down bridging trabecular remnants and the biomaterial may not be completely surrounded by bone unless it is encased in a trabecular strut (Fig. 8.1). In medullary sites, depending on the rate of material resorption/dissolution, bone may be formed on the implant surface initially but this is eventually remodeled so that the medullary cavity is re-established.15 There is limited evidence that methods of bone formation in the same implantation site may differ depending on the material implanted.

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Degradation rate of bioresorbable materials I P

B M

8.1 Reorganization of trabecular struts at 26 weeks following implantation of β tricalcium phosphate particles loaded with transforming growth factor β in a rat tibial model. Key: I: tibial implant, P: βTCP Particle, M: marrow, B: bone. Toluidine blue stained section. (Reproduced with permission and copyright© of the British Editorial Society of Bone and Joint Surgery.30)

For example Gao et al.16 reported trabecular-like bone formation in a diaphyseal sheep tibial defect in response to tricalcium phosphate compared to the ‘disseminated, patchy’ bone contact of bioglass and ‘buttressed’ bone contact of hydroxyapatite. The wide variation in macro and microstructure of these materials and in the way they dissolved and degenerated in the implantation site, however, make it extremely difficult to relate healing response directly to the biomaterial. Structure and surface characteristics, porosity, pore size and degree of connectivity of the implant affect surface area, diffusion of nutrients, cellular attachment, migration and revascularization.11 Therefore, unless these parameters are strictly controlled (an extremely difficult task), it is impossible to isolate the effect of the biomaterial itself to healing and resorption. Specialist models It is the purpose of an animal model either to simplify and control variables in order to isolate individual effects or to recreate the clinical environment as closely as possible. Therefore it may not always be appropriate to test a biomaterial in a segmental or cancellous defect. A number of models have been described which recreate specific features of the clinical environment; for example, the presence of a permanent implant (joint replacement surgery),17 osteochondral defects (full thickness damage to articular

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surfaces),18,19 patellar tendon defects,20 intervertebral disk reconstruction,7 spinal fusion8,9,21 and reconstruction of the orbital wall.22 These models test the ability of a biomaterial to serve a specific purpose rather than, or in addition to, examining osteoconductive properties and bioresorption per se.

8.2.3 Choosing the most appropriate model Although we have said above that one aim of an animal model is to recreate the clinical situation as closely as possible, in reality exact replication is not possible and some degree of compromise is necessary. It is vitally important that, whichever model is chosen is reproducible, reliable, well characterized and fit for purpose: beyond that, considerations of convenience, cost and acceptability to society tend to be deciding factors in the final choice. Species There are several good reviews which compare the closeness of biological23 and mechanical24 characteristics of several species with that of the human. No animal model will replicate the human situation due to obvious differences in locomotor, hormonal and physiological characteristics but some are more appropriate than others. It is well known that many species, rodents in particular, have a much increased metabolic rate compared to humans and this affects their rate of healing and susceptibility to infection. In addition, there are also species differences in the rates of bone remodeling which can directly affect the amount of bone formation and implant resorption. For example, rats have haversian systems similar to human cortical bone but their basic remodeling units have a much reduced remodeling capacity and, unlike in humans, non-unions in dog spinal models are rare.24 Skeletal maturity should also be a consideration as young animals heal faster and more consistently – see Festing et al. for details of maturation age and life span for commonly used experimental species.2 Size The size of the animal becomes important when considering the material bulk properties. If large granules are to be used or the implant has a complex 3D structure which is important to material design and handling properties, then a small animal may not be appropriate – a large animal model will allow use of a material more closely related to what will be commercially available. The experimental design of a study may also dictate the size of animal used as larger animals allow multiple implantation sites for paired sample analysis. Rabbits are the smallest animal from which

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sufficient iliac crest bone can be harvested to provide an autograft control group: below this, allograft must be used. Cost and convenience In reality, considerations of cost, availability of resources and convenience will play a large part in the decision of which model to use. Because small animals are cheaper to source and to house, they allow larger group sizes to be used which will increase the power of a study. Furthermore, if longterm material resorption is to be studied, the use of rodents will reduce the length of time required as their life span is relatively short.2 Using large animal models such as sheep and goats requires more resources (particularly in terms of housing) and they are more difficult to handle. Surgical procedures on these animals are also more complex, as is surgery in any animal that requires external fixation. Finally, considerations of acceptability to society are also important and geographical variations in the species of choice for a given condition are common.

8.3

Outcome measures

Measurements of material resorption in vivo may be direct, e.g. measures of osteoclast number and activity17,25 or indirect, e.g. the size and shape of material on X-ray.16 It is difficult to standardize exposure and position of radiographs and almost impossible to obtain quantitative data; therefore, most measures of resorption are performed on histological sections. Specimens can be processed for histology using either decalcified or undecalcified methodology. When retrieving specimens for decalcified histology, it is preferable to leave some of the soft tissue surrounding the bone intact as it aids cutting. Indeed, it is preferable to leave soft tissue in place whenever potential disruption of the defect site is a concern, regardless of the processing method. A formaldehyde-based fixative is preferable as it preserves tissues better than alcohol-based methods.26 Following processing and cutting, sections are stained to differentiate tissue types. A number of special stains exist which will specifically identify bone and osteoid (unmineralized bone) (Table 8.2). Alternatively, some researchers prefer to visualize the sections using back-scattered scanning electron micrographs.11,13,25 From these micrographs, resorption can be measured in a number of ways. Firstly, using the area of material remaining in the defect. Most often resorption is compared between groups but occasionally a time 0 group has been included in the experimental design25 which allows an absolute measure of the amount of material resorbed. It is usual that several different time points will be examined to determine initial resorption and long-term remodeling of the biomaterial. Depending on

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Table 8.2 Histological stains suitable for differentiation of bone tissue (Bancroft & Gamble, 200237 and Villanueva & Lundin, 198991) Stain

Bone/mineralized tissue

Osteoid

Nuclei

Villanueva Toluidine blue Goldner’s trichrome Stevenel’s Blue & van Gieson Alizarin red & methylene blue Haematoxylin & eosin von Kossa & van Gieson picro fuchsin Solochrome cyanin R van Gieson Sanderson’s Rapid Bone Stain & acid fuchsin

green unstained green red

red blue orange-red green

purple dark blue blue-grey blue

orange-red

blue

dark blue

pink/purple black

pink red

blue blue

blue red pink

red-orange red blue

blue blue blue

the biocompatibility and inherent dissolution properties of the material, a layer of bone may form very quickly on the material surface, preventing osteoclast access and resorption1,12 but, over time, the material will be remodeled. Longitudinal studies in experimental animals report varied results and have shown that 95% of a calcium phosphate implant is resorbed in anything from 26–86 weeks.27 Although the ideal situation would be material degradation matched by bone formation, it is not possible to use the amount of new bone formation in the defect site (bone area) as an indication of implant resorption as they are not always directly linked. For example, when bFGF was added to βTCP in a rat tibial defect, there was increased ceramic resorption compared to the control, but this was not matched by an increase in bone formation in this group.12 In general, and unless otherwise treated with an osteoinductive factor,28 bone forms in a centripedal fashion and ‘grows in’ from the margins of the defect. If the rate of material resorption is faster than that of bone formation, fibrous tissue will fill the space.29 Although macrophages have been seen around biomaterials that either resorb very quickly29 or are implanted in the form of small particles,25 the osteoclast is the main cell involved in biomaterial resorption. Osteoclasts, or multinucleated giant cells, have been identified many times on the surface of biomaterials11,25,30–32 (Fig. 8.2) and Gauthier et al. found an increase in osteoclast number and implant resorption between experimental groups,

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Degradation rate of bioresorbable materials

P

P 8.2 At 3 weeks, multinucleated giant cells (arrows) can be seen on the surface of βTCP particles (P) implanted in a rat tibial defect model. Toluidine blue stained section. (Reproduced with permission and copyright© of the British Editorial Society of Bone and Joint Surgery.30)

which they related to an increase in surface area.25 Therefore, measures of osteoclast number and activity can be indicative of implant resorption. Multinucleated giant cells/osteoclasts express tartrate-resistant acid phosphatase (TRAP), so to accurately count or assess osteoclast number it is preferable to stain for this enzyme directly.17,25,32 Trabecular measures are often not appropriate due to the presence of the implant. At longer time points, when substantial resorption of the material has occurred, trabeculae may reform and some researchers have commented upon, but not quantified, their appearance.11,14,30 The reappearance of trabecular structures depends upon the material and the degree of resorption. Yet even at longer time points they do not always reappear, as reported in a study investigating the resorption of HA-coated proOsteon 200R which noted continued absence of trabeculae at 24 weeks in a rabbit model.13

8.4

Histomorphometric measurements

Histomorphometry is considered to be the gold standard of bone tissue measurement and is routinely used to assess in vivo implant models.17,28,33,34 It is normally carried out on undecalcified bone sections that have been stained to identify the feature of interest, and provides a means to measure particular morphological features that are not normally assessed by bone histology (Table 8.2). Features such as cortical width, cortical porosity, cancellous bone volume, trabecular thickness and number are all assessed using histomorphometry. Histomorphometry of bone can be divided into the measurement of static or dynamic parameters and these can be acquired as primary measurements or derived from primary data. Static primary parameters measure the structural features of bone such as length and area. In evaluating

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bioresorbability, it is useful that some of these measurements can also be applied to the material under test and, in particular, to the bone/biomaterial interface. Dynamic parameters reflect the kinetics of bone, and measure features such as bone and mineral formation rates. Because histomorphometry is the only procedure that can measure bone architecture, modeling/ remodeling and turnover, it is invaluable when assessing bioresorbability and its effect on bone character. There are numerous parameters that can be measured using histomorphometry, many of which are predominantly useful to the clinician in diagnosing metabolic bone disease such as osteoporosis. This section will review only those indices that are likely to be of relevance to the researcher. It is recommended that the American Society for Bone and Mineral Research’s recommended system of nomenclature and methodology is used when assessing bone histomorphometry.35 In the past, bone histomorphometry was carried out manually, using a light microscope and appropriate graticule, but this method has largely been superceded by automatic and semi-automatic computerized image analysis. Commercially available software, such as Osteomeasure (OsteoMetrics, Inc., Decatur, GA. USA) or Bioquant Osteo II (Bioquant Image Analysis Corp., Nashville, TN. USA) have been developed to perform automatic bone histomorphometry, but these systems are more appropriate for the diagnosing clinician than the researcher. The National Institute of Health (NIH) has developed a free image analysis system called ‘Image’ that has proved useful to the researcher in assessing histomorphometry in a semiautomatic mode. Image for the Mac is downloadable free from http://rsb. info.nih.gov/nih-image/ while the PC version ‘Scion Image’ is available free from http://www.scioncorp.com/.

8.4.1 Static indices The effect of a biomaterial on bone, or vice versa, may be measured using static indices. Static indices that are measured directly from the tissue section are known as ‘primary’ indices or measurements. Four types of primary measurement are made – area, length (usually of a perimeter or boundary), distance between points or between lines, and number.35 Table 8.3 lists some commonly used indices; for the purposes of material analysis, this would also include residual material area but there is no standardized abbreviation for this measure. It should be noted that the indices listed in Table 8.3 (and 8.4, see later) are not exhaustive and represent the 2D parameter. With the application of sterology, histomorphometry can also be reported in 3D. However, that is beyond the scope of this section. For further reading on 3D histomorphometry and other indices, see Parfitt.35,36

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Primary index

Abbreviation

Area

Bone area Osteoid area Mineralized area Void area Marrow area Material area Tissue area

B.Ar O.Ar Md.Ar Vd.Ar Ma.Ar

Length

Affinty index* Mineralized perimeter Bone perimeter Osteoid perimeter Eroded perimeter Single labeled perimeter Doubled labeled perimeter Osteoclast perimeter

AI Md.Pm B.Pm O.Pm E.Pm sL.Pm dL.Pm Oc.Pm

Distance

Cortical width Osteoid width Mineralized width Trabecular width Wall width

Ct.Wi O.Wi Md.Wi Tb.Wi W.Wi

Number

Osteoblast number Osteoclast number Osteocyte number

N.Ob N.Oc N.Ot

T.Ar

* The length of bone directly apposed to the implant/total length of bone-implant interface ×100%92.

To make valid comparisons between subjects, common referents must be used. Referents are usually a clearly defined area or perimeter within the section such as tissue area/perimeter, bone area/perimeter or osteoid area/ perimeter.35 For example, to compare mineralized area between subjects, this could be expressed as a percentage of the total tissue area – Md.Ar/T.Ar (see Table 8.3 for definitions). The context of the measurement should determine the referent used and, in terms of material testing, the referent is often tissue area within the original defect margins.17,28 Primary measurements of width such as cortical or trabecular width have value in isolation and do not require a referent to make comparisons between subjects.

8.4.2 Dynamic indices Dynamic indices (Table 8.4) are secondary (derived) indices from primary measurements and reveal the rate of bone formation, mineralization and resorption. Mineral apposition rate and bone formation rate measurements

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Table 8.4 Dynamic measurements used in histomorphometry Dynamic indices

Unit

Formula and definition

Mineral apposition rate (MAR)

µm2/day

Ir.L.Wi/Ir.L.t (Inter-label width/ Inter-label time) Mean distance in µm between midpoints of fluorescent markers divided by number of days elapsed

Bone formation rate (BFR)

µm2/µm/day

MAR*(M.Pm/B.Pm) M.Pm = (dLPm +sLPm/2)/B.Pm Indication of new bone formation per unit of perimeter per day based on double and single label perimeters (other referents can be used)

Formation period (FP)

days

W.Wi/Aj.AR Aj.AR = MAR*(L.Pm*O.Pm) Life span of an osteoid surface

Erosion period (EP)

days

FP*(E.Pm/O.Pm) Life span of an erosion surface

require in vivo fluorochrome labeling at specific time points.37 Because the label incorporates into the bone at the calcification front, the dynamic rates can be calculated from the distance between fluorescing labels and this is a commonly used outcome measure in animal models.17,22,38 Secondary (derived) indices are also used to measure some non-dynamic structural parameters. Trabecular number, for example, is calculated according to the parallel plate model as (B.Pm/T.Pm)/Tb.Wi. (See Table 8.3 for definitions.35) A disadvantage of the histomorphometric approach is that it is destructive and, because of this, when designing experiments it is prudent to keep in mind the philosophy of the 3Rs. Another drawback is that tissue preparation is time-consuming and requires some skill to minimize wastage. Alternative methods of assessment commonly used are based on imaging strategies.

8.5

Imaging

The visualization of healing of bone defects and callus formation in experimental models has traditionally been assessed using plain X-ray radiography. This methodology can also be useful when evaluating the incorporation of a resorbable radio-opaque biomaterial. But problems of standardizing position of the limb, specimen distance, and exposure mean that X-rays can reliably provide only qualitative measurements. Furthermore, plain X-ray

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methods are limited by their low resolution, which excludes the detection of any histomorphometric parameters in a sample. Therefore, in most instances, radiographs are used to subjectively monitor healing of the defect28 or material resorption (described as ‘softening’ of the implant margins)13 at early, non-sacrifice, time points. Peripheral Quantitative Computed Tomography (pQCT) scanning is an X-ray system that is used clinically to assess bone mineral density and cross-sectional parameters of peripheral bones such as the tibia and radius. Tomographic processess allow 3D images to be constructed from a series of scanned sections. The development of research-dedicated pQCT machines has led to an increase in the use of pQCT in the assessment of bone architecture,39 defects,40 healing, and implant resorption41 in small animal models. pQCT has a resolution of 100–1000 μm, giving it a ∼×100 fold improvement over plain radiographs, within a field of view of 5–15 cm. The spatial resolution of pQCT is, however, vastly inferior to that of histomorphometry but it is reported to be reliable42 and, if used properly, can provide some accurate and reproducible histomorphometric data.39 While pQCT can separate cancellous and cortical bone parameters, it cannot reliably evaluate all the trabecular measurements available to histomorphometry, or provide data on dynamic characteristics such as MAR or EP. It also cannot provide information on the cellular aspects of bone, such as osteoblast or osteoclast number. As a methodology, it offers a number of advantages; it is quick, non-destructive, can be used in vivo, and is therefore suitable for longtitudinal studies.39 Micro-CT has a similar role to pQCT in the in vivo research field43–46 but has the advantage of having a much higher resolution. With modern machines delivering resolutions in the <10 μm range, all trabecular histomorphometry parameters can be assessed. Micro-CT has the disadvantage of being able to handle only small specimens, typically <100 mm dia. with corresponding fields of view.

8.6

Summary

The use of animals in research is an emotive subject and research into viable alternatives continues apace. In terms of material testing, it is most likely that animal research will be reduced in future by the use of mathematical and computational modelling and early results in this field have been reported.47 Although research into the production of reliable and predictive models is in its infancy, this area shows great promise. In the meantime, the variability and imprecise nature of some aspects of animal testing is being reduced by the production of international and national guidelines. ISO 10993, which deals with the biological evaluation of medical devices, has

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emerged over the past decade and, more recently, a veterinary specialty group (AOVET), in conjunction with the AO Research institute and the European Academy for the Study of Scientific and Technological Advance, has published the ‘10 golden rules’ for the conduction of animal studies in musculoskeletal research.48 This, in addition to the developments in imaging technology which allow quantitative measures to be obtained without sacrifice and/or destruction of the specimen, will mean that measuring the rate of material resorption in the future will not only become increasingly precise and reliable but more importantly, more humane.

8.7

References

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9 Modelling of the degradation process for bioresorbable polymers D. FARRAR, Smith & Nephew Research Centre, UK

Abstract: This chapter discusses attempts to quantitatively model degradation processes in bioresorbable polymers, focusing on hydrolytic degradation of polymeric biomaterials, particularly the polylactide/ polyglycolide family of polymers. Key processes of degradation are reviewed including diffusion of water, kinetics of polymer hydrolysis, changes in mechanical properties/crystallinity and late stage degradation including mass loss. The chapter then looks at the special case of surface erosion as opposed to bulk degradation. Finally, the modelling of temperature effects is covered, as this relates to the development of accelerated test methods for degradation. Key words: mathematical modelling, kinetics of hydrolysis, effect of temperature on degradation, mechanical properties, mass loss, bulk degradation vs. surface erosion.

9.1

Introduction

There are at least three good reasons to model the degradation of bioresorbable materials. Firstly, it can help in elucidating and understanding the underlying mechanisms of degradation. For example, observations suggest that the degradation of polylactide and polyglycolide is ‘autocatalysed’ by the acid end groups produced by chain scission; but we can really only determine the precise role of these acid groups by first constructing models for the hydrolysis reaction and then fitting experimental data to these. Secondly, modelling can help us to condense data and facilitate comparisons of different materials and test conditions. For instance, a model of hydrolysis can enable us to reduce the data from a study of molecular weight vs. time into one rate constant, which can more easily be compared with data from another study using different materials or temperatures. Similarly, temperature dependence can be reduced to a single ‘activation energy’. Thus, a large number of experimental data points can be condensed and compared. If these two reasons seem rather academic, the third reason for modelling is more practical. Degradation studies, particularly in real time, can be time consuming, taking months or even years for complete mass loss. This can make the development time for new materials or devices very lengthy. With 183

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an appropriate and validated model, however, it is possible to use accelerated test protocols and/or data from short-term tests to predict long-term performance. Given the benefits of modelling, there is a surprising lack of work in this area. Many published studies of degradation characterise and describe changes in properties such as molecular weight, crystallinity, strength and mass loss over time. However, there is rarely any attempt to link these aspects of degradation other than through qualitative models and hypotheses. In this chapter we shall review the literature on modelling of degradation and also suggest some potential approaches to modelling aspects of degradation that are not well covered in the published literature. We shall confine our review to the hydrolytic degradation of polymeric biomaterials, with particular focus on the polylactide/polyglycolide family of materials, since these are the most widely studied and most important from the point of view of use in commercial medical devices.

9.2

Overview of degradation processes for bioresorbable polymers

In the simplest case it is possible to think of the degradation of bioresorbable polymers, such as polylactide and polyglycolide, in terms of four processes: • • • •

Diffusion of water into the polymer Hydrolysis of the polymer chains and reduction in molecular weight Reduction in mechanical properties and changes in other physical properties such as crystallinity and glass transition temperature Production of water-soluble oligomers and mass loss

These processes occur more or less in the order given above, over an increasing timescale, though there is considerable overlap between them. So, for example, diffusion of water throughout the sample may take only hours or days, whereas complete mass loss may take months or even years. These processes, and the relationships between them, are shown in Fig. 9.1. Figure 9.2 shows some different qualitative models for different modes of degradation. These include surface erosion, bulk degradation and bulk degradation with autocatalysis. In surface erosion, degradation takes place primarily at the surface of the polymer and mass is lost gradually from the surface while the centre of the polymer remains relatively intact. This leads to a gradual erosion of the polymer from the surface inwards. In contrast, in bulk degradation, the polymer molecular weight reduces more or less

Modelling of degradation process for bioresorbable polymers Diffusion of water into polymer

Water uptake/ mass increase

Hydrolysis of ester bonds

Reduction in molecular weight

Crystallisation of low molecular weight polymer

Production of acid end groups (autocatalysis)

Production of water-soluble oligomers/ monomer Diffusion of oligomers out of polymer/ loss of acid end groups

185

Increase in crystallinity

Change in physical and mechanical properties

Swelling of polymer by osmosis

Preferential degradation/loss of amorphous regions Mass loss

9.1 ‘Flow diagram’ showing processes involved in hydrolytic degradation and the relationships between them. Shaded boxes refer to the most frequently measured properties. Dotted lines/boxes indicate processes applicable only to semi-crystalline polymers.

evenly throughout the whole of the specimen. Mass loss occurs only in the latter stages of degradation at which point the whole sample has lost its integrity. The third case is bulk degradation with autocatalysis. This is initially similar to the simple bulk degradation case but it is assumed that water-soluble breakdown products cannot escape from the specimen except at the surface and so are trapped in the centre of the material. Since, in this case, it is assumed that these acidic breakdown products catalyse the degradation, degradation proceeds more rapidly in the centre of the sample while a relatively less degraded skin forms on the surface. Eventually this

186

Degradation rate of bioresorbable materials a

b

c

Degradation time

Degradation products

Degree of degradation

9.2 Different modes of degradation: a) Surface erosion, b) bulk degradation, c) bulk degradation with autocatalysis.

ruptures and releases the breakdown products. These different modes of degradation, which will be discussed in more detail later, nevertheless share most of the features described above and shown in Fig. 9.1. In the following sections, we will look at how each of these processes can be modelled and how such models can be used to describe the overall degradation process.

9.3

Modelling of key processes

9.3.1 Diffusion of water Before any hydrolysis can occur, water has first to diffuse into the material. The mathematical description of diffusion is well known and is described in texts such as Crank.1 In one dimension, diffusion is described by Fick’s second law: ∂c ∂ 2c =D 2 ∂t ∂x

[9.1]

where c is the concentration of diffusant (in this case water), x is the space co-ordinate in the direction of diffusion, t is time and D is the diffusion coefficient. For radial diffusion into a cylinder, it can similarly be shown that the diffusion equation is:

Modelling of degradation process for bioresorbable polymers ∂c 1 ∂ ⎛ ∂c = rD ⎞ ⎝ ∂t r ∂r ∂r ⎠

187 [9.2]

where r is the radial co-ordinate. These equations can be solved analytically but the solutions quickly become quite complex for anything other than very simple geometries and boundary conditions. A wide range of standard solutions can be found in texts such as Crank1. The diffusion equations can also be solved using numerical methods. One of the simplest is the finite difference method. Using explicit methods of solution gives the following formula for concentration of diffusant as a function of position and time: c i , j + 1 = c i , j + R (c i − 1, j − 2c i , j + c i + 1, j )

[9.3]

using the notation shown in Fig. 9.3 and where R = Dδt/δx2. Similarly, for a cylinder: R {(2i + 1) c i + 1, j − 4ic i , j + (2i − 1) c i −1, j } i ≠ 0 2i i=0 + 4 R ( c1, j − c0, j )

c i, j +1 = c i, j + = c0, j

[9.4]

where R = Dδt/δr2. For simple geometries in one dimension, solutions can be generated using standard spreadsheet software. However, the most powerful approach to modelling diffusion is Finite Element Analysis (FEA). This requires specialised software packages, but many are commercially available. FEA can be used to model complex geometries and can be linked to Computer Aided Design (CAD) packages. Models can also be developed that can be combined with degradation and/or mechanical properties.

c0,0

c1,0

cN,0 ci,j–1

ci–1,j

ci,j

ci+1,j

ci,j+1

δr

δt

9.3 Notation system used in finite difference modelling.

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Degradation rate of bioresorbable materials

All these modelling approaches require values for the diffusion coefficient of water as part of the input. There are surprisingly few published measurements of diffusion coefficients for bioresorbable polymers, and those that can be found do not give particularly consistent values between different studies.2–5 A selection of published values for the diffusion coefficient of water in the most commonly used bioresorbable polymers is given in Table 9.1. There is need for more work in this area to give more definitive measurements of diffusion coefficients of bioresorbable polymers, as well as an understanding of how changes in diffusion coefficients relate to changes in properties such as chemical composition, crystallinity and molecular weight. So far, we have treated the diffusion of water into a bioresorbable polymer as a simple process following Fickian diffusion laws. It has been assumed that the diffusion coefficient, D, and the equilibrium water uptake or solu-

Table 9.1 Rate of hydrolysis (k2) at 37°C, diffusion coefficient (D) at 37°C and activation energies for diffusion (ED) and hydrolysis (Ek) for some bioresorbable polymers Polymer

k2 (day−1)

D (×10−8cm2s−1)

ED (kJ/mol)

Ek (kJ/mol)

Poly-L-lactic acid (semi-crystalline)

0.0067a 0.0169b 0.0115–0.0288c 0.0052d 0.0018e

8.04f 6.70–9.00g

49–53g

100.5d 101.4h

11.04f 5.10g 50.0g 1.07k

37g

Poly-L-lactic acid (amorphous) Poly-D,L-lactic acid

Poly-D,L-lactide-co-glycolide Poly-L-lactide-co-glycolide

0.0183–0.0447c 0.05i 0.0071j 0.023–0.074a 0.0288–0.0566b 0.029e

Polyglycolic acid

0.0031j 0.0034–0.0050a 0.0018e 0.125p

Polyglyconate B

0.101r

Polycaprolactone

a

62g

115l 114m 100n 96–102.4o

22.9f 48.0g

31g

0.015f 0.5q

Ref 13, b Ref 14, c Ref 15 – range of polymerisation conditions, d Ref 11, e Ref 9, Ref 2, g Ref 3 – 40°C, h Ref 49 – self-reinforced PLLA, i Ref 16, j Ref 17, k Ref 4 – 20% drug loaded, l Ref 50, m Ref 51 – 50 : 50 lactide : glycolide, n Ref 52 – 82 : 18 lactide : glycolide, o Ref 21 – 90 : 10 glycolide : lactide, p Ref 18, q Ref 5, r Ref 19

f

Modelling of degradation process for bioresorbable polymers

189

bility are constant. These assumptions are probably reasonable in the early stages of degradation (the first few days or weeks, depending on the polymer) but are certainly not true once the polymer starts to undergo significant degradation. As we shall see later, for a bulk degrading polymer, it is this initial period of water uptake that is most important. In particular, it is observed experimentally that, rather than levelling off as predicted by simple Fickian models of diffusion, the water uptake of the polymer increases dramatically once the polymer starts to degrade, and it tends to continue to rise throughout the degradation process.6,7 As the polymer absorbs water and degrades, the polymer chains become more mobile and the glass transition temperature, Tg, is lowered; these changes allow water uptake to increase. Furthermore, once water-soluble breakdown products are produced, water uptake is further increased by osmotic effects. Thus, the continuing breakdown of the polymer chains and consequent loss of integrity of the polymer, coupled with the increasing production of hydrophilic chain ends and water-soluble breakdown products, mean that water uptake continues to increase right up to the point where the polymer begins to completely break down. This complex situation for water uptake might lead one to believe that there might be consequent complexity in modelling hydrolysis, since water content of the polymer might be expected to play an important part in determining the rate of the hydrolysis reaction. In fact, this is not the case and, as we shall see, the reaction can be modelled very well assuming ‘pseudo-first-order’ kinetics in which the concentration of water is assumed to be constant. It has been proposed that one explanation for this is the existence of different water environments within the polymer such that only a portion of the water – termed the ‘bound water’ – is reactive; this has been confirmed by experimental studies.6

9.3.2 Kinetics of polymer hydrolysis It is generally accepted that the random chain scission of most polyα-hydroxy acids, such as poly(lactic acid), poly(glycolic acid) and their co-polymers, proceeds according to the reaction: ∼COO∼ + H2O → ∼COOH + HO∼ [9.5] The simplest model for the kinetics of this process assumes that the reaction of the ester bonds is simply proportional to time. If we assume that the number of polymer chains is equal to the number of ∼COOH end groups and that these end groups are the product of the hydrolysis of an ester bond, then the kinetics of the hydrolysis can be described by the following equation:8,9 dcend = k1′. cester . cw dt

[9.6]

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Degradation rate of bioresorbable materials

where cend is the concentration of end groups, cester is the concentration of ester bonds, cw is the concentration of water, k1′ is a rate constant and t is time. For bulk degradation, it is assumed that the concentration of water is even throughout the sample and constant over time. This latter assumption appears questionable given the observations of water uptake described in the previous section; nevertheless, for the reasons discussed there and as we shall see, it appears to hold. It is also assumed that, at least in the early stages of degradation, the concentration of ester bonds is effectively constant. On integrating, Equation 9.6 therefore gives: cend − cend ,0 = k1t

[9.7]

where cend,0 is the concentration of acid end groups at time t = 0 and k1 is a rate constant equal to k1′.cester.cw. If we now use the fact that cend is equal to the reciprocal of number average molecular weight, Mn so that: c end =

1 Mn

[9.8]

then Equation 9.7 becomes 1 1 − = k1t Mn Mn,0

[9.9]

where Mn,0 is the initial number average molecular weight at time t = 0. This equation is usually applied practically by plotting the reciprocal of Mn against time to obtain a linear plot with gradient equal to the rate constant, k1. However, most studies use a modified approach which takes into account catalysis of the hydrolysis reaction by the acid end groups. Since these acid end groups are themselves a product of the degradation reaction, this is often referred to as ‘autocatalysis’. The kinetics of the autocatalytic reaction are conventionally described by the following equation:9 dcend = k2′cester . cw .cend dt

[9.10]

Using the same assumptions as above, that cester and cw are constant and cend = 1/Mn, this becomes: dcend = k2. cend dt

[9.11]

from which it is possible to show that the change in molecular weight over time should progress according to: Mn = Mn,0 e − k2t

[9.12]

Modelling of degradation process for bioresorbable polymers

191

Again, this is usually practically applied to experimental data by plotting ln Mn against time to obtain a linear plot with gradient equal to −k2. It is important to note here that it is the number average molecular weight, Mn, which appears in Equations 9.9 and 9.12. This is because it is Mn that directly links to the concentration of polymer chain ends that are the product of the hydrolysis reaction. Many published studies use weight average (Mw) or viscosity average (Mv) molecular weight in plots of reaction kinetics, which is not strictly correct. In practice, the same linear trends are often seen because Mw is directly proportional to Mn if the form of the molecular weight distribution and polydispersity (equal to Mw/Mn) are constant. This is usually a reasonable assumption, at least in the early stages of degradation. Both models represented by Equations 9.9 and 9.12 have been used to represent the hydrolysis of bioresorbable polyesters. This is usually done by plotting either 1/Mn or ln(Mn) as a function of time. However, only a few papers compare the two models.8–12 It is usually found that the autocatalytic model gives the best fit under normal circumstances and, in the majority of published studies that use a model to analyse degradation results, this is the only one used. The value of such kinetic models and their use to analyse data lies not only in illumination of degradation mechanisms but also in the generation of degradation rate constants that can be used easily and quantitatively to compare different polymers. These rate constants can also, in principle, be used to model and predict the behaviour of other similar polymers without the need for lengthy degradation studies. Some published values for rate constants, k2, obtained at 37°C11,13–19 are summarised in Table 9.1. For these rate constants there is a wide range of values quoted for what are nominally the same polymers in terms of chemical structure. This variability probably reflects not only differences in measurement methodology but also wide differences in properties such as polymer molecular weights, purity, residual monomer levels and crystallinity. Reliable measurements need to be carried out on well-defined materials and further work needs to be done to understand the origin of the variability in rate constants between different studies. The autocatalytic degradation model proposed by Pitt, and represented by Equation 9.12, has been widely applied in many experimental studies. However, exceptions to the model have been reported and this has prompted some workers in the field to consider alternative kinetic models. One debate concerns the role of the acid end groups in catalysing the hydrolysis reaction. When Pitt proposed his model of acid-catalysed hydrolysis, he assumed that the reaction rate was simply proportional to the concentration of carboxylic acid end groups, cend. Thus, he was assuming that the acid group was ‘not ionised in the hydrophobic polymer bulk and functions by hydrogen bonding to the ester links’.9 However, an alternative that has

192

Degradation rate of bioresorbable materials

been proposed for acid catalysis is that the hydrolysis rate is proportional to the concentration of hydrogen ions, [H+], which is itself expected to be related to cend, since:

[ H + ] = (K.cend )1/ 2

[9.13]

where K is the dissociation constant of —COOH end groups. Lyu et al.20 used this assumption to derive an alternative kinetic equation: ⎛ 1⎞ ⎝ N⎠

1/ 2

⎛ 1 ⎞ =⎜ ⎝ N 0 ⎟⎠

1/ 2

1 1/ 2 + k2.5 . ( K .cester ) .cw . t 2

[9.14]

where N is the degree of polymerisation, N0 the degree of polymerisation at time t = 0, and k2.5 is a reaction rate constant. The model was tested against real data by plotting (1/Mn)1/2 vs. time – in which case a linear fit of the data would indicate validity of the model. In fact, Lyu et al. found that neither Equation 9.12 nor the alternative Equation 9.14 gave a simple fit for their data for the degradation of a poly(L-lactide-co-DL-lactide) copolymer. Instead, they proposed a two-stage degradation process with a particular molecular weight at which the rate of hydrolysis increased. Other studies have reported similar effects for poly(glycolide-co-lactide).21 Siparsky et al.22 used a slightly different approach in developing a new equation for autocatalysis, though they again made the distinction between the acid catalyst concentration (equal to (K.cend)1/2) and the acid product concentration, cend. They solved the equation: dcester 1/ 2 = k.cw.cester .(c.end.K ) dt

[9.15]

The solution they generated fitted their experimental data for PLA well, but it should be noted that the degradation experiment was carried out in solution, rather than in the solid state; it was therefore more likely that the acid end groups would be able to dissociate. They also found that the degradation of polycaprolactone was not self-catalysed in these conditions. In practice, whatever model of hydrolysis is used, it is often possible to obtain reasonable linear fits to experimental data, at least qualitatively; this is especially so if one is willing to consider degradation occurring in two or more stages with different rate constants.20 It is therefore important to exercise caution in drawing inferences about degradation mechanisms from limited data. There is still a need for more work to clarify precise mechanisms and models for the hydrolysis of bioresorbable polymers.

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193

9.3.3 Mechanical properties A great number of published studies measure changes in mechanical properties as a function of degradation time. However, only a small number of these studies have attempted to link this to molecular weight and reaction kinetics. Given that changes in mechanical properties are often those that are of interest when considering the performance of bioresorbable medical devices such as sutures, tacks, screws and plates that have at least some load-bearing function, it is surprising that more attempts have not been made to model them. As we have just seen, reasonably good models exist to predict molecular weight (specifically Mn) as a function of degradation time. The question, then, is whether we can link changes in properties such as strength, modulus and crystallinity to changes in molecular weight. In fact, many properties of polymers, including tensile strength, can be related to Mn by an equation of the form shown below,23 first proposed by Flory:

σ = σ∞ −

B Mn

[9.16]

where σ is the tensile strength, σ∞ and B are constants. It is therefore possible to take data from degradation experiments and plot measured tensile strength as a function of 1/Mn to obtain values of the constants σ∞ and B. Furthermore, it is possible to combine Equation 9.16 with an equation for hydrolysis kinetics such as Equation 9.12, to obtain a relationship between tensile strength and time:19

σ (t ) = σ ∞ −

B Mn,0 .e − kt

[9.17]

This approach has been adopted in a number of published studies11,19,21 but has not been successful, generally because Equation 9.16 was not shown to fit the mechanical data well. This is probably because it is not appropriate to apply the Flory equation to tensile strength data other than in the brittle failure regime where it has previously been applied. However, what these studies do show is that there is a definite, unique relationship between tensile strength and number average molecular weight. Studies by Deng et al.21 and Weir et al.11 measured degradation at different temperatures. Yet when tensile strength was plotted as a function of Mn, all data fell on a single curve, irrespective of degradation temperature. In the case of Deng et al., the following empirical relationship between strength and Mn was found by curve fitting: BSR = a + b ln MW

[9.18]

194

Degradation rate of bioresorbable materials

where BSR is the breaking strength retention, MW is the molecular weight (either Mn or Mw), a and b are constants. In contrast, Weir et al. found a linear relationship between strength and Mn. Similarly, Farrar and Gillson19 performed degradation of the same polymer with different initial molecular weight, and again found that there was a unique relationship between strength and Mn which was determined empirically by curve fitting. Clearly, in practice, many factors are going to determine absolute values of tensile strength and it is unreasonable to expect one parameter, such as molecular weight, to describe the polymer completely. However, what these studies show is that, when other parameters such as chemical structure and crystallinity are held constant, then changes in tensile strength can be related to changes in molecular weight surprisingly well, albeit through an empirically determined relationship. Given that one of the main drivers for the modelling of bioresorption is to reduce the amount of time required to characterise degradation, this is still a useful approach. Degradation studies carried out in real time at 37°C are often lengthy, but these studies show that accelerated testing carried out at higher temperatures can be used to predict tensile strength as a function of time, provided we can relate changes in molecular weight at the higher temperature to changes in real time. This is a subject we shall return to in a later section. Furthermore, it is not just changes in tensile strength that can be related to molecular weight. Deng et al. showed unique relationships between molecular weight and modulus and also between molecular weight and strain retention, while Weir et al. found relationships between Mn and mass loss, melting temperature and glass transition temperature. Thus, all these properties can be studied using accelerated test protocols.

9.3.4 Crystallinity The degree of crystallinity can have a significant effect on polymer properties such as modulus, strength, water sorption and degradation kinetics. Changes in crystallinity are frequently measured and reported in published studies of biodegradation of semicrystalline polymers such as polyglycolide, poly(L-lactic acid) and polycaprolactone.24 It is nearly always the case that the degree of crystallinity increases as the polymer degrades and this can be explained by two processes. Firstly, it is assumed that in semi-crystalline polymers the amorphous phase degrades faster than the crystalline. When mass loss begins it is therefore the amorphous component that is lost more readily and so crystallinity increases. However, there is also the possibility of crystallinity increasing by further crystallisation of the polymer due to annealing effects at the test temperature, or as the mobility of the polymer chains increases due to plasticisation by water and/or reduction in molecular weight. Changes in crystallinity of PGA have been investigated in detail

Modelling of degradation process for bioresorbable polymers

195

by Cameron et al.25,26 and these studies show the complexity of changes in crystallinity and microstructure as degradation proceeds. A qualitative model for these changes was proposed. Relatively simple quantitative models for changes in degree of crystallinity during hydrolysis of PET have been developed by Ballara and Verdu.27 These include changes due to both extraction of low molecular weight compounds and ‘chemicrystallisation’ due to chain scission. Such models, however, have not been applied to bioresorbable polymers, though it would certainly be of interest to do so since PET is a polyester like PLA and PGA and its hydrolysis shares many features with those polymers.

9.3.5 Late-stage degradation and mass loss If it is strength retention that largely determines the mechanical performance of a device, it is the rate of mass loss that largely determines the biological response. New tissue cannot start to replace the device until mass loss begins; and the rate of release of degradation products will determine their degree of accumulation in surrounding tissues and the likelihood of any potential inflammatory response as a result. Mass loss is frequently measured and reported in the literature, but to date, little if any attempt has been made to understand and quantitatively model the process by which it occurs. Perhaps this is not so surprising given the complexity of the problem. Many of the approaches used so far in modelling reaction kinetics and mechanical properties assume that degradation is in the early stages before significant mass loss has occurred, so that we can assume the concentration of ester bonds is constant, for example. Also, we are dealing with a distribution of soluble breakdown products including not only monomer but also dimer, trimer and higher oligomers. Each product may be produced at a different rate and have a different rate of diffusion out of the polymer. This is a complex situation that is not only difficult to model mathematically but also potentially requires a lot of input variables whose values need to be determined experimentally. Nevertheless, it is possible to take a simplistic approach to begin to look at distributions of oligomers produced by the degradation of a polymer. The molecular weight distribution resulting from random chain scission can be shown, from a statistical consideration, to be given by the expression:28,29 Nx =α (1 − α )

x−1

[9.19]

where Nx is the number fraction of species with degree of polymerisation x (x-mers) and α is the fraction of bonds that have undergone scission. Similarly it can be shown that the weight fraction of x-mers, Wx, is given by: Wx = xα 2 (1 − α )

x−1

[9.20]

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Degradation rate of bioresorbable materials

This distribution is known as the ‘most probable distribution’ or ‘the Flory most probable distribution’. Note that, in deriving Equation 9.20, it is assumed that the average degree of polymerisation, DP, approximates to 1/α. Lehrle et al. have used this approach to study the degradation (though not the hydrolysis) of poly(hydroxybutyrate).28 Plots of ln Nx and/or ln(Wx / x) vs. (x-1) should be linear if Equations 9.19 and 9.20 are valid. The results showed that random scission did not fully explain the results and instead suggested that scissions near the ends of the chains were favoured. We shall return to this idea for the case of hydrolysis shortly. Figure 9.4 shows how the distribution of oligomers might evolve over time for a bioresorbable polymer. In principle, if we knew the rate of diffusion of each x-mer in the polymer, and their solubility in water, we could model the rate of mass loss of the polymer due to the loss of these oligomers into solution. Work by Park30 suggested that there was a critical weight average molecular weight for water solubility of poly(D,L-lactide) oligomers in the range 1050–1150. This result suggests that, for this polymer, oligomers up to about seven or eight monomer units (7- or 8-mers) are water-soluble. If the water-soluble oligomers are assumed all to diffuse out of the polymer spontaneously (perhaps not so unreasonable an assumption given the lack of integrity and highly hydrated state of most bioresorbable polymers in the late stages of degradation), then the cumulative curve for production of 7- or 8-mers in Fig. 9.4b might be expected to approximately correspond with the mass loss for the polymer. At least qualitatively, it does look like many published mass loss curves, in as much as there is initially an induction period when mass loss is negligible; this is followed by the onset of mass loss, which soon becomes rapid and follows an approximately linear trend.24 Of course, this approach represents a gross oversimplification, since in practice the loss of the water-soluble components will influence the degradation kinetics through processes such as loss of acidic catalysing groups and changes in molecular weight distribution. In particular, the influence of the loss of catalysing degradation products has been demonstrated quite dramatically by Li, Garreau,Vert et al.31–34 who showed that degradation proceeded faster in the centre of large specimens, where degradation products were trapped, than in the outer surface, where they were able to escape. Ultimately, this led to the formation of a highly degraded centre contained within a less degraded shell of polymer. It has been suggested that rupture of this shell can lead to very rapid release of acidic breakdown products into the surrounding tissue. A further complication arises in the case of semicrystalline polymers where, as we have already seen, the amorphous regions of the polymer degrade more rapidly than the crystalline. This complicates the mass loss profile and will certainly cause the distribution of breakdown products to deviate from the simple analysis given above.

Modelling of degradation process for bioresorbable polymers

197

(a)

Weight fraction of x-mers

0.3

0.2

8-mer 7-mer 6-mer 5-mer 4-mer 3-mer 2-mer 1-mer

0.1

0 Time (arbitrary) (b) Cumulative weight fraction of all x-mers

1.0 0.9 0.8 0.7 0.6 0.5

8-mer 7-mer 6-mer 5-mer 4-mer 3-mer 2-mer 1-mer

0.4 0.3 0.2 0.1 0 Time (arbitrary)

9.4 Distribution of oligomers based on ‘most probable’ distribution assuming random chain scission. a) Weight fraction of oligomers. b) Cumulative weight fraction of all oligomers up to the highest indicated.

However, this simple model does give us some insights into features of the mass loss process. For example, in considering the Flory most probable distribution, the approximation was made that α = 1/DP = m/Mn, where m is the molar mass of one monomer unit. From this we can see that the distribution of low molecular weight breakdown products should depend only on Mn at a given time. Hence, the onset and progression of mass loss should depend only on Mn, independent of the initial molecular

198

Degradation rate of bioresorbable materials

weight and degradation conditions. This kind of relationship has been confirmed in at least one study,11 as we saw earlier. As a first approximation, the situation can be somewhat simplified by considering only the production of monomer. In studies of the degradation of PLA and PLGA copolymers, Hakkarainen et al.14 found that at the time of total mass loss, about 50% of the polymers had been hydrolysed to monomer. Therefore, we need to consider the rate of production of monomer vs. other oligomers. So far we have assumed the hydrolysis of these polymers to be a random chain scission process with main chain ester bonds being equally likely to be broken as bonds near the end of the chain. In fact, there is evidence that this is not the case for many bioresorbable polymers, including the lactide/glycolide family. Shih35,36 was one of the first to study random vs. chain end scission in bioresorbable polymers. In this work Shih also proposed models of the two mechanisms and a graphical method to distinguish the two modes of degradation. This method involves measuring and plotting the amount of monomer produced, m1 /C0 (where m1 is the concentration of monomer and C0 is the initial concentration of backbone bonds) vs. the degree of degradation, α. For random scission Shih showed that: m1 / C0 = α 2

[9.21]

whereas for the case of pure ‘unzipping’ or chain end scission: m1 / C0 = α

[9.22]

Shih analysed the hydrolysis of PCL and PLA in solution under acidic conditions and concluded that the degradation of PCL proceeded purely by random scission kinetics but, in the case of PLA, chain-end scission had a rate approximately 10-fold faster than random scission of internal bonds. Shih also analysed the hydrolysis of ‘neat’ PLA in sheet form in a constant humidity chamber (i.e. not in solution) and again concluded that unzipping was faster than random scission. In a very sophisticated mathematical model of the degradation of bioresorbable polymer microspheres, Batycky et al.37 constructed a model that took into account both random scission and chain-end scission. They modelled both molecular weight evolution and mass loss, as well as drug release, and compared the model to experimental data for PLGA microspheres. The results confirmed that a good fit was only possible by assuming both chainend and random scission to take place, with end-chain scission having a rate constant higher than random scission by a factor of approximately 10 000. Other studies38–40 have also confirmed the tendency of polylactide to degrade by unzipping from the chain ends and have proposed some possible chemical mechanisms and kinetic models. The impact of chain end scission on hydrolysis kinetics is probably negligible in the early stages of degradation when the concentration of chain ends is very low. In this early situation,

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199

random scission dominates, in turn generating more chain ends from which unzipping can take place. Then, in the later stages of degradation, chain end scission can become an important process and its effect will be to accelerate the production of soluble monomers and hence mass loss. Before leaving the subject of late-stage degradation and mass loss, it should also be pointed out that the discussion so far has focused on in vitro degradation, ignoring the effect of enzymes. This is generally regarded as being reasonable early on in the degradation process but it has been suggested that, in the later stages of hydrolysis, enzymatic degradation may be important, especially if small fragments of polymer undergo phagocytosis.41 This needs to be taken into account when considering the relevance of in vitro testing to the in vivo or clinical reality.

9.4

Modelling of surface erosion

9.4.1 Bulk degradation vs. surface erosion So far we have confined our analysis to bulk degradation, in which it is assumed that the polymer degrades more or less evenly throughout its volume. For example, we assumed throughout our analysis of reaction kinetics that the concentration of water was equal throughout the polymer. We shall now look a the case where this assumption is no longer true and degradation kinetics are fast, relative to water diffusion. In this case, diffusion of water can become the rate-limiting step and the degradation mechanism can change to one of surface degradation or erosion. A theoretical model was developed by von Burkersroda et al.42 that allowed prediction of whether a polymer would degrade by bulk or surface degradation. This model compared the time frames for polymer chain degradation vs. water diffusion to obtain a dimensionless number, ε, termed the ‘erosion number’, given by the expression: x λπ ln [ x ] − ln ⎡⎣ 3 Mn / NA ( N − 1) ρ ⎤⎦ 2

ε=

4 Deff

(

)

[9.23]

where Deff is the effective diffusivity of water into the polymer, 〈x〉 is the device dimension, λ is the degradation rate of the polymer, Mn is the number average molecular weight, N is the average degree of polymerisation, NA is Avogadro’s number and ρ is the polymer density. For high values of ε (>>1), the reaction with water is fast compared to diffusion, and surface erosion dominates. On the other hand, for low values (ε << 1) diffusion is faster than hydrolysis, and bulk degradation takes place. For ε = 1, the mechanism is equivocal and a critical polymer dimension was defined, Lcrit, where this condition is met. Devices larger than Lcrit thus undergo surface erosion

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Degradation rate of bioresorbable materials

whereas those smaller than Lcrit undergo bulk degradation. This model was used to explain why polymers such as polylactides tend to bulk degrade at normal dimensions while polyanhydrides tend to surface erode. An alternative analysis was undertaken by Lyu et al.43, who developed a mathematical model combining the simultaneous diffusion and reaction of water with a biodegradable polymer. In this model they derived an expression for a critical specimen thickness, Wc (similar to Lcrit above), below which the sample undergoes bulk degradation and above which its surface erodes. This was given by: 1 ⎞ WC = ⎛ ⎝ 1+ C′ ⎠

1/ 2

(D / k )1 / 2

[9.24]

where D is the diffusion coefficient of water, k is the degradation rate and C′ = CS /Cm where CS is the solubility of water in the polymer and Cm is the total amount of water consumed. A very similar result was derived by the author44 by using a finite difference model for diffusion of water, as given by Equation 9.4, to calculate the concentration of water, and by combining this with a reaction kinetics model, as given by Equation 9.10, to give molecular weight within a cylindrical specimen as a function of time and position. From these results it was shown that the cylinder had a critical radius for bulk vs. surface degradation, analogous to Lcrit and WC, given by (D/k)1/2. From these models it can be seen that whether a polymer undergoes bulk or surface degradation depends on the ratio of degradation rate to water diffusion coefficient. However, sample size is also important and, potentially, any polymer can undergo either mechanism depending on its size.

9.4.2 Surface erosion There have been several attempts to model surface erosion of polymers using computer simulation techniques. Gopferich et al.45–47 have used MonteCarlo methods to model this process, setting up a two-dimensional grid where each pixel represents a small part of the polymer and represents either a slow-eroding (crystalline) or fast eroding (amorphous) unit. Pixels could erode if they were in contact with an eroded neighbour, the probability of erosion occurring following a specified distribution, with different rates for fast- and slow-eroding pixels. The model was compared to experimental data for polyanhydrides and it was shown that it could successfully mirror the movement of erosion fronts, mass remaining and porosity for these polymers. In a similar approach, Zygourakis and Markenscoff48 developed a ‘cellular automata’ model of erosion and used this to study drug release from a matrix.

Modelling of degradation process for bioresorbable polymers

9.5

201

Temperature effects

One of the most practical applications of modelling is its use in analysing the effect of temperature on hydrolysis, particularly the use of elevated temperature as a means of accelerating degradation studies. Most physical and chemical processes are accelerated at higher temperatures and these include both the diffusion of water into the polymer and the subsequent hydrolysis of the polymer chains. The most frequently used model for the effect of temperature on both these processes is to assume that they follow an Arrhenius type behaviour. For diffusion this means that the diffusion coefficient, D, has a temperature dependence given by: D = D0 e − ED / RT

[9.25]

where ED is an activation energy, R is the universal gas constant, T is the temperature (in Kelvin) and D0 is a constant. The key parameter here is the activation energy, ED, since if we know this and the diffusion coefficient at one reference temperature, we can work out the diffusion coefficient at any other temperature. Diffusion coefficients for water into a range of bioresorbable polymers at a range of temperatures are given by Siparsky et al.3 who also gives the activation energies for water diffusion in these polymers. Values tend to fall in the range 30–65 kJ/mol and a selection of values is included in Table 9.1. A similar approach can be adopted for hydrolysis reaction kinetics and in this case the Arrhenius equation gives: k = k0 e − Ek / RT

[9.26]

where Ek is the activation energy for the hydrolysis reaction and k0 is a constant. Again, the key parameter that determines temperature dependence is the activation energy, Ek, and some values from the literature for common bioresorbable polymers are summarised in Table 9.1.11,21,49–52 It will be noted that the activation energies for hydrolysis fall within the range 96–115 kJ/mol. This is significantly higher than the activation energy range for diffusion of water which means that raising the temperature has a greater impact on accelerating hydrolysis than diffusion. Following our discussion of surface vs. bulk degradation, we therefore need to make sure that in using elevated temperature testing we do not move from a bulk to a surface degradation mechanism. By combining reaction kinetics models (such as Equations 9.9 or 9.12) with Equation 9.26 it is possible to show that the relationship between the degradation rates at two different temperatures is given by:52 ⎛ k⎛ ⎞⎞ ⎜ ⎜ − ⎟⎟ kT 2 = e⎝ R ⎝ T1 T2 ⎠ ⎠ kT 1 E

1

1

[9.27]

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Degradation rate of bioresorbable materials

where kT1 and kT2 are the reaction rates at temperatures T1 and T2. Using the range of values of Ek given above, it can be seen that every increase in temperature of 10°C will increase the degradation rate by a factor of roughly three to four, for typical temperatures used in degradation testing. While most studies find that the Arrhenius equation works well for the temperature ranges studied, there is some debate in the literature about testing above and below the glass transition temperature, Tg. Reed and Gilding53 reported that the effect of raising temperature on the degradation rate of PGA was greater above Tg than below it. Agrawal et al.51, studying the degradation of PLGA over a range of temperatures, found that activation energies were distinctly different for temperatures above and below Tg. They concluded that tests carried out at T > Tg should not be used to predict degradation behaviour at T < Tg. A similar observation was made by Lyu et al.20 for polylactide. They proposed an alternative temperature dependence equation for tests spanning the range above and below Tg. This was the Vogel–Tammann–Fulcher (VTF) equation: k = k0 e −(Ek / R(T −TS ))

[9.28]

where TS is a reference temperature (found to be 0°C for PLA). Lyu et al. found that this modified equation could be used not only to predict molecular weight change at 37°C based on higher temperature testing, but also mass loss and water uptake. In fact, they predicted that total mass loss for their PLA polymer would take 3–5 years at 37°C but this point was reached in just 2–3 days at 90°C. This shows the power of using elevated temperature, together with a suitable model, to accelerate degradation testing.

9.6

Future trends

As we have seen, the modelling of degradation of bioresorbable polymers is at a relatively early stage of development, particularly for anything other than the kinetics of the hydrolysis reaction. We can certainly expect this field of research to develop in the future, with more work being needed to fully understand and model all aspects of degradation. We have already suggested some approaches to modelling the less well-studied properties, such as changes in strength, crystallinity and mass loss. Most of the models described have been relatively simple and have been analytical in nature. In the future we will undoubtedly see the application of more advanced methods, such as computer simulation and modelling. For example, Karst and Yang54 recently used a molecular modelling and simulation approach to study the hydrolysis of blends of PLLA and PDLA. A similar approach has been used by Zhang et al.55 to predict the glass transition temperature

Modelling of degradation process for bioresorbable polymers

203

of polylactide as a function of molecular weight. We have already seen the use of cellular automata or Monte-Carlo techniques to model surface erosion of polymers, and mathematical simulations such as these may also be used to study hydrolysis reactions, particularly late-stage degradation and mass loss (see, for example, Staggs56,57 who has used analytical methods and Monte-Carlo simulations to model random and chain-end scission in thermal degradation). Such techniques will have an increasing role to play in the future.

9.7

Conclusion

We have seen how modelling the hydrolysis of bioresorbable polymers can help us to understand detailed mechanisms of the degradation process, to condense and rationalise experimental data and to shorten the development time for new materials. However, much more work is needed in this area and there is huge potential to apply more advanced modelling techniques and to carry out detailed experimental studies to test these models. If successful, the rewards will be a more complete understanding of bioresorption as well as the more efficient, rational and faster design and development of improved medical devices.

9.8

References

1 crank j, The Mathematics of Diffusion, Oxford, Oxford University Press, 1975. 2 yoon j-s, jung h-w, kim m-n, park e-s, ‘Diffusion coefficient and equilibrium solubility of water molecules in biodegradable polymers’, J App Poly Sci, 2000 77 1716–1722. 3 siparsky g l, voorhees k j, dorgan j r, schilling k, ‘Water transport in polylactic acid (PLA), PLA/polycaprolactone copolymers, and PLA/polyethylene glycol blends’, J Environ Polym Degrad, 1997 5 125–136. 4 smith r w, booth j, massingham g, clough a s, ‘A study of water diffusion, in both radial and axial directions, into biodegradable monolithic depots using ion beam analysis’, Polymer, 2004 45 4893–4908. 5 zaikov g e, ‘Quantitative aspects of polymer degradation in the living body’, JMS-Rev Macromol Chem Phys, 1985 C25 551–597. 6 schmitt e a, flanagan d r, linhardt r j, ‘Importance of distinct water environments in the hydrolysis of poly(DL-lactide-co-glycolide)’, Macromolecules, 1994 27 743–748. 7 wiggins j s, hassan m k, mauritz k a, storey r f, ‘Hydrolytic degradation of poly(D,L-lactide) as a function of end group: Carboxylic acid vs. hydroxyl’, Polymer, 2006 47 1960–1969. 8 helder j, dijkstra p j, feijen j, ‘In vitro degradation of glycine/DL-lactic acid copolymers’, J Biomed Mater Res, 1990 24 1005–1020.

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9 pitt c g and gu z, ‘Modification of the rates of chain cleavage of poly(εcaprolactone) and related polyesters in the solid state’, J Control Release, 1987 4 283–292. 10 weir n a, buchanan f j, orr j f, dickson g r, ‘Degradation of poly-L-lactide. Part 1: In-vitro and in-vivo physiological temperature degradation’, Proc Instn Mech Engrs Part H: J Engineering in Medicine, 2004 218 307–319. 11 weir n a, buchanan f j, orr j f, farrar d f and dickson g r, ‘Degradation of poly-L-lactide. Part 2: Increased temperature accelerated degradation’, Proc Instn Mech Engrs Part H: J Engineering in Medicine, 2004 218 321–330. 12 anderson j m, ‘Perspectives on the in vivo responses of biodegradable polymers’, in Hollinger J O, Biomedical Applications of Synthetic Biodegradable Polymers, Boca Raton, CRC Press, 1995. 13 cha y, pitt c g, ‘The biodegradability of polyester blends’, Biomaterials, 1990 11 108–112. 14 hakkarainen m, albertsson a, karlson s, ‘Weight losses and molecular weight changes correlated with the evolution of hydroxyacids in simulated in vivo degradation of homo- and copolymers of PLA and PGA’, Polym Degrad Stab, 1996 52 283–91. 15 zhang x, wyss u p, pichora d, ‘An investigation of poly(lactic acid) degradation’, J Bioact Compat Polym, 1994 9 80–100. 16 helder j, dijkstra p j, feijen j, ‘In vitro degradation of glycine/DL-lactic acid copolymers’, J Biomed Mater Res, 1990 24 1005–1020. 17 pego a p, poot a a, grijpma d w, ‘In vitro degradation of trimethylene carbonate based (co)polymers’, Macromol Biosci, 2002 2 411–419. 18 hardwick s, A study of the degradation of poly(glycolic acid), Smith & Nephew Internal Research Report, 2000. 19 farrar d f, gillson r k, ‘Hydrolytic degradation of polyglyconate B: The relationship between degradation time, strength and molecular weight’, Biomaterials, 2002 23 3905–3912. 20 lyu s, schley j, loy b, lind d, hobot c, sparer r, untereker d, ‘Kinetics and time–temperature equivalence of polymer degradation’, Biomacromolecules, 2007 8 2301–2310. 21 deng m, zhou j, chen g, burkley d, xu y, jamiolkowski d, barbolt t, ‘Effect of load and temperature on in vitro degradation of poly(glycolide-co-L-lactide) multifilament braids’, Biomaterials, 2005 26 4327–4336. 22 siparsky g l, voorhees k j, miao f, ‘Hydrolysis of polylactic acid (PLA) and polycaprolactone (PCL) in aqueous acetonitrile solutions: Autocatalysis’, J Environ Polym Deg, 1998 6 31–41. 23 ward i m, Mechanical Properties of Solid Polymers, Chichester, Wiley, 1983. 24 shawe s, buchanan f, harkin-jones e, farrar d, ‘A study on the rate of degradation of the Bioabsorbable polymer polyglycolic acid (PGA)’, J Mater Sci, 2006 41 4832–4838. 25 king e, cameron r e, ‘Effect of hydrolytic degradation on the microstructure of poly(glycolic acid): An x-ray scattering and ultraviolet spectrophotometry study of wet samples’, J App Polym Sci, 1997 66 1681–1690. 26 king e, robinson s, cameron r e, ‘Effect of hydrolytic degradation on the microstructure of quenched, amorphous poly(glycolic acid): An X-ray scattering study of hydrated samples’, Polym Int, 1999 48 915–920.

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27 ballara a, verdu j, ‘Physical aspects of the hydrolysis of polyethylene terephthalate’, Polym Degrad Stab, 1989 26 361–374. 28 lehrle r, williams r, french c, hammond t, ‘Thermolysis and methanolysis of poly(β-hydroxybutyrate): Random scission assessed by statistical analysis of molecular weight distributions’, Macromolecules, 1995 28 4408–4414. 29 lehrle r s, pattenden c s, ‘Effect of gamma irradiation on the MW distribution of polyisobutylene: Use of statistical logarithmic plots to assess GPC results’, Polym Degrad Stab, 1998 61 309–318. 30 park t g, ‘Degradation of poly(D,L-lactic acid) microspheres: Effect of molecular weight’, J Control Release, 1994 30 161–173. 31 li s m, garreau h, vert m, ‘Structure–property relationships in the case of the degradation of massive aliphatic poly-(α-hydroxy acids) in aqueous media, Part 1: Poly(DL-lactic acid)’, J Mat Sci: Materials in Medicine, 1990 1 123–130. 32 li s m, garreau h, vert m, ‘Structure–property relationships in the case of the degradation of massive aliphatic poly-(α-hydroxy acids) in aqueous media, Part 3: Influence of the morphology of poly(L-lactic acid)’, J Mat Sci: Materials in Medicine, 1990 1 198–206. 33 grizzi i, garreau h, li s, vert m, ‘Hydrolytic degradation of devices based on poly(DL-lactic acid) size-dependence’, Biomaterials, 1995 16 305–311. 34 li s, ‘Hydrolytic degradation characteristics of aliphatic polyesters derived from lactic and glycolic acids’, J Biomed Mater Res, 1999 48 342–353. 35 shih c, ‘A graphical method for the determination of the mode of hydrolysis of biodegradable polymers’, Pharmaceutical Res, 1995 12 2036–2040. 36 shih c, ‘Chain-end scission in acid catalysed hydrolysis of poly(D,L-lactide) in solution’, J Control Release, 1995 34 9–15. 37 batycky r p, hanes j, langer r, edwards d a, ‘A theoretical model of erosion and macromolecular drug release from biodegrading microspheres’, J Pharmaceutical Sci, 1997 86 1464–1477. 38 de jong s j, arias e r, rijkers d t s, van nostrum c f, kettenes-van den bosch j j, hennink w e, ‘New insights into the hydrolytic degradation of poly(lactic acid): Participation of the alcohol terminus’, Polymer, 2001 42 2795–2802. 39 van nostrum c f, veldhuis t f j, bos g w, hennink w e, ‘Hydrolytic degradation of oligo(lactic acid): A kinetic and mechanistic study’, Polymer, 2004 45 6779–6787. 40 belbella a, vauthier c, fessi h, devissaguet j-p, puisieux f, ‘In vitro degradation of nanospheres from poly(D,L-lactides) of different molecular weights and polydispersities’, Int J Pharmaceutics, 1996 129 95–102. 41 pitt c g, ‘Poly-ε-caprolactone and its copolymers’, in Chasin M, Langer R, Biodegradable Polymers as Drug Delivery Systems, New York, Marcel Dekker, 1990. 42 von burkersroda f, schedl l, gopferich a, ‘Why degradable polymers undergo surface erosion or bulk erosion’, Biomaterials, 2002 23 4221–4231. 43 lyu s, sparer r, untereker d, ‘Analytical solutions to mathematical models of the surface and bulk erosion of solid polymers’, J Polym Sci: Part B: Polymer Physics, 2005 43 383–397. 44 farrar d f, ‘Modelling the hydrolytic degradation of bioresorbable polymers’, Second International Conference on New Biomedical Materials, Cardiff, 5–8 April 2003.

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45 gopferich a, langer r, ‘Modeling of polymer erosion’, Macromolecules, 1993 26 4105–4112. 46 gopferich a, ‘Mechanisms of polymer degradation and erosion’, Biomaterials, 1996 17 103–114. 47 gopferich a, tessmar j, ‘Polyanhydride degradation and erosion’, Adv Drug Del Rev, 2002 54 911–931. 48 zygourakis k, markenscoff p a, ‘Computer-aided design of bioerodable devices with optimal release characteristics: A cellular automata approach’, Biomaterials, 1996 17 125–135. 49 pohjonen t, tormala p, ‘Hydrolytic degradation of ultra-high-strength selfreinforced poly-L-lactide. A temperature dependence study’, in Leung K S, Hung L K, Leung P C, ‘Biodegradable Implants in Fracture Fixation’, World Scientific Publishing, 1994. 50 buchholz b, ‘Accelerated degradation test on resorbable polymers’, in Plank H, Dauner M, Renardy M, ‘Degradation Phenomena on Polymeric Biomaterials’, Berlin, Springer-Verlag, 1992. 51 agrawal c m, huang d, schmitz j p, athanasiou k a, ‘Elevated temperature degradation of a 50 : 50 copolymer of PLA-PGA’, Tissue Engineering, 1997 3 345–352. 52 pietrzak w s, kumar m, eppley b l, ‘The influence of temperature on the degradation rate of LactoSorb copolymer’, J Craniofacial Surgery, 2003 14 176–183. 53 reed a m, gilding d k, ‘Biodegradable polymers for use in surgery – poly(glycolic)/ poly(lactic acid) homo and copolymers: 2. In vitro degradation’, Polymer, 1981 22 494–498. 54 karst d, yang y, ‘Molecular modelling of the resistance of PLA to hydrolysis based on the blending of PLLA and PDLA’, Polymer, 2006 47 4845–4850. 55 zhang j, liang y, yan j, lou j, ‘Study of the molecular weight dependence of glass transition temperature for amorphous poly(L-lactide) by molecular dynamics simulation’, Polymer, 2007 48 4900–4905. 56 staggs j e j, ‘Modelling random scission of linear polymers’, Polym Degrad Stab, 2002 76 37–44. 57 staggs j e j, ‘Modelling end-chain scission and recombination of linear polymers’, Polym Degrad Stab, 2004 85 759–767.

10 Influence of processing, sterilisation and storage on bioresorbability F. BUCHANAN and D. LEONARD, Queen’s University Belfast, Northern Ireland

Abstract: This chapter will focus on the influence of processing, sterilisation and subsequent storage on polymeric bioresorbable materials. Areas covered will include melt-processing (moulding, extrusion) and the main sterilisation and packaging techniques. This will incorporate factors such as heat, atmospheric oxygen, moisture, impurities, processing aids, additives and high-energy radiation on the structural and chemical stability of a bioresorbable polymer. Both synthetic polymers (e.g. polyesters) and naturally derived polymers such as starch, collagen and chitosan will be considered. Key words: injection moulding, extrusion, sterilisation, packaging, additives.

10.1

Introduction

This chapter will focus predominantly on polymeric bioresorbable materials, the rationale for this being that polymeric biomaterials are more likely to be influenced by the conditions encountered during processing, sterilisation and subsequent storage. For example heat, atmospheric oxygen, moisture, impurities, processing aids, additives and high-energy radiation processes all may influence the structural and chemical stability of a bioresorbable polymer. This applies to both synthetic polymers (e.g. polyesters) and naturally derived polymers such as starch, collagen and chitosan. Any processing or sterilisation procedure is likely to have a measurable effect on a bioresorbable polymer. It is therefore important that manufacturers are aware of any produced effects in the polymer and that these effects are understood and monitored or, if possible, predicted and planned for. For example, the use of gamma radiation to sterilise a bioresorbable polymer for medical application inevitably leads to a change in the molecular weight of the material, potentially through chain scission (reduction in molecular weight) or cross-linking (increase in molecular weight). This, in turn, influences the mechanical properties and bioresorption behaviour of the polymer. However, as long as this effect is known and accounted for during design, its consistency and predictability means that gamma radiation remains a viable method of sterilisation for bioresorbable polymers. 209

210

Degradation rate of bioresorbable materials

All medical devices will inevitably require shelf storage in a hospital or distribution centre for a significant time prior to use. The ‘shelf-life’ of the device needs to be sufficiently long to allow storage for practical time periods without the device’s expiration date being exceeded. This may involve the requirement for inert packaging of the device, or even storage at low temperature. To summarise, for successful exploitation of a bioresorbable medical device, the influence of processing and sterilisation must be consistent and predictable and it must be stable within its packaging for a realistic time period during shelf storage.

10.2

Processing techniques

Bioresorbable polymers for the majority of orthopaedic applications (e.g. screws, pins, plates, suture anchors) will generally be finally shaped by a melt-processing technique. These techniques can be grouped into three distinct categories. The following sections describe these categories.

10.2.1 Extrusion The significance of extrusion to polymer processing is such that it can be said that almost all thermoplastics, including those that are bioresorbable, are processed by extrusion at some stage of their manufacture.1 At its most basic, an extruder consists of a screw within a heated barrel. Polymer is fed from a hopper into one end of the barrel. The screw then conveys the polymer along the barrel where it is heated through conduction from the barrel itself and through shearing from the mechanical action of the screw. This heating acts to melt the polymer. Tapering of the screw thread means that the depth of the channel between the screw and the barrel is reduced as the polymer moves along the barrel. This acts to compact the polymer melt before it is forced through a shaped die at the end of the barrel. The polymer is then cooled with water or air to form a solid with a constant cross-section that closely resembles that created by the die. Variables that can be modified to control the extrusion process include screw design, barrel temperature, screw speed and cooling parameters. It is also common to use extruders featuring two or more screws when a higher level of blending or compounding is required.2–4 Extruders can also be used to produce polymer films, through either casting or film blowing. The former involves using a long thin die to extrude a film of polymer onto a cooling roller. The latter involves extruding a tube of molten polymer into which air is blown to expand the tube to the desired size before it is cooled, slit and flattened.1

Influence of processing, sterilisation and storage

211

10.2.2 Moulding A variety of polymer processing techniques fall into the category of moulding. They include injection moulding, thermoforming, rotational moulding, compression moulding and transfer moulding. The most significant of these is injection moulding. The fundamental process of injection moulding involves injection a polymer melt under pressure into a shaped mould cavity. In practice, the first part of an injection moulding machine closely resembles an extruder, as described in the previous section. However, in injection moulding the extrusion screw can be reciprocated forwards and backwards along the barrel. Also, where in extrusion the polymer melt is forced through a die, in injection moulding it is forced into a shaped and finite cooled mould. Injection moulding is a cyclic procedure that can be described as follows. At the beginning of the cycle, the screw is stationary in the forward position. In this position the screws is closing the gate to the mould. When the polymer in the mould from the previous cycle has solidified, the screw begins to turn, drawing more material into the barrel. As the gate to the mould is blocked, this has the effect of pushing the screw back to the rear position. This creates a volume of molten material between the gate of the mould and the tip of the screw. The extent of this volume can determine how much polymer is injected into the mould, or the ‘shot size’. When the desired volume has been obtained, the screw stops rotating. During this stage, the mould is opened allowing the previously formed material to be ejected. The mould is then closed again and the screw is pushed forward, injecting the polymer melt into the mould. The reduced temperature of the mould in comparison to the barrel causes the polymer to solidify. During solidification the screw is used to maintain pressure within the mould. At this stage the cycle is complete and a new one may begin. Factors used to control injection moulding include the barrel temperature, the shot size, the mould temperature and the injection pressure.2–4 Compression moulding and transfer moulding are more simplistic forms of injection moulding. In compression moulding a shot of solid polymer is placed within a mould, which is then heated and pressurised to form and shape the polymer melt. The mould is then cooled to solidify the polymer. Transfer moulding is similar, with the exception that the polymer is melted outside the mould before being introduced with a simple plunger driven system. In rotational moulding, polymer is placed in a mould, normally much greater in volume than the amount of polymer. The mould is then heated while being rotated. This causes the polymer to melt and coat the inside wall of the mould. The mould is then cooled, solidifying the polymer and causing the polymer to form a hollow shape the replicates the inside of the mould.

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Degradation rate of bioresorbable materials

Thermoforming involves shaping a heated sheet of polymer, normally produced through extrusion, with a mould. This can involve using a vacuum to draw the polymer into the mould, using pressure to force the polymer onto a mould or mechanically forcing the mould into the polymer.2–4

10.2.3 Fibre spinning Fibre-spinning is less commonly used than the aforementioned processing techniques; however it has been applied to bioresorbable polymers. The process involves drawing fine fibres of either melted or dissolved polymer into arrangements using gravity, pressure, mechanical force or electrical charge. This can be used to produce polymer fabrics or mesh-like scaffolds.

10.3

Processing-related degradation

The processing techniques outlined in Section 10.2 often involve the application of heat and mechanical shearing to the bioresorbable polymer involved. Unfortunately, both of these processes can lead to degradation of the polymer.5 A study by Weir et al.6 revealed that the molecular weight of PLLA decreased by 10% following melt extrusion at 200°C. It is therefore important to understand the susceptibility of the bioresorbable polymer to these forms of degradation before the material is processed. Only then is it possible to modify the process so as to reduce the amount of degradation which will occur. The following sections examine how thermal and mechanical degradation affect bioresorbable polymers.

10.3.1 Thermal degradation The thermal degradation of polymeric materials has been the subject of extensive study and reporting, due to its perceived importance. The book entitled Thermal Degradation of Polymeric Materials7 states that the thermal degradation of a polymer generally consists of depolymerisation or polymer-chain fragmentation. These processes are manifested as reductions in the mass and molecular weight of the polymer, which in turn lead to changes in other material properties normally associated with degradation. The chemical reactions involved normally proceed through three major pathways: •

Side-Group Elimination: In this case a side group is removed from the polymer leaving the polymer unstable and prone to further degradation reactions. • Random Scission: Similarly to hydrolytic initiated chain scission, in this process a free-radical forms at some point on the polymer backbone splitting the polymer into lower molecular weight material.

Influence of processing, sterilisation and storage •

213

Depolymerisation: In this case a free-radical mechanism acts to reverse the polymerisation process which formed the polymer, producing the constituent monomers or dimers.

The response of a polymer to the continual addition of thermal energy causes certain transitions in properties that, although occurring at widely varying temperatures in different polymers, are observed in almost all cases. The book Plastics: Materials and Processing3 lists these transitions as follows: •

Heat Distortion Temperature (HDT): At very low temperatures a polymer is a rigid solid due to the restricted movement of its component molecules. Initially, heating allows the molecules to move slightly in the form of vibration, rotation and twisting, but this produces little change beyond a slight thermal expansion. Eventually, a temperature is reached at which the polymer becomes more easily deflected by loading. It is also susceptible to creep-based distortion. This is the HDT. • Glass Transition Temperature (Tg): With continued heating from the HDT, the molecules of the polymer gain sufficient energy such that groups of them can be moved transitionally with respect to the other molecules. This leads to a large increase in the flexibility of the polymer. The temperature at which this occurs is known as the Tg. Unlike at the HDT, the transition that occurs at the Tg requires energy to be absorbed by the molecules and, as such, during this transition the applied thermal energy will not result in a rise in temperature. • Melting Temperature (Tm): With the continued addition of thermal energy beyond the Tg of a polymer, the component molecules begin to be capable of transitional movement on an individual basis. This allows the material to flow. This is the process of melting, and results in the polymer becoming a liquid. As with the Tg, the transition at the Tm requires the absorption of thermal energy into the molecules and therefore no temperature increase is produced during this time. • Decomposition Temperature (Td): When additional heat energy is supplied to the melted polymer, it produces increased movement in the component molecules in the form of vibration, rotation and twisting. Eventually, these motions become so energetic that the covalent bonds within the molecule can be broken. This process results in degradation of the polymer, often manifested as a reduction in molecular weight, the emission of gases or the production of new polymer species. The nature of the polymer being heated has a significant influence over these transitions. A completely crystalline polymer will not exhibit a Tg because the crystalline structure continues to restrict transitional movement of the molecules. Polymers that are ‘thermosetting’ have (after ‘curing’) a Td that is lower than their Tm and so will decompose before melting.

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Degradation rate of bioresorbable materials

However, it is always important to understand the various transitions that occur in a polymer when using it, as unexpected thermal transitions can severely compromise the efficacy of a polymeric product. The HDT of a polymer is normally determined by simple mechanical testing at varying temperatures. The Tg and Tm can be determined by Differential Scanning Calorimetry (DSC), which is based on monitoring the corresponding rise in temperature to applied energy. In this way the latent energy absorbed during these transitions can be observed. It is also possible to determine the Tg through Dynamic Mechanical Analysis (DMA), which continuously monitors the elastic modulus of a material during heating so as to observe the sudden increase in flexibility associated with the Tg. The Td is normally monitored with Thermogravimetric Analysis (TGA). TGA can be defined as the measurement of the change in weight of a substance as the temperature of its environment is varied in a controlled manner.5 The principal variables in TGA are weight, temperature and time. Early TGA focused on changes in weight over time at a constant temperature, known as isothermal testing,5 but it is now also common to monitor the loss in weight as the temperature continually increases, known as a temperature sweep. A typical temperature sweep TGA plot for the bioresorbable polymer PLA is shown in Fig. 10.1. The mass decomposition is shown as a dashed line, while the derivative of this is shown as a solid line. The derivative is used to determine the important transitions during decomposition. Analysis of the TGA plot in Fig. 10.1 determined that two distinct decomposition processes existed. The first occurs at 296°C and is a depolymerisation of the material involving residual catalysing tin. The second occurs at

TGA [mg]

DrTGA [mg/min]

0.00

370 °C 225 °C

–2.00

0.00

–0.10

326 °C

–0.20 –4.00

296 °C 350 °C –0.30 200.00

300.00 Temperature [°C]

10.1 TGA results for a PLA sample.8

400.00

Influence of processing, sterilisation and storage

215

350°C and is due to general decomposition of the polymer into a range of products including lactide, carbon oxides and acetaldehyde.8 Thermal degradation is a function of time as well as of temperature. Work by Yu et al.9 investigated the reduction in molecular weight over time of PLLA at temperatures ranging from 180°C to 220°C. It was found that the increased temperature caused a greater reduction in molecular weight, but that in all cases a time of approximately 2 hrs was required for the molecular weight to stabilise at the lower value. In practice, few synthetic polymers can maintain their properties for long periods of time at temperatures above 200°C4, so it is advisable to minimise processing times working at temperatures in this range. Sivalingam and Madras10 studied the thermal stability of the bioresorbable polymers PCL, PGA and PLA using TGA. It was reported that the thermal stability of these polymers, based on peak decomposition temperature, was 402°C, 360°C and 295°C respectively. Analysis showed similarities in the thermal degradation of the three polymers. At relatively low temperatures, thermal degradation proceeded through random chain scission. This affected the molecular weight of the material but led only to a mass loss of up to 1–2%. At relatively high temperatures, the dominant degradation mechanism became specific chain-end scission. This produced mass loss values of up to 70%. As well as causing the decomposition of a bioresorbable polymer, elevated temperature also accelerates the rate of hydrolytic degradation if moisture is present. Agrawal et al.11 found by elevating the temperature of the medium that a PLGA copolymer yielded significantly faster rates of degradation. It was hypothesised that this was because more energy was available for chain scission reactions and there was enhanced chain mobility. Li and McCarthy,12 while investigating the effect that heat has on the degradation of PLA in buffer solution, compared the rate at 37°C with that at 60°C. It was found that at 37°C the degradation was much slower. There has been a number of recent studies on the thermal degradation of PLA.10,13,14 This is generally considered to occur by a random chain scission mechanism. Assuming this mechanism, it has been possible to predict the degradation reaction and its reverse reaction using a mathematical model9 with the results being quite consistent with experimentally obtained data. Several papers have successfully used the Arrhenius relationship to predict the acceleration in degradation brought about by an increase in temperature.15,16 The Arrhenius approach models a physical degradation of interest as resulting from an activated chemical process, such that the rate is proportional to exp(−Ea /RT), where Ea is the Arrhenius activation energy, R is the gas constant, and T is the absolute temperature.3 When appropriate, this method predicts a linear relationship between the logarithm of the time

216

Degradation rate of bioresorbable materials

to a specified amount of material property change and the reciprocal of the absolute temperature16 as follows: ln k = −

Ea + ln A RT

where k is the rate constant and A is a constant. Determining the constant Ea for a material, through studying the degradation behaviour at specific temperatures, allows the prediction of the degradation behaviour at any specified temperature T, providing that the degradation process is still comparative in all other respects. Weir et al.17 compared the hydrolytic degradation of PLA at 37°C, 50°C and 70°C. The study verified that the Arrhenius relationship held true for the elevated temperatures, indicating that the degradation mechanism involved is similar for all the temperature increments. It should be noted, however, that the thermal degradation of PLA may depend on the amounts of residual zinc compounds remaining following the polymerisation process.13,14 The polysaccharide starch has been the focus of interest as a potential natural bioresorbable polymer for medical applications, including tissue engineering. It degrades in the body through enzymatic action but is also known to be a highly thermo-sensitive material, with potential to degrade during melt-processing. A study by Gomes et al.18 investigated the processing of this material by both extrusion and compression moulding and compared these techniques to solvent casting and in-situ polymerisation. It was reported that the melt-processed starches tended to take up less water during degradation studies and displayed lower levels of weight loss. The weight loss was found to be highest for the extruded materials and this was accounted for by the potential for thermo-mechanical degradation occurring during the extrusion process, which in turn led to easier breakdown of material associated with leaching of low molecular weight chains. A concurrent study by Sousa et al.19 found that reinforcing starch–polymer blends, in this case a starch with ethylene/vinyl alcohol (EVOH) plus hydroxyapatite (HA) fillers, increased the susceptibility of the material to undergo thermomechanical degradation during the injection moulding processing. This was considered to be a result of the high viscosity of the composite due to the more pronounced viscous heat dissipation during melt flow. This section has detailed how bioresorbable polymers decompose at specific temperatures, but can undergo accelerated hydrolytic degradation beneath this temperature. However, the latter process relies on the presence of moisture. Middleton and Tipton20 highlighted this by showing that a raised temperature which produced 50% degradation in a PGA sample that contained 0.02% water caused no significant degradation when the moisture content was reduced to 0.003%. The significance of these findings is that, when processing a bioresorbable polymer at high-temperatures, it is

Influence of processing, sterilisation and storage

217

imperative to maintain the dryness of the polymer in order that the level of induced degradation is minimised.

10.3.2 Mechanical degradation In addition to heat, degradation of polymers can be initiated by mechanical stress.2 In Aspects of Degradation and Stabilization of Polymers,5 Murakami listed some of the mechanical processes that may cause mechanical degradation. They include: • • • • • •

Milling Shearing Tension Compression Ultrasonic Impact

The induced degradation can be in the form of a physical change, such as crystallisation, flow or molecular orientation. It may also take the form of a chemical change, such as main-chain scission, side-group scission or even cross-linking. In the processing of bioresorbable polymers as outlined in Section 10.2, mechanical stress is often applied to the polymer in the form of shearing. However, heating is normally applied to the material at the same time. It is therefore difficult to attribute any produced degradation to solely mechanical effects. Rothen-Weinhold et al.21 compared the effects of extrusion (Section 10.2.1) and injection moulding (Section 10.2.2) on PLA. The study found that both processing techniques caused a reduction in the molecular weight of the polymer. The reduction caused by injection moulding was seen to be greater than that caused by extrusion. It is possible that the comparatively higher shear forces involved in injection moulding led to this additional degradation. However, it was also noted that the injection moulded material reached a higher temperature during processing. Another comparison of extrusion and injection moulding of bioresorbable polymer was carried out by Weir at al.22 In this case the investigated polymer was PCL. The study found that the molecular weight of the material was not affected by either technique, indicating that the mechanical shearing was not causing chain scission. However, it was observed that the crystallinity of the polymer, while unaffected by extrusion, was increased by injection moulding. Several studies have closely investigated the effect on bioresorbable polymers of mechanical forces other than shearing. In their study on the degradation of PGA under applied strain, Miller and Williams23 hypothesised that the predicted life of a biodegradable material may be altered by

218

Degradation rate of bioresorbable materials

the stresses it encounters. To test this hypothesis, they subjected the PGA-based suture material Dexon® to pre-imposed strains of 0.25% and 0.5% and placed the samples in solution at 37°C with a control sample. It was discovered that increasing the strain reduced the degradation time of the PGA. The finding was supported by in-vivo testing in rabbits. A similar study carried out by Deng et al.,24 investigated the effect of loading on the degradation of braids made of PLGA. This study found that the extent of external loading applied to the braids had no significant effect on their rate of degradation. It is possible that the disparity in the findings of these two studies is due to the crystallinity of the materials investigated. PGA is highly crystalline but PLGA is often amorphous. A study carried out by Yang et al.25 found that the application of loading to PLA could cause a significant change in the orientation and crystallinity of the material. It is likely that this would have an effect on the degradation rate. A review paper by Middleton et al.,20 which examined the various processing factors that affected bioresorbable polymers, concluded that high-shear rates for long periods of time caused degradation even at low temperatures. However, it was also noted that the significance of the thermal effects was such that at high temperature all other factors were inconsequential. This, in turn, has led to the practice of using mechanical energy in preference to thermal energy wherever possible when processing bioresorbable materials. An example of this is the use of twin-screw extruders to increase the mechanical input during processing.3

10.3.3 Structural stability The thermal history of bioresorbable polymers is important in determining their final structural properties, such as degree of crystallinity. Of the aliphatic polyesters, PLA would seem most sensitive to cooling rate. Following melt processing involving a relatively high post-processing cooling rate, PLLA appears almost transparent, with very low levels of crystallinity, as determined by differential scanning calorimetry (DSC).6 It is common practice to anneal PLLA after melt processing in order to achieve a relatively stable crystalline state. Figure 10.2 shows the appearance of PLLA after compression moulding (cooled on metal plate) and subsequent annealing in air (4 hours at 120°C). Crystallinity was found to increase from 12% to 43% on annealing. Figure 10.3 shows the appearance of PLLA after extrusion (quenched in water bath at 24°C) and a similar annealing process (4 hours at 120°C). Crystallinity for the extruded PLLA was found to increase from 20% to 40% on annealing.6 These findings indicate that it is important to ensure that the bioresorbable polymer is relatively stable in terms of crystallinity prior to application. Further studies in the authors’ laboratory have indicated that non-annealed PLLA can behave

Influence of processing, sterilisation and storage on bioresorbability 219 (a)

(b)

(c)

10.2 a) Supplied PLLA pellets, b) Compression moulded PLLA, c) Compression moulded annealed PLLA.6

(a)

(b)

10.3 a) Extruded PLLA, b) Extruded annealed PLLA.

220

Degradation rate of bioresorbable materials

unpredictably in terms of changes in strength during degradation by hydrolysis (in-vitro accelerated tests at 70°C). Rather than the strength continuously decreasing with time due to chain scission, as normally predicted by PLLA, the strength initially increases due to the increasing crystallinity caused by annealing in the aqueous environment. This is later followed by decrease in strength due to degradation.

10.4

Sterilisation

An important prerequisite for the use of any material in a medical application is that it can be sterilised before use.26 This applies to both materials used in-vivo and materials used to grow tissue in-vitro.27 The British Standard BS EN 556–1 : 2001, entitled ‘Sterilisation of medical devices – requirements for medical devices to be designated “STERILE” – Part 1: Requirements for terminally sterilised medical devices’, clearly sets out what is required of a sterilisation procedure: A sterile product item is one which is free of viable organisms. European Standards for medical devices require, when it is necessary to supply a sterile product item, that adventitious microbiological contamination of a medical device from all sources is minimised by all practical means. Even so, product items produced under standard manufacturing conditions in accordance with their requirements for quality systems for medical devices may, prior to sterilisation have micro-organisms on them, albeit in low numbers. Such product items are non-sterile. The purpose of sterilisation processing is to inactivate the microbiological contaminants and thereby transform the non-sterile items into sterile ones.

The standard goes on to state that, for a device to be designated sterile, the theoretical probability of there being a viable micro-organism present in the device or on it must be less than or equal to 1 × 10−6. The simplest and most cost-effective way to achieve sterilisation of a material is to use steam autoclaving. However, due to the chemical and structural sensitivity of bioresorbable polymers to heat and humidity, they cannot undergo this process.27 In general, the effects of sterilisation methods on polymer are not well compiled, nor is raw data easily accessed.28 The following sub-sections deal with the sterilisation techniques that can be successfully applied to bioresorbable polymers, and describe some of the recorded effects.

10.4.1 Ethylene oxide Exposure to the chemical ethylene oxide (EtO), the molecule of which is illustrated in Fig. 10.4, is a reliable, cost-effective method of sterilising most

Influence of processing, sterilisation and storage CH2

221

CH2 O

10.4 The ethylene oxide molecule.

polymeric devices.29 It is estimated that 48% of medical device sterilisation involves EtO.30 The chemical itself is a flammable, colourless gas at temperatures above 10°C. Other uses for it include the production of solvents, antifreeze, textiles, detergents and adhesives. It is known that EtO is toxic, and can cause breathing difficulties and blurred vision.31 As a result, great care must be used to ensure that all the gas is removed from a sterilised device before packaging and eventual use. This can often result in long aeration times, with research having indicated that up to two weeks may be required to fully remove residual EtO from a product.20 Several studies have indicated that EtO sterilisation has little effect on the physical properties of a bioresorbable polymer, especially when compared to radiation sterilisation techniques.6,30,32 However, recently, more serious questions have been asked about the safety of EtO sterilisation. The chemical has been proved to be a carcinogen and linked with leukaemia. It has also been shown to cause genetic and nerve damage, and potentially paralysis.31 America’s National Institute for Occupational Safety and Health has recommended that EtO be handled as a mutagen and potential carcinogen to humans and that unnecessary use be eliminated.29 For this reason, many companies are now looking for alternatives to EtO sterilisation for use with bioresorbable polymers, specifically those used for medical applications.

10.4.2 Gamma irradiation The sterilisation procedure with the second largest market share is γirradiation, with about 40% of medical devices being sterilised in this way.30 High-energy radiation has emerged as a preferred method of sterilisation due to its high efficiency, negligible thermal effects and the fact that, due to the penetrative capabilities of the radiation, articles may be packaged and sealed before sterilisation.26 γ-irradiation does not require post-sterilisation treatment of the samples, and can sterilise in a measurable, predictable manner.29 The radiation is produced from a radioactive source, typically Cobalt 60 (Co60), which produces γ-radiation with photon energies of 1.17 and 1.33 MeV. These values are five orders of magnitude larger than the average energy of a chemical bond.32 The radiation dose is delivered at a rate of between 5 and 25 kGy per hour33 with a standard sterilisation dose being between 25 and 40 kGy.34,35 The main drawback with this type of sterilisation, as referred to previously, is that it is known to induce

222

Degradation rate of bioresorbable materials

degradative molecular changes in polymers, in the form of chain scission and cross-linking.26,30 Both these processes alter the molecular weight of the polymer, decreasing and increasing it, respectively. A study by Nugroho et al.36 found that, when PLLA was exposed to γ-irradiation at a dose typical of sterilisation, the effect was chain scission dominated, producing a reduction in molecular weight. This means that the process is considered as degradation.37 There is a debate whether this is comparable to hydrolytic degradation, although reducing the molecular weight is likely to reduce the subsequent time required for the material to be resorbed. The study also found that, in semi-crystalline materials such as PLA, PGA and PCL, the chain scission occurred mainly in the amorphous regions of the polymer, and the crystalline regions degraded only at higher doses. Sterilisation with γ-radiation works through a process called Compton scattering (Fig. 10.5).33 In Compton scattering, the γ-radiation photons incident upon an atom within a polymer molecule. This causes the atom to emit a highly energised electron, or β particle. This electron initiates the release of more electrons from other atoms. These secondary electrons both bring about sterilisation, by inactivating any viable micro-organisms, and produce free radicals.35 In the Nugroho et al. study,36 electron spin resonance (ESR) testing showed that three types of radical were produced by γ-irradiation of PLA. They are shown in Fig. 10.6. Radical (I) is produced by abstraction of hydrogen from methane groups, while (II) and (III) are produced by cleavage of the carbon – carbon double bond. These radicals can react with one another to cause cross-linking. Alternatively, they can combine with available hydrogen or oxygen to become neutralised, making the net effect chain scission.37 A comprehensive study of the effects of gamma radiation on PLA was carried out by Nugroho et al.36 in 2001. In this study, powdered and solid PLA samples were irradiated by a Co60 source with doses of up to 1000 kGy. The samples were thoroughly characterised both before and after hydrolytic degradation. It was found that the Tm of the PLA decreased with increasing radiation dose, from approximately 165°C for the unirradiated material to approximately 125°C for the material irradiated with a dose of

Scattered photon

Incoming photon Atom

Scattered electron 10.5 Gamma-induced Compton scattering.

Influence of processing, sterilisation and storage

223

1000 kGy. The authors attributed this phenomenon to radiation-induced damage caused to the polymeric crystals. Samples that were irradiated in air were found to have lower Tm values than those irradiated in a vacuum, indicating that radiolytic degradation is accelerated by the presence of oxygen. Analysis of the molecular weight of the irradiated samples showed that, for doses of up to 200 kGy, chain scission occurred in the material causing the molecular weight to be reduced. For doses above 200 kGy, only a slight further reduction in molecular weight was recorded. The authors concluded that, even though chain scission occurred at a rate proportional to the received dose, in the range of 200–1000 kGy free-radical combination was causing cross-linking, which offsets the effects of the chain scission. Tensile testing of the samples revealed that doses of less that 30 kGy had little effect on the polymer’s strength. Doses of between 30 and 100 kGy sharply reduced the strength and elasticity of the samples. Beyond 100 kGy, little more reduction in the strength was observed. The authors theorised that, at doses of less than 30 kGy, chain scission occurred in the amorphous region only, which has little effect on the material’s strength. This is possibly because of the ‘cage-effect’ theory. This theory states that radicals are generated in all areas of the polymer; however, within the tightly packed crystalline structures, recombination of the radicals is much more likely than in the loosely packed amorphous regions. Therefore, chain scission in the polymer crystals is reversed. At greater doses, the crystalline regions of the polymer are affected, which produces a much more significant effect on the mechanical properties. The hydrolytic degradation study of the samples revealed that, contrary to previous findings, the gamma radiation caused samples to have lost less mass after a given time than the control samples. The authors attributed this to partial cross-linking of the polymer. A further explanation, not offered by the authors, is that the incorporation of enzymes into the degradation medium had an effect of the degradation process. ESR testing of the irradiated samples found the same radicals present as illustrated in Fig. 10.6. Although there is little published information on the effect of gamma sterilisation on natural polymers, a study by Noah et al.38 into the effect that gamma sterilisation had on the natural bioresorbable polymer collagen found that the radiation dramatically reduced the material’s resistance to biodegradation.

O

· C CH3 (I)

O

H

O

C

C·

C·

CH3 (II)

(III)

10.6 Radicals produced in PLLA by gamma radiation.36

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Degradation rate of bioresorbable materials

10.4.3 E-beam irradiation Electron-beam sterilisation involves using an electron accelerator to treat the material requiring sterilisation with a stream of highly energised electrons known as β-rays. Unlike γ-radiation, which is electromagnetic, β-rays consist of corpuscular particles.35 Table 3.1, from Woo and Sandford,33 gives a physical comparison of γ and β radiation. For applications such as sterilisation, high-energy commercial e-beams, operating in the 10 MeV range are used.39 e-beams offer the following advantages over γ-irradiation: • • • • • •

The dose rate can be as much as 10 000 times higher, leading to shorter processing times.33 There is no requirement for a nuclear source, meaning the process is more environmentally safe and acceptable to the public.35,39 Unlike γ facilities, e-beam generators can be switched off when not in use. There is better utilisation of energy as all the radiation is directed at the material to be sterilised. e-beam offers more control over the parameters of the sterilisation process. e-beam has been reported to be less damaging to the sterilised material.30

E-beam sterilisation also has several disadvantages. Due to the high dosage rate, a specific dose of e-beam radiation causes more heating in the material than the same dose of γ radiation.35 The corpuscular nature of e-beam radiation means that its penetrative depth is much less than for γ radiation. The depth to which the radiation penetrates is dependent on the kinetic energy obtained from the potential difference through which the electrons were accelerated, but it is unlikely that e-beam could be used for sterilising very thick devices.30,33 The main disadvantages that have hindered the development of e-beam sterilisation are that of the high capital costs for installing Table 10.1 Comparison of e-beam and gamma radiation (Woo and Sandford33)

Charge Rest mass Energy Velocity c = Speed of light.

Electron (β)

Gamma (Co60)

−1 9e-28 g

0 0

0.1–15 MeV 0.3–0.99 c

1.2 MeV c

Influence of processing, sterilisation and storage

225

e-beam equipment and the unreliability of the complex machinery involved.40 Recently, however, e-beam machines have been improved to give increased available energy, more compact design and improved reliability.33 These improvements make e-beam faster, easier and cheaper from an operational point of view, which when coupled with the lack of environmental risk from nuclear radiation means e-beam is more likely to be used in the future.35 Although the primary event of interaction with matter is different between γ and β radiation, the major interaction for both is Compton scattering (see Section 10.4.2 and Fig. 10.2). It is mainly the shower of secondary electrons that initiates the ionising events which activate numerous chemical reactions, many of which lead to degradation of the material.33 As the e-beam treatment of bioresorbable polymers with β radiation is a newer technique than treatment with gamma radiation, there is less literature available on the subject. Perhaps the most comprehensive work carried out in the area is detailed in a series of papers published by Loo et al.37,41,42 These studies extensively investigated the effects induced in PLA and PLGA by e-beam radiation. Doses of up to 500 kGy were investigated. The study found that both PLGA and PLA underwent chain scission when irradiated in this manner, leading to a lower molecular weight. The PLA was found to be more resistant to the radiation than the PLGA, a phenomenon attributed to its higher crystallinity allowing recombination through the ‘cage effect’. The extent of the reduction in molecular weight was found to be proportional to the received dose for all PLA samples, and for the PLGA samples treated with less than 200 kGy of radiation. Beyond this level, the molecular weight of the PLGA was observed to remain constant. The authors suggested that crystallinity was introduced into the PLGA at this stage, giving the material more resistance to radiolytic degradation through the ‘cage effect’. Both the Tm and Tg of the polymers were reduced by the β radiations. The reduced molecular weight of the polymers, coupled with the reduction in Tg, was found to promote crystallisation in both polymers. Hydrolytic degradation studies carried out on β irradiated samples revealed a similar trend to that observed following gamma irradiation, with the reduced molecular weight of the samples prior to immersion in the degradation medium leading to earlier mass loss. The predictability of the effects induced in these bioabsorbable polymers, established through repeated testing, led the authors to propose e-beam radiation as a potential tool to control their rate of hydrolytic degradation. While e-beam radiation clearly has an effect on the degradation of synthetic bioresorbable polymers, it does not necessarily affect natural polymers in the same way. A study by Silva et al.43 found that doses of up to 100 kGy of β radiation had no appreciable effect on the properties of the natural bioresorbable polymer chitosan.

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Degradation rate of bioresorbable materials

10.4.4 UV irradiation Treatment of bioresorbable polymers with UV radiation has been shown to be a cheap and effective method of achieving surface sterilisation.27 A study by Fischbach et al.27 investigated the effect that UV radiation doses appropriate for sterilisation had on a bioresorbable polymer used for tissue engineering purposes. The study found that doses sufficient in magnitude to achieve sterilisation could be applied to the polymer without causing significant degradation. However, increasing the dose beyond this level led to degradation-related changes in the surface properties of the polymer. UV irradiation is therefore a viable tool for the surface sterilisation of bioresorbable polymers, but care must be taken in order that an excessive dose is avoided.

10.4.5 Plasma The growing role of complex, delicate medical devices, coupled with the hazardous nature of EtO has given rise to an interest in plasma sterilisation. Cold plasma is a partially ionised gas at reduced pressure, comprising ions, electrons, UV photons, and reactive neutrals such as radicals, and excited and ground-state molecules. It has been proved capable of destroying microorganisms,44 and hence of being used for sterilisation. Low-temperature plasma technology had previously been used to modify the surface of biomaterials without changing their bulk properties. In the mid-1980s, Sterrad® became the first commercially available gas plasma sterilising system. It used hydrogen peroxide as the reactive gas, but was soon followed by Plazlyte®, which utilised peracetic gas vapour.45 Plasma sterilisation offers advantages over other sterilisation techniques that include shorter processing time, minimal effect on material dimensions and properties, no toxic or carcinogenic by-products and no disposal problems.30 However, many questions still remain with regards to its efficacy and potential hazards from possible modification of treated materials.44 Studies by Holy et al.,46 Ayhan,45 and Nuutinen et al.30 all found that plasma sterilisation had little effect on the material properties of bioresorbable polymers, but these studies did not assess the efficacy of the sterilisation. It should also be considered that plasma sterilisation could not be used on the same scale and for the same throughput as γ-irradiation.

10.5

Maximising shelf-life: Packaging and storage

There are several techniques that can be employed to minimise the degradation occurring in a bioresorbable material during storage of the material after processing. This section details some of these techniques.

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227

Several environmental factors can promote the degradation of a bioresorbable material, altering its subsequent performance. In addition to the presence of moisture, which was highlighted in Section 10.3.1, the presence of gases, the presence of acids or bases, the local temperature, the presence of micro-organisms and the presence of low-energy radiation such as ultraviolet light can induce degradation in both bioresorbable polymers and cements.1,2,47,48 To prevent unnecessary degradation of a product and to maximise the time it can be stored before use, steps must be taken to shield the product from these environmental factors. This normally involves the application of specific packaging and storage techniques.

10.5.1 Materials used in the packaging of bioresorbable materials When designing the packaging of any product, the aim is to achieve two main goals. The first is to contain the product in whatever form it is to be stored or transported. The second is to protect and preserve the product from physical, chemical or biological harm.49 Bioresorbable products do not need more physical protection than non-bioresorbable products. However, they do need particular protection from chemical and biological damage. This can be achieved through the appropriate use of various packaging technologies. In general, packaging is formed from50 a combination of paper, wood, glass, steel, aluminium or plastic. However, if the product is to be protected from moisture, oxygen and other chemical that may cause damage, it is important to use materials with appropriate barrier properties. Glass packaging has excellent barrier properties, being impermeable to gases and moisture, but is limited in practice by its weight and fragility. An example of glass packaging for bioresorbable materials would be the bottling of medicinal tablets. Aluminium foils provide excellent barrier properties while still allowing packaging to be light and flexible. An aluminium foil with a thickness of greater than 17 microns forms a barrier that excludes 100% of all gases.49 This foil can be included between two layers of polymer. Aluminium is also resistant to oxidation and can be easily formed and processed into appropriate packaging shapes. Also, the metal’s bright, glossy appearance is considered of benefit to the aesthetic of the packaging. Polymers are also commonly used to package bioresorbable materials. Approximately 60% of the polymer used in packaging is polyethylene (PE).50 Other commonly used polymers include polypropylene (PP), polystyrene (PS) and polyvinyl chloride (PVC). The barrier properties of these materials are quite different. PE forms a good barrier against moisture but is a poor barrier against hydrocarbon solvents. PS is a poor barrier against both moisture and gas. The barrier properties of PP vary greatly depending

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Degradation rate of bioresorbable materials

on the orientation of the polymer. PVC is an excellent barrier against oil and grease but can be easily permeated by moisture. Despite these varying properties, polymers are commonly used for packaging sensitive materials. One of the main reasons for this is that different polymers can be combined to provide the different barrier properties they offer whilst eliminating their shortcomings. A good example of this is ethylene/vinyl alcohol (EVOH), which is an excellent oxygen barrier but is damaged by water. Sandwiching a layer of EVOH between two layers of PE gives a packaging material that acts as a barrier against both oxygen and moisture.

10.5.2 Extending the shelf-life of packaged bioresorbable materials One measure of the efficacy of packaging is the shelf-life of the product it contains. The shelf-life can be defined as the length of time a product remains in an acceptable condition for use. Whilst it is largely dependent on the materials used, there are also six techniques that can be used to extend the shelf-life. They are: • • • • • •

Reduced temperatures Thermal processing Water reduction Chemical preservation Modified atmospheres Sterilisation49

Sterilisation has been previously discussed in Section 10.4. However, when packaging a bioresorbable material, it is important to consider what sterilisation processes may be applied. If EtO is to be used to sterilise the material, the packaging must be porous to gas but not to micro-organisms.50 This means that the materials will not be protected from atmospheric oxygen and moisture. The same can be said of packaging for plasma sterilisation. If a material is to be radiation sterilised, careful consideration must be given to any effect the radiation may have on the packaging. Polymers in particular can be adversely affected by radiation, reducing the shelf-life of the product they contain. The possible effects include cross-linking and chain scission along with the associated changes in mechanical properties, as well as discolouration of the polymer.50,51 Reducing the temperature of packaged bioresorbable materials is advisable as the temperature dependence of degradation processes means that cooling will minimise the induced degradation. However, maintaining a reduced temperature can be expensive and is not always practical. Thermal processing is used to extend the shelf-life of packaged materials by killing off any micro-organisms that may be present. However, the

Influence of processing, sterilisation and storage

229

temperature sensitivity of most bioresorbable materials means this is not applicable for their packaging. The negative effects of moisture on bioresorbable materials, both polymeric and ceramic, have been well established, along with the need to use packaging materials that form a good moisture barrier. In addition to this, it is important to eliminate as much moisture from the product as possible before packaging. It is also important to reduce the relative humidity of any gas present in the packaging. The inclusion of chemical preservatives with a packaged bioresorbable material can extend the shelf-life. This commonly takes the form of a desiccant salt, designed to absorb any residual moisture before hydrolysis can occur. It is also possible to reduce the effects of oxidation by including antioxidants or oxygen absorbers. One common form of oxygen absorber is fine iron powder.49 The addition of solid citric acid has been found to significantly extend the shelf-life of some bioresorbable ceramics.48 Modified Atmosphere Packaging (MAP) can also effectively extend the shelf-life of packaged bioresorbable materials. Most commonly, this would take the form of vacuum packing, in which the atmosphere is removed from the packaged product, eliminating both oxygen and moisture; or through inert atmosphere packing, in which the packaging is filled with a gas such as nitrogen that does not promote degradation.

10.6

Additives for reducing degradation

The premature degradation of bioresorbable materials can be avoided through the use of additives such as stabilisers and antioxidants.2 Generally, these additives work by interfering with the degradative process, or by minimising the cause of the degradation.4 In this section, the different types of additive that can be used are discussed.

10.6.1 Antioxidants Many of the degradation processes that occur in bioresorbable materials, particularly in polymers, involve free radicals. Therefore, if an agent is introduced to the material that is capable of combining with free radicals, and thereby stabilising them, the degradation process cannot proceed. Chemicals that can do this are known as antioxidants. The majority of antioxidants used as additives in polymers are either sterically hindered phenols or aromatic amines.4 These chemicals react quickly with any radicals present to eventually form non-radical products. The choice of antioxidant additive to use in a bioresorbable material depends largely on the material itself and what its intended final use is.

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10.6.2 Metal deactivators It has been mentioned that the presence of some metals can catalyse the degradation of bioresorbable materials. Compounds containing Fe, Co, Mn, Cu, Ce, or V can contribute to this effect. The contribution these metals make to degradation can be minimised if chemical additives are used that combine with the metal ions to form inactive complexes.4 These chemicals are called metal deactivators, and normally contain one of the elements N, O, S or P.

10.6.3 Light stabilisers The degradation caused by UV radiation can be minimised by the addition of various chemical additives. If the colour of the bioresorbable material is unimportant, the inclusion of an opaque additive such as carbon black can act to reduce UV-promoted degradation. More advanced additives include:4 • •

•

UV absorbers: These materials absorb UV radiation before dissipating the energy as heat or higher wavelength radiation. Quenchers: These materials contain chromophores which are capable of converting the UV radiation to heat or fluorescent and phosphorescent radiation. Hydroperoxide decomposers: These additives act to decompose hydroperoxides, which are an integral part of UV-initiated degradation.

10.6.4 Additives for reducing thermal degradation Section 10.3.1 detailed how heat contributes to the degradation of bioresorbable materials. Additives can be used to minimise the thermal degradation that occurs in a material. These additives work in two main ways, namely thermal stabilisation and process aiding. Thermal stabilisers are materials that preferentially absorb heat when they are included in a material. They commonly take the form of powdered inorganic materials such as limestone, talc and alumina.3 These materials are selected because of their large heat capacities. Processing aids act to reduce the temperature required to successfully process a material, normally by reducing the melting temperature.3 In polymers, processing aids often take the form of plasticizers. These are organic substances of low volatility that reduce the Tg and Tm of polymers, and reduce the viscosity of the polymer melt at a given temperature.4 The polymer can then be processed at a lower temperature, and will suffer less thermal degradation.

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It should be noted that both thermal stabilisers and processing aids can have a significant effect on the final properties of a bioresorbable material. Therefore, their use is not always appropriate. It should also be noted that, if a resorbable material is to be used in a medical application, the use of additives is strictly controlled to ensure that the patient’s safety is maintained, and that this may limit the possible use of additives.

10.7

Conclusion

In choosing to utilise bioresorbable materials, there is an associated requirement to consider the processing, sterilisation and storage requirements. The conditions involved with each of these can have a significant effect on the subsequent performance of the bioresorbable material. It is crucial that, before a bioresorbable material is selected, consideration is given to the many ways the material may be damaged during its processing, storage and application. It is equally important that the user is aware of the various methods available to minimise the damage induced. This chapter has outlined the various aspects relevant to these processes.

10.8

References

1 c. bastioli, Handbook of Biodegradable Polymers, Woodhead Publishing, Cambridge, 534 (2005). 2 r. j. crawford, Plastics Engineering (Butterworth-Heinemann, Oxford, UK, ed. 3rd, 1998). 3 a. brent strong, Plastics: Materials and Processing (Prentice-Hall, Inc., New Jersey, USA, ed. 2nd, 2000). 4 m. chandra, s. k. roy, Plastics Technology Handbook (Marcel Dekker, Inc., New York, USA, ed. 3rd, 1998). 5 h. h. g. jellinek, Aspects of Degradation and Stabilization of Polymers, Elsevier, Amsterdam, 690 (1978). 6 n. a. weir, f. j. buchanan, j. f. orr, d. f. farrar, a. boyd, Biomaterials 25, 3939 (2004). 7 k. pielichowski, j. njuguna, Thermal Degradation of Polymeric Materials (Rapra Technology Limited, Shropshire, UK, 2005). 8. f. d. kopinke, m. remmler, k. mackenzie, m. möder, o. wachsen, Polymer Degradation and Stability 53, 329 (1996). 9 h. x. yu, n. x. huang, c. s. wang, z. l. tang, J Appl Polym Sci 88, 2557 (2003). 10 g. sivalingam, g. madras, Polymer Degradation and Stability 84, 393 (2004/6). 11 c. m. agrawal, d. huang, j. p. schmitz, k. a. athanasiou, Tissue Engineering 3, 345 (1997). 12 s. li, s. mccarthy, Biomaterials 20, 35 (1999). 13 y. fan, h. nishida, y. shirai, y. tokiwa, t. endo, Polymer Degradation and Stability 86, 197 (2004/11). 14 h. abe, n. takahashi, k. j. kim, m. mochizuki, y. doi, Biomacromolecules 5, 1606 (2004).

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15 t. pohjonen, p. tormala, In: Biodegradable Implants in Fracture Fixation, Proceedings of ISFR Symposium 75 (1993). 16 j. wise, k. t. gillen, r. l. clough, Polym. Degrad. Stab. 49, 403 (1995). 17 n. a. weir, f. j. buchanan, j. f. orr, d. f. farrar, g. r. dickson, Proc. Inst. Mech. Eng. Part H J. Eng. Med. 218, 321 (2004). 18 m. e. gomes, j. s. godinho, d. tchalamov, a. m. cunha, r. l. reis, Materials Science and Engineering: C 20, 19 (2002/5/31). 19 r. a. sousa, j. f. mano, r. l. reis, a. m. cunha, m. j. bevis, Polym. Eng. Sci. 42, 1032 (2002). 20 j. c. middleton, a. j. tipton, Biomaterials 21, 2335 (2000). 21 a. rothen-weinhold et al., Eur. J. Pharm. Biopharm. 48, 113 (1999). 22 n. a. weir, f. j. buchanan, j. f. orr, d. f. farrar, Plastics, Rubber and Composites 32, 265 (2003). 23 n. d. miller, d. f. williams, Biomaterials 5, 365 (1984). 24 m. deng et al., Biomaterials 26, 4327 (2005). 25 x. yang, s. kang, y. yang, k. aou, s. l. hsu, Polymer 45, 4241 (2004). 26 o. kantoglu, o. guven, Nuclear Instruments and Methods in Physics Research, Section B: Beam Interactions with Materials and Atoms 197, 259 (2002). 27 c. fischbach et al., Surf Sci 491, 333 (2001). 28 plastics design library staff, The Effect of Sterilization Methods on Plastics and Elastomers, Plasties Design Library (Div. of Wm. Ardrew Inc.), Morris, NY (1994). 29 k. a. hooper, j. d. cox, j. kohn, J Appl Polym Sci 63, 1499 (1997). 30 j. p. nuutinen, c. clerc, t. virta, p. tormala, J. Biomater. Sci. – Polym. Ed. 13, 1325 (2002). 31 occupational health and safety administration, u.s. department of labour, OSHA Ethylene Oxide Factsheet (2002). 32 l. costa et al., Biomaterials 19, 659 (1998). 33 l. woo, c. l. sandford, Radiat. Phys. Chem. 63, 845 (2002). 34 k. gorna, s. gogolewski, Polym. Degrad. Stab. 79, 465 (2003). 35 l. montanari et al., J. Controlled Release 90, 281 (2003). 36 p. nugroho, h. mitomo, f. yoshii, t. kume, Polym. Degrad. Stab. 72, 337 (2001). 37 s. c. j. loo, c. p. ooi, y. c. f. boey, Polym. Degrad. Stab. 83, 259 (2004). 38 e. m. noah, j. s. chen, x. y. jiao, i. heschel, n. pallua, Biomaterials 23, 2855 (2002). 39 r. l. clough, Nuclear Instruments and Methods in Physics Research, Section B: Beam Interactions with Materials and Atoms 185, 8 (2001). 40 j. w. barnard, Nuclear Instruments and Methods in Physics Research, Section B: Beam Interactions with Materials and Atoms 208, 98 (2003). 41. j. s. c. loo, c. p. ooi, f. y. c. boey, Biomaterials 26, 1359 (2005). 42 s. c. j. loo, c. p. ooi, y. c. f. boey, Biomaterials 26, 3809 (2005). 43 r. m. silva, c. elvira, j. f. mano, j. san roman, r. l. reis, Journal of Materials Science–Materials in Medicine 15, 523 (2004). 44 s. lerouge, m. tabrizian, m. r. wertheimer, r. marchand, l. yahia, Biomed. Mater. Eng. 12, 3 (2002). 45 f. ayhan, h. ayhan, e. piskin, J. Bioact. Compatible Polym 13, 65 (1998). 46 c. e. holy, c. cheng, j. e. davies, m. s. shoichet, Biomaterials 22, 25 (2001). 47 a. gopferich, Biomaterials 17, 103 (1996).

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48 u. gbureck, s. dembski, r. thull, j. e. barralet, Biomaterials 26, 3691 (2005/6). 49 w. soroka, Fundamentals of Packaging Technology (The Institute of Packaging, ed. 3rd, 1999). 50 s. e. m. selke, Understanding Plastics Packaging Technology (Carl Hanser Verlag, 1997). 51 r. l. clough, k. t. gillen, g. m. malone, j. s. wallace, Radiation Physics and Chemistry 48, 583 (1996/11).

11 Influence of porous structure on bioresorbability: Tissue engineering scaffolds P. TOMLINS, National Physical Laboratory, UK

Abstract: A challenge for tissue engineers is to select a suitable material from which to manufacture a temporary porous, permeable scaffold such that it is biocompatible, degradable over an appropriate timescale, has a structure that can maintain an increasing density of cells and has acceptable mechanical properties. This chapter describes some of the key classes of materials that meet the biocompatibility and degradability criteria as well as discussing various routes to manufacturing scaffolds. Different approaches to characterising scaffold and for monitoring the degradation process are also described. Key words: tissue scaffold, degradation, porous, electrospinning, polysaccharides, proteins, particle leaching.

11.1

Introduction

The portfolio of biodegradable materials that has so far been investigated by tissue engineers is small, certainly less than twenty, the rationale being that these materials have been approved by regulatory bodies and are not obviously bio-incompatible with living tissues. Such materials, or combinations of them, are used to fabricate temporary structures or scaffolds to use in tissue engineering or regenerative medicine applications. These structures are highly porous and consist of a series of interconnected cavities or pores. The function of the scaffold, when seeded with cells, is to provide a suitable environment for cell attachment, proliferation and growth, i.e. provision of appropriate levels of nutrients and oxygen as well as removal of potentially toxic waste products. This is achieved through a network of interconnected pores that can span a very wide lengthscale, from nanometres to sub-millimetre. Providing quantitative measures of the distribution of pore sizes and their interconnectivity is challenging, although necessary, and is discussed in more detail below. Most scaffolds are manufactured from synthetic or naturally occurring polymers (Piskin, 1997) and form the basis of this chapter, although other materials, such as bio-glass (Sepulveda et al., 2002), natural inorganic materials such as coral (Gravel et al., 2006) hydroxyapatite and tricalcium phosphate (Burg et al., 2000; Kretlow and Mikos, 2007) are also used: The choice of scaffold material depends on the intended application. 234

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235

A variety of methods are used to produce scaffolds; these are briefly described below, and include moulding, weaving, knitting and solid freeform fabrication. The choice of manufacturing method used has a significant impact on scaffold structure in terms of controlling pore size, homogeneity, i.e. through-thickness consistency and the proportion of closed, blind-end or open pores (these terms are defined in Fig. 11.5) as well as its mechanical integrity over time. All of these factors can, depending on the material, influence the rate of degradation. A key necessity of a cell-seeded tissue scaffold is to maintain a level of mechanical integrity that is appropriate for a particular application for a period of time such that the loss of scaffold performance is compensated for by the growth of new tissue. Optimising this process is not trivial, given the number of potential variables that need to be considered, which include: • • • •

• •

The choice of material, e.g. type of polymer, homopolymer or copolymer, single material or blend. The initial molecular weight and molecular weight distribution of the polymer. The degree of crystallinity. The presence of residual stress in, for example, extruded structures or surface features such as oriented skins created during processing. (Crosssections of injection moulded objects typically have a skin-core texture. Polymer chains within the skin tend to be aligned whilst those in the core are randomly oriented.). The local environment in terms of pH mechanical loading and fluid exchange. The structure and geometry of the scaffold – wall thickness, overall dimensions and homogeneity.

Typically, scaffold degradation is monitored in vitro using destructive or invasive test methods that include detecting changes in the molecular weight distribution of the polymer, in mechanical performance (e.g. tensile modulus or strain to failure) and in mass. The temperature at which the test is carried out is also an important consideration, but most in vitro tests are carried out at 37°C with the exception of those designed to accelerate degradation.

11.2

Materials

Tissue engineers use both natural and synthetic polymers. Natural materials, such as collagen, chitosan and starch, may offer some advantages over synthetic polymers as they occur in tissues. However, from a quality control perspective, such materials are more likely to vary between suppliers, and

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batches, in terms of polydispersity, branching and purity, compared with synthetic polymers. Much of the reported literature in tissue engineering is focused on two polyesters, poly(lactic acid) and poly(glycolic acid) and their copolymers – materials that have been used since the 1970s as soluble sutures. Most synthetic polymers degrade by hydrolysis and are insensitive to enzyme attack, which in principle should minimize patient-to-patient variability (Yang et al., 2001). Griffith (2000), Yang et al. (2001) and Hayashi (1994) have reviewed the use of synthetic polymers in tissue engineering, hence only a brief overview of materials commonly used in tissue engineering will be given here Further, more detailed information on materials can be found in Part II of this book.

11.2.1 Polyesters

Time (months)

The polyesters, poly(lactic acid) (PLA) and poly(glycolic acid) (PGA), as well as their copolymers, have been used extensively as scaffold materials in tissue engineering (Chu, 1995; Törmälä et al., 1998). The degradation rate of the copolymer can be tuned by altering the relative proportions of PLA and PGA, respectively (Fig. 11.1) to change the degradation time from weeks to years, depending on the geometry of the structure and local environmental conditions. PLA and PGA degrade by hydrolysis of the hydrolytically unstable ester linkages within the polymer backbone to form lactic and glycolic acids, respectively. Glycolic acid can be secreted in urine or further degraded to produce serine. As both serine and lactic acid are precursors to pyruvic acid, they can enter the Kreb’s or tricarboxylic acid cycle,

7

7

6

6

5

5

4

4

3

3

2

2

1

1

0

0 0

PGA

25

50 Copolymer ratio

75

100 PLA

11.1 Changes in the proportion of glycolic acid to lactic acid will affect the in vivo half-life a PLGA copolymer (from Miller et al., 1977).

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eventually being metabolised to form carbon dioxide and water. Hence, the primary route for secretion of the PGA and PLA breakdown products is through respiration (Mohammadi and Jabbari, 2006; Cha and Pitt, 1990). This is an autocatalytic process, i.e. the bulk material degrades at a faster rate than the surface due to a build-up of carboxylic acid end groups. For semicrystalline polyesters, such as poly-L-lactide (PLLA), the degradation process effectively occurs in two stages (Chu, 1981; Cam et al., 1995). Initially, the process begins in the amorphous regions that surround the crystallites (Cam et al., 1995) before attacking the more resistant crystalline regions (Cha and Pitt, 1989; Agarwal et al., 1994; Gan et al., 1997). This second stage of degradation is associated with a reduction in the mass of the specimen. Poly(caprolactone), PCL has also been widely used as a scaffold material (Oh et al., 2007). Akin to PLA, PGA and their co-polymers, PCL also degrades into products that can be easily metabolised (Woodward et al., 1985; Pitt et al., 1981). The in vivo performance of PCL scaffolds depends on the location of the material and the application, as well as on the molecular weight of the polymer. High molecular weight, semi-crystalline PCL, for example has been shown to have a very long degradation time in vivo (Pernin and English, 1997). Hydroxyapatite has been used to reinforce PCL when used as a bone substitute: The processing methodologies required to produce these materials have been extensively reviewed by Baji et al., 2006. Poly(propylene fumarate), PPF, also has potential as a scaffold material for bone tissue engineering applications (Hedberg et al., 2005; Lee et al., 2006; Peter et al., 1997). This material, like other polyesters, degrades into nontoxic products, which in this case are fumaric acid and polyproplyene glycol (Timmer et al. 2001, 2003; Peter et al., 1997). It can be injected into an in vivo site and cross-linked in situ.

11.2.2 Polysaccharides Naturally occurring chitosan is a crystalline polysaccharide that is soluble in dilute acids. It consists of glucosamine and N-acetyl-D-glucosaminoglycan groups, a structure similar to glycosaminoglycan (Domard and Domard, 2001). Chitosan is positively charged at physiological pH and has a hydrophilic surface that encourages cell attachment, differentiation and proliferation (Matthew, 2002; Madihally and Matthew, 1999). Chitosan is used either on its own or as a composite for tissue engineering applications, including production of skin, bone and cartilage (Gravel et al., 2006; Gong et al., 2000; Cheng et al., 2003; Yamaguchi et al., 2003; Itoh et al., 2003; Freier et al., 2005; Wang et al., 2005; Khor and Lim, 2003; Nettles et al., 2002). It has also been extensively used in neural tissue engineering as it has been shown to promote adhesion, neurite outgrowth or nerve cells (Itoh et al., 2003;

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Gong et al., 2000). The material degrades through enzyme activity; for example, Tomihata and Ikada (1997) demonstrated that the rate of degradation in vitro, when immersed in lysozyme solution, was comparable to that seen in vivo. Polysaccharide hydrogels typically lack the rigidity to be scaffold materials per se, although alginates, for example, have been used with stiffer PLA to improve cell loading and retention within the scaffold (Caterson et al., 2001). Hydrogels manufactured from agarose have shown promise as scaffold materials for culturing chondrocytes (e.g. Aufderheide and Athanasiou, 2005; Benya and Shaffer, 1982; Lee et al., 2000) and for axonal regeneration (Stokols et al., 2006).

11.2.3 Proteins Collagen is the most widely used natural polymer in tissue engineering as it is a key structural component of extracellular matrix (ECM). Although some 25 types of collagen are known, it is the Type 1 and III collagens that tend to be used for tissue engineering applications such as skin, trachea and peripheral nerve regeneration (Yannas et al., 1982; Teramachi et al., 1997; Yamamoto et al., 1999). Although its inherent biocompatibility makes it an attractive material to use for tissue engineering applications, its poor mechanical properties have limited its usage. A solution to this problem is to reinforce the material. Hiraoka et al. (2003) used collagen sponge reinforced with a PGA fibre as a potential scaffold. Yamamoto et al. (1999) adopted a similar principle using a silicone stent. Collagen can be denatured by acidic or alkaline processing to form gelatins A and B, respectively. Gelatin has been extensively used as a porogen in the manufacture of scaffold structures (see below) or as a scaffold material in its own right (Murugan and Ramakrishna, 2006; Zhao et al., 2002). The ubiquitious proteoglycan, hyaluronic acid has been extensively used to culture cartilage, typically when coupled with collagen (Liao and Cui, 2004). Bone, cartilage and ligaments have been cultured using scaffolds manufactured from silk-based fibres (e.g. Uebersax et al., 2006).

11.3

Processing

Polymeric tissue scaffolds can be fabricated using a range of techniques that include electrospinning, solvent casting followed by particulate leaching, melt moulding, solid free-forming, gas foaming, freeze drying, self-assembly and drawing. Each of these methods has its pros and cons in terms of being able to predict the final structure, cost, reproducibility and applicability to a particular material system. Increasingly, the processing methods are being considered with a view to scaling up where factors such as repeatability, throughput and cost are significant.

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Manufacturing processes can be broken down into two groupings: those that produce foam-like structures and those that generate fibres. The latter can be used as random meshes or be further processed to form fabrics. These manufacturing processes have been extensively reviewed elsewhere, hence only brief details will be described here (e.g. Sachlos and Czernuszka, 2003).

11.3.1 Manufacture of foam-like structures Solvent casting followed by particulate leaching The commonest way of producing polymeric foam-like structures (Fig. 11.2) with irregular pore sizes and interconnectivity is to dissolve the polymer in an organic solvent. A porogen, usually sodium chloride, is then stirred into the resultant solution, which is cast into a mould and the solvent allowed to evaporate away (e.g. Kuo and Leou, 2006; Reignier and Huneault, 2006). The resultant ‘cake’ of porogen/polymer is then immersed in water to extract the porogen. This process can take several days, depending on the pore sizes of the scaffold and the geometry of the ‘cake’. Progress can be monitored by measuring the conductivity of the water, which will decrease with time (if water is changed frequently). The pore size in the resultant structure can, to some extent, be controlled by sieving the porogen, although the level of control is not high, especially as the interconnectivity of the pores is, at least in part, dependent on the hydroscopic nature of the sodium chloride. Glucose, paraffin and gelatin particles are also used as sacrificial porogens (e.g. Burg et al., 2000; Ubersax et al., 2006; Thomson et al., 1995; Holy et al., 1999). The level of porosity depends on the amount of porogen added to the mixture and, at lower porosities where contact between adjacent particles is limited, i.e. less than approximately 60%, there is an increased risk of

~30 μm

11.2 Some scaffold structures are very complicated, with window-like apertures appearing in the walls of the pores, as shown in this salt-leached polycaprolactone matrix.

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particles becoming isolated (Hou et al., 2003). This not only has an impact on the level of porosity determined from density measurements but also may have an adverse effect on the micro-environment within the scaffold, depending on the permeability of the scaffold to water and the diffusivity of the soluble porogen within the polymer matrix. Phase separation of mixtures Foams can also be produced by phase separating polymer mixtures or solutions induced by a change in temperature, the so-called ‘Temperature Induced Phase Separation’ or TIPS method. The polymer mixture method relies on one of the polymers being water soluble. The two polymers are mixed in the presence of a solvent which is then evaporated; the resultant solid is then heated to induce phase separation, rapidly cooled, and immersed in water. The pore size and interconnectivity is governed by the composition of the mixture and the extent of phase separation, which in turn depends on temperature and time. Using this method to control pore size and interconnectivity has been reported as being difficult (Nam and Park, 1999) due to the sensitivity of the structure to the processing conditions. Any solvent or unwanted polymer residues in the structure are additional problems (Reignier and Huneault, 2006). Removal of frozen water or other solvents, e.g. benzene, through freezedrying or lyophilisation, can be used to prepare scaffolds from materials that are susceptible to thermal degradation (Yannas et al., 1980; Chen et al., 2004; Liao and Cui, 2004; Nettles et al., 2002). Freezing produces solvent crystals that drive the scaffold matrix material into the interstitial spaces. Under partial vacuum, the crystals can be removed, leaving a porous matrix. Hiraoka et al. (2003), for example, prepared PGA fibre-reinforced collagen sponge using this method, freezing a mixture of collagen/ PGA and water at −20°C. The pore size can be controlled via the freezing rate and, depending on the material system, pH. Faster freezing rates produce smaller solvent crystals and hence smaller pores in the resultant scaffold (Dagalakis et al., 1980; Doillon et al., 1980). Other examples include the preparation of PLA/ nano-particulate hydroxyapatite/collagen (Liao and Cui, 2004) and PLGA/ collagen (Chen et al., 2004). Melt moulding using blowing agents or sacrificial porogens Gomes et al. (2002), for example, compression moulded polymer blends of starch with either ethylene/vinyl alcohol or cellulose acetate loaded with salt to produce porous scaffolds. The porogen was then leached out using water. This method is identical to that of solvent casting described

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previously, except that it is solvent free. This approach is useful in that it avoids the potential problems of residual solvent and the presence of large cavities that can form during solvent casting. However, the method is applicable only to thermally stable polymers. Reignier and Hunealt (2006) applied the method to a blend of PCL and water-soluble poly(ethylene oxide), PEO, that also contained salt grains as a porogen. This polymer mixture is only partially miscible and can be made to phase separate when heated. The resultant structure after soaking in water is of a microporous walled porous scaffold as both the PEO and salt grains are removed during the leaching process. Gomes et al. (2002) also extruded the same polymer mixture using blowing agents rather than water-soluble porogens. The blowing agents create pores within the matrix through degradation or volatisation at elevated temperatures and this overcomes potential problems associated with trapped porogen particles. The pore sizes reported by Gomes et al. are 10–500 μm for the moulded materials and 50–300 μm for the extruded materials. The porosities are similar, ranging from 40% to 50%. However, the key difference between these two processing routes is the degree of interconnectivity, which was deemed to be high for compression moulded material and poor for extruded material. Perhaps this observation highlights the importance of an adsorbed film of water in providing connectivity between hygroscopic porogen particles.

Super-critical fluid-based foams Gases such as carbon dioxide can be dissolved in polymers at high pressures to form fluids (Barry et al., 2004, Kim et al., 2006, Mooney et al., 1996). Controlled release of the pressure results in the formation of a porous structure. The pore sizes, pore distribution, heterogeneity of the structure and interconnectivity are obviously dependent on the processing conditions. However, it has been reported that the degree of interconnectivity is less than that produced by the methods described so far (Reigner and Huneault, 2006) and that there is a distinct skin–core effect, the porous core being encapsulated by a solid skin (Nam et al., 2000). Mooney et al., 1996 report that only 10–30% of the pores are interconnected. The main advantage of this approach is the benign nature of the carbon dioxide ‘solvent’ and the fact that it is carried out at low temperatures. This means that labile molecules, such as growth factors, and even cells, can be incorporated into the structure during manufacture. The gas foaming method has also been combined with particulate leaching (sodium chloride particles) to produce porous-walled porous PLGA and PLGA combined with nano-particulate hydroxyapatite scaffolds (Kim et al., 2006).

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11.3.2 Manufacture of fibres and fabrics Electrospinning The principle of this approach is to use a high voltage source to create an electrically charged jet of a polymer solution that is discharged from a fine tube or capillary. The solvent in the jet evaporates, leaving a charged nanofibre that can be simply collected as a random mesh on a grounded metal screen, Fig. 11.3 (Kim and Reneker, 1999). For the process to work, the force of the electric field must be greater than the surface tension of the polymer solution or melt. Electrical forces accelerate the resultant jet causing stretching of the fibre. Organic solvents with high vapour pressures evaporate during this time, which reduces both the diameter of the fibre and its velocity. Altering the processing conditions, e.g. the viscosity of the polymer solution, its conductivity (by adding salt), the voltage or field strength and the distance between the tip and collector produces fibres with diameters ranging from about 10 nm to 300 nm that can be kilometres in length (Frenot and Chronakis, 2003; Reneker and Chun, 1996). The fibres can be collected as random meshes (Kim et al., 2003) or formed into multifilament yarns by using a rotating collection ring, or wound around a rotating mandrel. Pham et al. (2006) and Murugan and Ramakrishna (2006, 2007) have reviewed the use of this technique in producing tissue scaffolds and the potential benefits of using nanoscale fibres in cell culture. These are thought to provide a better mimic of the structural features of extracellular matrix, facilitating cell adhesion.

650 μm 11.3 Fibrous meshes can be produced by techniques such as electrospinning (Kim and Reneker, 1999).

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Knits and weaves Industrial textile methods have been used to manufacture tissue scaffolds. Wang et al. (2005) created 3 mm diameter tubes from chitosan yarns (about 200 μm in diameter). These tubes were subsequently filled with chitosan solution poured around a matrix of acupuncture needles to form macrochannels that were used as conduits to guide neural growth. The purpose of the outer knitted sleeve was to enhance the mechanical rigidity of the structure: thin-walled tubes produced by other means, e.g. TIPS or dip moulding, are reported to collapse during transplantation (Yamaguchi et al., 2003; Itoh et al., 2003; Freier et al., 2005). Zhang et al., 2006 have followed a similar approach using knitted tube manufactured from chitosan yarn as the basis of a scaffold for use in blood vessel tissue engineering. Sahoo et al., 2006 have coated electrospun PLGA nano-fibres on to a knitted PLGA scaffold to create a large biomimetic surface for attaching bone marrow stromal cells with a view to creating tendon or ligament. Knitted fabrics appear to be better suited to tissue engineering than braided fabrics, which are reported to be susceptible to inadequate nutrient transmission, poor cell seeding, poor matrix production and may have insufficient mechanical strength (Karamuk et al., 1999; Van Eijk et al., 2004; Cooper et al., 2005). Non-woven meshes have also been used as scaffold materials (Aufderheide and Athanasiou, 2005) to provide an extended surface on which to culture cells, although their mechanical integrity is poor.

Solid free-form Translation of a computer-generated scaffold design to a real structure offers a degree of control and versatility that is not attainable by other methods. Various approaches have, and are being, developed, that include 3D printing (Lam et al., 2002; Dellinger et al., 2006; Park et al., 1998; Smith el al., 2007), laser sintering, modified ink-jet printing and fused deposition or solid preform fabrication (SFF). Co-printing of cells and scaffold material have also been explored, e.g. by Smith et al. (2007) and Kachurin et al. (2002). In principle, pores of any size, shape, interconnectivity and distribution can be produced, although currently the minimum pore size is relatively large at 50 μm. Unexpected shrinkage of the material, as well as uncontrollable surface texture imparted as a result of the processing method, can cause the detailed structure to differ from the intended design. This process has been used by Lam et al. (2002) to print a starch-based polymer with PCL, the mixture also contained poly(vinyl alcohol) powder. After melt processing the samples were immersed in water to remove both poly(vinyl alcohol) particles and starch, leaving a porous-walled, porous PLC scaffold.

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11.4

Characterisation of tissue scaffolds

Characterisation of tissue scaffolds is critically important to ensure consistency between samples and batches, as well as facilitating an understanding of the relationship between structure and performance. However, this is a very challenging topic (Tomlins et al., 2005). The pore sizes, especially in foam-like structures, might be expected to range from tens of nanometres to around one hundred micrometres. All pore sizes in this range can contribute to the performance of the structure as a tissue scaffold, as schematically indicated in Fig. 11.4. Given the number of different processing routes and the range of materials available, there is potential to create an almost unlimited number of different scaffold structures and geometries. The degradation behaviour of these materials can, and are, studied both in vitro and in vivo. In vitro, the effect of storing samples in an agitated or static environment can easily be explored. This is not the case with in vivo measurements, where the local environment can vary significantly depending on where the material is located. Developing suitable metrics to ensure comparability between different scaffolds produced within a batch or between batches is an area of continued research. A common approach is to determine the porosity of the scaffold. This is usually derived from a measurement of the density of the material with respect to that of a non-porous reference, such as a cast block of the same material. Porosity values can also be derived from image analysis, where the comparison is of the total pore area or total pore volume with respect to sample area or volume, respectively, Table 11.1. Porosity as a parameter is actually of limited value as it is simply a measure of the total

Frequency

Cell conduits?

Service conduits? 0.010

1

100

Length scale (mm) 11.4 Pore sizes ranging from nanometres to micrometres will contribute to scaffold performance by affecting nutrient and gaseous transport as well as cell movement. The form of this distribution has yet to be evaluated.

Table 11.1 Techniques used to investigate scaffold structures (redrawn from Tomlins et al., 2005) Generic approach

Generic technique

Information available

Specific examples

Image analysis

Microscopy

Pore shape, size and size distribution, porosity.

Scanning electron microscopy Transmission electron microscopy Optical microscopy Confocal microscopy Optical coherence tomography Optical coherence microscopy

Micro X-ray computer tomography Magnetic resonance imaging Terahertz imaging

Pore shape, size and size distribution, porosity. Pore shape, size and size distribution, porosity. Pore shape, size and size distribution, porosity.

Porosimetry

Porosity, total pore surface area, pore diameter, pore size distribution Median pore diameter (assuming cylindrical geometry), through-pore size distribution, permeability

Intrusion methods

Porometry

Mercury porosimetry

Capillary flow porometry

Direct measurement

Measurement of density

Porosity, pore volume

Buoyancy method Helium pycnometry

Molecular probes

Diffusion of markers

Permeability

Cyclic voltammetry Molecular diffusion

246

Degradation rate of bioresorbable materials

pore volume within the structure and does not provide any information on its distribution. Furthermore, porosity can be unduly influenced by defects such as one or more relatively large cavities within a sample, which commonly form, for example, during solvent casting followed by porogen leaching. Approximate pore dimensions are usually based on the analysis of scanning electron microscopy images (for example, Boccaccini et al., 2005; Neves et al., 2005; Kim et al., 2006; Maquet et al., 2003), although such data is also obtained from other imaging techniques including micro-computer tomography, magnetic resonance imaging and optical coherence tomography, Table 11.1. Depending on the technique and the complexity of the structure, pore dimensions are usually quoted as a mean diameter, with an often-large standard deviation, especially for data derived from scanning electron micrographs. The origin of this uncertainty reflects the difficulty in identifying wall boundaries and/or in using a single parameter to represent a range of non-uniform shapes. In some cases, two distinct populations of pore sizes have been identified. Boccaccini et al., 2005, for example, report the average dimensions of macropores within a scaffold to be about 100 μm; these coexist with interconnected micropores of 10–50 μm in diameter in PLGA foams doped with bioglass particles. Being able to quantify the size of pores is only part of the challenge. Porous materials, including tissue scaffolds, can contain three different pore types, as shown schematically in Fig. 11.5 (Jena and Gupta, 2002): (i) (ii)

(iii)

Enclosed pores that are isolated within a matrix. Open- or through-pores that connect to the outer surface of the scaffold and permit flow of liquid (or gas) from one side of a structure to the other. Blind-end pores that have contact with an exposed internal or external surface through a single orifice.

Through-pore

Enclosed pore Blind-end pore

11.5 Schematic representation of the different pore types that can exist within a tissue scaffold (after Jena and Gupta, 2002).

Influence of porous structure on bioresorbability

247

All of these pores have a role to play in tissue engineering; for example, closed pores can reduce the path length for diffusion of oxygen to the core of the scaffold, providing that it can pass through the wall of the scaffold. Through-pores are key to the function of the scaffold providing conduits for cell movement and solute diffusion. Blind-end pores have a less welldefined role, although their presence facilitates gaseous diffusion into the structure by reducing the path length. Similarly, depending on the material type and characteristics, the proportion of open, closed and blind-end pores can also influence the kinetics of degradation, an issue that is not addressed in the literature. The presence of these different pore types, whilst offering some advantages for cells, increases the complexity of the characterisation problem. Even simple measures of porosity based on density will be subject to more uncertainty if samples contain different combinations of these pore types, a difficulty that is exacerbated for tissue scaffolds manufactured from semicrystalline polymers, as the level of crystallinity within the scaffold and its homogeneity needs to be determined. Scaffolds can degrade through surface erosion, or from within, or by some combination of the two. In any case, the structure of the scaffold will have an influence on the rate of degradation through changes in both surface and cross-sectional area. The latter is particularly relevant for high surfacearea-to-volume ratio nanofibres that degrade by autocatalytic hydrolysis, as the hydrophilic oligomers produced will tend to diffuse out of the structure more rapidly compared to larger diameter structures.

11.5

Methods for monitoring the degradation of polymeric tissue scaffolds

In vivo assessment of the degree of degradation is typically assessed through histological examination (van Tienen et al., 2002). This approach is used to show evidence of fibrous encapsulation and tissue in-growth, and identify the presence of polymorphonuclear (PMN) leucocytes and/or macrophages. Repeated retrieval of implants followed by histological examination is used to track the fate of the scaffold and the process of tissue ingress. A practical problem with this approach is that any initial inflammatory responses induced by the implant are difficult to distinguish from those triggered by the surgery. Oh et al. (2006) have tracked changes in the molecular weight of implanted PLGA and PLGA/PVAlc scaffolds and found that the rate of degradation in vivo is faster than that in vitro. This phenomenon could be due to degradation being accelerated by foreign body giant cells, or by enzyme action (Coury et al., 1996), or due to a drop in the local pH initiated by the presence of acidic degradation products or by some combination of these factors.

248

Degradation rate of bioresorbable materials

The most commonly used methods for monitoring in vitro scaffold degradation are invasive and/or destructive methods such as: • • • • • •

Detecting changes in permeability Monitoring the time-dependent weight loss of samples Detecting changes in mechanical properties such as strain to failure Monitoring time-dependent changes in molecular weight Monitoring changes in geometry and morphology Monitoring the pH of the medium in which samples are immersed

These are discussed in more detail below.

11.5.1 Permeability Permeability measurements provide a useful measure of scaffold permeability that overcomes many of the complexities associated with quantifying pore size distributions and interconnectivity; however, it should be noted that the results are defect sensitive. Monitoring the flow of water through a scaffold in a given time can be used as a simple, quantifiable measure of scaffold permeability, providing it has sufficient mechanical integrity. This test is based on Darcy’s law and enables the intrinsic permeability to be determined (e.g. Lee et al., 2006). The test can be carried out using either a constant head of pressure or one that changes with time (Bowles, 1992). For a constant head of pressure, Darcy’s law states that the permeability, p of the material is given by p=

324.8Q (h ) ∗ 10 −10 r 2P

[11.1]

where Q is the flow rate through the material (ml.s−1), r is the radius of the implant, h is its height and P the height of the water column. Spain et al. (1998) have used this approach as a measure of quality control for PLGA based scaffolds. This non-destructive test is particularly useful for assessing scaffolds destined for in vivo measurements. Pêgo et al. (2003) have simply measured the flux of water through a triple stack of scaffolds discs (each disc was 25 mm in diameter and 4 mm thick). The water flux over a period of 30 minutes at a pressure of 0.1 bar was used to demonstrate the interconnectivity of the scaffolds, which for large pores was 30 ± 5 g/(min.cm2) compared with 47 ± 1 g/(min.cm2) for a single glass frit filter. Both approaches can be used as a measure of scaffold tortuosity and to assess the consistency of manufacturing processes.

Influence of porous structure on bioresorbability

249

11.5.2 Determination of mass loss The procedure for detecting changes in the mass of degradable polymers is straightforward. Samples are removed at specific intervals from the solution in which they are stored, dried under vacuum, or freeze dried, for a period of 24–48 h and then weighed (Spain et al., 1998; Wu and Ding, 2004). Changes in mass are rapidly detected in biodegradable polymers that decay by surface erosion, e.g. polyanhydrides and polyorthoesters (von Recum et al., 1995). The method is much less suitable for materials that degrade from within, such as the polyesters, where the mass remains essentially constant until the advanced stages of degradation are reached, as shown in Fig. 11.6. This phenomenon has been reported many times (e.g. Middleton and Tipton, 2000; Weir et al., 2004; Wu and Ding, 2004; van Tienen et al., 2002) and is a reflection of the fact that degradation occurs in the bulk of the polymer through repeated chain scission. Wahl et al. (2007) have studied the effects that different pore size distributions have on the degradation rate of collagen–hydroxyapatite scaffolds. Their results show that the rates are independent of changes in the pore size distribution, at 13.5%/hr and 16.4%/hr, respectively, in the presence of collagenase. Differences in the pore size distribution created during the manufacturing process were assessed by analysing SEM images. This analysis showed two distributions; one containing pore diameters that ranged from 50–400 μm and the other from 50–200 μm. The porosities in both cases are equivalent at 91.9 ± 0.2% as determined by mass and density relationships.

1.0

WIW0

0.8 0.6 0.4 = poly(D,L-lactide) = 85:15 PLGA = 75:25 PLGA

0.2 0.0 0

4

8 12 16 20 Degradation time (weeks)

24

28

11.6 The normalised mass of foam scaffolds, W/Wo, remains relatively constant until the advanced stages of degradation. Samples stored in PBS at 37°C (Wu and Ding, 2004).

250

Degradation rate of bioresorbable materials

11.5.3 Changes in mechanical properties Changes in the mechanical properties of degrading scaffolds are usually detected by measuring Young’s modulus in compression (Wu and Ding, 2004; Wang et al., 2006; Hedberg et al., 2005; Reignier and Huneault, 2006) or in tension (Weir et al., 2004), together with failure strain (Weir et al., 2004). Figure 11.7 shows changes in the compressive behaviour of PLGA (75 : 25) scaffolds as a function of degradation time (Wu and Ding, 2004). These tests, which were conducted on a 10 mm diameter cylinder, show that the modulus is maintained until some point at which the materials either yielded or fractured, the onset of yield occurring at stresses less than that required to produce 10% strain in less degraded material. In contrast, Hedberg et al. (2005) determined the compressive modulus and peak strength of PPF/PLGA porous foams and found no change in properties over a period of 26 weeks after immersion in PBS at 37°C (13 mm and 6.5 mm diameter foams). Other mechanical test methods have been used. Athanasiou et al. (1998) and Agarwal et al. (2000) have measured the creep of porous scaffolds induced by an indent with a tip diameter of 1.5 mm. The elastic modulus can be obtained from these measurements using the Boussinesq– Papkovitch equation. Wang et al. (2006) devised a suture pull-out test to assess the performance of knitted chitosan tubular structures. The structures were sutured between two nerve stumps and tested in tension until pull-out or epineuria occurred. The mechanical integrity of tubular structures can also be assessed by measuring their burst pressure. Zhang et al. (2006) reported a ∼50% drop in the average burst strength of chitosan/

0.5 (2 weeks)

σ(MPa)

0.4

σy

0.3 (8 weeks)

0.2 0.1 0.0 0.00

0.05

0.10

0.15

ε 11.7 Prolonged immersion of 75 : 25 PLGA porous scaffolds in PBS solution at 37°C causes the material to yield at a strain of about 0.055 (redrawn from Wu and Ding, 2004).

Influence of porous structure on bioresorbability

251

gelatin-coated chitosan tubes after a period of two months immersion in lyozyme solution. Comparable in vivo assessment of degradation showed roughly a 50% drop in burst strength, compared with non-degraded material. They concluded that the difference between in vitro and in vivo behaviour is due to the uncontrolled in vivo conditions and the in-growth of fibrous tissue. It can be argued that such measures are actually more useful than in vitro data since they represent actual in vivo performance.

11.5.4 Changes in molecular weight Monitoring time-dependent changes in molecular weight is recognised as a particularly sensitive measure of degradation. A number of mathematical models have been proposed to describe the continual reduction in the number average molecular weight, Mn that occurs due to hydrolysis. Anderson (1995) derived the following relationship: 1 1 = + k it M nt M no

[11.2]

where Mnt is the number average molecular weight at time t, Mno is the number average molecular weight at t = 0, and ki is a rate constant. This relationship is based on a statistical approach and assumes that there is no autocatalytic contribution to the degradation rate. Pitt and Gu (1987) have addressed this issue and shown that M ntt = M no e − k2t

[11.3]

This relationship accommodates the autocatalytic reaction of carboxylic acid endgroups that results from the hydrolysis of the ester, but assumes a constant concentration of ester and water and that the concentration of carboxyl end groups is given by 1/Mn. This relationship has been used to successfully model the degradation behaviour of solid compression moulded bars of PLLA to the point of mass loss (Weir et al., 2004). The molecular weight distribution of polymers is usually determined by using gas permeation chromatography. Wu and Ding (2004) found that the number and weight average molecular weight of PLGA scaffolds decreased exponentially with time through the degradation period, Fig. 11.8. The slope of the semi-log plot shown in Fig. 11.8 increases with increasing glycolic acid content, which Park et al. (1998) attributed to the increased hydrophilicity of the lower LA/GA molar ratios. Chain scission in polyesters continually reduces the molecular weight to a point where the oligomers become hydrophilic and are able to leach out of the polymer into the surrounding medium. The rate at which the oligomers leave the polymer will depend on the surface-area-to-volume ratio,

252

Degradation rate of bioresorbable materials = poly(D,L-lactide) = 85:15 PLGA

106

= 75:25 PLGA

Mn (g·mol–1)

105

104

103

0

4

8

12

16

20

24

28

0

4

8 12 16 20 Degradation time (weeks)

24

28

106

Mw (g·mol–1)

105

104

103

11.8 The number (Mn) and weight (Mw) average molecular weights for porous scaffolds manufactured from different copolymer compositions decrease with time during immersion in PBS at 37°C (redrawn from Wu and Ding, 2004).

with nanofibres offering the fastest route (Li, 1999; Li and McCarthy, 1999; Kim et al., 2003). Shah et al. (2006) have experimented using multicomponent blends of PLA with triblock co-polymers of PLA-b-PEG-b-PLA and PLGA to retard diffusion of the more degradable polymers. The function of the triblock copolymer is to retard diffusion of PLA. This system can be tuned to begin degrading, as monitored by changes in mass loss after a period of only a few days. The fibre diameter was reported as 500–800 nm with porosity measured by mercury porosimetry to be 70–75%.

Influence of porous structure on bioresorbability

253

11.5.5 Monitoring changes in pH Monitoring changes in the pH of the PBS or saline solution is commonly used to track the degradation of materials such as PLGA. The measured change in pH depends on how often the medium is changed, an action that can induce exchange of medium within the scaffold itself. Oh et al. (2006) explored the effect of this variable by comparing the pH of PBS changed on a daily basis over a period of 70 days with samples where the PBS was not changed over the same period of time. The pH remained stable at 7.4 when the buffer was changed daily but fell significantly after 14 days in the unchanged media. The PLGA (50 : 50) scaffold degraded faster in the low pH environment. In other studies, samples were stored in a large volume of buffer. For example, Holy et al. (1999) stored samples in buffer at a foam : buffer ratio of 1 : 1000 (w/v), which ensured that the pH remained constant throughout their degradation experiment.

11.5.6 Specimen geometry and morphology Changes in sample geometry and/or morphology have been used to assess scaffold degradation with differing degrees of success. The fundamental issue of how to quantify the complex, irregular structures that are typical of tissue scaffolds means that images obtained at different points along the degradation timeline can be only qualitatively compared. Measuring changes in the geometry of samples offers a potentially better route to quantifying the effects of degradation, but for materials such as PLGA, the water absorbance that occurs as degradation progresses can potentially complicate matters (e.g. Peter et al., 1997). Athanasiou et al. (1998), using a PLGA (50 : 50) copolymer scaffold prepared by the precipitation method, examined cylindrical samples that had porosities of 0%, 33 ± 8% and 75 ± 4%, respectively. After two weeks immersion in PBS at 37°C, they observed the presence of a small number of circular pores, approximately 300 μm in diameter in all three-sample types. These circular pores appeared to be a measure of degradation and contrasted with the irregularly shaped 20–500 μm equivalent diameter pores produced during the manufacturing process. After four weeks, the initially non-porous material collapsed whilst the porous materials maintained their form for more than six weeks. Others have also observed the appearance of small holes in the walls of PLGA (50 : 50) scaffolds, e.g. Oh et al. (2006). Yoon and Park (2001) report that PLGA 75 : 25 scaffolds degraded over a period of ten weeks showed no sign of physical degradation. In contrast, PLGA 50 : 50 started to physically degrade at day 21 and PLGA 65 : 35 at

254

Degradation rate of bioresorbable materials

day 63. The gradual collapse of the 75 : 25 PLGA scaffold continued from day 21 to day 49, beyond which the scaffold began to swell. Similar swelling behaviour occurred in 65 : 35 scaffolds beyond day 21. Yoon and Park (2001), postulate that the appearance of swelling or not is linked with hydrolytic degradation of the PLGA and water uptake which leads to swelling, the latter happening only after the degradation process has begun. The swelling phenomenon has been reported by others, e.g. Oh et al. (2006), Azevedo and co-workers (2003) have proposed that weight loss of material and water uptake provide a useful measure of degradation in PCL/HaP (hydroxyapatite) composites. This can be determined by thermogravimetric analysis. Baji et al. (2006) have shown that this approach can reveal significant differences in the degradation behaviour of blended and grafted compositions. Hedberg et al. (2005) were unable to detect any changes in the porosity of PPF/PLGA foams as measured by MicroCT (Micro computer tomography) and mercury porosimetry during a period of 26 weeks’ immersion in PBS at 37°C. They also report minimal changes (<6%) in the size of cylindrical samples (H = l3 mm, d = 6 mm) over the same time period. However, Agarwal et al. (2000) report that the porosity of PLGA (50 : 50) decreased by more than 10% over a period of two weeks and thereafter increased over the following month to a value in excess of the original porosity. Comparisons between different publications are not easily made, due to differences in material type, processing methods, structural detail, sample geometry and storage protocol. Table 11.2 lists some examples of the different strategies used with respect to changing buffer solutions during in vitro degradation experiments. The key factors are the buffer volume with respect to the sample geometry and whether the system is agitated or not during storage. Both these factors will influence the degradation behaviour and any comparisons of pH measurement made between different studies.

11.5.7 Sample conditioning Many of the biodegradable scaffolds being considered for tissue engineering applications are manufactured from relatively hydrophobic materials. Immersing these materials in buffer does not guarantee complete wetting of the structure, a problem that can influence measurement of porosity using methods based on Archimedes’ principle. A number of different strategies have been employed to overcome this issue: Wu and Ding (2004) proposed a method for pre-wetting scaffolds that is based on subjecting the specimen to alternating vacuum/liquid immersion. Lu et al. (2000) and Kuo and Leou (2006) have used ethanol to pre-wet porous foams. The solubility

Table 11.2 Different authors use different protocols for storing samples for in vitro degradation studies Material

Sample geometry/ characteristics

Frequency of buffer change

Reference

Poly(1,3 trimethylene carbonate) + DL-lactide Polym TMC copolymer

Cylinder H = 14 mm D = 8.4 mm

Once per month for the first three months, thereafter every two weeks.

Pego et al., 2003

PLGA (50 : 50)

Cylinder H = 5 mm D = 7 mm

Every 3.5 days for eight weeks

Agrawal et al., 2000

Cylinder H = 13 mm D = 6.5 mm

Every 8 hr for the first 48 hr then daily for the next 5 days. Weekly for the next 25 weeks

Hedberg et al., 2005

PLGA

Disc H = 2 or 5 mm D = 10 mm

Daily for the first week, then at day 10 and 14. For the remainder of the 10-week experiment the PBS was changed weekly.

Yoon and Park, 2000

PLGA

Disc

Solution not changed throughout 8-week period of experiment.

Spain et al., 1998

PLGA (75 : 25)

H = 10 mm Base area 0.4 cm2

Samples stored in excess buffer that was changed every two weeks for 22 weeks

Holy et al., 1999

PPF/Tricalcium phosphate

Cylinder H = 12 mm D = 6 mm

PBS changed every 8 hr for the first 2 days, daily until day 7 then weekly for remaining time frame

Peter et al., 1997

256

Degradation rate of bioresorbable materials

of ethanol in water makes subsequent wetting much easier. A downside of this approach is that ethanol can induce crazing or plasticise materials and therefore alter the degradation behaviour. Absorbed water in polymers can act as a plasticizer, causing a change in mechanical properties, enhanced ductility and a reduction in the glass transition temperature, parameters that are also used as a measure of degradation due to molecular weight changes. Siemann (1985), for example, has reported a decrease of 12°C in Tg after soaking in water for 6 hours. The immersion time is not important per se as the percentage of water absorbed will depend on the surface area/volume ratio of the sample as well on as the type of material. In that sense, nanofibre structures are more likely to reach early equilibrium than a large cylindrical sample that has dimensions of, for example, H = 10 mm and D = 8 mm, assuming that it is fully wetted out. The presence of a plasticising agent is likely to mask changes in the mechanical performance of materials that are due to changes in the molecular weight of the polymers and, as discussed above, will influence measurement of dimensional changes.

11.5.8 Modelling degradation Mohammadi and Jabbari (2006), among others, have used Monte Carlo methods to simulate the degradation of poly(lactide), PLA scaffolds. The acidic environment generated within the scaffold through degradation of PLA accelerates the rate of degradation compared with material located at the surface. In their model, Mohammadi and Jabbari assumed that there is a linear decrease in molecular weight with time, each chain forming two new ones per scission event (Gonzalez et al., 1999; Hakkarainen et al., 1996). This process continues until the chain length reaches five or fewer monomer units, at which point the oligomers become water soluble and therefore are able to leach out of the matrix (Gonzalez et al., 1999; Hakkarainen et al., 1996). Mohammadi and Jabbari’s simulations show that hollow structures form inside the scaffolds, over a period of time that depends on their initial porosity. The time period increases with decreasing porosity. They also note that the kinetics of this process are highly dependent on the specimen and experimental conditions, as described in the introduction to this chapter. Their model also shows that diffusion is not sufficient to remove the acidic oligomers from the porous structure. This has an implication on the degradation behaviour of cell-seeded scaffolds, where release of oligomers from the scaffold will be impeded by the presence of cells and adsorbed proteins, as well as those measurements made on mechanically stimulated scaffolds or cultures in stirred bioreactors, or the response to local in vivo conditions.

Influence of porous structure on bioresorbability

11.6

257

Conclusion

Assessing the impact that scaffold structures have on the degradation behaviour of polymers is not an easy task given the number of variables, e.g. • • • • • • • •

Pore size and interconnectivity Sample geometry Proportions of different pore types – open, closed or blind-end Storage conditions Starting molecular weight of the material Material type The processing conditions used to manufacture the scaffold The metrics used to assess degradation.

The balance between ‘internal’ degradation and surface erosion is determined by the active degradation mechanisms, i.e. hydrolyis, particularly autocatalytic reactions and surface-active enzyme activity. Characterisation of the initial structure and sample storage conditions play an important role in influencing degradation kinetics but are typically overlooked in many studies. This is particularly true for characterisation, as there are no robust protocols and/or methods for analysing porous materials over a lengthscale that ranges from nanometre to sub-millimetre. For commonly used polyesters, the fastest rate of degradation occurs in low porosity materials due to autocatalysis or those that contain high levels of lactide.

11.7

Acknowledgement

This work was funded by the UK Department of Trade and Industry’s ‘Measurement for Emerging Technologies’ programme, Project 1.3.

11.8

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12 Influence of clinical application on bioresorbability: Host response J. C. Y. CHAN,1,2 K. BURUGAPALLI,1 J. L. KELLY,2 and A. S. PANDIT,1 1National University of Ireland, Galway, Republic of Ireland. 2University Hospital Galway, Republic of Ireland

Abstract: This chapter begins by reviewing the host response pathways initiated by the implantation of a biomaterial and the mechanisms involved in restoring tissue stability and function. It then discusses both the host factors and material factors influencing the biodegradation of degradable biomaterials. In addition, the key mechanisms of in vivo biodegradation involved in the process of biomaterial elimination from the host are also presented. Key words: host response, biomaterial implantation, host factors in biodegradation, material factors in biodegradation, mechanisms of in vivo biodegradation.

12.1

Introduction

The biodegradation of biomaterials introduces more complexities and challenges in the host when compared to non-degradable biomaterials. Until the biomaterial is fully degraded and the implant area remodelled, the host has to respond to the challenges of a changing environment, brought about by the alteration in size, shape and surface topography of the degrading biomaterial. There is also the constant release of degradation products that need timely elimination. The resulting biomaterial particles or partially degraded polymers represent additional stimuli for host response. While these issues may produce needless difficulties for the intended application, degradable biomaterials have numerous advantages and eliminate the problem of chronic irritation to the host, as well as the potential for tissue remodelling and regeneration. Hence, biodegradable materials have gained wide recognition and hold promise for many biomedical and tissue engineering applications as temporary scaffolds for tissue replacement, or as delivery vehicles for cells, drugs, bioactive molecules or genes. 267

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The rate, extent and mechanism of biodegradation of a foreign material (biomaterial) are dependent on a wide range of host and biomaterial factors. This chapter primarily reviews the host factors influencing the biodegradation of biomaterials. The general mechanisms of in vivo biodegradation and elimination of biomaterials are also discussed. In addition, special consideration is given to biomaterial factors because they have significant influence on host response pathways and these factors are modifiable by the biomaterial designer.

12.2

Host response cascade

Surgical implantation (including minimally invasive injection) of a biomaterial invariably results in physical injury to the host. A host responds to such injury through the natural mechanisms of wound healing that involves haemostasis, inflammation, proliferation and maturation (fibroplasia/ regeneration). These mechanisms aim towards the restoration of tissue stability and original function. However, the presence of an implanted foreign material alters the normal healing process. In response, the host attacks and tries to eliminate the implanted foreign material from its system. The resulting host response pathways are complex and represent intricately regulated cellular and biochemical events. These host response pathways generated in response to the implantation of a biomaterial, referred to as host response cascade, are summarised in Fig. 12.1.

12.2.1 Haemostasis Haemostasis is the first phase in the host response cascade, where the damage to local vascular tissue (haemorrhage) caused by the implantation of a biomaterial triggers blood coagulation, initiating the wound healing process, resulting in the formation of a platelet plug and provisional fibrin matrix. These arrest haemorrhage at the site of the injury and fill the injured space or tissue deficit caused by the injury. This further mediates the subsequent inflammatory and wound healing responses (Steinhubl, 2007).

12.2.2 Inflammatory response The inflammatory phase ensues after injury and provides the immune system access to the injured site. Irrespective of the end result, the initial response is termed acute inflammation, which is characterised by oedema and inflammatory cells infiltration. The primary role of the acute inflammatory process is to protect the injured site from infection by mounting a rapid but relatively non-specific phagocytic response involving neutrophils. The neutrophils engulf and degrade infectious agents, cell debris and low molecular leachable molecules from implanted biomaterial.

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Biomaterial implantation

Tissue injury Wound healing response Haemostasis Protein adsorption Water absorption Implant degradation Inflammatory response

Acute inflammation Neutrophils

Persistent injury

Chronic inflammation

Resolution of injury

Lymphocytes Monocytes differentiate to macrophages

Proliferation of fibroblasts and endothelial cells Proliferation Fibroblasts Macrophages fusion to form FBGCs Frustrated phagocytosis Fibrous capsule formation Macrophages and FBGCs Biomaterial encapsulated

Secretion of ECM by fibroblasts and capillary formation Implant site remodelling Host tissue in-growth Implant degradation and resorption

Host tissue replacement/ regeneration

12.1 Host response cascade in response to the insertion of a biomaterial.

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If the inflammatory stimulus (biomaterial) is removed, healing will take place. However, if the stimulus persists or when the implant causes persistent mechanical injury, acute inflammation progresses to chronic inflammation. During chronic inflammation, the host attains a balance between continuing inflammatory stimulus (injury) and tissue repair. Macrophages, giant cells and lymphocytes are the predominant cells of the chronic phase and they play a critical role in the biodegradation of implanted biomaterials and also in the mediation of host responses via soluble factors and cytokines (Burugapalli and Pandit, 2007; Liang et al., 2004).

12.2.3 Proliferative phase The transition from inflammatory response to healing response is characterized by the decrease in inflammatory cell influx and the proliferation of fibroblast and vascular endothelial cells. This proliferative phase is characterized by modification and reorganization of the provisional matrix (initially laid down by the coagulation cascade) to form granulation tissue by fibroblasts (Raghow, 1994). Fibroblasts synthesise new collagen and other extracellular matrix components such as glycosaminoglycans, fibronectin and tenascin at the healing site to replace the fibrin/clot matrix and the injured tissue. In turn, the new granulating matrix produced by activated fibroblasts assists the migration of inflammatory cells, fibroblasts and endothelial cells to form new capillaries (angiogenesis) and strengthening the healing site. The new blood capillaries provide the necessary nutrients and oxygen for cellular activity. The resulting granulation tissue is an organised host tissue that integrates the implanted biomaterial with the host tissue.

12.2.4 Fibroplasia/remodelling Depending on the nature and location of the injury, the type of tissue injured and the age and health of the host, two general outcomes are possible after an injury. The first and the most common outcome is that the injured tissue/organ fails to regenerate, resulting in the formation of fibrous scar tissue (replacement) (Williams and Harding, 2003; Witte and Barbul, 2003). Secondly, resolution can occur, that is, the restoration of normal tissue architecture and tissue function (regeneration). Regeneration is usually observed in foetal tissue and rarely in adult tissues. In either case, the granulation tissue formed in the proliferative phase is remodelled and the degrading biomaterial is progressively replaced by host tissue. The final outcome of the host response cascade is either degradation of the implanted biomaterial and formation of a stable fibrous scar, partial regeneration of the host tissue or regeneration of a complete and functional tissue or organ.

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12.3

271

Host factors influencing biodegradation

The process of in vivo biodegradation and replacement of the implanted biomaterials follows a complex orchestration of interdependent events involving a wide range of soluble and cellular components in a tightly regulated environment. Each tightly regulated event can contribute significantly to the passive and/or active biodegradation of the implanted biomaterial. The non-cellular and cellular factors influencing biodegradation are summarised in Table 12.1 and Table 12.2, respectively.

Table 12.1 Non-cellular factors influencing biodegradability Non-cellular factors

Role in host response cascade

Exudates

• Platelet plug and thrombus arrest haemorrhage • Complement activation • Dilution of toxic chemicals • Diffusion of cytokines, chemokines and growth factors • Protein adsorption (opsonisation) on biomaterial • Wetting of biomaterial • Serum proteins, electrolytes and water mediate hydrolysis • Drainage of low molecular weight (degradation) materials and debris to lymphatic and blood circulation • Anaerobic metabolism • Lactic acid accumulation • Acceleration of hydrolysis by acidic pH and elevated temperature • Initiation of angiogenesis • Provisional fibrin matrix fills tissue deficit • Fibronectin and hyaluronic acid incorporated in fibrin matrix facilitates inflammatory cells migration and new blood vessel formation • Fibrin acts as a chemotactic factor • Fibrinogen, immunoglobulin G (IgG) and complement protein (C3b) coat biomaterial surface and other foreign materials • Coating predisposes foreign material for cellular interactions and phagocytosis • Control chemotaxis, growth, differentiation, regulation and cell trafficking • Orchestrate host response cascade • Provides a framework for cell adhesion, migration and interactions • Binds cytokines and growth factors and modulate cellular responses

Hypoxia, hypoglycaemia, acidic pH, elevated temperature

Coagulation

Opsonisation

Cytokines

Extracellular matrix

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Table 12.2 Cellular factors influencing biodegradability Cellular factors

Role in host response cascade

Platelets

•

Phagocytosis

•

Activated platelets crosslink in the presence of fibrinogen and thrombin to form platelet plug • First cells to release large quantities of cytokines • Initiate chemotaxis (cell recruitment from blood and surrounding tissue)

• • • • •

Neutrophils

• • • •

Phagocytes recognise opsonins and antigens on the surface of foreign materials Endocytosis by phagocytes isolates foreign particles within vacuoles called phagosomes Lysosome fuses with a phagosome to form phagolysosome Foreign materials are digested by enzymatic and oxidative degradation pathways within phagolysosomes Low molecular degradation products are usually exocytosed for eventual drainage through lymphatic and blood circulatory systems Antigen presenting cells (APCs) present degradation products on their cell membranes as antigens in association with MHC Class II molecules for recognition by T-helper lymphocytes First inflammatory phagocytes recruited to the injury site Cleanse the injury site of debris and foreign materials Short-lived cells, incapable of mitosis After phagocytosis, becomes apoptotic

Monocytes

•

Circulating monocytes once recruited into tissues will differentiate into macrophages

Macrophages

•

Active phagocytosis to clear the injury site of debris, foreign pathogen and particles Release cytokines that regulate host responses and immune cell trafficking Long-lived cells – often last as long as the implant exists in the host A subset differentiate into APCs Function as APCs by phagocytosing materials and migrating to lymph nodes, where they present the intracellularly processed antigens to T-helper lymphocytes

• • • •

Frustrated phagocytosis

• •

Failure of macrophages to degrade nonphagocytosable or large particles (>10 µm) (including non-degradable materials) Release of enzymes, acid and reactive oxygen species into the extracellular environment

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Table 12.2 Continued Cellular factors

Role in host response cascade

Giant cells

• •

Fibroblasts

• • •

Macrophages fuse to form foreign body giant cells (FBGC) More aggressive than macrophages in phagocytosis and capable of degrading larger biomaterial particles Proliferate at the injury site Deposit extracellular matrix (granulation tissue) in place of damaged or excised tissue to fill the tissue deficit and to replace degrading biomaterial Remodel provisional matrix and granulation tissue

Endothelial cells

•

Proliferate and form new blood vessels in response to angiogenic factors such as PDGF, VEGF and NO.

Lymphocytes and plasma cells

•

Modulate immune-related and general host response cascade by release of regulatory cytokines T-helper lymphocytes are the key cells in the development of adaptive immunity against foreign materials Activated B-lymphocytes differentiate into plasma cells Plasma cells are antibody producing cells Function as APCs by processing and presenting phagocytosed foreign materials, and activate costimulatory pathways with T-helper lymphocytes to develop antibodies against antigens derived from foreign materials (including biomaterials)

• • • •

12.3.1 Physical, chemical and non-cellular factors Exudates The injury created by the implantation of a biomaterial results in two types of exudates at the implant/injury site. Firstly, the haemorrhage exudates appear as a direct consequence of the damage to local vascular network. These exudates are composed of cells, plasma proteins and growth factors derived from the circulation. The injury activates platelets, coagulation, fibrinolytic, kinin and complement systems, facilitating the formation of thrombus from haemorrhage exudates. Thrombus formation helps to arrest haemorrhage. Secondly, the acute inflammatory exudates appear in response to dilation of blood vessels during acute inflammation, bringing in fluids, electrolytes and serum proteins (e.g. immunoglobulins, fibrinogen) to the injured site. The serum proteins usually adsorb on the biomaterial and

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mediate cellular reaction to the biomaterial. The increased permeability of endothelial walls also facilitates the migration of inflammatory cells such as neutrophils, monocytes and lymphocytes into the injured site. The complex composition of fluids, electrolytes, plasma proteins and cellular components is collectively known as the acute inflammatory exudate. The presence of increased fluid and electrolytes is thought to buffer or dilute any toxins at the injured tissue. Soluble chemical mediators released from inflammatory cells, plasma, and injured tissue diffuse into the exudate and orchestrate the acute inflammation and subsequent wound healing. The exudates are continuously absorbed through the lymphatic system and this delivers antigens to regional lymph nodes that may trigger specific immunity to foreign material, including the implanted biomaterial.

Hypoxia, hypoglycaemia, acidic pH and elevated temperature Damage to local vascular network reduces vascular perfusion resulting in hypoxia (low oxygen levels) and hypoglycaemia (low glucose levels) at the injury site. Hypoxia and hypoglycaemia, in turn, cause active cells at the injury site to switch to anaerobic metabolic pathways, resulting in an acidic tissue environment due to the accumulation of lactic acid (Williams and Harding, 2003). In addition, acute inflammation is often associated with elevated local temperature at the site of inflammation or systemically (e.g. fever, sepsis). Acidic pH, elevated temperature in association with water, electrolytes and other components of the haemorrhage and acute inflammatory exudates initiate the hydrolysis of labile chemical linkages of polymer chains of biodegradable implants. For instance, the reduction of pH in the microenvironment of collagen and the secretion of cathepsin from macrophages and neutrophils create an acidic pathway for collagen breakdown (Etherington et al., 1981; Burleigh et al., 1974). Dung et al. (1994) have demonstrated that pre-treating collagen with acid increases its susceptibility to other proteolytic enzymes.

Coagulation The coagulation pathways that promote haemostasis and wound healing are activated by several factors in haemorrhagic exudates. The intrinsic and extrinsic coagulation pathways convert fibrinogen to fibrin to form a clot matrix. Fibronectin, a cell adhesion molecule, and hyaluronic acid are incorporated in the clot framework (Williams and Harding, 2003), and can facilitate the attachment and migration of inflammatory cells and new blood capillary formation. This aids direct contact between the implanted biomaterial and cells or other soluble factors (Anderson, 2001).

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Opsonisation Opsonisation is the coating of the biomaterial or any other foreign material with soluble factors in the haemorrhage and inflammatory exudates. The opsonins include fibrinogen, immunoglobulin G and complement C3b fragments. The complement C3b fragment is a major opsonin that is released as a result of complement cascade activation by blood coagulation. The complement system comprises a series of circulating inactive proteins that release C3b, taking only minutes to coat materials or microorganisms (Blondin et al., 1996). The binding of opsonins to the surface of biomaterial labels the biomaterial or particle as non-host and hence, targets it for attack by phagocytes (Kopp et al., 2002; Tang et al., 1998a; Hunt, 2004).

Cytokines, chemokines and growth factors Cytokines, chemokines and growth factors (GF) are a group of multifaceted low molecular weight (<20 kDa) proteins secreted by a variety of cells in the host. They mediate, regulate and control a broad spectrum of fundamental cellular functions including growth, differentiation, activation and trafficking of cells. The overall effect on cells is determined by the action of these cytokine combinations. Hence, they play an important role in the host response and biodegradation of an implanted biomaterial via their action on the host cells. Cytokines can be divided into interleukins (IL), colony-stimulating factors (CSF), interferons (IF), chemokines, tumour necrosis factor (TNF) and they also include growth factors. The main function of chemokines (CC, C, CXC and CX3C) is to attract/recruit cells from the circulation to the site where they are secreted. Growth factors (IGF, VEGF, TGF; etc.) function by stimulating cellular proliferation and/or differentiation. Some other cytokines have mainly regulatory roles. Cytokines are usually secreted at very low concentrations, have short life spans and act over short distances. The same cytokine can be synthesised by different cell types and a single cytokine can act on different cell types. They can act on cells that secrete them (autocrine effect), on nearby cells (paracrine effect) or on distant cells (endocrine effect). Cytokines can act either synergistically (two or more cytokines acting together) or antagonistically (cytokines causing opposing effects) to orchestrate cellular responses in the host. In the following sections, the role of relevant cytokines will be discussed only briefly when considering host and biomaterial factors that influence biodegradation. Elaborate reviews on cytokines and their role in the immune system are available in detail elsewhere in the literature (Borish and Steinke, 2003; Rollins, 1997).

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Extracellular matrix Extracellular matrix (ECM) is the dynamic network of protein and carbohydrate components that fills the intercellular spaces. The ECM provides a mechanical framework that distributes the stresses of movement and gravity while maintaining the structural integrity of tissues and organs. The cells of the tissues adhere to components of the ECM and are dependent on ECM for water, ions, nutrients and motility. The ECM also plays an important part by acting as a medium through which chemical messengers (e.g. cytokines, chemokines and growth factors) and other molecules (e.g. O2 and CO2) can freely diffuse. The ECM has key roles in the orchestration of the host response cascade. ECM provides the geometric framework that facilitates cell migration and modulates cell–cell interactions, in addition to cell–surface receptor and integrin interactions for cell–cell and cell–ECM adhesion. ECM also binds specific cytokines and growth factors that modulate the cellular behaviour and phenotype at the healing site (Raghow, 1994). The dynamic cell–ECM interactions and cytokines constantly modify and regulate the injury site during tissue remodelling and regeneration.

12.3.2 Cellular factors influencing biodegradation Platelets Exposure of circulating platelets to collagen from the subendothelial/injured tissue extracellular matrix results in the binding of platelets to collagen, mediated by von Willebrand Factor (vWF, a blood glycoprotein) (Schmugge et al., 2003; Fressinaud et al., 1987). This binding causes platelet activation, leading to the secretion of their cytoplasmic granules and lysosomal contents. The activated platelets are the first cells to release large amounts of cytokines, including platelet factor 4 (PF4), platelet-derived growth factor (PDGF), transforming growth factor-α (TGF-α), transforming growth factor-β1 (TGF-β1), platelet-derived endothelial growth factor, 5hydroxytryptamine (5-HT) and adenosine diphosphate. These cytokines initiate the host response cascade (involving coagulation, haemostasis, chemotaxis and increased vascular permeability), facilitating the recruitment of inflammatory cells to the injury site to attack or degrade foreign materials and to participate in wound healing (Williams and Harding, 2003; Zarbock et al., 2007; Gawaz et al., 2005; Anderson, 2001). Phagocytosis Phagocytosis involves the engulfing and subsequent digestion of foreign materials (e.g. tissue debris and degrading biomaterial components) by

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phagocytic cells such as neutrophils and macrophages (Hunt, 2004; Anderson, 2001). The phagocytes recognise antigens or opsonins (plasma proteins that coat foreign materials) on the surface of the foreign materials/particles. They push out pseudopodia to envelop the particle, internalizing the particle in a vacuole called a phagosome. The phagosome fuses with a lysosome to expose the isolated foreign particle to enzymatic and oxidative degradation by a cocktail of lysosomal enzymes and oxidative agents. This cocktail of harmful agents in the phagolysosome (fused phagosome and lysosome) acts to digest and destroy the foreign material. Phagocytosis is the main mechanism by which the host breaks down and removes the bulk of implanted biomaterial components.

Neutrophils Neutrophils are the first inflammatory cells recruited to the injury site and dominate the acute inflammatory phase of the host response cascade. The main function of neutrophils is to remove foreign materials or pathogen from the injury site. They employ both intracellular and extracellular mechanisms of attack to destroy, inactivate or digest these foreign agents. Foreign materials with recognisable and/or opsonised surface features enhance neutrophil binding to these materials. The binding, in turn, activates phagocytosis that results in intracellular digestion of phagocytosed materials. However, materials with non-recognisable and non-opsonisable surfaces, as well as large-sized particles avoid phagocytosis by neutrophils. The non-phagocytosable materials are attacked by the extracellular release of enzymes and oxidizing agents from neutrophils. Both phagocytosis and extracellular release of enzymes and oxidative agents (e.g. reactive oxygen and free radicals) have significant implications on the biodegradation of implanted biomaterials. Neutrophils are short lived (hours to a few days) and are incapable of mitosis. Post-phagocytosis, the neutrophils become apoptotic. The dead neutrophils themselves become cell debris which are, in turn, cleared by macrophages. The recruitment of neutrophils from the circulation is mediated by a wide variety of cytokines secreted by numerous cells, including endothelial and epithelial cells, macrophages, monocytes, lymphocytes, platelet, neutrophils and parenchymal cells. The cytokines that recruit neutrophils include tumour necrosis factor-α (TNF-α), IL-1, IL-6, interferon-γ (IFN-γ), plateletactivating factor (PAF), leukotriene B4 (LTB4), complement protein C5a, IL-8, macrophage inflammatory protein-2 (MIP-2) and granulocytecolony stimulating factor (G-CSF) (Wagner and Roth, 2000; Kullberg et al., 1999).

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Monocytes Monocytes are the circulating cells that migrate to the injured site and differentiate into tissue macrophages. Their recruitment is activated by monocyte chemotactic protein (MCP-1). MCP-1 is reported to be produced by vascular smooth muscles cells in response to acute mechanical injury, and the number of monocytes infiltrating the injured tissue is directly correlated with the resulting macrophage numbers at the injury site (Schober and Weber, 2005; Muller, 2001). Macrophages Macrophages are the predominant cell type in the chronic inflammatory phase of the host response. They exist in small numbers during the early acute inflammatory response, but increase in numbers over time. Resident macrophages are the first to detect the presence of a foreign particle in tissues. There is also a population of tissue monocytes that can move to the injury site and become macrophages. The number of tissue macrophages is usually insufficient to remove the foreign particle. Therefore, they initiate a response to recruit additional phagocytes from the circulation into the local area. Circulating monocytes enter the tissue from the circulation and differentiate into macrophages. The primary role of macrophages is to remove debris, foreign pathogen and particles (including implanted biomaterials). They also release potent cytokines and hydrolytic enzymes, whose actions play a significant role in immune cell trafficking and the formation and remodelling of granulation tissue at the injury site (Raghow, 1994; Anderson, 2001; Park and Barbul, 2004; Bogdan and Nathan, 1993; Jean-Baptiste, 2007). In addition, a subset of macrophages, called antigen presenting cells (APC) can phagocytose foreign materials, digest them and present the digested fragments (antigens) on their cell membranes in order to sensitise the immune cells (lymphocytes) to the foreign material. Phagocytic activity of macrophages is similar to that of neutrophils, but macrophages are larger in size, live longer (due to their capacity for oxidative metabolism) and usually last in the wound area until the conclusion of the inflammatory and wound healing responses. Similar to neutrophils, activated macrophages are capable of producing reactive oxygen radicals that are involved in bacterial killing and damage to biomaterials. Enzymes such as proteases and hydrolases are also secreted to digest the damaged extracellular matrix, and may play a role in the breakdown of biomaterials. Macrophages also play a major part in the orchestration of the wound healing process by secretion of cytokines and growth factors. Platelet activating factor, interleukins and arachidonic acid metabolites mediate the

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acute inflammatory response. These soluble chemical factors function to regulate the macrophages own activities, as well as other inflammatory and host cells such as lymphocytes, fibroblasts, osteoblasts, osteoclasts and foreign body giant cells. In addition, secreted growth factors (e.g. PDGF, EGF, FGF) have important roles in blood capillary formation and cellular function of fibroblasts. Frustrated phagocytosis In general, particles greater than 10 μm are not phagocytosable by either neutrophils or macrophages. Failure to phagocytose large particles or biomaterials may result in ‘frustrated phagocytosis’. In this situation, phagocytosis does not occur. The failure to phagocytose these large particles due to discrepancy in dimensions, forces these phagocytes to release harmful proteolytic enzymes, acid and reactive oxygen species directly into the extracellular environment. The release of these enzymes and oxidative agents usually breaks down the surrounding healthy tissue causing more damage and debris at the wound site, and sustains the inflammatory process (Hunt et al., 1992). Sometimes, the macrophages are also capable of forming an enclosed environment between them and the implanted biomaterial, where they release their vacuole contents, thus targeting only the foreign material (Hunt, 2004). In such cases, macrophages are reported to be capable of reducing the local pH to a strong acidic pH of about three, which is a very hostile environment for most invading pathogens and some implanted biomaterials (Hunt et al., 1996). Giant cells During frustrated phagocytosis, if neighbouring macrophages come in contact with each other, they tend to fuse and form multinucleated foreign body giant cells (FBGC). Effectively, several macrophages fuse to combine forces in destroying the foreign material (McNally and Anderson, 2002; Hunt, 2004). These giant cells can be more aggressive in phagocytosis and are capable of phagocytosis of larger particles. Like macrophages, giant cells are also capable of producing and secreting large amounts of harmful enzymes and reactive oxygen radicals in an attempt to degrade or modify the biomaterial surface. The macrophages and giant cells are generally the long-term (end-stage) cell populations around an implanted biomaterial, and they usually persist as long as it takes to degrade the implant. Fibroblasts Fibroblast recruitment to the injury site is facilitated by the release of platelet-derived chemotactic factors including PF4, PDGF and TGF-β.

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Leukotriene B4 (LTB4), PDGF and TGF-β are secreted by monocytes and lymphocytes (Tang, 1998; Anderson, 2001). Their main function is to lay down the granulation tissue in place of damaged or excised tissue. Subsequently, fibroblasts remodel the granulation tissue. In the presence of a degrading biomaterial, fibroblasts participate in its remodelling and replacement/regeneration. Endothelial cells Hypoxia in the healing tissue stimulates the release of angiogenic factors such as PDGF and vascular endothelial growth factor (VEGF) from hypoxic cells, including monocytes and fibroblast. These growth factors influence the physiological function of endothelial cells, leading to their recruitment and involvement in the process of neovascularisation/angiogenesis (Knighton and Fiegel, 1989; Michiels et al., 2000; Manalo et al., 2005; Faller, 1999; Bates and Jones, 2003). Angiogenesis is essential for the supply of oxygen and nutrients, removal of waste, debris and degradation products from the injury site. Lymphocytes and plasma cells (immune responses) Lymphocytes are responsible for the immune responses observed with pathogen/foreign particles (including implants). They are classified into T lymphocytes/cells and B lymphocytes/cells depending on their specific functions. T cells are mainly responsible for cell-mediated (adaptive) immunity, while B cells for antibody-mediated (humoral) immunity. B cells also function as antigen presenting cells (APCs). T lymphocytes are involved in cell-mediated responses and are able to respond to antigen through their T cell receptor (TcR) via Major Histocompatibility Complex (MHC) molecules on cell membranes. Helper T cells express the CD4 (Cluster Designation) molecule on their surface, while cytotoxic T cells express the CD8 molecule. T cells do not recognise free antigen but they recognise antigen that is associated with MHC molecules. Class I MHC molecules are expressed on the surface of most cells, while Class II MHC molecules are restricted on the surface of immune cells such as monocyte/macrophages and B lymphocytes. CD4 helper T cells recognise antigen presented by Class II MHC molecules while CD8 cytotoxic T cells recognise antigen presented by Class I MHC molecules. The binding of TcR to the antigen–MHC complex is specific for both the antigen and MHC. Antigens (large proteins) are broken down into peptides inside the presenting cells and then associate with MHC molecules intracellularly before being transported and expressed on the cell surface. Endogenous

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antigens (viral proteins) are produced inside the cell and presented by Class I MHC, while exogenous antigens (e.g. derived from a bacterium) are phagocytosed, processed and presented by Class II MHC. B lymphocytes recognise antigen through immunoglobulin molecules on their surface and trigger a cellular response in the lymphocyte which eventually present the processed antigen to helper T cells in association with Class II MHC. Antigen-specific B cells can also differentiate into plasma cells which produce antibody against the specific antigen. Antibodymediated responses include: (a) neutralisation (by blocking binding on virus, bacteria, toxin or an enzyme), (b) agglutination (multivalent antibody binding to more than one antigen/ microbe particles which are then phagocytosed), (c) opsonisation (be effective opsonins for phagocytes to recognise the antigen/pathogen), (d) complement activation (enhance phagocytosis by C3b opsonin or generating pores known as membrane attack complex (MAC) on cell membrane that will result in cell lysis), and (e) antibody-dependent cell-mediated cytotoxicity. The role of lymphocytes in modulating a healing wound has been characterised (Boyce et al., 2000; Fishel et al., 1987; Schaffer and Barbul, 1998; Martin and Muir, 1990). Both lymphocytes and plasma cells are involved in immune responses and mediate antibody and delayed hypersensitivity responses. However, their roles and responses to biomaterials are generally elusive. Macrophages are known to play an important role in the immune system as antigen-presenting cells (APCs) for T lymphocytes in association with MHC Class II molecules. At the same time, T lymphocytes can secrete cytokines to further activate macrophages, increasing their ability as APCs. Lymphocytes appear to interact with biomaterials and they are known to adhere and respond to different surfaces (Bergman and Zygourakis, 1999; Yokoyama et al., 1986; Maeda et al., 1986). In addition, lymphocytederived cytokines (IL-4 and IL-13) induce fusion of biomaterial adherent macrophages into foreign body giant cells (DeFife et al., 1997; Kao et al., 1995). In vitro studies showed that lymphocytes can enhance monocyte/ macrophage adhesion and fusion, and their co-culture increases the proliferation rate of lymphocytes, possibly through solubility factors (Brodbeck et al., 2003). These findings provide evidence that both lymphocytes and macrophages interact with each other in the vicinity of a biomaterial. Some concerns have been raised regarding unintended immune response towards polymers intended for human use. CD4 and CD8 positive cells, indicating helper T cells and cytotoxic T cells, respectively, were observed

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within degrading polylactide implants (De Jong et al., 2005). Bovine collagen had been found to induce antibody production and clinically unwanted responses in humans (Cooperman and Michaeli, 1984a, b). Class II MHC expression has also been reported for a range of degradable and non-degradable biomaterials including polylactide polymer, hydroxyapatite, polymethylsiloxane, stainless steel and titanium (Petillo et al., 1994; De Jong et al., 2005; Santavirta et al., 1991; Torgersen et al., 1995). In addition, decellularised extracellular matrix has also been found to induce antibody in humans, peaking between two and six weeks (Ansaloni et al., 2007). Fibrosis/fibrous capsule formation In the presence of a biomaterial, fibrosis or a fibrous capsule forms around the implant. Fibrous encapsulation around the implant is considered normal but this response may affect the implant’s intended function. Several authors quote various histological features and dimensions of the capsule as an indication of biocompatibility or acceptable response. Histological and stereological analysis allow investigators to examine and compare various parameters of the host response to biomaterials (Garcia et al., 2007a). It is important to mention that, if the biomaterial could fulfil the intended function and there is no complication associated with the implant, it should be considered biocompatible. For instance, in the case of prosthetic breast implants, excessive capsule formation resulting in capsular contracture, pain and implant distortion is clinically undesirable (Siggelkow et al., 2003; Wong et al., 2006). In the case of a biodegradable implant, a thick fibrous capsule that is both acellular and avascular may prevent infiltration of host cells and capillaries into the implant, thereby preventing its degradation, or slowing down drug release (Anderson et al., 1981; Ratner, 2002). Multiple factors can influence the thickness of the surrounding fibrous capsule. Surface functionality of implants has been studied to ascertain how it may affect protein absorption, inflammatory reaction and fibrous capsule formation (Tang et al., 1998b). Tang et al. showed that —COOH surfaces induce minimal inflammatory cells and fibrous capsule formation (23 μm thick) when compared to —NH2 (150 μm) and —OH (250 μm) surface groups (Suska et al., 2007). Other investigators found that —CH3 groups induces a thicker fibrous capsule when compared to —COOH or —OH groups (Barbosa et al., 2006). In addition, polymer fibre diameter was shown to influence the degree of fibrous capsule thickness and macrophage density (Sanders et al., 2000). The pore sizes of implants can also affect the formation of fibrous capsule (Brauker et al., 1995; Sharkawy et al., 1997).

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12.4

283

Influence of site of implantation on biodegradation

In addition to the physical, chemical and cellular factors, the actual site of biomaterial implantation in vivo may also affect the biodegradation rate. For example, in a study of copolymers of poly(ethylene glycol) with poly(butylene terephthalate), subcutaneous implantations were shown to degrade faster than intramuscular implantations (Beumer et al., 1994). In another study of polydioxanone orthopaedic pins, the bioresorption in the medullary canal of rabbits was shown to be higher than intramuscular or subcutaneous implantations. Furthermore, implantations in mature rabbits were found to induce higher strength retention when compared to those in immature rabbits (Bhatia et al., 1994). Mechanical loading, sheer and friction, e.g. in bone implants, can accelerate in vivo degradation of biomaterials (Claes, 1992). As illustrated in Table 12.3, pH is a major factor that changes from site to site in the host. Hence, numerous studies have been performed to study the effect of pH on biodegradability. Sutures are the best examples for materials that come in contact with a variety of sites in a host. Their strength and degradability are shown to be affected by pH and hence different materials suit different site of in vivo application. It was shown that the degradation of polyglactin 910 mesh samples was significantly delayed in the presence of acidic (pH 5.0) urine, the presence of Escherichia coli bacteria in urine, or the presence of both factors. In contrast, the samples degraded faster in basic condition (pH 8.0) (Best et al., 1999). Similarly,

Table 12.3 Variations in body fluid pH from site to site in host body Location

pH

Stomach Enclosure between macrophages’ cell membrane and biomaterial surface Urine Lysosome Large intestine Duodenum Saliva Intracellular Interstitial (extracellular space) Bile Cerebro-spinal fluid Blood Small intestine

1.0 to 3.0 ≥3.0 4.5 to 6.0 5.5 5.5 to 7 6 to 6.5 6.5 to 7.4 6.8 7.0 7.0 to 7.7 7.3 7.17 to 7.35 7.0 to 8.0

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polyglactin 910 and polyglycolic sutures were found to degrade faster at a basic pH of 10.09 when compared to acidic or physiological conditions (Chu, 1981b, 1982a, b; Tomihata et al., 2001), even though some studies were conflicting (Stewart et al., 1990). Li and McCarthy also showed that the degradation of polyglactin is slower in an acidic pH of 3.7 than in neutral pH. They attributed this to the rate of water absorption. At neutral pH, water absorption was increased due to the osmotic pressure driving the buffer solution into the polymer matrix to neutralize acidic groups, which was not the case for degradation at pH of 3.7 (Li and McCarthy, 1999). Polyglactin 910 retained the highest tensile strength at physiological pH, while polyglycolic sutures retained the highest strength at acidic pH. Increased degradation of polyglycolic sutures in the presence of infected urine has been demonstrated in vitro (Sebeseri et al., 1975) and in vivo (Milroy, 1976), although some investigators argued that the strength was acceptable at these conditions (Hovendal and Schwartz, 1979). Polydioxanone sutures lost their breaking strength comparatively faster in both sterile and infected urine when compared to polyglycolic acid sutures at similar conditions (el-Mahrouky et al., 1987). PDS-II appeared to tolerate these conditions better (Schiller et al., 1993). Polydioxanone sutures degrade in approximately six months in rats’ skin (Molea et al., 2000). Even though both polyglactin and polyglycolide sutures were shown to be superior to catgut sutures (Sugimachi et al., 1978; Tian et al., 1994; Mizuma et al., 1977; Deveney and Way, 1977) in gastrointestinal surgery, polydioxanone sutures exhibited improved stability and best resistance against the harsh condition of pancreatic and bile fluid in pancreatic and upper intestinal surgery (Muftuoglu et al., 2004; Lin et al., 1993). In addition, polydioxanone has been proposed as the suture material of choice for colonic surgery as the suture degrades slowly over six months in a rat colonic anastomosis model (Houdart et al., 1986; Durdey and Bucknall, 1984). However, care should be taken when using polydiaxanone suture in acidic environments such as in the stomach, as it has been shown to lose significant strength at acidic pH (Tomihata et al., 2001). While non-degradable polypropylene suture retains its strength in all pH environments, non-degradable nylon showed the fastest loss of strength in acidic condition. The degradation of lactide-ε-caprolactone copolymer suture was not affected by pH (Tomihata et al., 2001) while poly(fumaricco-sebacic anhydride) copolymers’ degradation is accelerated at basic pH (Santos et al., 1999). The morphology and degradation of PLGA microspheres were also found to be influenced by acidic pH (Zolnik and Burgess, 2007). In the presence of bacteria such as Streptococcus mites, Escherichia coli and Staphylococcus albus, polyglycolide degrades slower in vitro. Similar results were reported when sutures were implanted subcutaneously in vivo and inoculated with bacteria (Williams, 1980). Proteus-infected urine

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appears to accelerate the degradation process of suture materials (Greenberg et al., 2004; el-Mahrouky et al., 1987; Hering et al., 1989). Furthermore, the rate of biodegradation can also be affected by the blood supply available for the tissue where it is implanted. Locations such as nucleus pulposus of intervertebral discs, articular cartilage of joints and corneal tissues are examples of avascular tissues and have low metabolic activity and oxygen requirements. In contrast, tissues such as the bone marrow, lung, liver, kidney, lymphoid organs, skin and hollow organs have higher blood supply and better perfusion that influence the environment for biodegradation. Implantation of biomaterials in the central nervous system (CNS) can be complicated. The CNS tissue is protected from the rest of the circulation by the blood–brain barrier (BBB). Its primary role is to protect the CNS from potential risks (e.g. infectious or chemical agents) from the blood, while still allowing some small molecules (e.g. O2, CO2, ethanol and steroid hormones) to reach the CNS. This restricted access to the CNS can cause complications and can affect the removal of degradation products from biodegradable materials implanted in the CNS in vivo (Fournier et al., 2006).

12.5

Influence of species and repeated implantation

Investigators must be aware that implantation in different strains and species may produce different results. For instance, phagocytosis of crosslinked dermal sheep collagen rarely occurred in mice but calcification was observed more often. When compared to rats, giant cell formation was limited but there was higher stroma formation. Within the rat strains, the differences in T lymphocyte response could also be appreciated (Khouw et al., 2000). When the same biomaterials are implanted in vivo for the second time (at a different site) in the same individual animal, an accelerated host response cascade is observed. The second challenge enhances the foreign body reaction as shown by an increase in the infiltration of cells, vascularisation, macrophages and giant cells. There was also increased in MHC Class II positive cells as well as accumulation of plasma cells in the vicinity of the implants studied. These responses were not specific to the types of materials studied (polyurethane, silicone and cross-linked dermal collagen implants) (van Luyn et al., 2001).

12.6

Adverse outcomes of biodegradable polymers

While biodegradable polymers have shown success in various clinical applications and are widely used today (Puccio et al., 2005; Franklin et al., 2002;

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Gupta et al., 2006; Frame et al., 2004; Oelschlager et al., 2006; Mostow et al., 2005), certain unfavourable outcomes have been observed in clinical practice since their introduction. Biodegradable polymers used in the fixation of fractures were found to be complicated by fracture displacement, sterile fluid accumulation and discharge (Bostman et al., 1990b, 1991; Frokjaer and Moller, 1992), loss of mechanical strength and sterile sinus formation (Hofmann, 1992; Bostman et al., 1990a). Rapid implant degradation in vivo was implicated in the undesired inhibition of bone regeneration (Winet and Hollinger, 1993), delayed angiogenesis and osteogenesis in rabbit tibia when PLGA was used as an orthopaedic implant (Winet et al., 1995). These copolymers were also associated with re-operations and unsatisfactory results, especially when used in unstable and comminuted fractures in ankles (Pelto-Vasenius et al., 1998). The analysis of fluid derived from PGA osteosynthesis implants suggested that there was lymphocyte activation and lymphocyte-mediated immunological response against the material (Santavirta et al., 1990). Some investigators have shown vigorous local foreign body reaction, synovial irritation and subsequent degeneration, leading to osteoarthritis and ankle arthrodesis (Bostman, 1998). Foreign body reaction and revision surgery were associated with PLGA and PLA implants for maxillary and mandibular reconstruction surgery (Landes et al., 2006). Seroma formations, disease recurrences and vascular pseudoaneurysm formation were observed with degradable collagen-based scaffolds (Breyer et al., 2007; Gupta et al., 2006; John et al., 2006; McCready et al., 2005).

12.7

Mechanisms of in vivo biodegradation

Two main chemical mechanisms that the host utilizes to degrade biomaterials/foreign materials are hydrolysis and oxidation. Hydrolysis, the scission of a covalent bond by water, is the main chemical means by which biodegradation of biomaterials is achieved by the host. The other mechanism, oxidation, involves the oxidative scission of covalent bonds by free radicals. Depending on the properties of the biomaterial and the location of implantation in the body, the chemical means for scission of covalent bonds may or may not be catalyzed by enzymes. Physical means of wearing and solubilisation by water further assist the hydrolysis and oxidation mechanisms of biodegradation. The implanted biomaterial continuously loses its mass until it has completely disappeared. Mass loss often occurs by solubilisation of low molecular weight degradation (covalent bond scission) products and continues until the polymer bulk is completely broken and disintegrated (Timmins and Liebmann-Vinson, 2003).

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12.7.1 Hydrolysis and enzymatic degradation Hydrolytic biodegradation in vivo starts with water contact and host proteins adsorption on the biomaterial. The hydrolysis is catalysed by acids, bases, salts and enzymes. Enzymatic catalysis of hydrolysis is generally specific to the specific chemical bonds and this includes enzymes such as proteases, esterases, lipases and glycosidases. Hydrolysis is further assisted by the uptake of water by the biomaterial, which depends on the hydrophilicity, morphology, type and concentration of hydrolysable bonds and porosity of the biomaterials. Studies on microbial and enzymatic degradation of polymers have attracted huge attention in the challenge to produce environmentally friendly polymers (Eldsater et al., 2000; Shimao, 2001; Lefebvre et al., 1994). Even though most synthetic biodegradable polymers are generally degraded by hydrolysis, some synthetic polymers are susceptible to enzymatic attack. The mechanisms and enzymes in humans are different from those found in the environment but some insights and comparisons can be made. Williams and Mort studied the ability of carboxypeptidase A, αchymotrypsin, clostridiopeptidase A and ficin enzymes to increase the rate of hydrolysis of polyglycolic acid polymer (Williams and Mort, 1977). Lipases have been shown to degrade polycaprolactone and polylactic acid polymers as a method for waste recycling (Takamoto, 2001; Takahashi, 2004; Li et al., 2003). In drug delivery strategy, lipase has been used to study the degradation properties of microparticles intended for drug delivery (Chen et al., 2000a, b; Wu et al., 2000). Pancreatic lipase is present in the gastrointestinal tract and hence, may affect the degradation and drug release from these polymers. Esterases are present in human plasma, the pancreas, mammary gland and human monocyte-derived macrophages (Brodt-Eppley et al., 1995; Li and Hui, 1997). There has been recent interest in the study of hydrolytic degradation of vulnerable polymers (e.g. polyurethanes, polycarbonates and polyethers) by esterases such as cholesterol esterase and carboxyl esterase secreted by monocyte-derived macrophages (Labow et al., 2002a). Recent studies have also shown that polymer chemistry and its degradation products may influence the type of esterase synthesised and secreted by these cells (Labow et al., 2005a, b). The susceptibility of polyurethanes to enzymatic degradation by cholesterol esterase was shown to be dependent on the degree of hydrogen bonding between the carbonate and urethane groups (Tang et al., 2001a, b), surface morphology (Tang et al., 2002), and protein binding (Tang et al., 2003). Serine proteases have also been shown to have hydrolytic activity towards polyurethanes (Labow et al., 2002b)

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Naturally derived biomaterials are subjected to enzymatic degradation in vivo. The mechanisms of biodegradation of natural biomaterials have not been methodically studied when compared to synthetic biomaterials because of the complexity of the physiological conditions involved. In addition, there is the established recognition that natural polymers are present and constantly being remodelled in the body and are therefore, biodegradable. Collagens are the major insoluble extracellular matrix protein in mammals. Collagen has a characteristic repeating sequence of three residues (Gly—X—Y) in its primary structure, where X and Y are often proline and its post-translational modification to 4-hydroxyproline, respectively. This results in a stable quaternary structure, with the alpha chain adopting a left-handed poly-Pro II-like helix, and the three chains wrap around each other to form a right-handed superhelical conformation, with the glycine residues at the core of the helix (Okuyama, 2006; Kotch and Raines, 2006). These strands are held together by hydrogen bonding and also covalent bonds. After post-translational modification, secreted collagen helices selfassemble to cross-linked microfibrils and eventually form long fibrils (Kadler et al., 1996). Up to 19 types of collagen have been reported (Lee et al., 2001). This process creates molecular structures with properties of obvious interest to bioengineers. Nature utilizes these exceptional properties of collagen as a scaffold to attach a number of binding proteins to specific sites (Vakonakis and Campbell, 2007). In vivo, collagen plays a primarily structural role as well as in tissue homeostasis, organ morphogenesis, embryogenesis, angiogenesis and wound healing (Werb and Chin, 1998; Cawston, 1996; Sternlicht and Werb, 2001). The triple helical conformation makes collagen resistant to most proteinases. Specific collagenases, members of the multigene family of matrix metalloproteinases (MMP), can cleave the collagen triple helical structure (Sternlicht and Werb, 2001). They are zinc-containing metalloproteinases and usually require calcium as a co-factor for full activity. MMPs are produced by many cell types including stromal cells, epithelial cells, macrophages and leukocytes (Sternlicht and Werb, 2001). Collagenase is produced from fibroblasts which interact with collagen (Welgus et al., 1980). In bone tissue, cathepsin is produced by osteoclasts and plays a role in bone resorption and bone erosion in rheumatoid arthritis (Garnero et al., 1998; Iwata et al., 1997). These collagenolytic enzymes cleave the alpha chains of collagen after Gly (Gly—Ile/Gly—Leu) in a particular sequence that is approximately three-quarters away from the N-terminus of the collagen molecule (Aimes and Quigley, 1995). There is evidence to show that, in order to cleave the collagen at this position, collagenase first binds and subsequently unwinds the triple helical

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structure prior to hydrolysing the peptide bonds (Chung et al., 2004). Once they are cleaved into the three-quarter and one-quarter fragments, they are degraded by gelatinases and proteinases, and denature at body temperature to smaller peptides and amino acids. Elastases and cathepsins are involved in the degradation of collagen with a higher extent of natural cross-links, based on their ability to digest non-helical telopeptide regions. Proteolytic enzymes of broad specificity, such as pepsin and trypsin, degrade the nonhelical peptides at either end of native collagen. The degradation of collagen-based biomaterials is frequently illustrated by their susceptibility to degradation in vitro by enzymes such as collagenase, cathepsin, pepsin and tripsin (Okada et al., 1992; Anderson, 1969). While these in vitro assays allow relative comparisons on the degradation profile of these biomaterial constructs, accurate prediction of their in vivo degradation behaviour is complex and difficult to achieve (Yannas, 1975). Other biodegradable materials are also subject to enzymatic degradation. Chitosan, for example, is sensitive to lysozymal enzyme produced by inflammatory cells (Haque et al., 2001; Varum et al., 1996). Gelatin (hydrolyzed collagen) is degraded by collagenase and various other proteases in vivo. In addition, starch based biomaterials are degraded by serum amylase and also by macrophages (Desevaux et al., 2002). Alternatively, alginate is susceptible to hydrolysis in vivo but also can be enzymatically degraded by bacterial alginate lyase into monomeric sugars (Sakakibara et al., 2002). Hyaluronans can be enzymatically degraded by hyaluronidase (Zhong et al., 1994; Avitabile et al., 2001).

12.7.2 Oxidative degradation Implanted biomaterial induces inflammatory responses whereby acute inflammatory cells (neutrophils) migrate to the location and are subsequently followed by monocytes that differentiate into macrophages. These phagocytic cells can be stimulated to increase their uptake of oxygen and produce huge quantity of superoxide anion by one electron reduction of oxygen at the expense of NADPH (nicotinamide adenine dinucleotide phosphate), a process known as respiratory burst (Babior, 1984; Bannister et al., 1982). The superoxide anion can then react with other molecules in the presence of metal ions to generate other harmful oxygen radicals such as hydrogen peroxide, singlet oxygen and free hydroxyl ions (Badwey and Karnovsky, 1980). These harmful species generated via oxidative burst (Karnovsky and Badwey, 1983) have been suggested to increase the degradation of biomaterials (Ali et al., 1994). High local concentrations of harmful oxygen intermediates, acid and lysosomal enzymes are produced in a closed space between macrophages and the implanted biomaterial (Wright and Silverstein, 1984).

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There is growing evidence to show that oxidative degradation occurs in vivo and plays an important role in the degradation of implanted biomaterials (Christenson et al., 2004, 2005). When the oxidative degradation of poly(etherurethane urea) elastomers films was initiated in vitro, exposure of the polymer to air subsequently resulted in faster degradation when compared to leaving the polymer in solution (Schubert et al., 1997). This indicates that the degradation process could only proceed with the availability of oxygen. This is further supported in another study by the observation that antioxidants can inhibit the degradation of polymer in vivo (Christenson et al., 2006). Furthermore, the implantation of synthetic polymers has been shown to induce host inflammatory cells to produce hydrogen peroxide (Krause et al., 1993). It has also been demonstrated that adherent leucocytes on explanted polyglycolic acid polymer mesh from the peritoneal cavities of rats produce hydrogen peroxide in vitro (Devereux et al., 1991a, b). In addition, hydroxyl radicals are involved in the degradation of polycaprolactone and polylactic acid polymers both in vitro and in vivo (Ali et al., 1993, 1994). Lee and Chu (2000) also demonstrated that superoxide ion-induced degradation can affect the mechanical, thermal and surface properties of absorbable sutures.

12.8

Material factors influencing biodegradation

In vivo biodegradation of implants is influenced not only by the host responses, but also by the material properties of the biomaterial. It is therefore crucial to understand these material factors in order to appreciate the influence of host responses on biodegradation. This is especially important because biomaterial scientists have been able to tailor the properties of the biomaterials to dictate certain desired host responses. Some of the key material factors that control the susceptibility of a biomaterial to degradation are outlined in this section. In practice, a degradable material is subjected to various factors at once, each of which eventually influences its overall in vivo response. Although it might be possible to control each biomaterial factor prior to implantation, the overall in vivo response should not be interpreted in isolation.

12.8.1 Particle size Particle size is a key factor for degradation in vivo. The host can mobilise and eliminate only particles of submicron sizes. Particles with small diameters (less than a few nanometres) are mostly exchanged through the capillaries, whereas large particles of up to a few tens of nanometres are absorbed into the lymphatic system. Larger particles (50–200 nm) are transported

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with the lymph and mostly trapped and phagocytosed in the lymph nodes (Imam and Killingsworth, 2005). Another means of transport for foreign/ waste particles in vivo is as phagocytosed materials by APCs. Following phagocytosis, the APCs are transported through the lymphatic vessels to the nearest lymph node, where they present the intracellularly processed antigen along with MHC Class II receptors to T-helper lymphocytes. Hence, the host has to degrade the biomaterial down to dimensions that can be transported from the implant site through the lymphatic and circulatory systems for their eventual utilization in metabolic pathways or to be eliminated directly.

12.8.2 Polymers and their degradable chemical linkages Most biodegradable biomaterials are polymers. The molecular mass describes the average molar mass of the polymer molecules. It essentially describes the polymer molecule’s size. Polymers are made of several monomeric units linked by covalent linkages to form macromolecular chains (branched or non-branched) with varied molecular masses. These polymer chains (associated by chemical linkages/cross-linking or by non-covalent associations as in copolymers) make up the solid material that is usually too big for the host to remove directly from its system. Therefore, polymers are degraded into smaller chains during degradation, which effectively lowers the molecular mass and degree of polymerisation (Arshady, 2003; Ratner et al., 1996). In a water-insoluble polymer, its degradation is the transformation of the insoluble polymer to water-soluble oligomer or monomer units (D’Emanuele et al., 1992). This feature of polymeric materials has often been used to control drug delivery. For example, Braunecker et al. were able to control the release of the drug theophylline from polyglycolide polymer by varying its molecular mass. They found that decreasing the molecular mass and increasing the pore size and volume increases the rate of drug release (Braunecker et al., 2004). This process essentially controls the nature of an application such as the kinetics of drugs release (Gopferich, 1997; Vogelhuber et al., 2001) or mechanical performance of a biodegradable implant (Watkins and Anseth, 2003; Muggli et al., 1999). In addition, polymer erosion has also been shown to control the rate of drug release. Biodegradability of biomaterials depends on the availability of degradable chemical linkages and their amounts and distribution within the bulk of the material (Timmins and Liebmann-Vinson, 2003). The various bonds that are potential for biodegradation are listed in Table 12.4. The susceptibility of the different bonds to hydrolysis is strongly dependent on their local microenvironment (pH, ionic strength, temperature) and availability of catalysts (acids, bases, salts, metal ions and enzymes). The relative rates

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Degradation rate of bioresorbable materials Table 12.4 Degradable chemical linkages in biomaterials that determine their biodegradability Polymer class Polyanhydride

Bond structure O

O

C O C O

Polyorthoester O C

O

Polyester

O

Polyurea

H O H

Polyamide

O H

C O

N C N

C N O

Polycarbonate

O C O O H

Polyurethane

O C N

Polyphosphazenes N P

Polyether

O

of susceptibility of different bonds listed in Table 4 are discussed elsewhere (Timmins and Liebmann-Vinson, 2003; Pierre and Chiellini, 1986, 1987a, b). Hydrolysis rates have been reported to be in the order of polyanhydrides > polyorthoesters > polyester > polyurea > polyamide > polycarbonate > polyurethane, based on their bond structure and energies.

12.8.3 Surface and bulk erosion Erosion can be defined as the loss of mass from a biodegradable polymer. The erosion mechanisms of biodegradable polymers can be divided into either surface erosion or bulk erosion (Tamada and Langer, 1993). In surface erosion (heterogeneous erosion), the pattern of material loss is from the external surface proceeding inwards, the rate of which is proportional to the surface area; and takes longer for thicker implants (Shieh, 1994). Hence, for a surface eroding polymer, the erosion period can be achieved by adjusting the dimensions and shape of the polymer. In bulk erosion (homogeneous erosion), the material is lost uniformly from the entire volume or

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cross-section of the polymer. The rate at which bulk erosion proceeds is dependent on the amount or volume of the material and this generally decreases as the process proceeds. The lifetimes for bulk-eroding materials are the same, irrespective of implant thickness (Tamada and Langer, 1993). However, it should be kept in mind that implant geometry and erosion conditions can significantly influence whether a biodegradable polymer undergoes either surface erosion or bulk erosion (von Burkersroda et al., 2002; Grizzi et al., 1995). Polyesters such as polylactic acid and polylacticco-glycolic acid polymers are considered to undergo bulk erosion, while polyanhydride undergoes surface erosion (Gopferich and Tessmar, 2002). Water-insoluble polymer needs to undergo degradation prior to erosion. In biodegradable polymer such as PLGA polyesters, the primary degradation pathway is by hydrolysis, although oxidative and enzymatic degradation may play a role (Williams and Mort, 1977). Degradation in PLGA polyesters involves chain scissions (chemical depolarisation) of ester bond linkages in the polymer backbone, leading to molecular fragmentation. This involves water uptake followed by hydrolytic attack by water molecules of the ester bonds in a random way according to the Flory principle, which postulates that all linkages have the same reactivity (Shih, 1995). Factors that have been identified to affect hydrolytic degradation include molecular mass, hydrophilicity, crystallinity, copolymer composition, morphology, autocatalysis, percolation phenomena, the presence of proteolytic substance or other excipients and annealing treatment (Browning and Chu, 1986; Loo et al., 2005; Vert et al., 1991; Joziasse et al., 1998). In bulk erosion, autocatalysis may occur, leading the interior of a large size polymer such as PLGA to degrade faster than the outer zone (Therin et al., 1992; Li, 1999). This is due to the degradation products of PLGA, carboxylic acids, which increase the acidity inside the polymer and catalyse the reaction (acid autocatalysis). The release of soluble acidic products from the interior of a large size polymer is slow, and therefore catalyses the hydrolysis in the interior so that it degrade faster than the exterior.

12.8.4 Additives and impurities The addition of bioactive agents (such as drugs, proteins, DNA, plasmids, viruses) and impurities (un-reacted monomers, cross-linking agents, chemical inhibitors, antioxidants, metal ions) affects the chemical environment surrounding the degradable linkages in the biomaterials. Hence they have significant influence on the rates of biodegradation of biomaterials. The presence of a basic drug, gentamycin, in biomaterials such as poly(DLlactic acid) (Vert et al., 1994) and interpenetrating networks based on acrylic acid and gelatin (Changez et al., 2004), was shown to reduce biodegradation rates. The incorporation of coral (CaCO3) particles in poly(DL-lactic acid) polymer slows the degradation, attributed to the buffering effect of calcium

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carbonate and to the presence of coral/polymer interfaces that facilitate ion exchanges with the external solution (Li and Vert, 1996). When another basic drug, caffeine, was used, a complex degradation effect was observed. The basic drug can react with carboxyl chain ends and thus reduce the acid autocatalysis and degradation rate. However, it could also cause base catalysis and thus accelerate the degradation (Li et al., 1996). The addition of low molecular weight oligomers accelerated the degradation of poly(DL-lactic acid), and this again was attributed to acid autocatalysis by the acidic groups (Mauduit et al., 1996). Autocatalysis has also been shown to affect drug release in microparticles (Siepmann et al., 2005). Using poly(fumaric-cosebacic) anhydride microspheres and loading them with 2% serum albumin results in faster degradation in vitro (Sandor et al., 2002). The type of metal ion can have an influence and has been shown to promote or inhibit the migration of neutrophils. Hujanen et al. studied the effects of chloride salts of zinc, copper, nickel, chromium, iron and gold on the migration patterns of neutrophils in vitro (Hujanen et al., 1995). The authors found that zinc, copper and nickel promoted directed migration (chemotaxis) of neutrophils, while no effects were seen with chromium or iron. In contrast, an inhibition in migratory activity was observed with gold ions (Hujanen et al., 1995). Another study demonstrated a similar migratory effect by copper and nickel ions on neutrophil locomotion (Hunt et al., 1992).

12.8.5 Hydrophilicity Material hydrophilicity is a key factor in biodegradation as water is the universal solvent in biological systems. Hydrophilicity is determined by the polymer composition as well as its morphology and is a measure of how well the material absorbs, swells or dissolves in water. Surface hydrophilicity involves how readily the material is wetted by water, while bulk hydrophilicity affects the extent of water entry into the bulk of the material. When a polymer degrades via hydrolysis, hydrophilicity can be a crucial factor in determining its degradation rate. In a tissue engineering scaffold, hydrophilicity can be critical in determining how well fluid and nutrients can diffuse into the scaffold, hence its ability to support cellular functions and scaffold remodelling. In drug delivery applications, it may affect the pharmacokinetics, such as the onset, rate and duration of drug delivery.

12.8.6 Crystallinity A polymer may be described as crystalline if it contains regions of threedimensional ordering on the level of atomic dimensions, usually arising from intramolecular folding or stacking of adjacent chains. Chains within

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the crystalline regions are tight and ordered, while chains within the non-crystalline (amorphous) regions have lower chain density, and are more random with a higher degree of free motion. The degree of polymer crystallinity can be expressed as mass fraction or volume fraction of crystalline material (Ratner et al., 1996). In general, the initial degree of crystallinity affects the rate of polymer degradation. As the extent of water hydration in crystalline regions is less than that in amorphous regions, crystalline regions are less susceptible to degradation. The amorphous regions of polymers undergo hydrolysis first because of their higher rate of water uptake. This leads to subsequent increased crystallinity in the remaining polymer, with decrease in mechanical strength and molar mass (Chu, 1985). Crystallinity of biodegradable polymers, in general, increases with degradation time before decreasing as the hydrolysis proceeds. Secondary crystallization has been proposed as the reason for peak crystallinity as degradation time proceeds (Chu, 1981a, 1985). As hydrolysis occurs mostly in the amorphous regions at earlier stages, polymer chain scissions cause the reduction in tensile strength due to a decrease in chain entanglements. The remaining intact chains in amorphous regions reorganized to a tighter and more ordered crystalline state, lead to higher crystallinity. As hydrolysis proceeds in these crystalline regions later, the crystallinity decreases thereafter. However, with the addition of exogenous agents such as a drug, the rates of degradation in the crystalline and amorphous regions can vary. When studying the rate of drug release from a polymer, conflicting results have been obtained (Tsuji, 2000; Li et al., 1990; Frank et al., 2005). In one study, polyglycolide polymers with a higher crystallinity release more theophylline drug at earlier stages of the degradation (Hurrell and Cameron, 2002). In another study, poly(lactide-co-glycolide) copolymer used to deliver lidocaine demonstrated that a crystalline matrix releases the drug at a slower rate when compared with its amorphous counterpart of similar molecular weight (Frank et al., 2005). Prediction of drug release profiles must take into account the nature of the drug, its osmotic effects and its contribution to polymer degradation, all of which may dictate the rate of water absorption by the polymer (Frank et al., 2005).

12.8.7 Surface chemistry The bulk of biomaterials research has been directed towards developing biocompatible surfaces. This is because surface chemistry (the presence or absence of ionic groups on the surface and the nature of the ionic groups) has been demonstrated to dictate the direction of the host response cascade. For instance, hydrophilic surfaces induce lower degrees of protein adsorption

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and cell adhesion, and hence lower degrees of inflammatory response, compared to the opposite effects observed with hydrophobic surfaces. Hydrophilic surfaces can significantly inhibit both the initial and long term monocyte adhesion, and also inhibit lymphocyte proliferation (MacEwan et al., 2005). Both anionic and cationic surfaces showed mild inhibition of monocyte adhesion and increased macrophage fusion, while cationic surfaces decreased lymphocyte proliferation and inhibited the activity of monocytes (MacEwan et al., 2005). Anionic surfaces were shown to induce apoptosis of adherent macrophages to a greater degree when compared with hydrophilic or cationic chemistries (Brodbeck et al., 2001, 2002b). It was also found that there was an inverse relationship between the level of apoptosis and the promotion of macrophage fusion to form giant cells on different surfaces (Brodbeck et al., 2001). Altering the surface chemistries therefore, could provide a strategy to avoid the harmful effects of adherent macrophages on biomaterials degradation. The influence of surface chemistries on the cytokine expression of monocyte/macrophage in vitro has also been studied. Hydrophilic and anionic surfaces increased IL-10 expression and decreased the expression of IL-8. IL-8 promotes inflammation by cellular activation and chemotaxis, while IL-10 down-regulates the activity of inflammatory cells and fibroblasts (Brodbeck et al., 2002a). These surfaces also inhibited monocyte adhesion and IL-4 mediated macrophage fusion into FBGC (Brodbeck et al., 2002a). TNF-α induced apoptosis in biomaterial adherent macrophages. This process was found to be both dose and time dependent. The ability of TNF-α to induce apoptosis was inhibited by IL-4 (Brodbeck et al., 2002c). It is suggested that IL-4 functions to protect biomaterial-adherent macrophages from TNF-α induced apoptosis. It is also suggested IL-4 plays a role in increasing TNF receptor shedding, thereby limiting the effect of TNF-α. The inhibition of apoptosis is also associated with increase in macrophage fusion to form FBGC (Brodbeck et al., 2002c). Significant research had been carried out investigating the influence of surface chemistry, adsorbed proteins, conformation of the adsorbed proteins and cell behaviour on the surface of implanted biomaterials (Shen et al., 2004). Protein adsorption is dependent on molecular weight, concentration in the environment and strength of the adhesion between the adsorbed protein and the biomaterial surface (Jenney and Anderson, 2000b; Green et al., 1999). The biomaterial surface properties such as hydrophilicity may determine the strength of such interactions and the subsequent interactions with the physiological environment (Slack and Horbett, 1992; Balasubramanian and Slack, 2002). In addition to their effects on adsorbed proteins and cellular response, the difference in surface properties has also been proposed to mediate complement system activation and accumulation of inflammatory cells on the surface of implants in vivo (Tang et al., 1998a).

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Studies found that IgG-coated surfaces induced a substantial and prolonged neutrophil production of reactive oxygen species. When a hydrophilic surface was used to bind IgG, a decreased neutrophil response was observed when compared to a hydrophobic surface (Liu et al., 1997). Pre-coating polymer surfaces with human IgG resulted in significantly higher levels of neutrophil activation compared to human serum albumin or fibrinogen. The presence of platelets markedly increased the reactivity of neutrophils on fibrinogen- and IgG-coated surfaces, but not on albumin-coated surfaces (Nimeri et al., 2002). The reactive oxygen species production was also higher in neutrophils on IgG-coated surfaces than in cells interacting with albuminor fibrinogen-coated surfaces (Nimeri et al., 2003; Wettero et al., 2003). Polyester terephthalate implants precoated with IgG activate human neutrophils in vitro, and also attract substantial numbers of phagocytes (especially neutrophils and macrophages) when implanted in mice (Tang et al., 1993). Another study revealed that IgG-coated surfaces attract the largest amount of cells when implanted subcutaneously in rats (Kalltorp et al., 1999). Surface chemistry can influence the adhesions of inflammatory cells and the formation of FBGCs (Anderson et al., 1999; Jenney and Anderson, 1999). Pre-adsorbed IgG has been shown to increase macrophage adhesion and FBGC formation (Jenney and Anderson, 2000a). Macrophages that were adherent to immobilized IgG on polystyrene showed augmented TNF-α secretion in response to bacterial lipopolysaccharide (Richard et al., 1999). Human IgG was shown to induce gene transcription and dosedependent secretion of IL-1 receptor antagonist (IL-1ra) and IL-8 in cultures of normal human monocytes. This IL-1ra and IL-8 induction by IgG was enhanced in the presence of autologous T lymphocytes (Ruiz de Souza et al., 1995). It was also demonstrated that immobilized IgG on polystyrene initiated the secretion of TNF-α, IL-6 and IL-8 (Foreback et al., 1997). In addition to IgG, pre-adsorption of α2-macroglobulin also facilitated macrophage fusion and FBGC formation on the RGD-modified surface (Anderson et al., 1999). FBGC formation, TNF-α release, and pro-coagulant activity of monocytes were higher for surfaces pre-adsorbed with IgG, while FBGC formation was lower on surfaces with adsorbed fibrinogen and fibronectin (Shen et al., 2004). Adsorbed vWF inhibits while adsorbed IgG promotes long-term macrophage adhesion (Jenney and Anderson, 2000b). Hunt et al. (1996) also studied the influence of surface charge by altering the sulphonate ionic groups on the polymer backbone of polyetherurethanes. Intramuscular implantation of these polymers demonstrated that the surface charge can influence the early phase of acute inflammatory response in vivo. It was shown that 20%-charged polymer elicited significantly less neutrophil infiltration at two weeks post-implantation when compared to the 10%, 30% or control polyetherurethanes polymers (Hunt et al., 1996).

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12.8.8 Degradation products Degradation products of collagen appeared to affect the subsequent host response such as bone remodelling in rat (Zhao et al., 1999, 2000), wound healing (Pilcher et al., 1997) and programmed cell death of amniotic fibroblasts (Lei et al., 1996). These degradation products also function as chemoattractants for recruitment of endothelial cells (Li et al., 2004) and can affect vascular smooth muscle cell migration (Stringa et al., 2000).

12.9

Biomaterial design parameters

Although various synthetic biomaterials have been used as tissue replacements, scaffolds for tissue engineering and adjuncts to surgery (such as suture materials), the perfect biomaterial remains elusive. Substantial progress has been made in designing synthetic biomaterials hoping to imitate the natural extracellular environment. This task is hindered by numerous challenges that are still to be solved. Current simplistic design of synthetic biomaterials lacks the compositional, spatial and functional complexities associated with the natural biological environment (Lutolf and Hubbell, 2005). At present, no synthetic biomaterial can fully restore the normal tissue architecture and function. Biologically derived tissues have been proposed for use as biomaterials for tissue engineering and clinical application. These scaffolds provide a natural complex microenvironment similar to the biological components in the host for cellular functions (Hodde, 2002). Decelluarised biologically derived biomaterials in use today are composed of the natural extracellular matrix (ECM) proteins and carbohydrate such as collagen, elastin, fibronection, growth factors and glycosaminoglycans (Courtman et al., 1994). The extracellular matrix, which is the non-cellular component of tissues, plays a wide range of structural and cellular functions in tissue and organ morphogenesis, in addition to maintenance and repair of injury. Examples include decellularised human dermis, porcine small intestinal submucosa, bovine and equine pericardium, cholecyst-derived extracellular matrix, porcine urinary bladder mucosa and bladder submucosa, human cadaveric fascia and human amniotic membrane (Bellows et al., 2007; Burugapalli et al. 2007; Hodde, 2002; Chang et al., 2002). These substrates offer a biological foundation and have been shown to be inductive for cell migration and attachment, subsequent proliferation and differentiation, and eventual tissue remodelling in vivo. ECM has been shown to possess receptors and inherent bioactive molecules that aid in cellular and tissue regulatory functions (Badylak, 2007). In addition to their inherent cytocompatibility, ECM also offers the desired shape and strength of the original tissues from which they are harvested and processed (Coburn et al., 2007; Gloeckner et al., 2000).

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The decellularised matrices act as porous scaffolds for cellular attachment, migration, proliferation and formation of new tissue.

12.9.1 Copolymers, interpenetrating polymer networks (IPN) and composites No single polymer/material can have all the desired properties for any particular biomedical application. Biomaterial scientists often combine two or more materials to tailor the desired properties for specific applications. Copolymers, IPNs and composites are examples of such materials. The complex compositions of these materials are additional variables that affect biodegradation. Miller et al. studied the effect of varying copolymer composition on the biodegradation of glycolide–lactide copolymers in bone and abdominal wall implantation models (Miller et al., 1977). The half-lives of 100% PGA and 100% PLA were five months and six months, respectively, while the half-life for 50 : 50 PLGA was one week. Other investigators showed that the degradation rate of PLA is significantly slower than its copolymer with glycolide (Schmitt et al., 1994; Hakkarainen et al., 1996). Albertsson and Liu (1997) also observed that various blendings of poly(trimethylene carbonate) and poly(adipic anhydride) can have an effect on in vitro degradation by serving as an accelerator of degradation. In addition, the type of monomer structure itself can influence degradation rate. Using poly[bis(pcarboxyphenoxy) alkane] anhydrides as an example, changing the alkane from a methyl to a hexyl group decreased degradation from 10−1 to 10−4 mg/ h/cm2 (Leong, 1985). Furthermore, the type of initiator/catalyst system used to manufacture the polymer can also influence the water uptake and hydrolysis rate of polylactic acid (Schwach and Vert, 1999). In interpenetrating polymer networks based on acrylic acid and gelatin, increases in gelatin content resulted in increases in degradation rate (Burugapalli, 2004; Changez et al., 2004).

12.9.2 Cross-linking Despite the apparent advantages, the use of biological scaffolds still poses a number of issues, specifically their rapid biodegradation in vivo, concerns regarding immunogenicity, and unintended or incomplete tissue responses. In order to overcome these concerns, cross-linking methods have been employed over the decades. Cross-linking plays a role in slowing the biodegradation as well as eliminating or reducing cross-species response to structural proteins of ECM scaffolds (Courtman et al., 2001). The success of cross-linking can be appreciated in the development of bioprosthetic heart valves from xenogenic tissues. The shift from

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formaldehyde treatment and the introduction of glutaraldehyde crosslinking of bioprosthetic heart valves is considered a major achievement in this field (Zilla et al., 2007). Cross-linking maintains the non-thrombogenic surface of bioprosthetic heart valves, reduces antigenicity and prevents degradation, thereby increasing their in vivo life. This form of bioprosthetic heart valve fixation is still being used today. Indeed, chromium tanning of collagen has been used for more than 100 years in the preservation of animal skins in the leather industry (Covington, 1997). The process involves ionic, covalent and hydrogen bonding between chromium species and collagen. This form of biomaterial stabilisation has also been utilised to produce chromic catgut as surgical suture materials (Van Winkle et al., 1975). Various other forms of cross-linking agents exist, mainly developed out of necessity to improve the performance of bioprosthetic heart valves and pericardial tissue in vivo, and to overcome failure associated with calcification and cytotoxicity. Cross-linking agents ranged from the well-known aldehydes (formaldehyde and glutaraldehyde) to newer bifunctional cross-linkeds such as polyepoxy compounds and hexamethylene diisocyanate. Other modern cross-linking agents include the zero-length cross-linkeds such as acyl azide and carbodiimides (Khor, 1997). Even though glutaraldehyde cross-linking is still being used to stabilise biological scaffolds, newer functional and possibly less cytotoxic cross-linking methods are being developed. These include microbial transglutaminase enzyme, herbal compound genipin, and multifunctional dendrimeric cross-linking (Garcia et al., 2007b; Duan et al., 2007; Chang et al., 2002; Chan et al., 2008). There is a move from using non-degradable bioprostheses intended to last beyond the patient’s life to biodegradable scaffolds, which are designed to be remodelled in vivo over time to repair damaged or diseased organs (Brody and Pandit, 2007; Badylak, 2007). Scaffold degradation should be controlled and occur in parallel to match the rate of tissue regeneration in vivo (Burugapalli et al., 2007). The biodegradable scaffold used for tissue engineering must also provide the necessary function until newly formed tissue can assume this function. In an idealised biodegradable scaffold, as host cells invade and populate the implanted scaffold, they secrete new extracellular matrix to replace the original scaffold that degrades at a predetermined rate. It has been shown that the rate of in vitro and in vivo degradation of a biological scaffold can be tailored by the degree of cross-linking (Burugapalli et al., 2007). Liang et al. (2004) examined the influence of the extent of cross-linking on tissue response and regeneration pattern by using acellular bovine pericardium cross-linked with genipin. In a rat subcutaneous implantation model, they found that non-cross-linked and minimally cross-linked acellular bovine pericardium were rapidly degraded before new tissue formation could occur. In contrast, in the moderately cross-linked and

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highly cross-linked scaffold, new collagen formations were observed, although in the latter, this process was limited to the outer layer of the scaffold due to its greater ability to resist degradation. When similar scaffolds were implanted as a vascular patch in a canine model (Chang et al., 2004), an intact layer of endothelium was found on the inner surface of the graft, together with host fibroblasts and smooth muscle cells in the acellular graft. This occurred to a lesser extent in the glutaraldehyde cross-linked acellular bovine pericardium used in the study. However, unintended intimal thickening and chondroid metaplasia were observed with both scaffold types. In another study, similar genipin-cross-linked acellular bovine pericardia were used to repair myocardial defects in a rat right ventricle model (Chang et al., 2005). Intimal thickening covered with endothelial cells was found on the endocardial surface. Mesothelial cells were also observed on the outer surfaces of the graft that is thought to resist adhesion. Smooth muscle cells, neomuscle fibres, neocollagen, neoglycosaminoglycans and neocapillaries were observed within the scaffold at four weeks post-implantation. Gelatin has been used as a biomaterial to repair peripheral nerve defects (Mligiliche et al., 1999; Gamez et al., 2004). Degradable scaffolds prepared from gelatin and cross-linked with genipin were used as a filling material in a tubular silicon chamber and served as extracellular matrix to guide nerve regeneration across a 10 mm sciatic nerve gap in rats (Liu et al., 2004). It was shown that the residue of the filling prevented the migration and elongation of axons (Verdu et al., 2002), and this was thought to be due to the extended degradation time of the gelatin filling (Bigi et al., 2002; Liu et al., 2004). When cross-linked gelatin was used to serve as tubular conduit (instead of as a filling material) to bridge similar nerve gaps, regenerated nerve fibres containing mostly unmyelinated axons with surrounding Schwann cells were found across the gap at six weeks when the conduit started to degrade and neovascularisation increased (Chen et al., 2005). There was also dense scar tissue in the outer area of the regenerated nerve. A thin layer of fibrous capsule was observed after subcutaneous implantation, together with the expected acute and chronic inflammatory responses. The concentration of the cross-linked was found to dictate the in vivo degradation when cross-linked gelatine–tricalcium phosphate scaffolds were implanted subcutaneously (Yao et al., 2004). Less cross-linked scaffold was associated with higher degradation while highly cross-linked scaffold exhibited minimal in vivo degradation. The use of higher amounts of crosslinked was also associated with increased implant fibrous capsule formation that was attributed to cytotoxicity of the cross-linked. When similar scaffolds were implanted to repair rabbits’ calvarial bone defects, the authors found that the scaffolds were gradually degraded and were replaced by new bone, even though this process was limited only at the bone–scaffold interface

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(Yao et al., 2005). The same composite cross-linked with glutaraldehyde induced a similar response (Chen et al., 1998). There was gradual degradation of the composite scaffold over time as the scaffolds were replaced by new bone, and acellular matrix was laid down from the edge of the defect centripetally. While the engineered scaffold may be crucial, other factors can accelerate the healing process. For example, the addition of hyperbaric oxygen therapy greatly increased new bone formation as shown by radiological and histomorphometry techniques (Chen et al., 2004). The presence of bioactive factors, such as BMP-2 in a degradable scaffold induced ectopic bone formation subcutaneously and intramuscularly, and appeared to accelerate resorption due to the action of osteoclasts (Liang et al., 2005; Takahashi et al., 2005; Kato et al., 2006; Yoneda et al., 2005).

12.10 Conclusion Biodegradable materials offer numerous advantages and have been developed as the solution to some of the problems associated with nondegradable synthetic biomaterials. Synthetic biodegradable biomaterials have the benefit of controllable chemistries and variables that can target degradation profiles and physical properties to suit each specific clinical application. Natural biodegradable biomaterials, alternatively, provide a cytocompatible substrate similar to host extracellular matrix to enhance wound healing, tissue remodelling and regeneration. While the host response towards implantation of a biomaterial follows a general pathway towards wound healing and tissue homeostasis, certain unwanted responses can be altered by modifying the chemistry and physical properties of the degradable biomaterial. As novel biodegradable materials are being developed and assessed, further knowledge and research data will help to improve our understanding on how a biodegradable material can be best utilised synergistically with host responses to facilitate tissue regeneration, while avoiding unwanted effects in the host, and allow timely elimination.

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winet, h., hollinger, j. o. and stevanovic, m. (1995) Incorporation of polylactidepolyglycolide in a cortical defect: Neoangiogenesis and blood supply in a bone chamber. J Orthop Res, 13, 679–89. witte, m. b. and barbul, a. (2003) Repair of full-thickness bowel injury. Crit Care Med, 31, S538–46. wong, c. h., samuel, m., tan, b. k. and song, c. (2006) Capsular contracture in subglandular breast augmentation with textured versus smooth breast implants: A systematic review. Plast Reconstr Surg, 118, 1224–36. wright, s. d. and silverstein, s. c. (1984) Phagocytosing macrophages exclude proteins from the zones of contact with opsonized targets. Nature, 309, 359–61. wu, c., jim, t. f., gan, z., zhao, y. and wang, s. (2000) A heterogeneous catalytic kinetics for enzymatic biodegradation of poly([epsilon]-caprolactone) nanoparticles in aqueous solution. Polymer, 41, 3593–7. yao, c. h., liu, b. s., hsu, s. h. and chen, y. s. (2005) Calvarial bone response to a tricalcium phosphate–genipin cross-linked gelatin composite. Biomaterials, 26, 3065–74. yao, c. h., liu, b. s., hsu, s. h., chen, y. s. and tsai, c. c. (2004) Biocompatibility and biodegradation of a bone composite containing tricalcium phosphate and genipin cross-linked gelatin. J Biomed Mater Res A, 69, 709–17. yannas, i. v., burke, j. f., huang, c. and gordon, p. l. (1975) Correlation of in vivo collagen degradation rate with in vitro measurements. Journal of Biomedical Materials Research, 9, 623–8. yokoyama, m., nakahashi, t., nishimura, t., maeda, m., inoue, s., kataoka, k. and sakurai, y. (1986) Adhesion behavior of rat lymphocytes to poly(ether)poly(amino acid) block and graft copolymers. J Biomed Mater Res, 20, 867–78. yoneda, m., terai, h., imai, y., okada, t., nozaki, k., inoue, h., miyamoto, s. and takaoka, k. (2005) Repair of an intercalated long bone defect with a synthetic biodegradable bone-inducing implant. Biomaterials, 26, 5145–52. zarbock, a., polanowska-grabowska, r. k. and ley, k. (2007) Platelet-neutrophilinteractions: Linking hemostasis and inflammation. Blood Reviews, 21, 99–111. zhao, w., byrne, m. h., boyce, b. f. and krane, s. m. (1999) Bone resorption induced by parathyroid hormone is strikingly diminished in collagenase-resistant mutant mice. J Clin Invest, 103, 517–24. zhao, w., byrne, m. h., wang, y. and krane, s. m. (2000) Osteocyte and osteoblast apoptosis and excessive bone deposition accompany failure of collagenase cleavage of collagen. J Clin Invest, 106, 941–9. zhong, s. p., campoccia, d., doherty, p. j., williams, r. l., benedetti, l. and williams, d. f. (1994) Biodegradation of hyaluronic acid derivatives by hyaluronidase. Biomaterials, 15, 359–65. zilla, p., brink, j., human, p. and bezuidenhout, d. (2008) Prosthetic heart valves: Catering for the few. Biomaterials. Feb; 29(4): 385–406. zolnik, b. s. and burgess, d. j. (2007) Effect of acidic pH on PLGA microsphere degradation and release. J Control Release, 122, 338–44.

13 Scaffold and implant design: Considerations relating to characterization of biodegradablity and bioresorbability D. W. HUTMACHER, Queensland University of Technology, Australia, and C. X. F. LAM, National University of Singapore, Singapore Abstract: Bioresorbable polymer-based systems and scaffolds are widely used in biomedical applications, particularly in the field of orthopaedics and bone tissue engineering. This chapter investigates the versatility of polycaprolactone-based scaffolds which allow optimised material properties and design to accomplish the intended regenerative strategy and bioresorption after achieving its purposes. Critical in vitro and in vivo investigations outline the homogeneous degradation mechanism via hydrolysis and functional performance; with respect to scaffold design and the dynamics of material evolution as it degrades. Key words: biomaterial, hydrolytic degradation, in vitro and in vivo degradation, tri-calcium phosphate, polycaprolactone, composite scaffolds.

13.1

Introduction

Typical applications and research topics of polymeric biomaterials include tissue replacement, tissue augmentation, tissue support, and drug delivery. In many cases the body needs only the temporary presence of a device/ biomaterial, in which instance biodegradable and certain partially biodegradable polymeric materials are better alternatives than biostable ones. Recent treatment concepts based on scaffold-based tissue engineering principles differ from standard tissue replacement and drug therapies as the engineered tissue aims not only to repair but also regenerate the target tissue. Cells have been cultured outside the body for many years; however, it has only recently become possible for scientists and engineers to grow complex three-dimensional tissue grafts to meet clinical needs. New generations of scaffolds based on synthetic and natural polymers are being developed and evaluated at a rapid pace, aimed at mimicking the structural characteristics of natural extracellular matrix. Currently, the design and fabrication of biodegradable synthetic scaffolds are driven by four material categories: (i) common, clinically established polymers, including polyglycolide (PGA), polylactides (PLA), and polycaprolactone 319

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(PCL), (ii) novel di- and triblock polymers, (iii) newly synthesized or studied polymeric biomaterials, such as polyorthoester, polyanhydrides, polyhydroxyalkanoate, polypyrroles, polyether ester amides, elastic shape-memory polymers, etc., and (iv) biomimetic materials, supramolecular polymers formed by self-assembly, and matrices presenting distinctive or a variety of biochemical cues. Several scaffold materials have been investigated for tissue engineering applications including polymers of synthetic origin such as polyglycolide and polylactide, and natural polymers such as collagen and chitin. In the 1990s, scaffold design and fabrication were directly based on commercially available bioresorbable implants and devices. More recent techniques were developed based on modern polymer processing and textile technologies. In the latter scaffold fabrication techniques, the viscous behaviour of the polymers above their glass transition or melting temperatures, and their solubility in various organic solvents, were two important characteristics that dictated the type of fabrication process used. This work has been reviewed elsewhere (Agrawal et al., 1997; Widmer et al., 1998). Aliphatic polyesters, such as PLA, PGA, and PCL, are by far the most commonly used synthetic polymer materials for manufacture of scaffolds and medical devices. As it is beyond the scope of one chapter to review the entire field, this chapter will focus on describing the biodegradation phenomena in detail and discussing comprehensive data sets of scaffolds/devices made of polycaprolactone for tissue engineering applications.

13.2

Biodegradation and bioresorption

The success of a treatment by tissue engineering or regenerative medicine pivots about its focus to restore, maintain, or improve tissue or organ function (Langer and Vacanti, 1993; Hutmacher, 2000); ideally, complete regeneration and elimination of any consequential foreign materials introduced would be necessary to restore the injury to its original and natural state. Therefore, knowledge of the degradation and bioresorption mechanism of the material of choice is critical for the success of a tissue-engineered construct (TEC), especially for orthopaedic applications where the controlled transfer of loading or function to the newly regenerated tissue as the scaffold degrades is essential. For this reason and necessity, the development of various bioresorbable materials, both natural and synthetic, has bloomed over the last decade. Polymer degradation and resorption is a complicated process that entails numerous dynamics and is influenced by many elements, which ultimately shape the chemical and physical properties. Technically, all polymers, regardless of their chemical structure or origin, degrade under appropriate conditions. The term ‘non-degradable polymers’ should be taken to indicate

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polymers that do not degrade during use or until a very long time after use (decades to centuries), rather than not to degrade at all (Gopferich, 1996). Also, earlier in 1995, Li and Vert (1995) defined ‘biostable’ materials as materials that have been used for a long time for both permanent and timelimited applications. While there were many debates over the accurate definition of polymer degradation, the term should remain relevant not only to tissue engineers and material scientists, but also to chemists, physicists, environmentalists, etc. It was generally established that degradation of a polymer is a chemical reaction and this results in a change in the polymer properties (Ottenbrite, 1992). However, many disputes have been on the specifics of this chemical reaction and whether the change in property produces a ‘good’ (enhancement) or ‘bad’ (deterioration) polymer quality. Fundamentally, degradation is a ‘chemical reaction’ involving a cleavage of the main backbone chain. Thus, authors have adopted Gopferich’s definition (Gopferich, 1996), ‘polymer degradation is defined as the chemical reaction resulting in a cleavage of main-chain bonds, producing shorter oligomers, monomers, and/or other low molecular weight degradation products’. This degradation principally proceeds via three known pathways, degradation by biological agents, by a chemical route or by physical means, as highlighted in Fig. 13.1. Hence, degradation would entail the production of lower molecular weight by-products; the descriptive term then, which is closely associated with degradation to account for the passage of these low molecular weight products, is ‘erosion’. Gopferich (1996) describes the process of erosion as the loss of material due to monomers and oligomers leaving the polymer. Biodegradation is a common term often misunderstood and misplaced. Grammatically, ‘bio’ means in relation to a living organism. Hence, ‘biodegradation’ would mean degradation mediated or partially mediated by a biological system (Ottenbrite, 1992; Vert, 1992). The organisms here would include bacteria, microorganisms from various sources (e.g. garden compost,

Degradation (Cleavage of the main polymer backbone chain)

Biological agents

Chemical processes

Physical processes

Microbial action Enzymes released Acid/alkali released

Oxidation Hydrolysis Photolytic

Pyrolytic (thermal) Mechanical Electromagnetic radiation

13.1 Classification of mechanisms of polymer degradation.

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sewage sludge), as in the biodegradation of industrial and commercial waste and packaging materials. These organisms are directly responsible for the chemical and physical attack on the polymer backbone chains, and include enzymes and acids released by a living organism due to interaction. It is erroneous, especially in circumstances involving larger living organisms such as animals and humans, to conclude that the instrument and mechanism of degradation within the body are solely ‘biodegradation’. The breakdown of polymer chains within such a system is evidently more complicated; it can also be incited by another or a combination of chemical processes (e.g. hydrolysis) or physical pathways (e.g. thermal, mechanical) as shown in Fig. 13.1. The final phase of biodegradation is when the polymer is reduced to simple basic elements and compounds, such as water and carbon dioxide (Vert, 1992). In the biomedical field, the main concern for a TEC scaffold material is that, after it degrades and breaks down, the degradation by-products produced should ‘disappear’ without a trace, thereby eliminating any foreign material in the host. In 1992, a bioabsorbable polymer was defined as a polymer that can be assimilated by a biological system (Ottenbrite, 1992) and the process of elimination of the initial foreign material with no residual side effects (Vert et al. 1992). The action of cells (such as phagocytosis), unlike bacteria and microbial action, do not actually participate in the backbone chain breaking of the polymers, but rather in the assimilation of monomer or oligomers units. In conclusion, the concept of bioresorption should encompass the degradation of the polymer system and the final elimination or metabolisation of the degraded by-products from the body. Generally, polymers differ in the rate and mechanism of degradation, which is determined by the energy required to break the bond and the location of the bond. Polymers with strong covalent bonds in the backbone (such as CJC), and with no hydrolysable groups, require long times for degradation. The agents of degradation could be heat, electromagnetic radiation (visible light, UV, γ-radiation), chemicals (water, oxygen, ozone, halogens) and biological agents (enzymes), or any combination of the these. Many biopolymers are susceptible to degradation due to the effect of water – hydrolysis. These include some naturally occurring polymers, such as polysaccharides and proteins, as well as some synthetic polymers, particularly condensation polymers with hydrolysable linkages such as amide, ester, urea and urethane (Li and Vert, 1995). Thus, the main mechanism of degradation for poly(hydroxy esters) such as polyglycolide (PGA), polylactide (PLA) and polycaprolactone (PCL) is simple chemical hydrolysis of the hydrolytically unstable backbone (Pitt et al., 1981; Pitt, 1990; Vert et al., 1992; Li, 1999).

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While degradation accounts for the backbone chain scission, there are two distinct erosion phenomena, namely bulk and surface erosion, which depict loss of the cleaved oligomers. Bulk erosion is generally observed for large-sized devices made with high molecular weight aliphatic polyesters (Li and Vert, 1995). It was discovered that the rate of random hydrolytic ester cleavage was faster at the interior of the device than at the surface, which was likely to be caused by an autocatalytic effect. Subsequently, an onset of weight loss and a decrease in the rate of chain scission occurs when the degradation by-products (oligomers and monomers) become small enough to diffuse out of the polymer surface (Gopferich, 1996; Li, 1999), while surface erosion typically involves the hydrolytic cleavage of the polymer backbone only at the surface (Ginde and Gupta, 1987; Gopferich, 1996). The formation of hollow cores during the degradation of poly(D,L lactic acid) pins has been explained by the formation of acidic residues, which self-catalyze the further degradation of the polymer (Vert et al., 1991; Grizzi et al., 1995). This autocatalysis generally occurs in the core of the devices where the acidic by-products cannot be neutralised by the outside medium. The formation of hollow cores can lead to a burst of acidic residues from the device. The acid burst of degrading implants made of poly(D,L-lactic acid) or other hydroxy acid based polymers has been associated with late inflammatory responses observed in vivo (Bostman, 1992). Surface erosion, which is characterised by the loss of material at the surface, occurs when the rate at which water penetrates the device is slower than the rate of conversion of the polymer into water-soluble materials. Surface erosion results in the device thinning over time while maintaining its bulk integrity, including geometry. Polyanhydrides and polyorthoesters are classical examples of polymers that undergo this type of erosion. The advantage of surface eroding polymers is the predictability of the erosion process (Gopferich et al., 1995). In a nutshell, whether a polymeric device undergoes (or is dominated by) bulk or surface degradation, depends on the diffusion–reaction phenomena striking a balance between water diffusion and reaction rate (Grizzi et al., 1995; Gopferich, 1996).

13.3

Hydrolytic degradation of polycaprolactone

Poly(ε-caprolactone) (PCL) is a semi-crystalline linear aliphatic polyester belonging to the poly(ω-hydroxy esters) family (Kimura, 1993). PCL exhibits several unusual properties not found among the other aliphatic polyesters, such as an exceptionally low glass transition temperature (Tg) of −60°C, low melting temperature (Tm) of ∼60°C and high thermal stability with decomposition temperatures (Td) of ∼350°C (Engelberg and Kohn, 1991). PCL can be prepared by ring-opening, addition polymerization using

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Degradation rate of bioresorbable materials O O

(CH2)5

C

n

13.2 Repeating molecular structure of polycaprolactone.

stannous 2-ethyl hexanoate or anhydrous stannic chloride as catalyst (Perrin and English, 1998a). Its molecular structure has a repeating molecular unit of five non-polar methylene groups (CH2) and a single relatively polar ester group (Fig. 13.2). Molecular weight is controlled by the addition of chain control agents, which usually are water, primary alcohols, amines, or some hydrogen active compounds. The mechanical properties of solid PCL (Mw = 44 000) had been reported as: tensile strength 16 MPa, tensile modulus 400 MPa and elongation at break 80% (Engelberg and Kohn, 1991). Having a low Tg and existing in a rubbery state at room temperature, PCL is one of the most flexible and easy to process biomedical polymers. Its high thermal stability means that it can be processed at temperatures above melting without significant molecular weight loss (Hakkarainen, 2002). Although it has a slow degradation rate, its clinical potential, thermal stability and pliability has allowed PCL to be extensively researched and tested for many fabrication methods for biomedical applications, including the Fused Deposition Modeling (FDM). Regarded as a non-toxic and tissue-compatible material, one of the earliest commercial usages of PCL was as a drug delivery device. Capronor®, a one-year implantable subdermal contraceptive device (Ory et al. 1983; Darney et al., 1989) was approved by the FDA (Pitt, 1990). Extensive in vitro and in vivo biocompatibility and efficacy studies were also performed leading to the introduction of the Monocryl® monofilament sutures, a copolymer of PGA (75%) and PCL (25%) (Bezwada et al., 1995). Recent suture research has led to the development of a P(LA/CL) suture, where PCL (20%) was used as a copolymer with PLA (80%) to enhance the degradation resistance and profile of the latter (Tomihata et al., 1998, 2001).

13.3.1 Evaluation methods Several analytical methods have been used to evaluate the degradation of PCL scaffolds. Details of these methods are described elsewhere (Lam, Hutmacher et al., 2008) but in essence they are: (i) (ii)

Gravimetric – mass loss (erosion) of the scaffolds determined on an electronic balance. Porosity – determined using a gas pycnometer;

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(iii)

Morphology – gold-sputtered and observed on a scanning electron microscope (SEM), at 15 kV (Phillips XL30 FEG). (iv) Mechanical properties – compression tests conducted an Instron 4302 Material Testing System in accordance with ASTM D695–96. In vivo samples tested on an Instron 5548 Uni-axial Testing System (Micro Tester INSTRON 5848). All specimens were compressed at a rate of 1 mm/min. Compressive stiffness and yield stresses were evaluated. (v) Crystallinity – determined using differential scanning calorimetry (DSC) (Pyris 6 DSC from Perkin-Elmer). Samples were scanned from 20 to 80°C at a ramp rate of 5°C/min, using nitrogen purge. The crystallinity fractions calculated were based on an enthalpy of fusion value of 139.5 J/g for 100% crystalline PCL (Pitt et al., 1981). (vi) Molecular Weights – determined using gel permeation chromatography (HPLC-GPC). Statistical analysis was performed on all the quantitative results using the Student’s t-test for comparing means from two independent sample groups. A confidence level of 95% was used (p < 0.05).

13.4

Hydrolytic degradation of medical polycaprolactone (mPCL) versus, research polycaprolactone (PCL)

mPCL and PCL scaffolds were fabricated into porous three-dimensional (3D) scaffolds using the FDM 3000 rapid prototyping system from Stratasys Inc. (Hutmacher et al., 2001; Lam et al., 2007). Table 13.1 lists the materials used to fabricate polymeric and composite scaffolds of mPCL and PCL. In vitro hydrolytic degradation was carried out in phosphate buffered saline (PBS) at the physiological temperature of 37°C. In vitro accelerated degradation studies were carried out in sodium hydroxide (NaOH) maintained at 37°C throughout the experiment (Lam et al., 2007).

Table 13.1 Materials used for fabrication of mPCL-based and PCL-based scaffolds Materials

Abbrev

Source

Research grade polycaprolactone (440744) Medical polycaprolactone (B6003-2) β-Tri-calcium phosphate

PCL

Aldrich Chemical Company Inc.

mPCL

DURECT Corporation

TCP

Shanghai Rebone Biomaterials Co Ltd

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13.4.1 Long-term hydrolytic degradation of mPCL and PCL scaffolds in PBS The in vitro hydrolytic degradation of mPCL and PCL scaffolds in PBS at 37°C was conducted for a long-term period of up to 60 months. For the study, mPCL scaffolds were prepared into dimensions of 4 × 4 × 4 mm3 while PCL scaffolds were prepared into 6.5 × 6.5 × 13.5 mm3 samples. Though the sample dimensions are different, the degradation mechanism would essentially be the same as the architecture and strut sizes remained the same. The mPCL scaffolds were pre-treated with 5 m sodium hydroxide (NaOH) for 12 hours to render the surface more hydrophilic, a typical preparation for scaffold used for cell culture and in vivo studies. One group of PCL scaffolds was left untreated while the other group was treated with 5 m NaOH for up to 3 days and then washed; this was designed to study the thorough effects of pre-treatments. The scaffolds were then submerged in individual tubes with 10 ml of PBS (pH 7.4) with screw-caps tightened and maintained at 37°C in an incubator. At the end of selected time-points, the scaffolds were removed and rinsed thoroughly with de-ionised water, dried and placed in an oven at 35°C for 12 hours for further drying. After 26 and 41 months, the overall gross morphology of all the mPCL and PCL scaffolds showed no visual differences compared to the day the experiment started. Upon closer examination of the physical surface characteristics (up to 12 months), the scaffold strut surface remained smooth and defect free. However, at 26 and 41 months, the surface was slightly coarser and numerous micro-cracks had appeared (Fig. 13.3). The microcracks observed on the mPCL and treated PCL scaffolds were present but not obvious against the already corrugated and rough surface. Good fusion of the struts at the surface was observed before the degradation commenced and throughout the 26 and 41 months. Closer examination of Fig. 13.4 (b and d) reveals that the detachment of scaffolds parts was neither localised nor specific to the expected joints and stress concentration regions, such as the strut joints and layer interface. Indeed, from the SEM micrographs, it was observed that the breaking off regions and interface encompassed the whole fusion region with no distinction between struts of different layers. This demonstrates excellent union between the layers and struts from the FDM process, which produced the scaffold as though it was made from one bulk material. Our previous published results also show that the well-fused and integrated scaffold strut network would hold, even through thinning of the strut thickness through surface degradation (Lam et al., 2007). The overall rate of mass loss was extremely slow. The mPCL scaffolds lost an average of 12.2% in 45 months while untreated and treated PCL scaffolds lost an average of 15.3% and 18.4% in 60 months, respectively.

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Before degradation

OPZB

15kV

200μm

x55

(a) mPCL scaffold at 0 mth (×55).

(b) Untreated PCL scaffold at (c) Untreated PCL scaffold at 0 mth (×64). 0 mth (×500).

(d) Treated PCL scaffold at 0 (e) Treated PCL scaffold at 0 mth (×500). mth (×64). After degradation

(f) mPCL scaffold at 26 mths (×30).

(g) mPCL scaffold at 26 mths (h) Untreated PCL scaffold at (×500). 41 mths (×30).

(i) Untreated PCL scaffold at 41 mths (×500).

(j) Treated PCL scaffold at 41 (k) Treated PCL scaffold at mths (×30). 41 mths (×500).

13.3 Gross morphology of in vitro degraded mPCL and PCL scaffolds observed by SEM.

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(a) Untreated PCL scaffold at 41 mths, with excellent fusion between the layer interface (×500).

(b) Treated PCL scaffold at 41 mths, with scaffold periphery broken off. Breaking off regions non-specific to joint areas (×100).

(c) Treated PCL scaffold at 41 (d) Untreated PCL scaffold at mths, with struts broken 41 mths, with scaffold during handling (×200). periphery broken off (×200). 13.4 Excellent fusion of scaffold struts from the FDM fabrication process, broken struts and broken off scaffolds parts with an intact cross-section even after 41 months of in vitro PBS degradation.

The mass loss profile of the mPCL and untreated groups proceeded gradually, whereas the treated PCL group showed a faster mass loss spike from 0 to 3 months, then a similar gradual behaviour thereafter (Fig. 13.5). While loss of material would typically correlate to porosity increase, the porosity fluctuations of the mPCL and PCL scaffolds after 45 and 60 months, respectively, were less than 5% (Fig. 13.6). The mPCL and untreated PCL scaffolds exhibited an initial gradual increase in mass loss and the rate of mass loss appeared to have increased after the 40-week period. One possible reason for this mass loss rate is due to fragments of the brittle scaffold falling off in the tube used for the degradation study, which was observed for all the samples. The pre-degradation of 3 days with NaOH, was thought to have caused the initial spike in mass loss at 3 months for the treated PCL scaffolds. The treatment not only increased the surface area, it also left cleaved chains and oligomers exposed which were easily hydrolysed and removed during the first 3 months; after removal of those susceptible chains, the surface was similar to the untreated group. Thus, the rate of mass loss was the same thereafter. This suggests a similar mechanism of degradation regardless whether the scaffold was untreated or treated. A more intense pre-treatment simply eroded a thin

Scaffold and implant design 0

PCL untreated PCL treated mPCL

Mass loss (%)

5

10

15

20

0

10

20

30 40 Time (months)

50

60

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13.5 Mass loss profile of PBS degraded mPCL and PCL scaffolds.

80

Porosity (%)

70

60

50

PCL untreated PCL treated mPCL

40

30

0

10

20

30 40 Time (months)

50

60

70

13.6 Porosity changes of PBS degraded mPCL and PCL scaffolds.

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Degradation rate of bioresorbable materials

surface layer of polymer; the other physical and chemical attributes were not significantly influenced. The average molecular weights of the mPCL scaffolds were observed to decrease gradually by 60% (Mw) and 72% (Mn) after 26 months. The average molecular weights of the PCL scaffolds fell by about 6% (Mw) and 7% (Mn), for both untreated and treated, after 12 months. By 41 months, they were reduced by about 86% (Mw) and 93% (Mn) of the initial molecular weights. It was observed that both molecular weights decreased gradually in unison. It was accepted from the literature that the hydrolytic degradation of poly(α-hydroxy esters) could proceed via surface or bulk degradation pathways. What determines the way in which degradation takes place is the diffusion–reaction phenomenon (Li et al., 1990; Grizzi et al., 1995; Gopferich, 1996). Surface degradation and erosion involves the hydrolytic cleavage of the polymer backbone only at the surface (Ginde and Gupta, 1987; Gopferich, 1996). This occurs when the rate of hydrolytic chain scission and production of oligomers and monomers, which diffuse into the surroundings, is faster than the rate of water intrusion into the polymer bulk. Typically, this results in thinning of the polymer device over time without affecting the molecular weight of the internal bulk of the polymer; the molecular weight would usually remain unchanged over the degradation period. Polyanhydrides and polyorthoesters are classical examples of polymers that undergo this type of erosion. The advantage of surface eroding polymers is the predictability of the erosion process (Gopferich et al., 1995). This system of erosion is desirable for drug delivery devices as the release of drugs can be related to the rate of erosion. Bulk degradation occurs when water is able to enter the whole polymer bulk, and results in hydrolysis throughout the entire polymer matrix. Random hydrolytic chain scission takes place and produces an overall reduction in molecular weight for the polymeric device. If equilibrium is reached for the diffusion–reaction phenomena, water is able to diffuse into the polymer, hydrolyse the chains, and the oligomers or monomers are able to diffuse out, and the erosion occurs gradually. When this diffusion–reaction balance for bulk degradation is upset, the degradation mechanism can lead to internal autocatalysis (Li and Vert, 1995). A single chemical reaction is said to have undergone autocatalysis if the reaction product is itself the catalyst for that reaction. In this case (for biodegradable polyesters), the hydrolysis by-products are carboxly and hydroxyl end groups. The oligomers and carboxyl end groups at the surface of the polymeric device can easily diffuse into the media as compared to the by-products at the centre. This resistance to diffusion depends mainly on the chemistry of the polymer and the thickness (Grizzi et al., 1995), both fundamental factors influencing the diffusion–reaction equilibrium. New

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carboxyl end groups are formed as each ester bond is cleaved, building up the concentration of carboxylic acid. A gradient of acidity forms and degradation is faster internally than at the surface, due to autocatalysis. An outer layer of less degraded macromolecules is formed, which controls swelling and osmotic exchanges. As degradation advances, the autocatalysis becomes larger internally, creating two distinct masses, a higher molecular weight skin layer and a degraded lower molecular weight interior. At this stage, the degradation mechanism could be characterised by a bimodal molecular weight distribution. When the inner oligomers become small enough, they diffuse rapidly through the outer layer into the medium; an onset of weight loss and a decrease in the rate of chain scission occurs (Gopferich, 1996; Li, 1999). A higher molecular weight hollowed structure is formed. The rapid release of these oligomers and carboxylic acid can result in adverse tissue reactions and inflammation in vivo (Bostman, 1992; Bergsma et al. 1993. 1995). From this study, after 26 and 41 months, the molecular weights of the mPCL and PCL scaffolds were observed to decline gradually by up to 93% of the original molecular weights, while mass loss was only less than 12%. This could well indicate bulk degrading kinetics for the mPCL and PCL scaffolds. It was found that autocatalytic hydrolysis is likely to result and develop when ester hydrolysis occurring at the chain ends (one or both) is faster than at the internal ester bonds (Shih, 1995). The main culprit for autocatalytic hydrolysis is the carboxylic acid end group. It exponentially multiplies if chain scission is selective of the chain ends. Shih (1995) found that the hydrolysis of poly-(D,L-lactide) was not random under acidic conditions; rather, the rate of chain-end scission was 10-fold faster than the rate of random scission of the internal chain bonds. PCL samples instead just underwent random hydrolytic scission. These results were likely attributed to the distance between functional groups, such that the distance of the carbonyl and alkoxyl was shorter in poly (D,L-lactide) than the functional groups of PCL. Therefore, with generation of the hydroxyl groups and especially the acid groups from the first degradation steps, these can catalyse later hydrolysis reactions. Although the PCL scaffolds degraded via the bulk degradation pathway, there was no obvious evidence of internal autocatalysis. This was evident from the mono-modal molecular weight results by GPC (data not shown) and intact cross-section of the scaffold struts (Fig. 13.4). No hollowed-out structure was detected. The stiffness of the mPCL scaffold decreased over time, while the stiffness of the PCL scaffolds (both untreated and treated) was observed to have increased about 1-fold at the end of 41 months. The yield stress of mPCL scaffolds was maintained initially, but after 6 months, it decreased to onetenth of the initial yield strength. The yield stress of the PCL scaffolds was observed to increase slightly over the initial period (at 3 months for the

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untreated and at 6 months for the treated), then decrease thereafter. The initial yield stress for both PCL groups was about 0.8 MPa and the final yield stress of the brittle scaffolds at 41 months was about 0.1 MPa. The mean compressive stiffness and yield stress for the scaffolds are presented in Figs 13.7 and 13.8.

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The scaffold fabrication process did not alter much of the polymers’ chemical and physical properties. The PCL scaffolds at the start of the experimental study were robust constructs, still retaining a certain rigidity combined with elasticity. However, after 26 months, throughout the whole scaffold structure it was brittle, as revealed from the compressive test by the stress–strain profile during the compressive loadings (data not shown). The scaffold also easily crumbled into many fragments once a small force was applied. For the PCL scaffolds, the yield stress decreased 8fold while stiffness increased, evidently indicating a brittle structure. This resultant brittleness at the end of 41 months could be attributed to the random hydrolysis which reduced the entire molecular weight of the polymer (Pitt et al., 1981). Material science demonstrates that low molecular weights polymers are more brittle than high molecular weight polymers, as long polymer chains can easily stretch and slide over one another to allow the polymer to be stretched and possess better plasticity/ductility. Short chains are like multiple defects, which are susceptible to brittlefailure; they break or crumble due to their inability to interlock well with each other when the material is loaded. The brittleness of the scaffold was also experienced during the handling of the scaffold while drying and preparing for the SEM viewing. Some of the scaffold struts at the periphery (Fig. 13.4c) were broken off during handling and manipulation of the samples during preparation for SEM examinations (Fig. 13.4 b and d). The mean crystallinities of the mPCL scaffolds were observed to decrease gradually from the start, whereas those of the PCL scaffolds were observed to increase slightly after 3 months, then remain unchanged and decrease slightly near the end of the 41-month period. The overall crystallinity decrease was about 9% and about 3% over the 26- and 41-month periods for the mPCL and PCL scaffolds, respectively (Fig. 13.9). The slow degrading kinetics of PCL are primarily due to its chemical functional groups, crystallinity and molecular weight. These basic factors gives the PCL scaffold its hydrophobicity, resistance to water and hydrolysis. Hydrophobic PCL has a high water-contact angle of about 80° before NaOH treatment and about 60° after (Htay et al., 2004). Additionally, it is recognised that hydrolysis of PCL is considerably slower than that of the polyesters such as PGA and PLA because of its crystallinity and high olefinic character; particularly its chemical structure of five methylene groups which confers a hydrophobic characteristic (Perrin and English, 1998b; Albertsson and Varma, 2002). In the field of drug delivery, the slow rate of release of proteins from PCL microspheres due to the high crystallinity and hydrophobic nature of PCL has also been reported (Lu and Lin, 2002). This hydrophobic nature of PCL is probably the chief determinant for its slow and long hydrolysis response, which is likely to stretch beyond

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5 years in vitro and beyond 2 years in vivo, based on our current studies. Copolymerising and blending causes the overall crystallinity to decrease by increasing the accessibility of ester linkages, thus enhancing the rate of hydrolysis. Hence, PCL copolymers and blends can be prepared having a wide range of physical properties and degradation rates (Perrin and English, 1998b).

13.4.2 Long-term hydrolytic degradation of composite mPCL/TCP scaffolds in PBS The long-term in vitro hydrolytic degradation of mPCL/TCP composite scaffolds in PBS at 37°C was conducted for up to 45 months. The mPCL was uniformly blended with submicron-sized TCP, 20% by weight, and fabricated via the FDM technology (Lam et al., 2007). For the study, mPCL/ TCP scaffolds were prepared in dimensions of 4 × 4 × 4 mm3. The distribution of TCP was assessed to be uniform and homogeneous using a microComputered Tomography (μ-CT) machine (Skyscan1076, Belgium) and Contrast Imaging and Tomography (PCIT) beamline at the Singapore Synchrotron Light Source (SSLS) (data not shown). At the end of selected time-points, the scaffolds were removed and rinsed thoroughly with de-ionised water, dried and placed in an oven at 35°C for 12 hours for further drying.

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After 26 months of degradation in PBS at 37°C, the overall gross morphology of the mPCL/TCP remained similar to the morphology at 0 months. As the scaffolds were NaOH treated, no obvious micro-cracks similar to the PCL scaffolds were observed (Fig. 13.10). However, more pits were observed on the strut surface, which were likely due to dissolution or erosion of the TCP particles (Fig. 13.10c). At the fractured surfaces, large pits averaging 100 μm in size were observed in the scaffold struts, which were probably due to the increased hydrolytic attack at TCP sites. As water invades the polymer matrix, it encounters the TCP particles dispersed within and this increases the interfacial surface area for hydrolysis; eventually the TCP dissolves first. Thus the hydrolytic attack of PCL/ TCP is considered to take place over a much larger surface area and volume with the presence of hydrophilic TCP. Although the regions of attack are possibly increased compared to a homopolymer mPCL scaffold, the TCP was uniformly distributed and not interconnected, hence the separated pitting was observed. As hypothesized, the TCP particles accelerated water diffusion into the polymer by physically acting as ‘conductors’ and channels for water diffusion and enhanced hydrolysis through their hydrophilic nature. Similar to the earlier study, the composite mPCL/TCP scaffold exhibited an extremely slow mass loss rate; it lost about 12% of mass after 45 months (Fig, 13.11). Its mass loss trend mimicked that of the treated PCL scaffolds closely. The mPCL/TCP scaffold also had an initial spike and a gradual slope thereafter, which also correlated with the porosity trend (data not shown). The spike in this case was likely due to the TCP content in the polymer matrix, which allowed the PBS to infuse and hydrolyse the polymer matrix much faster initially. The region around the TCP is likely to be more amorphous, as TCP particles could obstruct crystallisation and resulted in smaller or thinner regions of crystalline lamellae. Although throughout the

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degradation period the melting temperature of the PCL/TCP was constant at around 62°C, it was slightly lower than PCL (65°C); this could indicate less ‘perfect’ crystallisation (Lam et al., 2007). Numerous pittings were observed on the mPCL/TCP composite scaffold only, which were probably due to these more amorphous regions brought about by incorporation of TCP (Fig 13.10 c and d). Additionally, the disintegration of TCP would escalate the initial mass loss as well. The average molecular weights of the mPCL/TCP scaffold decreased gradually by 79% (Mw) and 85% (Mn) after 26 months, the graphical profile being analogous to its mPCL counterpart, although the molecular weight changes were statistically significant. This would indicate a slightly advanced degradation system for the composite scaffold. All the scaffolds retrieved after 26 months felt more brittle when handled, resulting in the breaking off of some of the scaffold bits at the periphery. This extreme brittle outcome of the scaffold can be explained by the drastic reduction of molecular weight. The mPCL/TCP composite scaffold displayed better mechanical properties, stiffness and yield strength than mPCL scaffolds. Due to the TCP reinforcement within the polymer matrix, it also appeared more rigid and inflexible and less translucent, similar to earlier reports (Lam et al., 2007). The compressive stiffness of mPCL/TCP plunged 59% after 3 months, then continued on a gradual decrease subsequently; by 26 months it had dropped by 74%. During 26 months, the yield stress of mPCL/TCP decreased gradually by 50% (Fig. 13.12).

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The crystallinity of the mPCL/TCP scaffold was observed to continually decrease gradually by 12% up to about 6 months and increase thereafter (Fig 13.13). This fluctuation demonstrates the dynamics of a polymeric material property over time. The initial crystallinity decrease was probably

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due to the enhanced hydrolysis brought about by the TCP incorporation as ascribed earlier. As the polymer chains strive for a lowered thermodynamic energy balance, an environment temperature (of 37°C) was sufficient to enable the chain mobility to achieve a lower energy level (Lam et al., 2008). Continued hydrolysis of polymer chains further increased its chances of crystallization by reducing the length of each chain and therefore increasing their mobility.

13.4.3 Accelerated hydrolytic degradation of PCL scaffolds in alkali media In 1992, Vert and co-workers (1992) addressed the bioresorption issue, stating that it is a process of elimination of the initial foreign material with no residual side effects. The action of cells (such as phagocytosis) does not actually participate in the backbone chain breaking of the polymers, but rather the removal of monomer or oligomer units produced in a later stage of degradation. As observed from the earlier section, degradation under simulated physiological conditions, polycaprolactone-based scaffolds showed only a 20% weight loss and about 90% molecular weight loss after more than 4 years. Theoretically, it is possible to increase the hydrolysis reaction of polyesters to study the hydrolytic degradation in a shorter period. Although accelerated hydrolysis may not be a direct correlation to the actual physiological environment, this procedure enables comparison of different scaffolds and materials, in particular evaluation of the degradability of novel materials undergoing similar degradation kinetics and erosion mechanisms (Buchholz, 1992). Accelerated degradation studies in an alkaline medium have previously been reported by a small number of groups for polymeric films and devices (Cam et al., 1995; Li et al., 2000; Htay et al., 2004; Lam et al., 2007). Therefore, using an accelerated system, researchers can gain insight into a polymer’s degradation behaviour, morphological and chemical changes in a more acceptable time frame. However, it should be noted that numerous factors affect the degradation mechanism and an accelerated system might not be able to fully simulate a proper long-term study, such as in this current case. The accelerated degradation study using alkaline media was conducted using sodium hydroxide (5 m) to accelerate the hydrolysis reaction. Scaffolds measuring 6.5 × 6.5 × 13.5 mm3 were submerged in individual tubes containing 10 ml of NaOH, with their screw-caps tightened and maintained at 37°C in an incubator. PCL scaffolds were removed at weekly timepoints, up to 6 weeks, after which time the remnants were not retrievable. At the end of the respective time-points, the scaffolds were removed, rinsed thoroughly with de-ionised water, dried and placed in an oven at 35°C for 12 hours.

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13.14 Gross morphology of PCL scaffolds degraded in NaOH observed by SEM.

The scaffold gross morphology and architecture (struts and bars) were observed to physically reduce in size over time, while the pore size and porosity increased. Shrunken scaffolds of various extents were retrievable for up to 4 weeks: By 5 weeks, they had completely disintegrated into pieces and only fragments were recoverable. After 6 weeks, only fragments and gel-like remnants were observed (Fig. 13.14). The crystallinity was observed to increase sharply during the first week and a slower and gradual (not statistically significant) increase was seen thereafter (Fig. 13.15). The crystalline melting temperature (65°C) was constant over the degradation period. Thermal analysis showed a 10% increase in crystallinity in the first week, which was probably due to the combined effect of the removal of amorphous regions and the annealing effect at 37°C. Also, as the polymer chains were continually cleaved, a process of re-crystallisation of the shorter chains could now occur. The crystallinity possibly neared saturation after 1 week, and only gradually increased to 78% at 5 weeks; the highest crystallinity reported for PCL was 80% (Pitt et al., 1981). Throughout this time, the melting temperature of the polymer remained unchanged, indicating that the change in the bulk crystalline properties of the PCL was not significant. The average molecular weights (Mw and Mn) of the alkali-degraded PCL scaffolds remained relatively constant over the 5 weeks of degradation. The Mw averaged around 143 700 and Mn averaged around 95 500. While many degradation studies (Pitt et al., 1981; Grizzi et al., 1995; Li, 1999) have reported a molecular weight decrease, the molecular weight seemed unaffected over time. These values were in

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agreement with the previous results, as degradation took place only at the surface of the struts, with any cleaved chains resulting in lower molecular weight products being easily dispersed into the medium. Thus, when the remaining scaffold was analysed, the unaffected and undegraded bulk characteristics prevailed, an important trait of the surface degradation mechanism. The rate of mass loss was found to be non-uniform throughout (Fig. 13.16). There was a linear and steep degradation profile observed up to 4 weeks (87% mass loss), and a slower degradation rate thereafter. Incremental changes in porosity were observed to be consistent and linear throughout the 4 weeks (Fig. 13.17). Mass loss reflects the onset of erosion. The characteristics of bulk erosion (bulk degradation kinetics) are based on mass loss occurring in the more advanced period of degradation, whilst for surface erosion, mass loss occurs from the beginning. Thus, the results of this study have stronger inclination toward a surface degradation mechanism. The initial steep mass loss was due to the chain cleavage of amorphous polymer regions, which were more susceptible to attack than crystalline regions. Although crystalline regions were typically attacked secondarily, the degradation of crystalline areas did proceed simultaneously, but at a slower rate (Li and Vert, 1995). The swift mass loss observations suggest that, since alkali-mediated degradation of the amorphous regions proceeded at a extremely fast rate, the remaining crystalline regions held together by these amorphous segments could also be dislocated from the main polymer bulk. The results of this study support

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the hypothesis that the mechanism of degradation occurred via erosion of the amorphous polymer regions, with associated fragmentation and erosion of the crystalline regions near the surface. After the rapid mass loss rate spanning 4 weeks and after the scaffolds had lost about 87% of the original mass, the mass loss rate decreased significantly. This development, based on an understanding of degradation kinetics, would suggest that the scaffolds left after 4 weeks were likely to

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be composed of highly crystalline components, up to about 77% crystallinity, which were highly resistant to hydrolytic attack. In the surface degradation mechanism, the diffusion of medium into the polymer matrix is slower than the hydrolysis at the surface. As discussed in the earlier sections, PCL is a hydrophobic polymer so this seems to be in agreement with surface attack and moreover explains the inability/slow rate of diffusion into the polymer construct, resulting in a surface degradation kinetic (Li and Vert, 1995; Perrin and English, 1998a; Albertsson and Varma, 2002). Mechanical properties declined gradually up to 1 week, then there was a more drastic drop up to 2 weeks, which gradually declined thereafter (Figs. 13.18 and 13.19). More recently, Zein et al. (2002) and Lam et al. (2002) have demonstrated the dependence of mechanical properties on scaffold porosity when designing and fabricating scaffolds via rapid prototyping. Logically, as porosity increases (loss of material bulk), its mechanical properties would deteriorate correspondingly. This decline has been found to follow a power law relationship (Gibson and Ashby 1997). The high dependence of mechanical properties on scaffold porosity had also been reported by Thomson et al. (1995) and de Groot et al. (1997). The compressive mechanical properties and integrity of the PCL scaffold was not proportional to the mass loss or crystallinity increase during the degradation period; more extreme mechanical property losses resulted only after two weeks. The general understanding was that, as more material was removed, the weaker the structure would become. Therefore, PCL scaffolds could be treated up to 1 week in NaOH to attain a porosity of about 75%

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and a crystallinity of 70%, without critically affecting the mechanical strength and molecular weight. The degradation of PCL scaffolds in an alkaline medium showed a surface degradation and erosion mechanism, contrary to the long-term PBS study. Hydrolytic degradation of polyesters, such as PCL, is based on diffusion– reaction phenomena between water diffusion and rate of reaction (chain cleavage). Also, while accelerated studies theoretically simulate hydrolysis, they cannot completely replace the need for a proper physiological degradation study.

13.5

In vivo degradation of polycaprolactone-based scaffolds

From a clinical perspective, polymers fall into three use categories defined as; (A) extracorporeal uses; (B) permanently implanted devices; and (C) temporary implants for various clinical uses. Categories (A) and (B) generally require biostable materials. Since a degrading implant does not require surgery for its removal, degradable polymers are especially advantageous in those applications where the implant is needed only temporarily. In Category (C), there are degradable sutures, barriers for surgical adhesions, polymeric scaffolds for cell or tissue transplants, temporary vascular grafts, temporary small bone fixation devices and implantable drug delivery systems. Thus, the main applications of degradable

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polymers for clinical uses can be classified into the following three groups: (a) temporary scaffold or support; (b) temporary barriers; and (c) drug delivery devices.

13.5.1 Six-month in vivo degradation of PCL scaffolds The influence and effects of in vivo degradation were studied using PCL and PCL/TCP scaffolds of 5 × 5 × 2.5 mm3 dimension for up to 6 months. The scaffolds were prepared into three groups; namely, (i) PCL scaffold group treated in 5 m NaOH for 12 hours, as per typical preparation for in vivo work; (ii) pre-degraded PCL scaffold (PD-PCL) group treated for 7 days in 5 m NaOH, to study of effects of pre-degradation and (iii) PCL/TCP group, left untreated. The scaffolds were implanted equally into two sites, subcutaneous (SC) and intramuscular (IM), in a rabbit model. A total of 18 scaffolds were implanted in each rabbit of the six mature New Zealand White rabbits (∼4 kg) used. Housing and feeding was done according to standard animal care protocols, at the Animal Housing Unit, Faculty of Medicine, National University of Singapore. A qualified surgeon performed the implantation, details of which can be found elsewhere (Lam et al., 2008). At specific time points of 3 and 6 months, three rabbits were sacrificed and the scaffolds retrieved. Retrieved scaffolds were treated in 10 ml of undiluted Clorox® solution (5.25% solution of sodium hypochlorite, NaOCl) for 20 minutes, which was sufficient for the removal of the integrated tissue but did not affect the polymer properties (Lam et al., 2008). The scaffolds’ gross morphology and architecture were observed not to have been significantly affected even after 6 months in vivo. Only some PD-PCL scaffolds were crushed, with struts detached. The pre-degraded state and in vivo factors were likely causes for the damage. Under magnified view, micro cracks were observed on the surface of the PCL and PCL/TCP groups (Fig. 13.20), irrespective of the implant sites. The surface texture of the PD-PCL was already very corrugated, thus any resultant micro cracks were not able to be detected (Fig. 13.21). At both 3 and 6 months, numerous empty micro pits and pores were observed on the struts of the PCL/TCP scaffolds; some pitting had partial TCP exposed. The exposed areas of the bioactive TCP particles were covered with needle-like fibres, which could be due to degradation or mineral nucleation effects of biological and cellular interaction (Fig. 13.22). The averaged mass losses of the PCL, PD-PCL and PCL/TCP were 2.8%, 5.1% and 7.3%, respectively. Although not statistically significant (p < 0.05), the averaged mass losses were observed to increase in the order: PCL < PD-PCL < PCL/TCP, irrespective of implant site (Fig. 13.23). Minimal mass

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13.20 SEM view of micro crack on PCL scaffold surface (32000×) after 6 months in vivo degradation (SC). Inset: Overview of PCL scaffold gross morphology (64×).

13.21 SEM overview of PD-PCL scaffold surface texture (2000×) after 6 months in vivo degradation (IM). Inset: Explanted PCL, PD-PCL and PCL/TCP scaffolds (left to right).

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13.22 SEM view of pittings on PCL/TCP scaffold surface (2000×) after 6 months in vivo degradation (SC). Inset: Close-up on degraded TCP particle with crystalline fibres exposed (8000×).

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loss of PCL scaffolds could be attributed to the high molecular weight and hydrophobicity, while the PD-PCL was more corrugated, had a larger surface area and, possibly with exposed chain ends, eroded slightly faster, comparable to earlier PBS degradation results (Lam et al., 2007). Obviously, PCL/TCP was designed to degrade faster with more hydrolytic intrusions of hydrophilic TCP and, together with loss of the TCP particles, supplemented the mass loss. It was observed that PCL scaffolds degraded in vitro in PBS had a similar mass loss rate to the in vivo PCL. This suggests that the biological environment and cells could have minimal or no effect. Literature has suggested that in vitro and in vivo PCL drug-delivery systems have similar degradation rates and has concluded that physiological enzymatic involvement was not a significant factor (Pitt et al., 1981; Sun et al., 2006). It was observed that chain scission of PCL was not accompanied by loss of low molecular weight PCL fragments until the molecular weight (Mn) had decreased to around 5000. Similarly, in vivo studies with rats reported that PCL degradation first proceeded with non-enzymatic bulk hydrolysis, and a transient initial inflammatory response occurred only for the first 2 weeks. After 9 months, only when the molecular weight had reduced to about 5000 did a loss in mass emerge, and subsequently the PCL fragmented. Concurrent intracellular degradation of low molecular weight PCL (Mn 3000) powders of 53–500 μm showed rapid degradation and absorption within 13 days inside the phagosomes of macrophage and giant cells. The sole metabolite was ε–hydroxy caproic acid (C6H12O2) (Woodward et al., 1985; Pitt, 1992). Mechanical stiffness of the in vivo scaffolds was evaluated non-destructively. The mean compressive stiffness of the different groups, PCL, PD-PCL and PCL/TCP were again independent of implant site. Due to the predegradation, the PD-PCL scaffolds exhibited the lowest stiffness, of about 3.5 MPa, throughout the 6 months, compared to the PCL scaffolds with an initial stiffness of 6 MPa, which eventually increased to about 9 MPa measured at 3 and 6 months. Prior to the degradation study, the reinforced PCL/ TCP had the highest stiffness of 17 MPa; over time it decreased and was maintained at about 12 MPa after 6 months. Composite PCL-based scaffolds would not only increase the functional mechanical properties compared to PCL scaffolds, they would accelerate the degradation rate. Also, the presence and degradation of the bio-ceramic material could neutralise the acidic by products from polymer degradation (Li and Vert, 1996). The average crystallinity of all the scaffolds increased over 6 months and again showed no dependence on the implant site (Fig. 13.24). However, as expected, PD-PCL had the smallest increase, which averaged a modest 1%. Due to the pre-degradation most of the amorphous regions was already hydrolysed away leaving much less polymer for re-crystallization. PCL scaffolds saw crystallinity increase 5% after 6 months; more crystalline regions

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would resist deformation better. The greatest increase of 7% was with the PCL/TCP scaffolds; the TCP allowed swifter intrusion of water to hydrolyse polymer chains and these resultant shorter chains could re-crystallise faster. Also, while initial TCP in the polymer matrix could impede crystallisation, the dissolution or degradation of TCP over 6 months removed this impedance, and the elimination of the TCP particles enabled higher chain mobility. Collectively, water is one of the main factors of crystallinity increase, as water uptake into the polymer will lower the Tg, and increase the crystallisation rate (Gopferich, 1996). The crystallinity results concurred with the stiffness of the scaffolds, as a higher crystallinity would confer the structure enhanced stiffness, as seen from Fig. 13.25. The dynamic effects of polymer degradation could be observed from the drop in stiffness of the PCL/TCP scaffolds, which was due to erosion of the reinforcement TCP particles, creating defects in the scaffold structure. However, the re-crystallization effect halted the stiffness drop and maintained it for 3 to 6 months; it could possibly have reduced the rate of stiffness decrease during the first 3 months as well. The combined effects of the elevated temperature (37°C), increased number of shorter polymer chains, and higher amount of water absorbed, increased the re-crystallisation rate. While, for the PCL scaffolds hydrolysis played a minimal role, re-crystallization enhanced the mechanical stiffness (Fig. 13.25). Thus, the mechanical

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properties of PCL-based scaffolds exhibited a strong relationship with their crystallinity, as long as the structural integrity and microfilament uniformity were not breached. The average molecular weights (both Mw and Mn) of the in vivo degraded scaffolds remained unchanged over the 6 months of implantation. These results agree with findings in the literature (Pitt et al., 1981; Woodward et al., 1985; Sun et al., 2006). The histological sections of the in vivo scaffolds were stained with Giemsa and are presented in Fig. 13.26. Giemsa is a general stain for inflammatory cells or defence cells (such as leucocytes). Leucocytes (granulocytes and lymphoid cells) would normally be stained red/orange while mast cells would be coloured purple (granules). Erythrocytes (red blood cells) would appear grey to yellow, while other cells’ cytoplasm would appear in varying shades of light blue with the nuclei appearing blue to violet. Collagen and muscle would appear pale pink. Macroscopically, no swelling or pathological signs were observed in any of the groups and the animals appeared healthy. Red spots could be observed on the surface of the scaffolds extending inwards, indicating blood vessel formation. There was good infiltration of host tissue into all the scaffolds and no inflammatory elements were observed over the six months. Regardless of the implant site, adipose-like cells filled the PCL scaffolds completely at three months, but at six months, tissue remodelling resulted

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13.26 Histological sections of 6-month scaffolds implanted at subcutaneous and intramuscular sites. All the SC scaffolds were observed to be filled with only fibroblastic cells and exhibited a higher number of blood vessels (arrows) than the IM site scaffolds. All the IM scaffolds were observed to be filled with adipose cells and an increasing number of fibroblastic cells since 3 months. After 6 months, no generalised inflammatory activities were detected.

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differently between scaffolds implanted intramuscularly and subcutaneously, which were partially and completely filled with fibroblastic cells, respectively. At both time points, the number of vessels detected was consistently higher in the subcutaneous scaffold, which was probably due to the nature of the cell-types in the different sites. As muscle fibres are usually arranged in the same direction and orientation as their function, they are unlikely to invade the scaffold that disrupts their arrangement. Also, the blood vessels usually run parallel and along these fibres, and the fibres did not invade the scaffolds; only a small number proliferated into the scaffold to nourish the adipose-like cells. In vivo, skeletal muscles (or myocytes) are more capable of hypertrophy to adapt to metabolic demands rather than mitotic division (hyperplasia). Thus, it was not expected to find myocytes within the scaffolds, and it is not fully understood why the scaffolds were filled with adipose-like cells. Adipose cells and fibroblasts are the two general cell populations that proliferate most easily, even under low nutrient levels. The presence of the adipose-like cells could be related to the metabolism at the particular site. The PD-PCL scaffolds seemed to elicit fibroblastic proliferation faster than normal as they were completely invaded with fibroblastic cells already at three months, probably due to increased porosity and surface roughness. The PCL/TCP subcutaneous scaffolds were invaded with fibroblastic cells at both time points, while the intramuscular scaffolds were occupied with adipose-like cells at three months and with both adipose-like cells and fibroblastic cells at six months. Similarly, more blood vessels were detected in the subcutaneous scaffold.

13.5.2 Two-year in vivo degradation of mPCL scaffolds The long-term influence and effects of in vivo degradation were studied with mPCL scaffolds of 15 mm diameter implanted into critical-sized defects in the rabbit calvarial; the preparatory work and surgery can be found elsewhere (Schantz et al., 2003a, b). After two years, the rabbits were sacrificed and the scaffolds retrieved. The calvarias were trimmed to size and scanned on a custom-built uCT (Jones et al., 2004a, b). Upon examination, the mPCL scaffolds were observed to be still in place and intact after 2 years, similar to findings from the in vitro PBS study. No fibrous encapsulation was observed in all the samples. Closer examination revealed that neo-calcified regions had replaced some regions of the mPCL scaffold. (Fig. 13.27a) Image analysis using the uCT revealed that the mineralised regions were of density similar to the surrounding bone. The calcified regions in the scaffold not only invaded the scaffold pore spaces, viewed as the complementary architecture of the scaffolds, but some regions completely replaced the scaffold struts (Fig. 13.27b). From this specimen

352

Degradation rate of bioresorbable materials Bottom

Top

Soft tissue / PCL scaffold R

R

L

(a) Gross overview of explanted rabbit calvarial. Implanted scaffolds observed to be intact with portions being replaced by calcified matrix.

Bone within defect region

L

(b) Micro-CT images of rabbit calvarial with scaffolds in the defect. Some calcified (high density) regions had taken up the complementary architecture of the scaffolds, while some regions of the scaffold were completely replaced by the calcified matrix. From this specimen, calcification invaded mainly from the bottom of the right defect and from the top of the left defect.

13.27 Gross overview and uCT analysis of the same rabbit calvarial explanted after two years in a calvarial defect.

(Fig. 13.27), it could clearly be observed that calcification occurred mainly at the bottom on both the right and left defect site and a greater volume detected at the top of the left defect site. Molecular weight analysis revealed that there was a 63% and 74% decrease in Mw and Mn, respectively, after two years. Also, crystallinity decreased by about 4%. Histological examinations are currently being carried out.

13.6

Conclusions

The usage of bioresorbable polymers in biomedical applications, and especially in orthopaedics, is growing, from screw-plate systems to bone scaffolds. It is obvious that their chief advantage might also be their greatest disadvantage: bioresorbability. While initial mechanical properties can be tailored to meet the requirements, degradation kinetics must be carefully understood as eventual functional support will hinge on this as the polymer degrades. While most literature describes only chemical changes, we have demonstrated mechanical properties changes as the crystallinity remodels and molecular weights fall. Degradation entails other phenomena than just

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‘elimination’ of the polymer: crystallinity and molecular weight changes are dynamics which affect the eventual rate of degradation, discharge of acidic by-products, stiffness changes and participation in intracellular degradation, the eventual metabolism of the polymer. Therefore, it is crucial for the study to be conducted using actual implant architectural parameters. Although accelerated systems may give a preview of the degradation pathway of polymers, our studies show this theoretical simulation cannot fully decipher all the kinetics as would an actual long-term in vitro degradation study or an in vivo one. Eventually, only in fully understanding the comprehensive degradation characteristics of the polymeric device can it be appropriately designed and considered for successful application.

13.7

References

agrawal, c. m., k. a. athanasiou, et al. (1997). Biodegradable PLA-PGA polymers for tissue engineering in orthopaedics. Porous Materials for Tissue Engineering 250: 115–28. albertsson, a. and i. varma (2002). Aliphatic polyesters: Synthesis, properties and applications. In Degradable Aliphatic polyesters. A. Albertsson, Springer: 1–40. bergsma, e. j., f. r. rozema, et al. (1993). Foreign body reactions to resorbable poly(Llactide) bone plates and screws used for the fixation of unstable zygomatic fractures. J Oral Maxillofac Surg 51(6): 666–70. bergsma, j. e., w. c. de bruijn, et al. (1995). Late degradation tissue response to poly(L-lactide) bone plates and screws. Biomaterials 16(1): 25–31. bezwada, r. s., d. d. jamiolkowski, et al. (1995). Monocryl suture, a new ultra-pliable absorbable monofilament suture. Biomaterials 16(15): 1141–8. bostman, o. m. (1992). Intense granulomatous inflammatory lesions associated with absorbable internal fixation devices made of polyglycolide in ankle fractures. Clin Orthop Relat Res (278): 193–9. buchholz, b. (1992). Accelerated degradation test on resorbable polymers. In Degradation Phenomena on Polymeric Biomaterials. H. Planck, M. Dauner and M. Renardy. Berlin; New York: Springer-Verlag: 67–76. cam, d., s. h. hyon, et al. (1995). Degradation of high molecular weight poly(Llactide) in alkaline medium. Biomaterials 16(11): 833–43. darney, p. d., s. e. monroe, et al. (1989). Clinical evaluation of the Capronor contraceptive implant: Preliminary report. Am J Obstet Gynecol 160(5 Pt 2): 1292–5. de groot, j. h., h. w. kuijper, et al. (1997). A novel method for fabrication of biodegradable scaffolds with high compression moduli. Journal of Materials Science – Materials in Medicine 8(11): 707–12. engelberg, i. and j. kohn (1991). Physico-mechanical properties of degradable polymers used in medical applications: A comparative study. Biomaterials 12(3): 292–304. gibson, l. j. and m. f. ashby (1997). Cellular Solids: Structure and Properties. Cambridge University Press.

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ginde, r. m. and r. k. gupta (1987). In vitro chemical degradation of poly(glycolic acid) pellets and fibers. Journal of Applied Polymer Science 33(7): 2411–29. gopferich, a. (1996). Mechanisms of polymer degradation and erosion. Biomaterials 17(2): 103–14. gopferich, a., d. karydas, et al. (1995). Predicting drug-release from cylindric polyanhydride matrix discs. European Journal of Pharmaceutics and Biopharmaceutics 41(2): 81–7. grizzi, i., h. garreau, et al. (1995). Hydrolytic degradation of devices based on poly(DL-lactic acid) size-dependence. Biomaterials 16(4): 305–11. hakkarainen, m. (2002). Aliphatic polyesters: Abiotic and biotic degradation and degradation products. In Degradable Aliphatic Polyesters. A. C. Albertsson. Berlin, Springer: 113–38. htay, a. s., s. h. teoh, et al. (2004). Development of perforated microthin poly(epsiloncaprolactone) films as matrices for membrane tissue engineering. J Biomater Sci Polym Ed 15(5): 683–700. hutmacher, d. w. (2000). Scaffolds in tissue engineering bone and cartilage. Biomaterials 21(24): 2529–43. hutmacher, d. w., t. schantz, et al. (2001). Mechanical properties and cell cultural response of polycaprolactone scaffolds designed and fabricated via fused deposition modeling. J Biomed Mater Res 55(2): 203–16. jones, a. c., b. milthorpe, et al. (2004a). Analysis of 3D bone ingrowth into polymer scaffolds via micro-computed tomography imaging. Biomaterials 25(20): 4947–54. jones, a. c., a. sakellariou, et al. (2004b). Investigation of microstructural features in regenerating bone using micro computed tomography. J Mater Sci Mater Med 15(4): 529–32. kimura, y. (1993). Biodegradable Polymers. In Biomedical Applications of Polymeric Materials, T. Tsuruta, T. Hayashi, K. Kataoka, K. Ishihara and Y. Kimura, Boca Raton, CRC Press: 164–90. lam, c. x. f., d. w. hutmacher, et al. (2008). Evaluation of polycaprolactone scaffold degradation for 6 months in vitro and in vivo. J Biomed Mater Res A (in press). lam, c. x. f., x. m. mo, et al. (2002). Scaffold development using 3D printing with a starch-based polymer. Materials Science & Engineering C-Biomimetic and Supramolecular Systems 20(1–2): 49–56. lam, c. x. f., s. h. teoh, et al. (2007). Comparison of degradation of PCL & PCL-TCP scaffolds in alkaline dedium. Polymer International 56(6): 718–28. langer, r. and j. p. vacanti (1993). Tissue engineering. Science 260(5110): 920–6. li, s. (1999). Hydrolytic degradation characteristics of aliphatic polyesters derived from lactic and glycolic acids. J Biomed Mater Res 48(3): 342–53. li, s. and m. vert (1995). Biodegradation of aliphatic polyesters. In Degradable Polymers: Principles and Applications. G. Scott and D. Gilead, Chapman & Hall 43–87. li, s. and m. vert (1996). Hydrolytic degradation of the coral/poly(DL-lactic acid) bioresorbable material. J Biomater Sci Polym Ed 7(9): 817–27. li, s. m., x. h. chen, et al. (2000). Hydrolytic degradation of PCL/PEO copolymers in alkaline media. J Mater Sci Mater Med 11(4): 227–33. li, s. m., h. garreau, et al. (1990). Structure property relationships in the case of the degradation of massive aliphatic poly-(alpha-hydroxy acids) in aqueous

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media.1. Poly(Dl-lactic acid). Journal of Materials Science – Materials in Medicine 1(3): 123–30. lu, c. h. and w. j. lin (2002). Permeation of protein from porous poly(epsiloncaprolactone) films. J Biomed Mater Res 63(2): 220–5. ory, s. j., c. b. hammond, et al. (1983). The effect of a biodegradable contraceptive capsule (Capronor) containing levonorgestrel on gonadotropin, estrogen, and progesterone levels. Am J Obstet Gynecol 145(5): 600–5. ottenbrite, r. m. (1992). Discussion of degradation terminology, In Biodegradable Polymers and Plastics. M. Vert, J. Feijen, A. Albertsson, G. Scott and E. Chiellini. Cambridge (England), Royal Society of Chemistry ••: 73–92. perrin, d. and j. english (1998a). Polycaprolactone. In Handbook of Biodegradable Polymers. A. Domb, J. Kost and D. Wiseman, Harwood Academic Publishers: 63–77. perrin, d. and j. english (1998b). Polyglycolide and polylactide. In Handbook of Biodegradable Polymers. A. Domb, J. Kost and D. Wiseman, Harwood Academic Publishers: 3–27. pitt, c. g. (1990). Polycaprolactone and its copolymers. In Biodegradable Polymers as Drug Delivery Systems. M. Chasin and R. Langer. New York, Marcel Dekker 71–119. pitt, c. g. (1992). Non-microbial degradation of polyesters: mechanisms and modifications. In Biodegradable Polymers and Plastics. M. Vert, J. Feijen, A. Albertsson, G. Scott and E. Chiellini. Cambridge (England), Royal Society of Chemistry: 7–19. pitt, c. g., f. i. chasalow, et al. (1981). Aliphatic polyesters.1. The degradation of poly (epsilon-caprolactone) in vivo. Journal of Applied Polymer Science 26(11): 3779–87. schantz, j. t., d. w. hutmacher, et al. (2003a). Repair of calvarial defects with customised tissue-engineered bone grafts – II. Evaluation of cellular efficiency and efficacy in vivo. Tissue Engineering 9: S127–S139. schantz, j. t., s. h. teoh, et al. (2003b). Repair of calvarial defects with customized tissue-engineered bone grafts – I. Evaluation of osteogenesis in a threedimensional culture system. Tissue Engineering 9: S113-S126. shih, c. (1995). Chain-end scission in acid-catalyzed hydrolysis of poly(D,Llactide) in solution. Journal of Controlled Release 34(1): 9–15. sun, h., l. mei, et al. (2006). The in vivo degradation, absorption and excretion of PCL-based implant. Biomaterials 27(9): 1735–40. thomson, r. c., m. j. yaszemski, et al. (1995). Fabrication of biodegradable polymer scaffolds to engineer trabecular bone. Journal of Biomaterials Science – Polymer Edition 7(1): 23–38. tomihata, k., m. suzuki, et al. (2001). The pH dependence of monofilament sutures on hydrolytic degradation. J Biomed Mater Res 58(5): 511–8. tomihata, k., m. suzuki, et al. (1998). A new resorbable monofilament stature. Polymer Degradation and Stability 59(1–3): 13–18. vert, m. (1992). Biodegradable Polymers and Plastics. Cambridge (England), : Royal Society of Chemistry. vert, m., s. li, et al. (1991). More about the degradation of La/Ga-derived matrices in aqueous media. Journal of Controlled Release 16(1–2): 15–26. vert, m., s. m. li, et al. (1992). Bioresorbability and biocompatibility of aliphatic polyesters. Journal of Materials Science – Materials in Medicine 3(6): 432–46.

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14 Drug release from bioresorbable materials M. WESTWOOD and D. S. JONES, Queen’s University of Belfast, Northern Ireland

Abstract: Over recent years the use (and hence the number) of biodegradable polymers for drug delivery applications has dramatically increased. These materials offer particular promise for implantable drug delivery systems as the drug delivery system does not require removal at the termination of drug release. The physicochemical properties, and, in particular, the rate of degradation, of biodegradable polymers directly affect the resultant rate of release from these systems. This chapter provides an overview of the different types of biodegradable pharmaceutical polymers, the mechanisms of drug release from these polymers and, in addition, provides an overview of the drug delivery applications of biodegradable polymers. Key words: polyester, polyanhydrides, erodible hydrogels, biodegradation, drug delivery.

14.1

Introduction

Biodegradable polymers were initially used in the formulation of oral dosage forms as a coating to reduce the rate of dissolution of the tablet (Baker, 1987). However, more recently their use as bioactive implantable drug delivery systems has become their major application within the pharmaceutical and related sciences. Biodegradable polymers are widely used for the controlled delivery of therapeutic agents, including drugs, peptides and genes (Fulzele et al., 2007). Furthermore, these polymers may be developed to deliver the agent at a specific target site at defined rates (Park et al., 2005). Since a degradable implant does not need to be removed surgically from the body after the conclusion of therapy, degradable polymers are useful in situations where the device is implanted for a finite period of time (Kohn et al., 2004). Degradable polymers are also useful for the delivery of large molecules (proteins and peptides), which may be unable to diffuse through non-degradable delivery systems (Park et al., 1993a). Whilst they are a relatively discrete class of polymers, the pharmaceutical 357

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importance of biodegradable polymers is steadily increasing and the development of new biodegradable polymers and bioactive devices are under current worldwide consideration. This chapter aims to provide an overview of both the chemical and degradation properties of pharmaceutically approved biodegradable polymers and, in addition, to describe examples of their use as drug delivery systems. Most polymers, regardless of their chemical structure, will eventually degrade under the appropriate conditions (Göpferich, 1996; Lam et al., 2007). The time frame of this degradation is the important factor in the use of degradable polymers as drug delivery systems. Polymer degradation is a chemical reaction that results in a change in polymer (mechanical and chemical) properties over time. Degradation occurs by cleavage of the bonds of a polymer to produce shorter oligomers, monomers; or other low molecular weight degradation products (Göpferich, 1996). A polymer is considered to be biodegradable if the degradation is due to environmental factors. The most common chemical process by which biodegradable polymers degrade is hydrolysis; however, oxidative and enzymatic processes may also cause degradation (Kohn et al., 2004). The enzymatic degradation process is mostly observed in naturally occurring polymers developed from polysaccharides and proteins (Göpferich, 1996). Polymers that are degraded via a passive hydrolytic mechanism are preferred due to a more reproducible in vivo response. To do so, the need for specific molecules to facilitate degradation is obviated (Katti et al., 2002). Degradation of a polymer specifically refers to the chemical process by which bonds are cleaved, whereas polymer erosion occurs whenever mass is lost from the polymer. Erosion is a physical process and polymer degradation contributes (but not exclusively) to erosion (Larobina et al., 2002). Degradable polymers may be classed according to their main mechanism of erosion; surface-eroding polymers, which degrade preferentially at the surface of the device (heterogeneously) and bulk-eroding polymers, which degrade throughout the polymer matrix (homogenously). However, a combination of these two mechanisms frequently occurs (Baker, 1987). The rate of drug release depends mainly on intrinsic polymer properties, such as the type of hydrolytically unstable bond, water permeability, morphology and molecular weight. Examples of degradable polymers that are currently being investigated for drug delivery applications include: polylactic acid and polyglycolic acid (and their copolymers), polycaprolactone, poly(ortho esters) and erodible hydrogels. The next section describes the physicochemical (and degradation) properties that are relevant to their use as drug delivery systems.

Drug release from bioresorbable materials

14.2

359

Examples of biodegradable pharmaceutical polymers

14.2.1 Poly(glycolic acid) and poly(lactic acid) and their copolymers Polyglycolic acid (PGA) and polylactic acid (PLA) are two of the most commonly utilised degradable polymers within the pharmaceutical sciences. Their preparation involves a condensation reaction at high temperatures, via ring opening of the appropriate cyclic lactone (Merkli et al., 1998) (Fig. 14.1). Ring opening polymerisation is a method that is typically used to develop hydrolytically sensitive polymers (Nair and Laurencin, 2007). The polymerisation reaction of the cyclic diester is accelerated with Zn- or Sn- based catalysts. Degradation of PGA and PLA results in the formation of naturally occurring non-toxic substances (lactic acid and glycolic acid, respectively) that are readily eliminated from the body as carbon dioxide and water (Sawhney et al., 1993; Park et al., 2005). PGA is the simplest linear aliphatic ester (Fig. 14.1a), and it exhibits a highly crystalline structure (45–55% crystalline) (Kohn et al., 2004; Nair and Laurencin, 2007). PGA has a very high tensile strength and very low solubility in organic solvents. Due to the additional methyl group, PLA (Fig. 14.1b) is a more hydrophobic molecule than PGA. PGA is achiral; however, PLA exists in two steroisomeric forms that can result in four morphologically different polymers. However, only three of these forms are used in practice; the two stereoregular polymers, poly(d-lactide) DPLA and poly(l-lactide) LPLA, which produce semicrystalline polymers, and the racemic form (a)

(b)

14.1 Ring opening polymerisation of selected cyclic lactones to give (a) poly(glycolic acid) and (b) poly(lactic acid).

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Degradation rate of bioresorbable materials

poly(d,l-lactide) DLPLA, which always produces an amorphous polymer. LPLA is preferred over DPLA as d-lactic acid is not readily metabolised by the body (Park et al., 1993b). LPLA is approximately 40% crystalline, with high tensile strength (Nair and Laurencin, 2007). Amorphous DLPLA displays a much lower tensile modulus, which makes it more suitable for drug delivery applications. All polyesters (including PLA, PGA and their copolymers) possess an ester functional group in the polymer backbone that is prone to hydrolysis, i.e. in the presence of water (and in particular, at extremes of pH). This results in the generation of polymer fragments and monomers with one carboxylic acid and one hydroxyl end-group (Gupta and Kumar, 2007; Edlund and Albertsson, 2003). The hydrolysis reaction is catalysed by both Lewis acids and bases (Li, 1999). With respect to bioactive systems, the degradation rate is critical for determining the release rate of a drug from the system. The ability of water to access the labile bonds is one of the main factors in controlling (and determining) the rate of degradation (Xu et al., 2006). As a result, the hydrophilicity of the polymer backbone is significant in determining the degree of water penetration into the bulk of the polymer. Consequentially, PGA will degrade over a shorter time period than the more hydrophobic PLA (Khon et al., 2004). Water content in PGA during hydrolysis can rise to 42%, while in PLA it is limited to only 2% (Park et al., 1993c). Water uptake into PLA and PGA is faster than the rate of degradation and therefore degradation occurs homogenously in the bulk of the material. As degradation progresses, bond cleavage results in the formation of more carboxylic acid end groups on the polymer fragments. The rate of hydrolysis is therefore accelerated in a mechanism known as autocatalysis. Fragments that are formed inside the matrix are trapped and a low pH environment in the centre of the material results. Soluble polymer fragments close to the polymer surface are free to diffuse out of the matrix, leaving a porous surface with a higher MW than the centre (Hurrell and Cameron, 2001). Following implantation in vivo, these polymers are exposed to an aqueous medium at pH 7.4. This results in the neutralisation of the surface carboxyl groups and contributes further to the pH difference between the surface and interior of the material (Li, 1999). Therefore autocatalysis is more predominant in the centre of the matrix than at the surface. The rate of degradation is further increased by the build-up of water-soluble degradation products within the bulk of the polymer, which results in the osmotic uptake of water into the matrix. These processes result in a homogenous bulk-erosion behaviour, i.e. degradation occurs in the complete cross-section of the polymer matrix; however, it will occur at a faster rate in the centre of the sample (Li, 1999; Lee et al., 2003). The molecular weight of PLGA has been shown to decrease exponentially for a period of time (detected as

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decreased intrinsic viscosity) before weight loss is observed in the polymer (Kenley et al., 1987). This confirms the bulk-erosion mechanism of these materials: if only surface erosion was occurring, weight loss of the material would be detected prior to the reduction in molecular weight. Bulk degradation of PLA, PGA and their copolymers occurs in three stages: (i)

Incubation; the time required for water penetration into the polymer (ii) Induction; hydrolytic degradation of the polymer molecules occurring with no significant weight loss, (iii) Erosion; at a critical MW, dissolution/release of the polymer fragments occurs, with the start of the polymer erosion process (Lee et al., 2003). The critical molecular weight (MW) reached by polyesters prior to weight loss is between 2000 and 5000 (Baker, 1987). These mechanisms are shown schematically in Fig. 14.2 (Hurrell and Cameron, 2001). The morphology of the polymer also influences the degree of water penetration. Due to the densely packed structure within the crystalline regions of LPLA and DPLA, they are more resistant to hydrolysis than amorphous DLPLA (Maurus and Kaeding, 2004). Hydrolysis of the ester bonds will begin in the amorphous regions of the polymer. The molecular weight will decrease rapidly in this time period; however, the degree of crystallinity may increase because chain cleavage allows the polymer chains greater freedom to move, resulting in recrystallisation. Only when most amorphous regions have been degraded will the hydrolysis occur in crystalline regions. Zong et al. (2006) studied the changes of crystalline and morphological parameters during in vitro degradation of PGA homopolymers and PGA-co-PLA copolymers using wide angle X-ray diffraction and small angle X-ray scattering techniques. The MW and polydispersity decreased

Degradation products H2O

(i)

(ii)

(iii)

14.2 Schematic of bulk erosion: (i) water diffuses into the polymer matrix (incubation); (ii) degradation products build up in the centre of the material (induction); (iii) soluble polymer fragments diffuse through the outer, lessdegraded shell when a critical molecular weight is reached (erosion) (Hurrell and Cameron 2001).

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Degradation rate of bioresorbable materials

significantly during the first two weeks of degradation. However, the mass degree of crystallinity increased rapidly within the first two weeks. The initial polymer morphology does not have a significant effect on the degradation times. Samples of PGA that were prepared by either quenching or annealing developed similar degrees of crystallinity during the induction period and degraded at the same rate (Hurrell and Cameron, 2002). However, theophylline release from the materials was affected due to uneven distribution of the drug in the annealed sample. The copolymers of PLA and PGA (PLGAs) have been widely employed as their degradation rate can be engineered by changing the ratio of the polymers and/or their molecular weight (Park et al., 2005). PLGAs are less crystalline than the pure homopolymers, which greatly decreases the time of degradation (Miller et al., 1977). The rate of degradation of PLA can be increased in a controlled manner with copolymerisation with up to 50% PGA. Increasing the concentration of PGA further increases the degradation time. For example, it has been reported that the half-life of the polymers and copolymers in tissue decreased from 5 months for PGA homopolymer to a minimum of 1 week with 50 : 50 PGA-PLA copolymer and rapidly increased to 6.1 months for PLA homopolymer (Fig. 14.3). The molecular weight of the polymer is a further factor in determining the rate of degradation; low MW polymers will have a higher content of carboxylic acid groups at the chain ends, which accelerate the rate of degradation (Li, 1999; Park et al., 2005). The induction time required prior to the onset of erosion in DLPLA has been shown to increase with an increase MW (Pitt et al., 1981). The rate of bond cleavage is not constant, as due to

T½ Months

6

4

2

0 0

PLA

100

100

PGA Copolymer ratio

0

14.3 Half-life of PGA and PLA homopolymers and copolymers implanted in rat tissue. (Miller et al., 1977).

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autocatalysis it increases exponentially and the rate of cleavage of the first few bonds is very slow compared to the hydrolytic cleavage of the final bonds. Therefore, the MW effect on the rate of hydrolysis is mainly relevant at the beginning of degradation. The size of the device also affects the rate of degradation of polyesters. For example, Grizzi et al. (1995) examined the degradation rates of varioussized devices made of DLPLA and reported that the degradation rate was significantly dependent on the size of the device. In particular, as the thickness of the device was increased, the rate of degradation was increased. This was accredited to the formation of the more stable outer layer of the device during degradation with the resultant autocatalysis in the interior of the device. A sample with very small dimensions degrades at a slower rate as no build up of polymer fragments can occur in the interior of the device. Tamada and Langer (1992) investigated the erosion of thin and thick PLGA disks and showed that both geometries required the same time to fully degrade, the thicker disc displaying twice the degradation rate of the thinner sample. Degradation occurred uniformly across the whole material; a sample of twice the volume produced degradation products at twice the rate as the thinner sample, with the same overall erosion time. For bulk-eroding polymers, the volume of the device, and not the surface area, controls the rate of degradation. The incorporation of drugs into the polymer, an important consideration in the design of pharmaceutical systems, may also affect the rate of degradation. The interaction between the polymer and drug is a critical parameter in this respect. As the vast majority of drugs are either acids or bases, the inclusion of these compounds may modify biodegradation of polyesters. If the predominant effect of the drug is base catalysis, then degradation is accelerated; however, if the main effect is to neutralise carboxylic end groups, then degradation may be slowed down (Li, 1999). Non-ionised drugs can either speed up or retard the polymer degradation rate. A hydrophilic drug may increase water uptake and speed up the degradation rate, while a hydrophobic drug can have the opposite effect.

14.2.2 Poly(ε-caprolactone) Poly(ε-caprolactone) (PCL) (Fig. 14.4) is an aliphatic polyester that undergoes degradation through an autocatalytic hydrolysis reaction (Baker, 1987; Duda et al., 1998). This polymer is prepared by a catalysed ring-opening polymerisation of ε-caprolactone through an anionic, cationic, free radical, or coordinate polymerisation (Jiang et al., 2003). PCL is a semicrystalline polymer with a very low glass transition temperature (−60°C to −70°C) (Park et al., 1993d) that has been investigated as a biodegradable polymer since the 1970s (Jiang et al., 2003). The hydrophobicity and high crystallinity

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Degradation rate of bioresorbable materials

14.4 Poly(ε-caprolactone).

of this polymer result in very slow degradation and it is therefore suitable for long-term delivery of therapeutic agents, for periods greater than one year (Park et al., 2005; Sun et al., 2006). This polymer also displays a bulkerosion mechanism with an initial induction period in which a reduction in MW of the polymer is observed. Only when a critical MW is reached is a significant weight loss observed (Merkli et al., 1998). Pitt et al. (1981) studied in vivo degradation of PCL in rabbits and rats for periods up to 213 weeks. Degradation proceeds by random hydrolytic cleavage of the ester linkages. The MW decreased exponentially with degradation time until the MW was about 5000, at which point erosion began. The autocatalytic effect of hydrolysis in PCL is probably due to the increase in hydrophilicity of the polymer as the chains are cleaved, producing hydrophilic chain ends. Hydrolysis produces ε-hydroxycaproic acid, which may be further fragmented by phagocytosis. PCL is known to be permeable to drugs whose molecular weight is less than 400 D (Pitt et al., 1979). The high permeability of polycaprolactone and its copolymers, coupled with a long controllable induction period prior to polymer weight loss, enables the development of delivery devices that are based mainly on diffusion-controlled delivery of the drug (Pitt et al., 1979). An increase in PCL crystallinity reduces the polymer permeability and can significantly reduce drug diffusion with a reduction in drug release rate. Due to the slow degradation time, the microenvironment inside PCL during degradation is not acidic. This allows the incorporation of acid-sensitive therapeutic agents into PCL, which may not be incorporated into PLA or PGA (Sastre et al., 2004). For many drug-delivery applications, the degradation of PCL is too slow to directly influence drug release. Therefore, to increase the rate of degradation, PCL has been copolymerised with other more hydrophilic cyclic esters, including L-LA and γ-butyrolactone (Duda et al., 1998; Jiang et al., 2003). The properties of these copolymers can be controlled to exhibit various degradation rates and permeability behaviours. The effect of copolymerisation on the degradation rate is mainly due to the reduction of crystallinity of the material (Park et al., 1993b). Manipulation of the degradation rate of PCL may also be achieved by addition of acidic or basic additives to devices fabricated from this polymer, thus facilitating polymer degradation (Pitt and Gu, 1987). For example, the incorporation of oleic acid caused an increase in the rate of degradation that was proportional to the amount of

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acid added. An increased rate of degradation was also obtained with the addition of decylamine, as the amine functional group can launch a nucleophilic attack on the ester bond. Lin et al. (1994) incorporated primary, secondary and tertiary alkylamines into PCL. Primary alkylamines caused the most rapid degradation, whereas the effect of the tertiary alkylamines was not significant. PCL typically exhibits a low tensile strength and high flexibility (Wang et al., 2005; Duda et al., 1998).

14.2.3 Poly(ortho esters) The stability of a water-sensitive drug from a bulk-eroding polymer may be compromised by the ingress of water prior to erosion and release. A surfaceeroding polymer does not expose the water-labile agent until release occurs, thus maintaining the stability of the drug (Katti et al., 2002). Drug release from a bulk-eroding polymer is not well controlled and release is dominated by Fickian (non-linear) diffusion (Heller, 1993). Therefore, extensive investigations into the development of polymers that can erode by a predictable surface hydrolysis process to control the rate of drug release have been carried out. Depending on the surface erosion rate and device design, the device may release the drug at a constant rate (Heller and Barr, 2006). A representation of a surface-eroding polymer is shown in Fig. 14.5. Hydrolysis is initiated with the penetration of water into the polymer bulk. If hydrolysis is slow compared with diffusion, the complete bulk of the polymer matrix is affected, as is detected in PLA and PGA (Göpferich, 1997). However, if the rate of hydrolysis is faster than the rate of diffusion into the polymer, degradation is related to a surface erosion process. Poly(ortho esters) (POEs) are examples of polymers that exhibit surface erosion, due to their high hydrophobicity and low rate of water uptake (Park et al., 2005). These systems undergo degradation through hydrolysis of an acid-labile ortho ester linkage (Zignani et al., 1997), which is more susceptible to hydrolysis than an ester linkage. However, the polymer backbone of POEs is sufficiently hydrophobic to prevent water penetration

Surface degradation 14.5 Schematic representation of idealised surface erosion.

366 O

Degradation rate of bioresorbable materials O R

O

O

O O

n

POEI

R

O O O R

O CH2 C

n

O

O CH (CH2)4

O O O

O

(H) CH3 O

O O

CH

C

n

POE III

POE II O O

n

R'

O O

O O O R

n

POE IV 14.6 Various generations of poly(ortho esters).

and therefore initial hydrolysis is confined to the surface. They have been designed in order to achieve zero-order (linear) release. Since their first development in the late 1970s, poly(ortho esters) have evolved through four generations of biodegradable polymers (POE I, POE II, POE III, and POE IV), which are shown in Fig. 14.6 (Heller et al., 2002). POE I was first synthesised by the transesterification of diethoxytetrahydrofuran with a diol, under a high vacuum at high temperatures (Zignani et al., 1997). The degradation product of hydrolysis is γ-butyrolactone, which is rapidly converted to γ-hydroxybutyric acid (Heller et al., 2002). To prevent this acid causing an uncontrolled autocatalytic degradation reaction, a basic salt, e.g. sodium carbonate, is included in the matrix. The unfavourable synthesis conditions, and need for a basic excipient in the matrix, limit the applications of this polymer (Heller and Barr, 2006). POE II synthesis is relatively simple compared with POE I. It is prepared by the addition of a diol with a diketene acetal (initially 3,9-bis(ethylidene)2,4,8,10-tetraoxaspiro(5,5)undecane) at ambient temperature (Heller et al., 1980). The reaction is highly reproducible, the molecular weight of the polymer can be easily controlled by changing the stoichiometry (Heller et al., 2002), and the properties of the polymer (cross-linked or linear) can be altered by the choice of diol used in the synthesis (Zignani et al., 1997). Furthermore, the need for a basic excipient is eliminated. The hydrolysis products are initially neutral and only slowly convert to an acid once they have diffused away from the polymer. Therefore, no acid autocatalysis takes place. However, this polymer is extremely hydrophobic; low water uptake by these POEs (0.3 to 0.75%) results in a very stable polymer with a very long degradation time (Fig. 14.2) (Heller, 1993). To achieve desirable degradation rates, acid excipients have been added to the polymer, with only limited success due to the diffusion of the acid out of the polymer (Ng et al., 1997; Heller et al., 2002).

Drug release from bioresorbable materials

367

Both POE I and POE II are solid polymers at room temperature, whereas POE III has a flexible backbone and is semi-solid at room temperature. It may even be a viscous liquid, which allows the possibility of its being injected into the body (Heller et al., 2002; Kohn et al., 2004). The viscous nature enables the incorporation of therapeutic agents into the polymer matrix at room temperature without the need for solvents. Therefore, sensitive therapeutic agents such as proteins can be loaded into the polymer. These polymers exhibit a good relationship between polymer degradation and drug release, and they allow the incorporation of drugs through a simple mixing process at room temperature (Zignani et al., 1997). However, the synthesis is not as simple as for POE II, requiring long reflux times, and does not produce reproducible results (Heller et al., 2002). The fourth generation of poly(ortho esters) (POE IV) are a modification of POE II. To increase the rate of hydrolysis to a desirable rate and achieve greater control of drug release, short polymer segments based on lactic or glycolic acid are incorporated into the polymer backbone. An acid diol is prepared by the reaction between a diol and either lactide or glycolide. The exact structure of the latent acid diol is not important in controlling the erosion rate – the total concentration of α-hydroxy acid segments within the polymer is the determining factor (Heller et al., 2002). Figure 14.7 shows that the rate of erosion (as a function of weight loss) is well 100

Weight loss (%)

80

60

40

20

0 0

10

20

30

40 50 Time (days)

60

70

80

14.7 Weight loss as a function of poly(glycolic acid) dimer content for a polymer prepared from 3,9-diethylidene-2,4,8,10-tetraoxaspiro (5.5) undecane, transcyclohexanedimethanol glycolide (tCDM/Gly) and trans-cyclohexanedimethanol (tCDM). 25/75 tCDM-tCDM/Gly (■), 50/50 tCDM-tCDM/Gly (□), 25/75 tCDMtCDM/Gly (•), 90/10 tCDM-tCDM/Gly (°). 0.05 M phosphate buffer, pH 7.4, 37°C (Ng et al., 1997).

368

Degradation rate of bioresorbable materials

controlled by the concentration of glycolic acid segments incorporated in a POE IV (Ng et al., 1997). The hydrophilicity of the polymer will determine the effect of the latent acid on erosion rates. The concentration of latent acid has little effect on very hydrophobic polymers, while hydrophilic polymers are very sensitive to the presence of latent acid (Heller et al., 2002). The degradation of the LA or GA segments results in the formation of polymer fragments with carboxylic acid chain ends that lower the pH and catalyse the ortho ester hydrolysis. Further cleavage results in the formation of free lactic or glycolic acid, which is the main factor in catalysing the hydrolysis of the ortho ester bond. Hydrolysis of ortho esters generates the diol, or mixture of diols, used in the synthesis and also pentaerythritol dipropionate, which is followed by further hydrolysis to produce pentaerythritol and propionic acid (Heller et al., 2002). In addition, the hydrophilic fragments generated increase water uptake into the surface layers with a further increase in hydrolysis. These processes lead to a constant rate of weight loss from the polymer and also linear release of hydrolysis products, which indicate that the polymer is eroding predominantly through surface degradation. These polymers are able to control drug release for relatively short periods of time, which can vary from days to weeks (Einmahl et al., 2001). The mechanical properties of this group of polymers may be easily altered by varying the selection of diols used. By using flexible diols in the synthesis, these polymers may be formulated as viscous liquids, which can then be administered via injection. Increasing the proportion of a more rigid diol results in an increase in the glass transition temperature and allows the preparation of solid materials (Heller and Barr, 2004; Ng et al., 1997). Investigations into the biocompatibility of POE IV have demonstrated the biocompatibility of the polymer (Heller and Barr, 2004). Therefore, of the four different generations of POEs, POE IV is considered to be the biomaterial with greatest potential, having not only a scalable synthetic procedure, but also the ability to provide well-controlled release profiles for a wide range of pharmaceutical agents. The rate of drug release can also be controlled by the alkyl chain length of the diol incorporated in the polymer (Sintzel et al., 1998).

14.2.4 Polyanhydrides Many polyanhydrides have also exhibited an advantageous surface erosion behaviour. Polyanhydrides (Fig. 14.8) are one of the most hydrolytically unstable polymers developed for controlled drug delivery purposes. Polyanhydrides consist of a backbone of sparingly soluble monomeric carboxylic diacids joined by highly reactive anhydride linkages (Katti et al., 2002; Jain et al., 2005). Hydrolysis of the linkages results in the formation of two carboxylic acids.

Drug release from bioresorbable materials O *

O

C O

369

C R

n

14.8 General structure of polyanhydrides.

Polyanhydrides have been synthesised by ring opening polymerisation, interfacial polymerisation, and bulk melt polycondensation. Only the last method produces polymers with high molecular weight (Domb et al., 1993). The presence of aliphatic and aromatic groups in the monomer confers a hydrophobicity to the polymer; water is prevented from penetrating the network and erosion therefore occurs on the surface (Larobina et al., 2002). Aliphatic polyanhydrides may degrade in a number of days, whereas more hydrophobic aromatic polyanhydrides may take years to fully erode. Copolymerisation of the two, leads to a range of intermediate degradation rates depending on the aliphatic–aromatic composition (Kohn et al., 2004). Increasing the hydrophobicity of the aliphatic component can also decrease the degradation rate. An increase in the number of carbons in an aliphatic polyanhydrides from 1 to 6, resulted in a 1000-fold reduction in the erosion rate (Leong et al., 1985). This highly desirable feature allows the development of a drug delivery system that can provide the release of drugs for a specific length of time through simple changes in the polymer backbone. Hydrolysis of the anhydride bond is base catalysed and the pH of the surrounding environment has a profound effect on the rate of degradation. In a very acidic environment, many polyanhydrides will degrade very slowly (Leong et al., 1985). Also, since polyanhydrides degrade into carboxylic acids, at a low pH these degradation products are in their un-ionised form, with low solubility, and therefore polyanhydrides degrade more rapidly in basic media than in acidic media (Jain et al., 2005). Aliphatic polyanhydrides are highly crystalline, with poor mechanical properties. This is improved through copolymerisation with an aromatic polyanhydride (Jain et al., 2005). Polyanhydrides properties can be easily altered; introduction of an imide group improves their mechanical properties, while the presence of polyethylene glycol (PEG) groups increases hydrophilicity with increased erosion rates and rapid drug release (Kumar et al., 2002). Polyanhydrides have been synthesised from unsaturated fatty acids, which are hydrophobic natural body components, and decrease the degradation rate and drug release from the system. The initial MW of the polymer has been shown to have a varying effect on the rate of degradation. D’Emanuele et al. (1992) showed an early induction period in the degradation of polyanhydrides, prepared by a hot melt method, during which weight loss from the material was slow. The induction period occured until a critical MW was reached. An increase in MW of the

370

Degradation rate of bioresorbable materials O C

O O

(CH2)3

O

C

O O

C

O (CH2)8

n

C

O m

14.9 Poly(1,3-bis(carboxyphenoxypropane-co-sebacic acid).

initial polymer increased the length of the induction period and therefore delayed the erosion process. Other investigations have shown the degradation rate of polyanhydrides to be independent of initial MW. However, these systems were more porous than the melt-cast film, which facilitates the uptake of water, resulting in very rapid hydrolysis and degradation of the polymer to the critical MW for onset of erosion (Dang et al., 1996). The most extensively studied poly(anhydride) is 1,3-bis(carboxyphenoxypropane) (PCPP) (Fig. 14.9), copolymerised with the more hydrophilic sebacic acid (SA). Hydrophobic homopolymer PCPP and the copolymer PCPP-SA (85% PCPP) have shown constant degradation rates over a period of 8 months (which could be extrapolated over a period of three years). Copolymerisation of PCPP with 80% SA decreases this to a few days; an increase in the degradation rate by 800 times (Leong et al., 1985). A feature of the erosion of PCPP-SA, and many other polyanhydrides, is the creation of erosion zones in which the polymer is undergoing erosion (Tamada and Langer, 1992; Göpferich and Tessmar, 2002). This is a layer of degraded and highly porous material, which is separated from the nondegraded polymer by an erosion front. This moves at a constant rate from the outside to the inside of the material. The structure of the erosion zone is dependent on the composition and morphology of the polymer. As PCPPSA erodes, the aliphatic SA is eroded initially and leaves a porous CPP structure which erodes at a slower rate. A highly porous erosion zone may also be detected in semi-crystalline polyanhydrides. Crystalline regions of the polymer are more resistant to erosion than amorphous regions, and can maintain a porous polymer skeleton in the erosion zone over a substantial period of time. As the polymer components do not erode at the same rate, this can have an effect on the rate of drug release. If microphase separation occurs in the polymer, the drug can preferentially partition into one of the phases and release will be dependent on the release profile of that monomer (Shen et al., 2002). The release of hydrophobic p-polyaniline from PCPP-SA, 20 : 80), which displays microphase separation, has been reported to be dependent on the erosion of the CPP component into which it preferentially dissolves (Larobina et al., 2002). The release of the hydrophilic dye acid orange from the same system was much faster, with the drug depleted after only 50% of the polymer had eroded (Mathiowitz and Langer, 1987). For homopolymers and homogenous copolymers, this is not the case; the

Drug release from bioresorbable materials

371

drug is uniformly dispersed throughout the matrix and release is dependent on the rate of degradation. Polyanhydrides have also been derived from fatty acids, including erucic acid (fatty acid) dimer (FAD) (Göpferich and Tessmar, 2002). These materials have a highly hydrophobic character with a slow degradation rate. When copolymerised with SA, materials with desirable degradation profiles and mechanical properties can be achieved. An erosion zone also develops during the degradation of these polymers. Depending on the SA : FAD ratio, the erosion zone can consist of a semi-solid layer, which can slow the diffusion of drugs from the device. High proportions of SA in the copolymer result in a porous erosion zone with a higher rate of release. Release of a hydrophilic dye was found to be reduced as the concentration of FAD was increased, due to the formation of a gel-like barrier. As polyanhydrides are composed of low-solubility monomers, the elimination of the initial degradation products (diacid monomers) via solubilisation may not be a rapid process (Katti et al., 2002). Aliphatic monomers derived from fatty acids will be extensively metabolised in the β-oxidation pathway and eliminated from the body mainly as carbon dioxide. The excretion of hydrophobic aromatic monomers may take longer and occurs in the urine and faeces, with the monomer unchanged or conjugated to an amino acid. Investigations into the toxicity of polyanhydrides have shown that they display excellent biocompatibility with no evidence of inflammatory reaction (Laurencin et al., 1990).

14.2.5 Erodible hydrogels The defining characteristic of all hydrogels is their ability to swell, because their hydrophilic, cross-linked structure ensurs that large amounts of water are retained in their swollen structure without dissolving (Park et al., 1993c). Hydrogels display good biocompatibility due to their high water content which minimises frictional irritation to surrounding tissue and results in a very low interfacial tension. These systems are not suitable for the controlled release of low-MW hydrophilic drugs as they diffuse out of the system rapidly. However, they have been successful in the delivery of very hydrophobic drugs, such as steroids, and large molecules including peptides and proteins, which are trapped within the hydrogel until it degrades (Baker, 1987). Some of the previously discussed degradable polymers expose the drug to an acidic microenvironment or degradation products that can affect the stability of the agent, especially proteins and peptides (Guse et al., 2006). Hydrogel materials have been investigated to overcome this problem. In order to be removed from the body, the degradation of a hydrogel should result in the solubilisation of the network. Solubilisation of a hydrogel can occur in three ways (Baker, 1987; Park et al., 1993d):

372 (i) (ii) (iii)

Degradation rate of bioresorbable materials the polymer backbone may contain the bond susceptible to hydrolysis or enzymatic degradation, and the polymer fragments are soluble the susceptible bond can reside in the cross-links between polymer chains the cleavage may occur between the backbone and attached pendent groups. This third system will result in the solubilisation of the network only if the hydrogel is physically cross-linked and the removal of the pendent groups increases the hydrophilicity of the polymer.

These three systems are described schematically in Fig. 14.10. Type (i) erosion results in the formation of small, soluble fragments which are easily eliminated from the body, and it is therefore the most desirable mechanism of degradation. The synthetic degradable polymers previously discussed degrade in this manner; however, they are hydrophobic and do not form hydrogels on their own. By preparing copolymers, block copolymers and interpenetrating networks with a hydrophilic polymer, degradable hydrogels may be obtained (Park et al., 1993a). The incorporation of ester, ortho ester or anhydride labile moieties into a hydrophilic polymer backbone resulting in a hydrolysable hydrogel has been described (Cadee et al., 1999; Heller et al., 1990; Allcock, 1990). For example, many degradable polyesters, including PLA, PGA and PCL have been modified with hydrophilic poly(ethylene glycol) (PEG) to produce degradable hydrogels (Shim et al., 2006). Jeong et al. (1999) synthesised a degradable triblock copolymer

(i)

(ii)

(iii) 14.10 Schematic representation of hydrogel degradation via (i) cleavage of the polymer backbone, (ii) cleavage of the crosslinks, (iii) cleavage with removal of pendent groups (adapted from Park et al., 1993).

Drug release from bioresorbable materials

373

of PEG-PLGA-PEG. This copolymer shows sol-to-gel transition as the temperature is increased. This thermoreversible copolymer is useful for drug delivery applications because it can be injected as a free-flowing solution at room temperature, and upon injection into the body, the copolymer becomes a gel at body temperature, to form a degradable drug depot. The use of natural polymers to form degradable hydrogels has also been extensively studied. Cross-linked proteins and polysaccharides may also be degraded through cleavage of the backbone chain by hydrolytic and enzymatic mechanisms. Polysaccharides which have been cross-linked to form hydrogels include chitosan, dextran and gelatin (Lee and Yuk, 2007). Polysaccharides which contain carboxyl groups (e.g. chondriton sulphate and hyaluronic acid) may be derivatised with cysteine methyl ester and crosslinked to form hydrogels under mild oxidative conditions. Polymers undergoing type (ii) erosion are typically non-degradable water-soluble polymers such as PEG, polyacrylamide, poly(vinyl alcohol) and N-vinylpyrrolidone, from which networks are formed with unstable crosslinks. The resultant material is highly hydrophilic and therefore suitable only for the delivery of macromolecules or highly hydrophobic agents. These monomers have been cross-linked with small molecules including N,N′-methylenebisacrylamide, which degrades via hydrolysis when the cross-link concentration is low (<1%) (Kamath and Park, 1993), and aromatic azo cross-links which are degraded by bacterial enzymes in the colon (Shantha, 1995). Enzymatically degradable oligopeptides, poly(amino acids) and polysaccharides have also been employed as cross-linking agents (Kurisawa et al., 1998; Yoshida et al., 2003). Type (iii) erosion of a hydrogel does not decrease the MW of the polymer; therefore, unless the original weight is below the limit for clearance from the body, the polymer cannot be used as a degradable bioactive implant (Park et al., 1993a). These systems may be used for the release of low molecular weight drugs, as release is controlled through the rate of cleavage of the susceptible bond. Hydrogels in this category include poly(L-glutamic acid), poly(hydroxyalkylglutamines), and copolymers of N-(2-hydroxypropyl)methacrylamide with the p-nitrophenyl esters of N-methacryloylated oligopeptides; all have been studied for the attachment of low MW drugs.

14.3

Mechanisms of drug release from biodegradable polymers

Drug release from monolithic degrading systems in which the drug is uniformly dispersed throughout the polymer matrix is considered in this section. Drug release is dependent on both the rate of diffusion of the drug and also the rate of degradation of the biodegradable matrix. As discussed, two different mechanisms of polymer degradation can occur, bulk and

374

Degradation rate of bioresorbable materials

surface degradation, which directly affect the mechanism of drug release from the system. In bulk eroding polymers, the predominant method of drug release is the diffusion of the drug prior to or occurring concurrently with degradation. In a surface degrading polymer, drug release is determined by both the rate of diffusion and the rate of degradation.

14.3.1 Degradation-controlled systems In degradation-controlled systems, the rate of degradation is faster than the rate of diffusion. As only one mechanism of drug release needs to be considered, one of the main advantages of degradation-controlled systems is the predictability of drug release. Drug release is dependent on both the shape and the surface-to-volume ratio of the device, and therefore the reduction in the size of the degrading device over time must be considered. Surface-eroding polymers, such as polyanhydrides and poly(ortho esters) have highly labile bonds and therefore degrade quickly, making them ideal candidates for this type of release. Assuming surface-degrading polymers erode at a zero-order rate and that the rate of drug release is proportional to the surface area of the system, the following general equation has been derived for eroding spheres and cylinders (both with radius a), and slabs (with thickness 2a) (Hopfenberg, 1976): Mt /M∞ = 1 − (1 − kot / Coa)

n

[14.1]

where Mt and M∞ are the cumulative amounts of drug released at time t and at infinite time, respectively, Co is the initial drug concentration, n is the shape factor representing spherical (n = 3), cylindrical (n = 2) or slab (n = 1) geometry. The shape factor defined above (in Equation 14.1) by Hopfenberg, has, in subsequent work, been applied to other geometrical forms such as halfspheres and squares (Karasulu et al., 2000). From Equation 14.1 it is clear that only slabs display zero-order release. Cylinders and spheres display a reduced rate of drug release with time. For a slab geometry this becomes (Equation 14.2); Mt /M∞ = kot /(Coa)

[14.2]

Physicochemical and geometrical factors affecting drug release from surface-eroding polymers may be summarised as follows: (i) (ii) (iii)

the rate of drug release is directly proportional to the initial drug concentration (i.e. the drug loading) the duration of drug release is directly proportional to the size of the device the rate of drug released is proportional to the total area of the device, which remains constant only for a thin slab.

Drug release from bioresorbable materials

375

Another more detailed model that may be used to describe surface erosion was developed by Cooney (1972), albeit to originally describe the dissolution of tablets. In a similar fashion to Hopfenberg, the erosion (‘dissolution’) of the polymer is described as a zero-order process occurring at the surface of the device and release rates are proportional to the total area of the device. However, a further step is considered in this process, namely the release of the degradation product through a stationary solvent front at the device surface. The surface erosion of the polymer is therefore described in two steps: detachment of an atom or molecule from the solid surface, followed by diffusion of this agent through a diffusional boundary layer into the bulk fluid. For polymers with low erosion rates, surface detachment tends to control the overall dissolution rate. For a cylinder of initial length Lo and initial diameter Do, erode according to Equation 14.3 in which f designates the fractional dissolution (drug release) at time t (relative to t = 0) and K is a rate constant f =

(Do − 2Kt )2 + 2 (Do − 2Kt ) (Lo − 2Kt ) Do 2 + 2 Do Lo

[14.3]

When Lo /Do approaches zero (slab geometry), a constant relative dissolution rate is obtained.

14.3.2 Diffusion-controlled systems In diffusion-controlled systems, the rate of drug diffusion is greater than the rate of degradation. With this in mind, Heller and Baker (1980) developed a mathematical model for these systems that was based on the Higuchi equation for describing diffusion-controlled release (where the initial drug concentration is assumed to be much greater than the drug solubility in the polymer). The Higuchi equation for non-degradable slabs is: Mt = A( 2 DC sCot )

1/ 2

[14.4]

Since DC s = P( permeability) ,

[14.5]

Mt = A( 2 PCot )

[14.6]

1/ 2

where Mt is the drug released at time t, A is the surface area of the device, P is the permeability of the polymer to the drug, Co is the initial concentration of the drug, Cs is the solubility of the drug in the polymer and D is the diffusion coefficient. The drug release is proportional to t , and the

376

Degradation rate of bioresorbable materials M

Diffusion/erosion controlled

Diffusion controlled

0

t1/2

14.11 Theoretical drug release from a polymer slab by diffusion, and by combined diffusion and erosion.

release rate is reduced continuously over time due to a reduction in the drug concentration. In many bulk-degrading polymer systems the drug permeability will increase over time as the porosity of the material is increased by increasing degradation. This can compensate for the normal reduction in rate of drug release that occurs in non-degrading polymers over time and may even result in constant or increasing drug release rates over time (Faisant et al., 2002). Figure 14.11 shows a representation of the release from a diffusioncontrolled (Higuchi model) and a diffusion- and erosion-controlled release system. If it assumed that the increase in permeability is inversely proportional to the number of intact bonds in the polymer, then, P / Po = N / N − Z

[14.7]

where N is the initial number of bonds and (N − Z) is the number of bonds remaining after Z cleavages of the polymer chains. Po is the original permeability and P is the permeability after Z cleavages of the polymer chains. If it is further assumed that the rate of cleavage follows first-order kinetics, then dZ /dt = k ( N − Z )

[14.8]

where k is the first-order rate constant. Then, by integration, ln [ N /( N − Z ) ] = kt

[14.9]

Equations 14.7 and 14.9 can be combined and substituted into equation 14.6, to give Mt = A( 2 Poe ktCot )

1/ 2

[14.10]

Drug release from bioresorbable materials

14.4

377

Drug delivery applications of biodegradable polymers

As well as delivering the therapeutic agent in a controlled manner, another advantage of biodegradable delivery systems is the elimination of the requirement for surgical removal of the device after the conclusion of therapy (Danckwerts and Fassihi, 1991). An implantable drug delivery device is, by necessity, a temporary device, as the drug may be required only for a finite period of time, or the device will eventually be depleted of the drug. The removal of polymer implants can be more traumatic for the patient than insertion, and therefore biodegradable implants may improve patient acceptance and compliance (Sun et al., 2006). Biodegradable polymers can also be used to form semi-solid injectable implants or they can be used to encapsulate particular systems as micro and nanoparticles that can be injected into the body. In the following subsections, examples of the use of biodegradable polymers for the controlled delivery of therapeutic agents are detailed. The chosen examples have been selected to illustrate the utility of biodegradable polymers as controlled release matrices.

14.1.1 PLA, PGA and PLGA as drug delivery systems Osteomyelitis is a deep bone infection that can occur after knee or hip surgery. Treatment will often require surgery followed by antibiotic therapy, and it is difficult to achieve a sufficient concentration of antibiotic at the site through systemic administration. Fluconazole (FLU) is used in the treatment of fungal osteomyelitis and prolonged systemic treatment can result in an increased risk of adverse effects. Rods of DLPLA and LPLA containing FLU were inserted in the bone marrow of 30 rabbits. The DLPLA polymers showed a sustained near zero-order release for a period of 12 weeks with potential as a treatment in fungal bone infections (Soriano et al., 2006). A problem associated with the insertion of a stent is a possible foreign body reaction that may ultimately result in the reobstruction of the artery; this process is known as restonosis (Jain et al., 2005). Paclitaxel eluting stents have been shown to greatly reduce these events. A drug-eluting stent will be composed of the stent base and a polymer drug carrier. The use of PLGA as the paclitaxel-loaded polymer, loaded into small holes on the side of a stent, was investigated with promising early human-study results (Serruys, 2005). Another early clinical study on a stent coated with PLA loaded with everolimus (an anti-proliferation drug for the treatment of restensosis) has also produced encouraging results (Grube et al., 2004).

378

Degradation rate of bioresorbable materials

Leuteinising hormone-releasing hormone (LHRH) agonists, e.g. leuprolide acetate (LA), are used in the treatment of prostate cancer, as they can effectively suppress serum testosterone levels. Initially, inconvenient daily injections were required for treatment; however, depot formulations of LA have been developed to provide a sustained release. LA-2550 (22.5 mg LA) is a degradable system composed of D,L-lactide-co-glycolide polymer dissolved in N-methyl-2-pyrrolidone. This is mixed with LA and subcutaneously injected. Upon contact with body fluids, it coagulates to form a solid depot of LA, which controls release for a period of 3 months and maintains therapeutic suppression of testosterone (Chu et al., 2002). These systems have been approved by the FDA and are marketed as ‘Eligard’ (Sartor, 2003). Other PLGA implants for parenteral use, which are commercially available to deliver LHRH agonists, are ZoladexTM (goserelin) and DepotProfactTM (buserelin) (Kissell et al., 2002). Depending on the indication, controlled release occurs over a time-span of 1–3 months. Another in-situ forming implant based on PLA/PGA has been investigated, to provide sustained release of tumor necrosis factor (TNF) receptors in the treatment of chronic inflammatory diseases. In this, PLGA (75 : 25) was dissolved in the solvent glycofural and the protein receptor (sp55-R) dispersed in this solution. The mixture was injected subcutaneously in mice bearing TNF-producing tumours. The biodegradable system was able to provide long-lasting protection against TNF (Eliaz et al., 2000). Microparticles of PLGA loaded with 5-fluorouracil (5-FU), an antineoplastic drug, have been intracranially injected in patients with the brain cancer gliobestoma (Benoit et al., 2000). Local delivery of anticancer drugs is highly beneficial due to the large range of side effects that can occur with systemic administration and also due to the short half-life of several of the drugs. 5-FU has a half-life of only 15 minutes and therefore controlled release of this drug is very desirable.

14.4.2 Poly(ε-caprolactone) as a drug delivery system Microparticles of PCL loaded with 5-FU have also been investigated. Drug release was examined in vivo, after subcutaneous injection into rats. 5-FU was detected in the plasma after 18 days. Drug release was controlled by diffusion as the microparticles had not degraded at this time. No harmful inflammatory response to the particles was found and they had begun to degrade after a three-month period (Sastre et al., 2004). In a different application, paclitaxel, an inhibitor of microtubuli formation, was mixed into melted poly(D,L-lactide)-block-poly(ethylene glycol)block-poly(D,L-lactide) copolymers to obtain paste formulations, which could be inserted at the site of surgery. These pastes are administered at tumor resection sites where a controlled release of paclitaxel may destroy

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remaining tumor cells and prevent local recurrence of disease. The in vitro release of paclitaxel from these paste formulations with 20% drug loading was controlled by polyethylene oxide content over a period of up to 2 months by a combined diffusion and polymer erosion mechanism. The pastes were shown to be efficient at inhibiting tumour growth in mice (Zhang et al., 1996). Injectable calcium phosphate has been used as a bone substitute for osteoconduction purposes. To avoid infection in the implantation site, PCL microparticles encapsulating vancomycin were incorporated into the calcium phosphate. Release was detected for 7 days in vitro. Release was either diffusion-controlled or degradation-controlled, depending on the method of preparation of the particles. Copolymers of PCL and the polyester δ-valerolactone (VL) have been prepared to increase the degradation rate of PCL. Release of estramustine from a small cylinder of the copolymer, implanted in rats, was dependent on the degradation of the polymer and, after an initial burst release lasting 1 week, drug release was constant up to week 20 of the study (Kazumichi et al., 1991). Microspheres of PCL-DLPLA were investigated for the controlled release of β-estradiol, a hormone used in the treatment of menopausal symptoms, including the increased risk of developing osteoporosis. An almost constant rate of release from the microspheres was observed over a period of 70 days in vivo. The eye is suited for the implantation of polymer system for local delivery of therapeutic agent, due to easy accessibility and also to the blood–retinal barrier, which isolates the eye from the main circulation (Bourges et al., 2006). Many degradable systems have been developed for the controlled release of ocular drugs. Local delivery of β-blockers to the eye is very beneficial to avoid the cardiovascular side effects. PCL nanocapsules have been loaded with carteolol for ophthalmic use, and have shown a reduction in systemic side effects in comparison to an aqueous solution of the drug (Marchal-Heussler et al., 1993). Borhani et al. (1993) describe a delivery system based on combinations of different MW PCL. Increasing the proportion of high MW PCL to low MW PCL decreased the rate of degradation of the system and release of 5-FU was slowed down. These devices showed promising in vivo results for the treatment of proliferative retinopathy. The treatment of chronic uveitis with cyclosporin A is impeded by poor penetration of the drug into the eye and serious side effects with systemic administration. Implants composed of PLGA-co-PCL were investigated in vivo and were found to inhibit inflammation in experimental uveitis by controlling cyclosporine A release, with no systemic or local toxicity (Dong et al., 2006).

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PCL has been employed as a controlled release matrix for the improved treatment of periodontal disease (Medlicott et al., 1996, 1999). In these studies, films of PCL containing the chosen antimicrobial agent (chlorhexidine) were prepared by solvent evaporation, the rate of drug release from the implant in vitro being dependent on various parameters, e.g. the salt form of drug. The implant was then adhered to the tooth adjacent to the gingival and the concentration of drug within the salivary pellicle determined as a function of time. The authors illustrated that a constant (zeroorder) release of drug occurred and was super-minimum inhibitory for the duration of the study. In a related application, the authors have examined the use of PCL as a medical device biomaterial. Interestingly, it was shown that the development of encrustation (in vitro) on the surface of PCL was lower than on non-biodegradable polymers, the result being attributed to the degradation of the polymer (Jones et al., 2002). In a follow-up study, the authors illustrated the bacterial anti-adherence properties of PCL that had been impregnated with the hydrophilic antimicrobial complex poly(vinylpyrrolidoneiodine) (Jones et al., 2005). These results were attributed to the ability of PCL to act as a controlled-release matrix for the antimicrobial species.

14.4.3 Poly(ortho esters) as drug delivery systems Solid POE I has been investigated for the delivery of homosulfanilimide to treat infected burns, for the delivery of the opioid antagonist naltroxone, and also for the delivery of the contraceptive levonorgestral. However, due to the uncontrolled autocatalysis of these systems, they were not investigated further (Heller et al., 2002). Solid POE II has been tested in animal models for the delivery of several drugs including 5-fluorouricil (5-FU) for the treatment of cancer, naltrexone, and norethisterone as a contraceptive (Einmahl et al., 2001). Intensive research into a responsive delivery system that can release insulin in response to glucose concentration for the treatment of diabetes has been carried out. One method has been loading a degradable, highly acid-sensitive POE II with insulin and surrounding this with a hydrogel containing immobilised glucose oxidase (Heller et al., 1990). To increase pH sensitivity, N-methyldiethanolamine was incorporated to produce a tertiary amine-containing POE. Glucose is oxidised to gluconic acid within the hydrogel, which results in accelerated polymer erosion and increased insulin release. Semi-solid POE III, which is an injectable system, has been investigated for the delivery of tetracycline (for treatment of periodontal disease), 5-FU and metoclopramide. However, only POE IV has been considered for possible commercial applications (Heller et al., 2002).

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Two formulations of semi-solid POE IV suitable for injection are under development – hydrophobic material developed from 1, 10-decandiol and a more hydrophilic POE based on triethylene glycol (TEG). As the more hydrophilic POE will degrade faster, it is useful for applications where the drug is needed for a short period of time. A formulation of TEG loaded with bupivicaine analgesic to be applied at the site of surgery is being investigated. This would reduce the need for high doses of systemic painrelief which is accompanied with side effects such as drowsiness and constipation. The anaesthetic is released in a controlled manner over a period of two days. The more hydrophobic decandiol has been investigated for treatment of periodontal disease. The concentration of an antimicrobial in the periodontal pocket is significantly increased if a sustained delivery device is placed at the site in comparison to antimicrobial rinses or systemic antibiotic therapy. Tetracycline was loaded into the POE semi-solid formulation and near zero-order release was observed over a period of 2 weeks. Early human trials have shown promising results, with high tolerance of the implant, and appropriate tetracycline concentration levels were maintained for 11 days if the implant was retained.

14.4.4 Polyanhydrides as drug delivery systems In some clinical situations, removal of an implant is not possible; in sitespecific delivery to the brain, to treat Parkinson’s disease for example, the carrier polymer must be biodegradable in nature (Senthilkumar et al., 2007). The first polyanhydride-based drug delivery system to enter clinical use was for the delivery of the chemotherapeutic agent, bis-chloroethylnitrosourea (BCNU). The implant, marketed as Gliadel®, is made from PCCP:SA (20 : 80) and is implanted intracranially to patients suffering from glioblastoma multiformae, a usually lethal form of brain cancer. SeptacinTM implant is a polyanhydride consisting of poly(euric acid dimer) (EAD) and SA in a 1 : 1 ratio, loaded with 20% w/w gentamicin for the treatment of osteomyelitis. The polymer is moulded into beads to be placed at the site of surgery. Results from a human in vivo study showed high local concentration of gentamicin at the required site and very low systemic exposure (Li et al., 2002). Controlled local delivery of anaesthesia post-surgery has also been investigated. The desired period of pain-relief could be a period of 1 day to 1 week, and polyanhydrides are ideally suited to deliver a sustained dose over a period of one week. Devices composed of PCPP-SA, 1 : 4) containing bupivicaine HCl, dibucaine HCl and dibucaine free base were prepared. Sustained release for a period of 3–14 days was observed, depending on the drug incorporated. When implanted near the sciatic nerve in rats, reversible neural blockade was observed for 4 days (Masters et al., 1993). PCPP-SA

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loaded with heparin and cortisone was implanted in the eye to inhibit the neovascularisation of a corneal tumour. The drugs were released simultaneously and release lasted three weeks, the time in which complete degradation had occurred. The neovascularisation process was inhibited during this time (Langer et al., 1985).

14.4.5 Hydrogels with degradable backbone Biodegradable hydrogels based on water-soluble polyesters containing double bonds were cross-linked by copolymerisation with vinylpyrrolidone. Cleavage occurred in the ester linkages in the backbone of the polymer that had been modified by vicinal carboxylic acid groups, with the generation of soluble poly(ethylene glycol) and poly(N-vinylpyrrolidone) (Heller et al., 1983). To increase the hydrolysis of the polyester linkage, electronwithdrawing diacids were included in the polymer backbone. By altering the type of diacid included in the polymer and also the cross-linking density, the rate of degradation could be controlled. The protein bovine serum albumin (BSA) was incorporated in the hydrogel during polymerisation. The release of BSA was determined by the rate of hydrolysis. Sawhney et al. (1993) synthesised a triblock PLA-PEG-PLA block copolymer with photopolymerisable acrylate end groups (Fig. 14.12). Three structural domains in the macromer each play a role in the resulting hydrogel function: the PEG backbone gives the hydrogel its hydrophilicity, the PLA linkages provide biodegradability, and the polymerisable end groups allow for the formation of a cross-linked network. These gels degrade upon hydrolysis into PEG, lactic acid and oligo(acrylic acid). The ratio of PEG to PLA segments can be used to manipulate a number of physical properties, from the permeability of the gel to its degradation rate. Continuous release of protein (BSA) from the gel was achieved for a period of two months. The release was degradation controlled initially; however, as degradation progressed, a reduction in cross-linking density increased the porosity of the hydrogel and the rate of release was increased by diffusion out of the more porous hydrogel. Kamath and Park (1995) produced chemically cross-linked dextran hydrogels; dextran was functionalised to introduce double bonds, which were

O H2C

C H

C

CH3 O O

C H

C

H2 H2 C C O

O m

n

O

CH3

C

C H

O O

C m

C H

14.12 Biodegradable poly(ethylene glycol)-b-poly(lactide) macromer.

CH2

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cross-linked upon exposure to γ-radiation. Invertase was incorporated into the material prior to polymerisation. Enzymatic degradation of the polymer backbone occurred in the presence of dextranase. The rate of degradation from the network could be controlled by the MW of the dextran monomer; larger molecules produced hydrogels with a higher swelling capacity and therefore they degraded faster. Release of the protein invertase from the network was dependent on the presence of dextranase. A degradable hydrogel based on a semi-IPN structure composed of PCL and cross-linked PEG was developed by Cho et al. (1999). The degradation of the linear PCL chains was enhanced with increasing concentration of PEG in the network. Constant release of clonazepam loaded in the network was achieved for 45 days in vivo. Injectable in situ-forming hydrogels have been prepared from a mixture of an alginate solution and calcium carbonate in the presence of PLGA microspheres loaded with a hormonal growth factor (Insulin-like Growth Factor I, IGF-I). IGF-I increased the proliferation rate of MG-63 cells, which could be useful for conformal filling of bone defects (Luginbuehl et al., 2005). Hydrogels cross-linked by formation of stereocomplexes between enantiomeric polymers have been investigated. de Jong et al. (2001) grafted dextran with L-lactic acid or D-lactide acid which, when combined in aqueous solution, form a physically cross-linked hydrogel. Degradation occurs via the ester linkage in the lactate graft or via the cleavage of the lactide graft from the dextran backbone. These polymers may be premixed with protein prior to the gelation process. Release from the hydrogel was dependent on the size of the protein and also the dispersity of the lactate grafts.

14.4.6 Hydrogels with degradable cross-links PEG-based hydrogels have been synthesised with degradable thioetherester links by mixing unsaturated PEG-acrylates with nucleophilic PEGthiols (Elbert et al., 2001). BSA was incorporated in the hydrogel prior to polymerisation, the cross-linking reaction being self-selective and therefore not involving the protein molecules. As the linkage is hydrolysed, the cross-linking density is reduced and release of albumin occurred. Release rates were modified by changing the degree of functionality of the PEG monomer. Zero-order release was obtained over a four-day period from the tetra-functional PEG-hydrogel. Degradable hydrogels were prepared by conjugate addition of PEGmultiacrylate to dithiothreitol in the presence of human growth hormone (hGH). It was necessary to precipitate hGH with linear PEG or Zn2+ in order to protect the hormone during the polymerisation process. Precipitation of the hormone with Zn2+ also increased the stability in the hydrogel

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and delayed release by slowing the dissolution of the agent. Release was controlled by changing the MW and degree of functionality of the PEG acrylate. Zero-order release kinetics were achieved in vitro (van de Wetering et al., 2005). Degradable hyaluronic acid (HA) hydrogels were synthesised by photopolymerisation of vinyl group modified HA in combination with a diacryloyl PEG–poly(propylene glycol)–PEG tri-block copolymer (Pluronic) (Kim and Park, 2002). Pluronic copolymers are thermally responsive due to the formation of micelles at increased temperatures, and the hydrogel is therefore thermally responsive. The water uptake capacity continuously decreases with increasing temperature, indicating that the association of the Pluronic component occurs within the network and results in a reduction in water uptake capacity of the HA/Pluronic hydrogels. These hydrogels degrade due to the hydrolysis of an ester linkage present in the structural unit of the di-acryloyl Pluronic component. The erosion of the hydrogel occurrs much faster at higher temperatures; this is proposed to be due to the exposure of the ester linkage at higher temperatures, due to the micellisation of Pluronic. Release of recombinant human growth hormone (rhGh) was mainly dependent on the erosion of the hydrogel, and proceeded at a faster rate at 37°C than at 13°C. Aromatic azo bonds are cleaved in the colon by bacterial azoreductase. Therefore, cross-links composed of aromatic azo groups should degrade in this area of the gut. Hydrogels composed of hydroxyl ethyl methacrylate (HEMA) copolymerised with methacryloyloxy azobenzene were prepared (Shantha, 1995). The hydrogel was pH sensitive and did not swell in simulated gastric fluid (acidic), and drug release was minimal. Drug was released in simulated intestinal fluid in the absence and in the presence of azoreductase producing bacteria. The degree of swelling was higher and drug release increased compared to the acidic environment; however, without the bacteria present in the release media, drug release (5-FU) occurred from the surface only. In the presence of the enzyme, the cross-links were cleaved with a much greater rate of release. Zero-order release of 5-FU was achieved over a period of 4 hours from the degrading hydrogel. This hydrogel could therefore find an application in the oral delivery of sensitive drugs to the colon.

14.4.7 Hydrogels with degradable pendent groups The in vivo release of norethindrone from a biodegradable steroid–polymer conjugate was examined (Zupon et al., 1983). Norethindrone was first derivatised and reacted with poly-N-(3-hydroxypropyl)-L-glutamine. This was administered to rats by subcutaneous injection. The in vivo release of steroid remained fairly constant, though it showed a decrease over the

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9-month investigation. This biodegradable drug–polymer conjugate has potential applications for the controlled delivery of norethindrone, which could have applications in long-term contraception. Cross-linked hydrogels were prepared by the copolymerisation of methoxydiethoxyethyl methacrylate and the p-nitrobenzoate ester of 2hydroxyethyl methacrylate (Pitt and Shah, 1995). This hydrogel can degrade through the hydrolysis of the p-nitrobenzoate ester group, which results in a loss of hydrophobic p-nitrobenzoate from the network. This increases the hydrophilicity of the hydrogel with an increased water uptake followed by an increase in the rate of hydrolysis. Pendent group cleavage occurs via firstorder kinetics, the increased hydrophilicity of the network compensating for the fall in hydrolysis rate, and zero-order release of p-nitrobenzoate was achieved. However, the system is not biodegradable as permanent crosslinks between the polymers remain. Side chain substituted copolymers of N-isopropyl acrylamide and acryloxy succinimide were prepared (Shah et al., 1997). These polymers exhibited a lower critical solution temperature (LCST), dependent on the copolymer ratio. As the succinimide side chain was hydrolysed, the hydrophilicity of the polymer increased with a resultant increase in the LCST and increased rate of hydrolysis. Above the LCST, the polymers were insoluble. However, as hydrolysis proceeded, the LCST was raised and the polymer became soluble. This type of system could be used to develop an injectable delivery device at room temperature that would form an insoluble drug depot in the body (LCST greater than room temperature but below body temperature).

14.5

Conclusions

This chapter has presented an overview of the type of biodegradable polymers that are used pharmaceutically, their degradation, and their application as drug delivery systems. The chemical diversity of biodegradable polymers offers many opportunities for the design of drug delivery systems whose drug release kinetics may be engineered to match the demands of the clinical condition under treatment. It is anticipated therefore that the importance of biodegradable polymers as drug delivery systems will continue to evolve in the near future.

14.6

References

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Index

Absolok 118 absorbable materials, fabrication of 13, 16 absorbers, UV 230 accelerated degradation tests 139–40 acetylation, degree of 71 acid acrylic 299 alginic 81 autocatalysis 294, 368 burst 323 catalysing groups 196 catalysis 192 end groups 183, 190 pH 274 acidophil 23 ACP (amorphous calcium phosfate) 101, 102 activation energy 183 acute inflammatory response 35, 36–7, 270, 279, 289 cells and 289 exudates 273–4 phase of 277 process 268 additives 229–31, 293–4 Adriamycin 118 AFM (atomic force microscopy) 155, 156–7 agranulocytes 23, 24–5 albumins 27

alcohol 151 aldehyde 151 alginates 11, 75–82, 238, 289 applications 79–82 brown algae species 77 matrix 82 physicochemical properties 78 primary functions of 78 structure 77 wound dressing products 79 alginic acid 81 algorithms, ‘skyscraper’ 158 aliphatic polyanhydrides 369 aliphatic polyesters 48–54, 118–19, 120, 121, 320 aliphatic polymer chains 150–1 degradation 54–60 monomers 371 alkali media 338–43 alkylamines 365 allograft control group 170 allylic groups 150–1 alternative pathway 32–3 aluminium foils 227 American Cancer Society 17 American Society for Bone and Mineral Research 173 amide 148 amine 151 amorphous calcium phosfate (ACP) 101, 102

393

394

Index

amorphous phase polymer degradation 196 semi-crystalline polymers and 194 amorphous regions 340 degradation and 237 radicals and 223 water hydration and 295 anhydride 148 bond, hydrolysis of 369 Animal Housing Unit (University of Singapore) 344 animal research 3Rs of 161, 162–3 ‘10 golden rules’ 177 reduction in 161, 162–3 see also in vivo test methods anion 289 anionic surfaces 296 antibody mediated response 281 anticoagulants 29 antigen presenting cells (APC) 278, 280–1, 291 antigenicity 10 antioxidants 229, 290 AO Research Institute 177 AOVET (veterinary specialty group) 177 apoptosis 296 aromatic cycles 150–1 aromatic polyanhydride 369 Arrhenius equation 201–2 Arrhenius relationship 215–16 ‘Aspects of Degradation and Stabilization of Polymers’ (Murakami) 217 atmosphere modification, shelf-life and 228 atomic force microscopy (AFM) 155, 156–7 autocatalysis 8, 237, 360 degradation model (Pitt) 191 host response and 293, 294 hydrolysis and 247 implant design and 323, 330–1 modelling 183–6, 190–2

autoclaving, steam 16 autograft control group 170 B lymphocytes 25, 34, 280–1 basic fibroblast growth factor (bFGF) 167, 171 basophils 24 behavior prediction, in vivo 107–8 β-rays 224–5 bFGF (basic fibroblast growth factor) 167, 171 bioabsorption, defined 146 bioactive agents 293 biodegradation 320–3 cellular factors 272–3, 276–82 defined 44, 118, 145, 321–2 host factors and 271–2 implantation site and 283–5 in vivo mechanisms 286–90 material factors 290–8 non-cellular factors 271, 273–6 bioerosion, defined 146 Biofix 118 bioglass 234, 246 biological environment 19–38 biomaterial interactions 33 blood 22–8 chronic inflammation 37 fibrin formation 31–2 future trends 37–8 hostile 19–22 injury, host response 34–6 plasma protein cascades 28–31 triple response, acute inflammation 36–7 biological tests, in vitro 145–59 degradation methods 146–52 resorbability assessment 152–4 surface analysis 154–9 biologically derived tissues 298 biomaterials ancient civilisations and 43 design 298–302 interactions 33, 173 naturally derived 288

Index studied polymeric 320 synthetic 298, 320 biomedical applications alginates 79–82 chitin/chitosan 73–5, 76 biomimetic materials 320 biopolymers 322 bioprosthetic heart valves 299–300 Bioquant Osteo II software 173 bioresorption 320–3 defined 67–8, 118, 146 biostable materials, defined 321 bis-chloroethylnitrosourea (BCNU) 381 blind-end pores 246–7 blood 22–8 cells and plasma 22 clots 31 extracellular pH and 22 formed elements 23 ionic concentrations 27, 147 protein content 27 blood brain barrier (BBB) 285 body temperature 22 bonds, hydrolysis and 291–2 bone architecture 164, 167 /biomaterial interface 173 tissue differentiation 171 tissue measurement 172 Boussinesq-Papkovitch equation 250 bovine serum albumin (BSA) 382, 383 bradykinins 34–5 British Standard for sterilization 220 bulk degradation 8, 49–53, 184–6, 330 with autocatalysis 184–6 erosion 9, 292–3, 323, 340, 363 stages of with PLA/PGA 361 vs suface erosion 199–200 burst pressure 250–1 by-products degradation 135, 322–3 hydrolysis 330 non-degradable 10

395

cage-effect theory 223, 225 calcification 285, 300 calcium alginate 79–80, 82 calcium carbonate 105 calcium metaphosfate (CMP) 105 calcium phosfates (CaPs) 95–9, 99–102, 107–8, 379 amorphous (ACP) 101, 102 ceramics 152 compounds 96 doped 98, 104–5 implants 171 in vivo use 98 calcium pyrophosfate (CPP) 105–6 calcium silicate 106 calcium sulfate dihydrate (CSD) 97 calcium sulfate hemihydrate (CSH) 106 calcium sulfates 106–7 capillary zone electrophoresis (CZE) 120, 135–7, 138 Capronor 118 carbon 150–1 carbon dioxide 241 carbonate 148 carboxyl end groups 330–1 casting of polymer films 210 solvent 216, 240 cathepsins 289 cationic surfaces 296 cellophane 84 cells acute inflammatory 289 adhesion 242 blood 52 dentritic 25 immobilized 78, 82 inflammatory multinucleated giant 152 mast 24 plasma 280–2 red blood (RBC) 23 Schwann 301 type/culture 153–4 vascular endothelial 270, 280

396

Index

see also fibroblasts; foreign body giant cells (FBGCs); giant cells cellular automata model 200, 203 cellular factors 272–3, 276–82 non- 271, 273–6 cellulolysis 87 cellulose 83–7 degradation 87 hemodialysis application 85–7 microcrystalline 84 oxidized regenerated 10 structure 69 synthetically modified (SMC) 86 cellulose acetate 85 membranes 84 cellulose triacetate (Hemophan) 85 Celsus 35 central nervous system (CNS) 285 ceramics 95–109 biological degradation of 151–2 bone substitutes 97 degradation paths 146 future trends 108 in vivo transformation 102–4 in vivo use 98 kinetics 101 modelling resorption 107–8 resorbable 5–6 solubility 99–101 chain scission 209, 222–3, 251, 293 modelling and 183, 189 random 195–6, 198–9, 212 random hydraulic 330 chemical factors 273–6 linkages, polymers 291–2 preservation, shelf-life and 228–9 chemicrystallization 195 chemistry, surface 295–7 chemokines 275 chitin 68–75 α-chitin 71 applications 70, 73–5 β-chitin 71 γ-chitin 71 physicochemical properties 71–2 structure 69

chitosan 10, 68–75, 237, 243, 289 applications 70, 73–5, 76 hydrogels and 373 physicochemical properties 71–2 structure 69 toxicity 72–3 chitosan-alginate polyelectrolyte complex (PEC) membranes 74 chronic inflammation 37, 270 citric acid, solid 229 classical pathway 32 clot retraction 31 Cluster Designation (CD) molecule 280 coagulation 28–31, 274 cascade 28 initiation of 29–30 pro/anti- 29 collagen 10, 238, 270, 288–9, 298 chromium tanning of 300 /PGA mixture 240 scaffolds 249 colony-stimulating factors (CSF) 275 common pathway 31 complement cascade 28, 32 composite 299 composite mPCL/TCP scaffolds 334–8 gross morphology 335 mass loss profile 336 compression mechanical degradation process 217 moulding 211, 216 Compton scattering 222 Computer Aided Design (CAD) 187 computer simulation techniques 200, 202–3 computer tomography (CT) 60, 176 micro- 176, 246, 254, 334 contact microscopy 155–8 contact profilometry 155–6 control allograft group 170 autograft group 170 materials 163–4 copolymers 118–19, 132, 135, 252, 299 coral 234

Index cortical width 174 costs 170 critical-sized defect 164 cross-linking 222–3, 299–302 agents, chitosan and 75 chemical 10 hydrogel 383–4 molecular weight and 209 proteins 373 crystal phagocytosis 152 crystalline polymers 213 crystalline regions 340, 361, 370 degradation and 237 mechanical properties and 223 water hydration and 295 crystallinity 184, 194–5, 218, 294–5 changes 132, 364 polymer morphology and 56–7, 132, 361–2 scaffold material 325, 333–4, 336–7, 339–40, 342–3, 347–9 crystallization, secondary 295 crystals, fragmentation of phagocytosed 152 CSD (calcium sulfate dihydrate) 97 CSH (calcium sulfate hemihydrate) 106 cuprammonium process 85 Cuprophan membrane 85 cytokines 34–5, 275, 277, 278, 281, 296 cytotoxicity 300–1 Darcy’s law 248 DCP (dicalcium phosfate) 95, 99, 101, 102–4, 107–8 DCPD (dicalcium phosfate dihydrate) 95, 99–101, 102–4, 107–8 Decapeptyl 118 decomposition temperature (Td) 213 definitions, degradation terms 3–4, 44, 118, 145 degradation accelerated tests 139–40 autocatalysed 183 biological methods 146–52

397

by-products 118, 135, 322–3 -controlled systems 374 defined 44, 118 enzymatic 120, 123–5, 287–9 four stage 53–4, 184–6 hydrolitic 147–50, 185, 323–5, 325–43 late-stage 195–9 mass loss 184–5 mechanisms 4–5, 9 media 55–6 model of heterogeneous 51 oxidative 150–1, 289–90 pH-induced autocatalytic 8 polymer 43 processing-related 212–20 products 298 study protocol 120 terminology 3–4, 145 variables 235, 257 see also bulk degradation; erosion dentritic cells 25 depolymerisation, in thermal degradation 213 Depot-Profact™ 378 Design of Animal Experiments – Reducing the Use of Animals in Research Through Better Experimental Design (Festing) 163 designer alginates 78 destructive methods, degradation monitoring 248 devices see medical devices Dexon 44, 59, 60, 118, 218 dextrans 10 hydrogels and 373 diabetes mellitus 73 diacids 47–8 dialysis, cellulose and 85 diblock polymers 320 dicalcium phosfate (DCP) 95, 99, 101, 102–4, 107–8 dicalcium phosfate dihydrate (DCPD) 95, 99–101, 102–4, 107–8 dietary chitosan 72 ‘dietary fiber’ 87

398

Index

differential scanning calorimetry (DSC) 120, 133–4, 214, 218 diffusion of water 184, 186–9 diffusion-controlled systems 375–6 diffusion-reaction phenomena 51, 330, 343 diols 47–8 dissolution kinetics 97 dissolution-precipitation reaction 97 DLPLA (poly(d,l-lactide)) 360–3, 377, 379 DLPLGA (poly(DL-lactide-co glycolide)) 52 DNA 82, 86–7 dressings, reasons for applying 73 drug release 357–85 applications 377–85 controlled 82 delivery systems 67–8, 70–1, 74–5, 83–4, 291, 347 devices 7, 118, 344 mechanisms of 373–6 pharmaceutical polymers 359–73 rate of 295 strategy 287 drugs, encapsulation of 82 dynamic characteristics 176 dynamic mechanical analysis (DMA) 214 dynamic static indices 174–5 e-beam irradiation 224–5 egg-box model 80–1 elastases 289 electrolytes, water and 20–2 electron microscopy environmental scanning (ESEM) 120, 123–5, 126, 127 scanning (SEM) 154–5, 158, 249, 333, 339, 345, 346 spray injection mass (ESI-MS) 120, 135, 137 transmitting (TEM) 154–5 electron spin resonance (ESR) testing 222–3 electron-beam irradiation 224–5

ELISA (enzyme linked immunosorbent assay) 159 encapsulation, drug 82 enclosed pores 246–7 endocytosis 6 endophagosome envelope membrane 152 endothelial cells 270, 280 environmental factors 227 in vivo 8–10 see also biological environment environmental scanning electron microscope (ESEM) 120, 123–5, 126, 127 enzyme linked immunosorbent assay (ELISA) 159 enzymes 6, 9, 11 biological tests and 150–1 cellulose and 87 collagen and 289 drug release and 358, 383 host factors and 277–8 in vivo degradation and 287–9 processing and 216, 238 sterilization and 223 eosinophils 23–4 episcopic illumination 158 erodible hydrogels 371–3 erosion 8, 358 bulk degradation stage 361 defined 321 ‘number’ 199 see also degradation; surface erosion erucic acid (fatty acid) dimer (EAD) 371, 381 erythrocytes 23 Escherichia coli 283–4 ester 148 bonds 189–90, 198, 293, 331 cleavage, hydrolytic 323 esterases 287 ethanol 256 ether 149 ethics, animal research and 162–3 ethyl cellulose 84 ethylene oxide (EtO) 60, 220–1

Index gas 16 molecule 220 sterilization and 228 ethylene/vinyl alcohol (EVOH) 216, 228 European Academy for the Study of Scientific and Technological Advance 177 evaluation methods 324–5 Excebrane 86 extracellular fluid component (ECF) interfacial reactions 20–1 ionic cencentrations 20, 147 extracellular matrix (ECM) 238, 242, 270, 276, 288, 298 scaffolds 299 extracellular pH, blood and 22 extracorporeal polymer use 343 extrinsic pathway 30 extrusion 210, 216, 217 exudates 273–4 fabric manufacture 242–3 fabrication, absorbable materials 13, 16 Faculty of Medicine (University of Singapore) 344 feedback loops 30 Festing, M. 163 fiber dietary 87 manufacture 242–3 spinning 212 fibrin formation 31–2 fibroblasts 270, 279–80, 280 fibroplasia 270 fibrosis 282 fibrous capsule formation 282 fibrous tissue formation 31 film blowing, of polymers 210 Finite Element Analysis (FEA) 187 Flory equation 193 Flory most probable distribution, oligomers and 196–7 Flory principle 293 fluconazole (FLU) 377 fluorescent microscopy 158–9

399

foam-like structures, manufacture 239–41 Food and Drug Administration (FDA) 12, 72, 84, 324, 378 foreign body giant cells (FBGCs) 6, 26, 27, 37, 281 formation 296–7 multinucleated 279 formed elements, plasma and 22 fractal algorithm 158 fractures 286 fixation 68 freezing 240 fused deposition 243 Fused Deposition Modeling (FDM) 324, 334 fabrication process 326, 328 gamma radiation 16, 209, 221–3, 224, 383 gas foaming method 241 plasma sterilising system 226 Gel Permeation Chromatography (GPC) 57, 331 gelatin 10, 238, 289, 299, 301 hydrogels and 373 gelling, alginates and 78, 80–1 generally-regarded-as-safe (GRAS) 84 geometrical factors, drug release and 374 geometry, specimen 253–4 giant cells 270 foreign body (FBGCs) 6, 26, 27, 37, 281 formation 285, 296–7 multinucleated 152, 171–2, 279 Giemsa stain 349 glass packaging 227 glass transition temperature 184, 202, 213, 368 Gliadel® 381 glycolic acid (glycolide – GA) 47, 57 graft donor sites 80 granulation tissue 270, 278, 280 granulocytes 23–4

400

Index

GRAS (generally-regarded-as-safe) 84 gravimetry 120, 324 growth factors (GF) 163, 275 basic fibroblast (bFGF) 167, 171 Insulin-like (IGF-I) 383 platelet derived (PDGF) 276, 279–80 vascular endothelial (VEGF) 280 H-NMR (Nuclear Magnetic Resonance Spectroscopy, proton based) 57, 120 haemorrhage exudates 273, 274 haemostasis 28–31, 268 haversian systems 167, 169 heat distortion temperature (HDT) 213–14 hemodialysis 84, 85–7 Hemophan 85 Higuchi equation 375 histamines 34–5 histology 170–1 sections 349–50 histomorphometry 170–2, 172–5 host response 267–302 biodegradation factors 271–82 biomaterial design 298–302 cascade 268–70 implantation site 283–5 in vivo biodegradation mechanisms 286–90 material factors 290 polymer adverse outcomes 285–6 species, influence of 285 to injury 34–6 hostile environment 19–22 human digestive system 72 human growth hormone (hGH) 383 hyaluronans 289 hyaluronic acid (HA) 238, 384 /Pluronic hydrogels 384 hydrogels 238 with degradable backbone 382 with degradable crosslinks 383–4, 385 with degradable pendent groups 384–5

erodible 371–3 systems of solubilisation 371–2 hydrogen bonding 288 hydrogen peroxide 290 hydrolysable polymers 147–8 hydrolysis 4 of anhydride bond 369 biodegradation and 48–9 biopolymers and 322 by-products 330 crystalline regions and 295 modelling and 183–4, 186–7, 198–9 PET 195 polycaprolactone (PCL) 198 polymer kinetics 189–92 rate of 188 temperature and 201 hydrolytic degradation 147–50 accelerated 338–43 aliphatic polyesters 48–9 host response and 287–9, 291–2 of medical polycaprolactone (mPCL) 325–43 of polycaprolactone (PCL) 323–5, 325–43 hydroperoxide decomposers, UV and 230 hydrophilic oligomers 247 hydrophilic surfaces 296 hydrophilicity 4, 294, 368 hydrophobic materials 254 hydrophobicity 4 hydroxyapatites (HA) 100, 105–7 fillers 216 nano-particulate 240 /PCL composite 254 precipitated (PHA) 95, 99–100, 102–3 scaffolds 234, 237, 249 sintered (SHA) 95, 97, 99–100, 108 hydroxyl ethyl methacrylate (HEMA) 384 hypoglycaemia 274 hypoxia 274 image analysis 158, 173, 244, 351–2 imaging techniques 175–6, 246

Index imide 148 immune system 146, 280–2 immunoglobulins 27, 297 immunological reactions, alginate and 78 implantation repeated 285 sites 164–8, 283–5 implants 343 alginate and 78 biomaterials and 43 calcium phosfate 171 ceramic 6 degradation stages 5 design/application 7–8 hydrolytic degradation 51 metal 68 multifunctional degradable 8 plasma protein interactions 33 sterilization 16 thickness 293 impurities 293–4 in vitro in vivo correlation 56 see also biological tests, in vitro; physicochemical tests, in vitro in vivo behavior prediction 107–8 biodegradation 286–90 calcium phosfates (CaPs), use of 98 environment 8–10 in vitro correlation 56 mPCL degradation 351–2 PCL degradation 344 transformation, ceramics 102–4 in vivo test methods 161–77 histomorphometry 170–2, 172–5 imaging 175–6 models 162–70 incubation, bulk degradation stage 361 induction, bulk degradation stage 361 inflammation acute 35, 36–7, 268, 270, 277, 279, 289 chronic 37, 270 mediators 34–5

401

inflammatory multinucleated giant cells 152 injection moulding 211, 216, 217 injury, host response 34–6 innate immunity 34 Insulin-like Growth Factor I (IGF-I) 383 interferon (IFN) 35, 275 interleukins (IL) 35, 275 interpenetrating polymer networks (IPN) 299 intracellular fluid (ICF) 20 ionic concentrations 20 intrinsic pathway 30–1 invasive methods, degradation monitoring 248 ionic concentration blood 27, 147 intracellular fluid (ICF)/extracellular fluid (ECF) 20, 147 ions 150 iron powder 229 γ-irradiation 16, 209, 221–3, 224, 383 kinetics ceramics 101 polymer hydrolysis 189–92 ‘pseudo-first-order’ 189 Kolff, Professor W.J. 84 Krebs cycle 48, 236 labile linkages 118 Lactomer 118 laser profilometry 155 laser sintering 243 late-stage degradation 195–9 leaching, particulate 241 leukocytes 23, 247 extravasation 36 light-stabilisers 230 lipases 287 loading conditions 164, 167 lower critical solution (LCST) temperature 385 low-temperature plasma technology 226

402

Index

Lupron Depot 118 luteinizing hormone-releasing hormone (LHRH) 378 lymph nodes 294 lymphocytes 24–5, 270, 280–2 B 25, 34, 280–1 T 25, 280–1, 285 T-helper 291 macrophages 6, 171, 247 biological factors and 25–6, 32, 37, 150–1 fusion 296 host response and 270, 278–9, 281 ‘tissue resident’ 26 magnetic resonance imaging (MRI) 60, 246 Major Histocompatibility Complex (MHC) 280–2, 291 manufacturing methods 235 mass loss 195–9, 324, 340–1, 346 determination of 249 mPCL scaffolds 331 PCL scaffolds 329 mast cells 24 material control 163–4 factors 290 shape and thickness 153–4 see also scaffold material matrix metalloproteinases (MMP) 288 Maxon 47, 60, 118 MCP (monocalcium phosfate) 91 MCPM (monocalcium phosfate monohydrate) 99 mechanical factors degradation 217–18 stiffness 347 tests 120, 137–9 mechanical properties 193–4, 325 changes in 250 of PCL 342 reduction in 184 stiffness 347

mechanisms degradation 4–5, 9 drug release 373–6 in vivo biodegradation 286–90 polymer degradation classification 321 media alkali 338–43 degradation 55–6 medical devices 67 biomaterials and 43 processing 60 medical polycaprolactone (mPCL) 325–43 gross morphology 327 in vivo degradation 351–2 long-term hydrolytic degradation 326–34 porosity changes 329 melt processing 210–22, 243 melting temperature (Tm) 213 metal deactivators 230 methoxy poly(ethylene glycol) (mPEG) 52 micro-computer tomography 176, 246, 254, 334 microcracks 345 milling, mechanical degradation process 217 models, heterogeneous degradation 51 models, in vivo 162–70 described 164–9 pretest considerations 162–4 resorption 107–8 selected features of 165–6 selection of 169 specialist 168–9 models, for polymers 183–203 cellular automata 200, 203 degradation processes 184–6 future trends 202–3 key processes 186–99 reaction kinetics 201 surface erosion 199–200 temperature effects 201–2 tissue scaffolds 256

Index Modified Atmosphere Packaging (MAP) 229 modified ink-jet printing 243 moisture, packaging and 229 molecular modelling 202 molecular weight (MW) 57–8, 118, 184, 325 changes in 251–2, 364 chitosan 72 degradation 132, 133 dextran monomer 383 distribution 195–6 distribution changes 125–7, 130 gamma irradiation and 222–3 hydrogel erosion and 373 modelling and 184, 193–4, 194–5, 198 mPCL scaffolds 330, 331, 336, 352 mPCL/TCP scaffolds 336 PLA 217 PLGA 360 of polyanhydrides 369–70 polymers with low 362–3 scaffolds and 349 trimodal distribution 130 monitoring methods 247–56 monocalcium phosfate (MCP) 99 monocalcium phosfate monohydrate (MCPM) 99 Monocryl 60, 118, 324 monocytes 24–5, 25–6, 278, 280 monomers 198 aliphatic 371 generation/collection 4 Monte Carlo methods 200, 203, 256 morphological changes 132, 133 morphology 325 most probable distribution, oligomers and 196 moulding 211–12 mPCL see medical polycaprolactone (mPCL) mPEG (methoxy poly(ethylene glycol)) 52 multinucleated giant cells 152, 171–2, 279

403

National Centre for the Replacement, Refinement and Reduction of Animals in Research 162 National Institute for Health (NIH) 173 National Institute for Occupational Safety and Health (USA) 221 natural polymers 67–88, 235–6 alginates 75–82 cellulose 83–7 chitin/chitosan 68–75 naturally-derived materials 10–11 neutrophils 6, 24, 32, 150–1, 277 migration of 294 nitrile 149 nitrogen purging 16 non-cellular factors 271, 273–6 notation system, FEA 187 Nuclear Magnetic Resonance Spectroscopy, proton based (H-NMR) 57, 120 number-average molecular weight (Mn) 125 nylon 284 obesity 72 octocalcium phosfate (OCP) 95, 99–100, 101, 102–4 oligomers 184, 251, 323, 330–1 distribution of soluble 195–7 open pores 246–7 Opsite 74 opsonisation 275 optical coherence tomography 246 optical profilometry 155 oral dosage 357 orthopaedic applications 5, 210 bioresorbable materials 68 metallic 67–8 osteoclasts 6, 152–3, 159, 162, 167, 171–2 Osteomeasure software 173 osteon/haversian architecture 167 β-oxidation pathway 371 oxidative agents 277 oxidative degradation 150–1, 289–90

404

Index

packaging 59–60, 226–9 parallel plate model 175 particle size 290–1 particulate leaching 241 pathways alternative 32–3 classical 32 common 31 extrinsic 30 intrinsic 30–1 Payen, A. 83 PCL see polycaprolactone (PCL); poly(ε-caprolactone) (PCL) PDLA poly(D-Lactic acid) 44 PEG (polyethylene glycol) 44, 86, 369, 382, 383–4 Peripheral Quantitative Computed Tomography (pQCT) 176 permanently implanted devices 343 permeability, polymer 248 PET (polyethylene terephthalate) 195 pH 279 acidic 274 calcium phosfates and 107–8 defined 22 -induced autocatalytic degradation 8 of media 55 variations in 9–10, 12, 53, 253, 283–4 PHA (precipitated hydroxyapetites) 95, 99–100, 102–3 phagocytic steps 152 phagocytosis 6–7, 276–7 crystals, fragmentation of 152 frustrated 279 pharmaceutical applications 70 biodegradable polymer importance 357–8, 359–73 calcium alginate 82 chitin/chitosan 73–5, 76 pharmaceutics 70 pharmacology 117 Phase Contrast Imaging and Tomography (PCIT) 334 phenol 151 phosfate buffer solution (PBS) 326–34, 334–8

phosphonate 149 Phusiline 118 physical factors 273–6 physicochemical factors, drug release and 374 physicochemical tests, in vitro 117–40 methods 121–40 protocol 120 Pitt, C.G. 191 pittings 346 plasma 22 cells 280–2 sterilization 226, 228 plasma protein cascades 28–31 implant interactions 33 plasmin 31 plasminogen activator 31 plaster of Paris (calcium sulfate hemihydrate) 106 plasticisation 194, 256 ‘Plastics: Materials and Processing’ (Brent Strong) 213 platelet derived growth factor (PDGF) 276, 279–80 platelets 23, 26–7, 276, 297 plugs 28–9 Plazlyte 226 PLGA see poly(lactide-co-glycolide) (PLGA) PLLA see poly(L-Lactic acid) (PLLA) Pluronic/HA hydrogels 384 poly(1,3-bis(carboxyphenoxypropaneco-sebacic acid) (PCPP:SA) 370, 381 poly(α-hydroxy acids) 44–7, 189 homopolymer properties 46 poly(amides) 13 poly(amino acids) 12–13 polyanhydrides 12–13, 48, 200, 293 drug release and 368–71, 381–2 structure 369 surface erosion and 323, 330 polycaprolactone (PCL) 47 accelerated degradation 338–43 chain scission and 222

Index drug release and 363–5, 378–80 gross morphology 327, 339 hydrolysis and 198, 322, 323–5, 325–43 /hydroxyapetite composite 254 in vivo degradation 344 long-term hydrolytic degradation 326–34 mechanical degradation and 217 molecular structure 324 physicochemical tests and 118, 124, 126–7, 130–1, 139 porosity changes 329 as scaffold material 237, 239, 241, 243 structure 364 synthetic scaffolds and 319–20 thermal stability of 215 see also medical polycaprolactone (mPCL) polycondensates 48 polycyanoacrylate 149 poly(d,l-lactide) (DLPLA) 360–3, 377, 379 poly(D-Lactic acid) (PDLA) 44, 359–61 polydioxanone 12 sutures 284 poly(DL-lactide-co glycolide) (DLPLGA) polymers 52 polyesters 47–8, 236–7, 293 aliphatic 48–54, 118–19, 120, 121 polyethylene (PE), as packaging material 227–8 polyethylene glycol (PEG) 44, 86, 369, 382, 383–4 polyethylene oxide (PEO) 241 polyethylene terephthalate (PET) 195 polyglactin 284 poly(glycolic acid – caprolactone) (PGACL) 135–7, 138 polyglycolic acid (PGA) 11–12, 44–7, 118, 287 chain scission and 222 changes in crystallinity 194 degradation rate 58, 59–60, 202

405

drug release and 359–63, 377–8 half-life 55, 299, 362 hydrolysis and 322, 333 mechanical degradation and 217–18 mechanical properties 55–6, 137 modelling 183–4, 195 monofilament sutures 324 as orthopaedic implant 286 reaction 53–4 as scaffold material 236–7, 238, 240 self-reinforced (SR-PGA) 58 synthetic scaffolds and 319–20 thermal stability of 215 weight loss 367 polyglycolic sutures 284 polyglycolide see polyglycolic acid (PGA) polyglyconate 47 poly(β-hydroxybutyrate) 11 poly(L-Lactic acid) (PLLA) 44, 46, 52–3, 237 crystallinity and 218–20 degradation behavior 251 gamma irradiation and 222–3 processing degradation and 212 poly(l-lactide) (LPLA) 359–60, 377 poly(lactic acid -caprolactone) (PLACL) copolymers 119 polylactic acid (PLA) 11–12, 44–7, 118, 382 chain scission and 222 degradation rate 56, 57, 58, 59–60, 252 drug release and 359–63, 377–8 e-beam irradiation and 225 half-life 55, 299, 362 hydrolysis and 216, 322, 333 mass loss and 198 mechanical degradation and 217–18 modelling and 183–4, 192, 195, 200 as orthopaedic implant 286 as scaffold material 236–7, 238, 240, 256 stereocopolymers 118, 121–5, 127, 129–30, 132–4, 138–40

406

Index

sutures and 324 synthetic scaffolds and 319–20 TGA of 214 thermal stability of 215 poly(lactic acid/glycolic acid) (PLAGA) copolymers 50–4, 57, 118–22 passim, 126, 128, 134–40 passim polylactide see polylactic acid (PLA) polylactide-co-glycolide (PLGA) 47, 240–1, 246, 383 copolymers 56 degradation rate 202, 362–3 drug release and 377–8 e-beam irradiation and 225 half-life 236, 299 hydrolysis and 293 mass loss 198 mechanical degradation and 218 molecular weight 247, 251–2, 360 as orthopaedic implant 286 particles 27 PLGA-co-PCL implants 379 as scaffold material 243, 248, 249, 250, 254 temperature and degradation 215 water absorbance and 253–4 polymerisation, in situ 216 polymers 4–5 adverse outcomes 285–6 aliphatic 54–60 bioabsorbable, defined 44 biological degradation of 147–51 bioresorbable, defined 44 characteristics/properties 14–15 chemical linkages 291–2 composition 54–5 crystalline 213 degradation 43 degradation mechanisms classification 321 degradation paths 146 drug release applications 385–7 films 210 further information 17 interpenetrating networks (IPN) 299

morphology 56–7, 132, 253–4 pharmaceutical 359–73 semicrystalline 196 tissue scaffolds 247–56 transitions in properties 213 use categories defined 343 see also models for polymers; natural polymers; synthetic polymers polymorphonuclear (PMN) leukocytes 247 polyorthoesters (POEs) 12–13, 365–8 drug release and 365–8, 380–1 structure 366 surface erosion and 323, 330 poly-p-dioxanone (PDS) 47 polypropylene (PP) packaging material 227 suture 284 polypropylene fumarate (PPF) 47–8, 237, 250, 254 polysaccharides 10, 71, 75, 83, 237–8, 373 polysebacic acid (PSA) 48 polystyrene (PS), as packaging material 227 polyurethanes 287 polyvinyl alcohol (PVAlc) 44 as scaffold material 247 polyvinyl chloride (PVC), as packaging material 227–8 pore size distribution 249 types 246 porosity 324, 340–2 changes 329 values 244, 246–7 PPF (polypropylene fumarate) 47–8, 237, 250, 254 precipitated hydroxyapatites (PHA) 95, 99–100, 102–3 primary static indices 173 processing techniques 210–12 degradation and 212–20 temperature reduction aids 230–1 procoagulants 29 profilometry 155–6, 157

Index proliferative phase 270 prostaglandins 34–5 protein adsorption 33, 153 blood content 27 plasma cascades 28–31 plasma/implant interaction 33 proteinase K 120, 124 proteins 238 complement 32 prothrombin activator 29 proton nuclear magnetic resonance (1H NMR) 120, 134–5, 136 PSA (polysebacic acid) 48 Pseudomonas lipase 120, 124 PVAlc (poly(vinyl alcohol)) 44, 247 quenchers, UV and 230 radiation sterilization 228 radiographs 176 random chain scission 195–6, 198–9 in thermal degradation 212 reaction kinetics models 201 reaction mechanisms chemistry 48–9 erosion fronts 53–4 reactive nitrogen species (RNS) 150 reactive oxygen species (ROS) 150 recombinant human growth hormone (rhGH) 383 recrystallization 347, 361 red blood cells (RBC) 23 referents 174 refinement, in animal research 161, 163 regeneration, after injury 270 replacement after injury 270 in animal research 161, 162 resorbable ceramics 5–6 resorption assessment 152–4 modelling 107–8 process 6–7 respiratory burst 289 rhenanite 107

407

Rhizopus arrhizus lipase 124 Rhizopus delemer lipase 124 RNS (reactive nitrogen species) 150 ROS (reactive oxygen species) 150 rotational moulding 211 ‘roughage’ 87 sample conditioning 254–6 geometry/morphology 253–4 size 58–9, 200 storage protocols 255 SAXS (Small Angle X-ray Diffraction) 57 scaffold degradation in vitro 248 structure and 257 scaffold fabrication process 333 techniques 320 scaffold material 319–53 biodegradation/bioresorption 320–3 crystallinity profile 333–4 medical polycaprolactone (mPCL) 325–43 polycaprolactone (PCL) 323–5, 325–43 polycaprolactone-based scaffolds 343–52 stiffness profile 332, 337 yield stress profile 332, 337 scanning electron microscopy (SEM) 154–5, 158, 249 environmental (ESEM) 120, 123–5, 126, 127 implant design and 333, 339, 345, 346 scanning probe microscopy 156–8 Scanning Tunnelling Microscopy (STM) 155 Schwann cells 301 Scion Image analysis system 173 sebacic acid (SA) 370 SEC (size-exclusion chromatography) 57, 120, 125–31 selectins 36 semicrystalline polyesters 237

408

Index

SHA (sintered hydroxyapatites) 95, 97, 99–100, 108 shape, of material 153–4 shearing, mechanical degradation process 217 shelf-life 210 degradable polymers 16 maximising 226–9 shot size, in moulding 211 side-group elimination, in thermal degradation 212 sintered hydroxyapatites (SHA) 95, 97, 99–100, 108 size animal 169–70 exclusion chromatography (SEC) 57, 120, 125–31 particle 290–1 skyscraper algorithm 158 Small Angle X-ray Diffraction (SAXS) 57 smart materials 17 SMC (synthetically modified cellulose) 86 solid free-form fabrication (SFF) 243 solubility 97, 99–101 solvent casting 216, 240 somatic gene therapy 82 species animal 164, 169 influence of 285 specimen see sample sports medicine 68 SR-PGA (polyglycolic acid (PGA), self-reinforced) 58 stability, structural 218–20 stains 170–1, 349 standards British 220 ISO 163, 176 Stanford, E.E.C. 75–6 Staphylococcus albus 284 starch 216, 289 static indices 173–5 steam autoclaving 16 sterilization 59–60, 220–6

implants 16 shelf-life and 228 standard techniques 59 Sterrad 226 stiffness profile 332, 337, 348–9 mechanical 347 PCL scaffolds 342 STM (Scanning Tunnelling Microscopy) 155 storage 226–9 conditions 16 sample protocols 255 Streptococcus mites 284 stroma formation 285 structural stability 218–20 studied polymeric biomaterials 320 sulfonamide 149 sulfonate 149 supramolecular polymers 320 surface analysis 154–9 chemistry 295–7 degradation 49–53 functionalization techniques 33 roughness 153, 154, 158 texture 345 surface erosion 8–9 drug release and 365, 368, 374–5 host response and 292–3 implant design and 323, 330, 340 modelling 184, 186, 199–200 surgery 117 sutures 283–5, 324 synthetic biomaterials 298, 320 synthetic polymers 11–13, 14–15, 43–61 aliphatic polyesters 48–54 aliphatic polymer degradation 54–60 further information 61 host response and 287, 290 medical devices, processing 60 widely used 44–8, 117 synthetic scaffold materials 235–6, 320 T lymphocytes 25, 280–1, 285 -helper 291

Index tartrate-resistant acid phosphatase (TRAP) 172 temperature body 22 effects on hydrolysis 201–2 elevated 274 reduction, shelf-life and 228–9 Temperature Induced Phase Separation (TIPS) 240 temporary factors barriers, degradable polymers as 344 implants 343, 377 scaffold, degradable polymers as 344 tension, mechanical degradation process 217 test methods see biological tests, in vitro; in vivo test methods; models; physicochemical tests, in vitro test protocols, accelerated 184 tetracalcium phosfate (TetCP) 99, 106 therapeutic applications, permanent 117 thermal analysis 339 thermal degradation 212–17 major pathways 212–13 ‘Thermal Degradation of Polymeric Materials’ (Pielichowski) 212 thermal processing, shelf-life and 228–9 thermal stabilisers 230–1 thermoforming 212 thermogravimetric analysis (TGA) 214–15 thermosetting 213 thickness of implants 293 of material 153 of samples 154–5 thioester 148 3D printing 243 thrombosis 29 thrombus formation 273 through pores 246–7 tissue factor 30 resident macrophages 26

409

tissue engineered construct (TEC) 320 scaffold material 322 tissue engineering 16–17, 67, 216, 226 principles 319 synthetic polymers and 117 tissue engineering scaffolds 234–57 characterisation of 244–7 investigation techniques 245 materials 235–8 methods for monitoring 247–56 processing 238–43 toxic effects 9 trabecular factors measures 172 number 175 struts 168 width 174 transesterification reactions 132 transfer moulding 211 transitions in polymer properties 213 transmitting electron microscopy (TEM) 154–5 TRAP (tartrate-resistant acid phosphatase) 172 triblocks 382 copolymers 372 polymers 320 tricalcium phosfate 234 tricalcium phosfate (TCP) 334–8 α-TCP 99, 101, 102–4 β-TCP 99–100, 101–3, 106, 107, 171–2 particles 347–8 tricarboxylic acid cycle 236 triethylene glycol (TEG) 381 trimethylene carbonate (TMC) 47 triple response, acute inflammation 36–7 TRIS (tris hydroxymethyl aminomethane) buffer 107 Tritirachium album 124 tumour necrosis factor (TNF) 35, 275, 296–7, 378 Ultra High Molecular Weight Polyethylene (UHMWPE) 145, 146

410

Index

ultrasonic, mechanical degradation process 217 ultraviolet (UV) radiation 226, 230 University of Singapore 344 urethane 148 vacuum drying 16 vascular factors endothelial cells 270 endothelial growth factor (VEGF) 280 flow, alterations 35–6 permeability 36 vasoconstriction 28 Vicryl 59, 60, 118 Virchow, R. 35 visual examination 121–3 vitamin E 86 Vogel-Tammann-Fulcher (VTF) equation 202 Vroman effect 33 waste recycling 287 water

electrolytes and 20–2 purification industry 72 reduction, shelf-life and 228 weight loss 137, 139 polyglycolic acid (PGA) 367 weight-average molecular weight (Mw) 125 wound closure 10 healing 80, 268 management 73–4 wound dressings alginates and 79–80 general classification 73 products 79 X-ray diffraction (XRD) 120, 132–3 X-ray radiography, plain 175–6 X-rays 170 yield stress profile 332, 337 Young’s modulus 250 Zoladex 118, 378