Degradable Polymers for Skeletal Implants [1 ed.] 9781617283574, 9781606924266

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Copyright © 2009. Nova Science Publishers, Incorporated. All rights reserved. Degradable Polymers for Skeletal Implants, Nova Science Publishers, Incorporated, 2009. ProQuest Ebook Central,

Copyright © 2009. Nova Science Publishers, Incorporated. All rights reserved. Degradable Polymers for Skeletal Implants, Nova Science Publishers, Incorporated, 2009. ProQuest Ebook Central,

Copyright © 2009. Nova Science Publishers, Incorporated. All rights reserved.

DEGRADABLE POLYMERS FOR SKELETAL IMPLANTS

No part of this digital document may be reproduced, stored in a retrieval system or transmitted in any form or by any means. The publisher has taken reasonable care in the preparation of this digital document, but makes no expressed or implied warranty of any kind and assumes no responsibility for any errors or omissions. No liability is assumed for incidental or consequential damages in connection with or arising out of information Degradable Polymers for Skeletal Implants, Ebook Central,that the publisher is not engaged in contained herein. Nova This Science digital Publishers, documentIncorporated, is sold with2009. the ProQuest clear understanding

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DEGRADABLE POLYMERS FOR SKELETAL IMPLANTS

PAUL I.J.M. WUISMAN AND

THEODOOR H. SMIT Copyright © 2009. Nova Science Publishers, Incorporated. All rights reserved.

EDITORS

Nova Science Publishers, Inc. New York

Degradable Polymers for Skeletal Implants, Nova Science Publishers, Incorporated, 2009. ProQuest Ebook Central,

Copyright © 2009 by Nova Science Publishers, Inc. All rights reserved. No part of this book may be reproduced, stored in a retrieval system or transmitted in any form or by any means: electronic, electrostatic, magnetic, tape, mechanical photocopying, recording or otherwise without the written permission of the Publisher. For permission to use material from this book please contact us: Telephone 631-231-7269; Fax 631-231-8175 Web Site: http://www.novapublishers.com NOTICE TO THE READER The Publisher has taken reasonable care in the preparation of this book, but makes no expressed or implied warranty of any kind and assumes no responsibility for any errors or omissions. No liability is assumed for incidental or consequential damages in connection with or arising out of information contained in this book. The Publisher shall not be liable for any special, consequential, or exemplary damages resulting, in whole or in part, from the readers’ use of, or reliance upon, this material. Any parts of this book based on government reports are so indicated and copyright is claimed for those parts to the extent applicable to compilations of such works.

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Independent verification should be sought for any data, advice or recommendations contained in this book. In addition, no responsibility is assumed by the publisher for any injury and/or damage to persons or property arising from any methods, products, instructions, ideas or otherwise contained in this publication. This publication is designed to provide accurate and authoritative information with regard to the subject matter covered herein. It is sold with the clear understanding that the Publisher is not engaged in rendering legal or any other professional services. If legal or any other expert assistance is required, the services of a competent person should be sought. FROM A DECLARATION OF PARTICIPANTS JOINTLY ADOPTED BY A COMMITTEE OF THE AMERICAN BAR ASSOCIATION AND A COMMITTEE OF PUBLISHERS. LIBRARY OF CONGRESS CATALOGING-IN-PUBLICATION DATA Wuisman, Paul I. J. M. Degradable polymers for skeletal implants / Paul I.J.M. Wuisman and Theo M. Smit. p. cm. Includes index. ISBN H%RRN 1. Polymers in medicine. 2. Polyesters--Biocompataibility. 3. Orthopedic implants. I. Smit, Theo M. II. Title. R857.P6W85 2009 610.28'4--dc22 2008046727

Published by Nova Science Publishers, Inc. Ô New York

Degradable Polymers for Skeletal Implants, Nova Science Publishers, Incorporated, 2009. ProQuest Ebook Central,

DEDICATION

Copyright © 2009. Nova Science Publishers, Incorporated. All rights reserved.

In memory of professor Paul I.J.M. Wuisman MD, PhD 18 August 1956 - 25 July 2007

Degradable Polymers for Skeletal Implants, Nova Science Publishers, Incorporated, 2009. ProQuest Ebook Central,

Copyright © 2009. Nova Science Publishers, Incorporated. All rights reserved. Degradable Polymers for Skeletal Implants, Nova Science Publishers, Incorporated, 2009. ProQuest Ebook Central,

CONTENTS Preface

ix

About the Editors

xi

Part I: Basic Science and Engineering

1

Chapter 1

The Molecular Structure of Degradable Polymers Carmen Scholz

Chapter 2

Time-Dependent Failure in Load-Bearing Polymers. A Potential Hazard in Structural Applications of Polylactides Leon E. Govaert, Tom A.P. Engels, Serge H.M. Söntjens and Theo H. Smit

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Chapter 3

Poly(lactide)s and their Copolymers: Physical Properties and Hydrolytic Degradation Hideto Tsuji

3

21

41

Chapter 4

Composites Based on Degradable Polymers K. E. Tanner

71

Chapter 5

Product Realization: The Processing of Bioabsorbable Polymers G. Lawrence Thatcher

93

Chapter 6

Sterilization of Biodegradable Polymers Tuija Annala and Minna Kellomäki

123

Chapter 7

Surface Properties of Degradable Polymers Theo G. van Kooten and R. Kuijer

139

Chapter 8

Biodegradation and Autocatalysis of Polylactides Sudhir S. Chakravarthi and Dennis H. Robinson

163

Chapter 9

Biological Testing of Degradable Polymers in Vivo O.M. Böstman

177

Part II: Clinical Applications of Degradable Implants Chapter 10

Fracture Repair with Bio-Resorbable Implants C.J. van Manen and M. van der Elst

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195 204

viii

Contents

Chapter 11

Fibula Regeneration after Vascularized Fibular Graft Harvesting Arthur de Gast, Hay A.H. Winters and Paul I.J.M. Wuisman

211

Chapter 12

Degradable Polymers in Cranio-Maxillofacial Surgery U. Eckelt

229

Chapter 13

Absorbable Materials in Shoulder Surgery Lennart Magnusson, Jüri Kartus and Lars Ejerhed

241

Chapter 14

Degradable Polymers in Hand Surgery Abigail R. Hamilton, Chaitanya S. Mudgal and Jesse B. Jupiter

253

Chapter 15

Degradable Polymers as ACL Substitutes Erica D. Taylor and Cato T. Laurencin

267

Chapter 16

Fixation of ACL Grafts with Degradable Polymer Screws Jon Olav Drogset

281

Chapter 17

Degradable Polymers in Meniscus Reconstruction Eric L. W. de Mulder,Gerjon Hannink, Tony G. van Tienen and Pieter Buma

295

Chapter 18

Bioabsorbable Implants: Cervical Spine Mark Dumonski, Kern Singh and Alexander R. Vaccaro

313

Chapter 19

Degradable Polymers in the Lumbar Spine T. U. Jiya

321

Part III: Innovation and Future Developments

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Chapter 20

Chapter 21

337

Areas of Applications and Limitations for Degradable Polymer Implants Kurt Ruffieux

339

Perspectives and Possibilities for Degradable Polymers for Skeletal Implants Ying Deng, Najmuddin Gunja and Kyriacos A. Athanasiou

351

Contributors

363

Index

371

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PREFACE When skeletal structures or tissues fail due to trauma or disease, additional support is required to take over the mechanical function of the structures involved. Traditionally, skeletal implants are made of metal, but they essentially have a temporary function: once healing is achieved, their removal is desired both from both the clinical and biomechanical point of view. This consideration motivated the development of degradable implants, which have the evident advantage over metal devices that they degrade over time and thus eliminate the necessity of retrieval operations. In addition, the healing process may be stimulated by the successive loss of their mechanical properties, corresponding with increased loading on the healing tissues. Degradable polymers, however, have their own drawbacks and pitfalls. They are much less strong than metals, and their degradation products have been reported to cause inflammation and other tissue reactions. Another important issue is that degradables should behave in a predictive way: they should maintain mechanical integrity for a certain period of time, and then degrade slowly in order to avoid tissue reactions. However, it now appears that degradable polymers behave differently at different implantation sites. Furthermore, the visco-elastic nature of polymers under sustained mechanical loading may lead to untimely implant failure, which needs to be considered when designing the implant. Finally, the quality of the polymer strongly depends on the raw material, the processing techniques, the thermal and loading history, sterilisation, and other parameters. So, a thorough understanding of polymer chemistry and mechanics, as well as surgical techniques and in vivo conditions of the implant are required to successfully apply degradable polymers for skeletal implants. The purpose of this book is present an integrated overview of the properties of degradable polymers and their application in skeletal surgery. The book will consist of three parts: A. Basic Science and Engineering; B. Clinical application of degradable implants; and C. Innovation and future developments. All chapters are written by experts with a substantial track record on the specific subject. This book on Degradable Polymers for Skeletal Implants was an idea of my friend and colleague Paul Wuisman, who passed away on 25 July 2007, only 50 years old. Paul was a visionary surgeon and an inspiring leader, determined, disciplined, and demanding. He aimed to deepen our understanding of skeletal tissue biology and to create a scientific basis for the treatment of his patients. Paul suggested the development of degradable spinal cages in 1997 and a cervical cage was introduced clinically in April 2007. Throughout his career he challenged his students and colleagues and motivated them to think beyond their boundaries.

Degradable Polymers for Skeletal Implants, Nova Science Publishers, Incorporated, 2009. ProQuest Ebook Central,

x

Theodoor H. Smit

With his dedication and positive attitude, he touched the lives of many. He will sorely be missed by his patients and by his colleagues and friends from all around the world, who had the privilege to live and work with him. May he rest in peace.

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Theodoor H. Smit, MSc, PhD 2008 in Amsterdam, The Netherlands

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ABOUT THE EDITORS

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Prof. Paul I.J.M. Wuisman MD PhD (1956-2007) studied Medicine at the University of Utrecht (Netherlands) from 1975-1982. He fulfilled his orthopaedic residency at the Trauma Center in Linz (Austria) and at the Departments of Orthopaedic Surgery at the Universities of Würzburg and Münster (Germany). He registered as an orthopaedic surgeon in 1988. From 1988-1990 he was a research fellow in the fields of infection diseases and orthopaedic oncology at the University of Florida, Gainesville (U.S.A.). He completed his PhD-theses on Giant Cell Tumor of Bone in 1988 and on osteosarcoma in 1991. In 1991 he became Head of the Department of Orthopaedic Surgery at the Westfälische Wilhelms-Universität in Münster (Germany) and since 1994 he was professor of the Department of Orthopaedic Surgery at the Vrije Universiteit Medical Center Amsterdam. He was co-founder and first president of the Skeletal Tissue Engineering Group Amsterdam (STEGA). His main research interests were in the areas of the Spine, Stem Cell Technology and Biomaterials. Paul Wuisman passed away on 25 July 2007. Theodoor (Theo) H. Smit, MSc, PhD (1965), holds a Master's Degree (ir.) in Mechanical Engineering (main: micro-technology) from the Technical University Delft (Netherlands) since 1989. He completed his PhD-study on the mechanical significance of the trabecular bone architecture in a human vertebra at the Technical University Hamburg-Harburg (Germany) in 1996. Since 1996, he is affiliated to the Department of Physics and Medical Technology of the Vrije Universiteit medical center, Amsterdam (Netherlands), since 2007 as associate professor in Skeletal Physics and Tissue Engineering. His current topics of research are related to bone remodelling, spinal fusion, degradable polymers, spine (Bechterev, scoliosis), tissue engineering, and mechano-biology. He is the founding director of the Skeletal Tissue Engineering Group Amsterdam (STEGA foundation, 1998-2007) and currently acts as its president.

Degradable Polymers for Skeletal Implants, Nova Science Publishers, Incorporated, 2009. ProQuest Ebook Central,

Copyright © 2009. Nova Science Publishers, Incorporated. All rights reserved. Degradable Polymers for Skeletal Implants, Nova Science Publishers, Incorporated, 2009. ProQuest Ebook Central,

Copyright © 2009. Nova Science Publishers, Incorporated. All rights reserved.

PART I: BASIC SCIENCE AND ENGINEERING

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In: Degradable Polymers for Skeletal Implants Editors: P.I.J.M. Wuisman and T. H. Smit

ISBN 978-1- 60692-426-6 © 2009 Nova Science Publishers, Inc.

Chapter 1

THE MOLECULAR STRUCTURE OF DEGRADABLE POLYMERS Carmen Scholz* Department of Chemistry, University of Alabama in Huntsville, AL, USA

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ABSTRACT This chapter describes the chemistry of degradable polymers with a special emphasis on polyesters. The chemistry of esters in general is reviewed and provides an introduction to the understanding of the processes that lead to the formation of polyesters and their subsequent hydrolytic degradation. The success of biomedical implants depends largely on four factors: their surface properties and interaction with adjacent tissue, their overall biocompatibility, their medically unobjectionable degradation, secretion and/or resorption and their mechanical strength. Polyesters, in particular poly(lactide), poly(glycolide) and copolyesters thereof have been proven to best fulfill the demands set forth by orthopedic implants and a multitude of commercial products is available. Synthesis, structureproperty relationships, biocompatibility and biodegradation behavior of poly(lactide), poly(glycolide) and their copolyesters are comprehensively discussed. Other polyesters, such as poly(hydroxyalkanoates) and poly(caprolactone) are briefly described. The chemical complexity of poly(urethanes) offers a plethora of physical properties and combined with a rather sufficient biocompatibility makes this class of polymers an interesting candidate for novel biomedical materials.

INTRODUCTION Mankind has used polymeric materials in medical application for thousands of years. Natural polymers, such as cotton, a polysaccharide, silk and wool, poly(amino acid)s, also known as proteins, have been used as wound dressings after impregnating them with extracts from medicinal plants. Wood, which consists of lignin, a complex, crosslinked polyphenol and polysaccharides, was used to splint broken bones. By the 19th century the manufacturing *

301 Sparkman Dr., Huntsville, AL 35899, [email protected].

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Carmen Scholz

of new synthetic material became feasible. In fact, the realm of synthetic polymers started in part through accidents (In 1839, Charles Goodyear’s experiments with natural rubber and a spilled sulphur pot led to the discovery of the vulcanization process) and in part by focused research (In 1860, John Wesley Hyatt sought a material to replace ivory billiard balls and developed celluloid, which later on was also used as dentures and worked well as long as the wearer would abstain from hot food). A manifold of new materials was generated but the molecular structure of these novel materials remained unknown. The concept of macromolecules, that is, long, thread-like molecules with molecular weights of several 10,000’s to several million grams per mole was developed by Staudinger [1] in the 1920’ies, and it took 10+ years, during which he encountered ridicule and discouragement, until this novel idea of gigantic molecules was accepted by the scientific community and earned him a Nobel Prize in 1953. Modern research into polymeric materials was driven by two main forces: (i) to improve and enhance the properties of existing materials and (ii) to generate new materials to satisfy an increasing demand caused by technological progress. Polymeric materials or ‘plastics’ as they became known, quickly conquered the materials market. They were initially designed to replace structural materials, therefore they are characterized by toughness, durability, high impact and tensile strength (Kevlar, a chemical relative of Nylon is used in bullet-proof vests), increased performance, improved reliability and the ability to adjust their properties to match intended uses. Plastics show virtually no corrosion and age slowly. The aspect of intended depolymerization or degradation came only later into focus.

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BIODEGRADABLE POLYMERS The development of degradable polymers for biomedical applications is tightly intertwined with the development of degradable plastics developed for the packaging industry and one-time use articles. More than 100 million tons of plastics are produced every year worldwide [2] and concerns about degradability became prominent in the 1980’ies as environmental awareness grew in the face of ever-growing landfills and omnipresent plastic litter. While the packaging industry has turned to degradable materials only very recently, often due to legislative pressure, degradable materials were immediately met with sincere interest by the biomedical sector. The idea of synthesizing materials that degrade within the environment of a human body, preferably according to a pre-determined time-frame and possibly aided by the biochemical actions of mammalian cells became increasingly appealing to the biomedical community. Degrading sutures that do not require removal, bone implants that first replace bone fractions that were lost due to disease or accident, stimulate bone growth and subsequently degrade, drug-carrying surface coatings for permanent implants, drug delivery vehicles that degrade after their drug cargo has been delivered to target cells, and recently delivery vehicles for genes that encode for therapeutic proteins and for RNA sequences that oppress the transcription of harmful proteins – all rely on biocompatible polymeric materials that degrade or at least can be secreted after their mission is fulfilled. Biodegradable polymers can be obtained from two different sources: (i) natural polymers can be used as generated in nature after purification or can be further derivatized to render their properties. Nature produces several classes of polymers: polysaccharides (cellulose,

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The Molecular Structure of Degradable Polymers

5

chitin, starch), poly(amino acid)s (proteins and peptides), polyphenols (lignin), polyolefins (natural rubber) and polyesters (polyhydroxyalkanoates). All of those materials are the result of enzymatically catalyzed reactions. It is one of the fundamental principles in nature that what is produced by one enzyme can be degraded by another one. In terms of natural polymers; the action of any polymerase is counteracted by a depolymerase. Therefore, any material of natural origin is inherently biodegradable, thus the balance of growth and decay is kept. Whether the depolymerases are expressed in the target environment, e.g. human tissue, is a different subject. Typically, mammalians do not express depolymerase enzymes as they are found in soil microorganisms that bioremediate compostable, biodegradable (packaging) materials; (ii) polymers can be synthesized by converting the fractionation products obtained in petroleum refineries into monomers by applying ordinary methods of organic synthesis. These monomers are then converted into polymers using conventional polymerization techniques. Polymer science uses addition polymerizations, which employ either radical or ionic techniques and step polymerizations, which use condensation reactions. Those basic polymerization techniques yield the multitude of synthetic plastic materials that we are familiar with: polyesters, polyamides, polyolefins, polyurethanes, polycarbonates, polyanhydrides and other specialty polymers. This chapter will concentrate mainly on one class of polymers: polyesters, as they provide the properties that are needed in orthopedic applications in terms of strength, durability and degradability. Some considerations will be given to polyurethanes as they have garnered some consideration in the field of biomedical polymers.

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THE CHEMISTRY OF ESTERS Figure 1 shows an ester group, it is characterized by a carbon atom that is bound by a double bond to one oxygen atom (carbonyl structure) and by a single bond to another oxygen atom. The carbon atom and the singly bound oxygen atom are covalently bound to the rest of the molecule, indicated by R and R’. A prominent example for a small molecule ester is shown in Figure 2.

O R

C

O

R'

Figure 1. An ester bond connects two molecular fragments R and R’ to form a new molecule, which is called an ester.

Figure 2. Isopentyl acetate, an ester found in banana oil gives bananas their typical fragrance.

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Carmen Scholz

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Figure 3. The formation of an ester from an acid and an alcohol. One molecule of water is also produced in this condensation reaction.

Figure 4. Mechanism of the hydrolytic ester degradation. The nucleophilic attack of a hydroxyl anion leads to the cleavage of the ester bond and yields an alcohol and an acid.

Esters often have a pleasant fragrance or taste, the molecule shown in Figure 2 gives bananas their typical fragrance. Other examples for esters produced by plants are: butyl butanoate and methyl butanoate (pineapple), benzylbutanoate (rose), methyl salicylate (wintergreen), benzyl acetate (oil of jasmine). Synthetic esters are commonly used as food additives and fragrances. Esters are produced by the condensation reaction of an acid with an alcohol under release of one molecule of water; see Figure 3 for a general esterification reaction. Reversing this reaction, also called ester hydrolysis (the reaction of an ester with water) or saponification reaction (base catalyzed reaction of an ester with water), degrades the ester, see Figure 4, yielding the acid and alcohol from which the ester was originally synthesized. Hydrolysis reactions are typically catalyzed by acids or bases. As it is apparent in Figure 4, the presence of water is crucial for the degradation of any ester (also polyesters). Esters do not degrade in the absence of water. Water dissociates in hydroxyl anions and protons, which combine with another water molecule to form a hydronium ion, H3O+. As shown in Figure 4, the hydroxyl anion attacks the partially positively charged carbonyl carbon of the ester bond, resulting in

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The Molecular Structure of Degradable Polymers

7

the formation of a tetrahedral intermediate state (1). The tetrahedral state is not stable and the carbonyl group (C=O) is reformed by cleaving the C – O bond, thus yielding the acid and the alcoholate (2). The alcohol is regenerated by the reaction of the alcoholate with the proton that was generated by the dissociation of water (3). In summary, the alcohol and acid from which the ester was formed, are obtained at the end of the hydrolysis reaction. Other factors that determine the degradability of an ester and the rate at which this reaction occurs in polyesters are discussed below.

THE CHEMISTRY OF POLYESTERS

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Molecules that contain a multitude of ester bonds along a polymeric chain are called polyesters, see Figure 5. Depending on the structure of the monomer(s) used, one of two general polyester structures, shown in Figure 5a and 5b, are obtained in polyester syntheses. Poly(lactic acid) is represented by structure 5a, and poly(ethylene terephthalate) is a typical example for structure 5b. The repeat unit is shown in brackets and the subscript “n” denotes how many times this repeat unit is in average present in the polymer. The number of repeat units is also called the average Degree of Polymerization, DP. The synthesis of polyesters is more complex than the synthesis of small molecule esters. In order to guarantee the formation of a macromolecule, the monomers must be bifunctional, thus allowing for a polymer chain with multiple ester linkages to develop. Polyesters can be synthesized from heterobifunctional monomers or from diacids (or diacid derivatives) and diols. Heterobifunctional monomers, see Figure 6, carry the alcohol as well as acid function, which are necessary for the formation of an ester bond, within one molecule and are called hydroxyl carboxylic acids. Polyesters are formed by the repeated self-condensation reaction of the hydroxyl carboxylic acid and water must be continuously removed to achieve a high degree of polymerization.

a.

O O

R'

O

C

O R

C

O O

R'

O

C

O R

C O

R'

n b. Figure 5. Structural elements of polyesters, Figure 5a depicts a polyester synthesized from a heterobifunctional monomer, Figure 5b depicts a polyester synthesized from an diol and a diacid or diacid derivative.

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8

a. Figure 6a. A hydroxyl carboxylic acid, a heterobifunctional monomer for a polyester synthesis.

O 2 HO

R

C

O OH

O C

R

R

C

O + 2 H2O

O

b. Figure 6b. Dimerization of the hydroxyl carboxylic acid to yield a lactide.

O

O HO

C

R

C

O

O OH

+

HO

R'

OH

C

R

C

O R'

O

n

+

(n-1) H2O

Figure 7. Schematic depiction of the formation of a polyester from a diacid and a diol. This is a condensation reaction and the small molecule, here water, must be constantly removed in order to achieve high molecular weights.

O O

CH* CH2

C

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CH3

n

a.

O O

CH* CH2 CH3

O

C

O n

CH* CH2 CH2 CH3

C m

b. Figure 8.Bacterial polyesters; Figure 8a. structure of poly(β-hydroxybutyrate), a homopolymer, Figure 8b. poly(β-hydroxybutyrate-co-valerate), a random copolymer.

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Typically, polymerizations of these monomers are conducted after two monomers have been combined to form a dimer, Figure 6b. This activated ring-form of a heterobifunctional monomer is more reactive than the hydroxyl carboxylic acid shown in Figure 6a and readily undergoes ring-opening polymerizations. Two homobifunctional molecules, see Figure 7, are used to synthesize polyesters from diols and diacids, or diacid derivatives, respectively. As the equilibrium reaction greatly disfavors the formation of polyesters, rather complex technical processes, which will not be discussed here, are employed to ensure the formation of high-molecular weight polyesters. Bacterial polyesters are another type of polyesters that conquered some attention in the biomedical sector. These polyesters, chemically known as poly(β-hydroxyalkanoate)s, with poly(β-hydroxybutyrate), PHB, Figure 8a, as its best known representative, are produced enzymatically by bacteria in the form of inclusion bodies for the purpose of internal carbon and energy storage. Poly(β-hydroxybutyrate-co-β-hydroxyvalerate), PHBV, is another common bacterial copolyester, Figure 8b. The carbon source, for instance glucose, is converted to pyruvate by glycolysis. The polymerization involves a cascade of three enzymes: β-ketothiolase, encoded by the phbA enzymes A and B, transforms the acetyl-CoA into acetoacetyl-CoA, the acetoacetyl-CoA reductase, encoded by the phbB enzyme, transforms acetoacetyl-CoA into β-hydroxybutyryl-CoA, which is finally polymerized by the PHB synthase, encoded by the phb C enzyme [3,4].

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POLYMER TERMINOLOGY A polymer is defined as a material that (i) undergoes swelling before dissolution when in contact with a solvent, (ii) forms highly viscous solutions at rather low concentrations, (iii) has a non-uniform particle size distribution and (iv) does not undergo dialysis. The most important physical characteristics of polymers are: the molecular weight and the glass transition temperature. Contrary to small molecules, the molecular weight of polymers is nonuniform. The glass transition is a property unique to polymers. The molecular weight of polymers is usually reported in Dalton (Da); one Dalton equals one gram per mole. A molecular weight of 100,000 Da indicates that one mole (3.023 x 1023 molecules or polymer chains) weigh 100,000 g or 100 kg. All reported molecular weight data are averages. Depending on the polymerization technique, molecular weight distributions can be very broad (radical polymerizations) or narrow (anionic and living polymerizations). The Polydispersity Index, PDI, is a measure of the broadness of molecular weight distribution. The closer the PDI value is to unity the narrower is the distribution in molecular weight, meaning all chains are approximately of the same length. A narrow, monodisperse molecular weight distribution is typically advantageous to the processing of the material, and a necessity when synthesizing self-assembling structures, for instances for drug delivery purposes. The glass transition temperature, Tg, is defined as the “onset temperature of molecular motion”. At this particular temperature, the polymer transitions from a glassy to a rubbery state for amorphous polymers or a thermoplastic state for semi-crystalline polymers. It is the temperature at which the chains have enough energy to slide by one another in a reptationlike motion as response to stress or strain. Since not all polymers have a melting point, some decompose before melting and amorphous polymers transition gradually from the rubbery

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state to a gum-like state and finally to a liquid, the Tg is more commonly used than the melting point when describing the physical properties of a polymer. The tacticity of a polymer describes the sequencing of asymmetric carbon atoms along the polymer chain. Three possibilities exist: atactic polymers contain no regular sequence in the asymmetric carbon atoms; isotactic polymers are characterized by the presence of only one symmetry type for the asymmetric carbon (both, R/S, and D/L notations are used), see Figure 8a and b for an example for isotactic polymers, all asymmetric carbon atoms in PHB and PHBV are in R-configuration; syndiotactic polymers are characterized by an alternating configuration of the asymmetric carbon atoms. Only enzymatic techniques, as in the synthesis of bacterial polyesters lead to absolutely fault-free isotactic polymers. Chemical approaches have been developed to yield highly isotactic or highly syndiotactic polymers. Polymers are either amorphous, that means no long-range order can be observed in X-ray crystallography investigations or semi-crystalline with varying degrees of crystallinity, which can be observed and quantified by wide-angle X-ray scattering studies. Unlike small molecules, polymers are never completely crystalline. Semi-crystalline polymers crystallize by chain folding which yields lamellar structures. The slower the crystallization, either out of solution or melt, the higher is the degree of crystallinity. Thus, the cool-down rate in polymer processing determines the degree of crystallinity and thus the physical and degradation properties of the resulting material. On the other hand, quenching a polymer melt always yields an amorphous polymer, by freezing-in the three-dimensional arrangement the chains had in the melt, a disorganized state that is called a random coil. Whether a polymer is able to crystallize depends on its chemical structure. In general, polarity, isotacticity and regularity enhance a polymer’s tendency to crystallize. Atactic and branched polymers are amorphous.

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BIOCOMPATIBILITY OF BIODEGRADABLE POLYMERS Before considering individual biopolymers, the issue of biocompatibility should be addressed. Biocompatibility must always consider two concerns: (i) the effect of an implant on the surrounding tissue and (ii) the effect of surrounding tissue on the implant. For any material to be considered as an implant material or other biomedical material, such as drug delivery vehicle, it has to be biocompatible. Biocompatibility can be defined in several ways, and is also target-organ dependent, but the smallest common denominator includes the following properties and requirements: the material must not be toxic and must not induce inflammatory responses, degradation products must be non-toxic and resorbable or secretable. In addition, the degradation time should match the healing time, and the mechanical properties of the biomaterial should match the mechanical properties of the tissue it is replacing. Finally, the material must be sterilizable. The surface of any biomedical material is most crucial as it is in immediate contact with the surrounding tissues and fluids and is exposed to the immediate chemical and physical response of the host tissue [5]. While the bulk properties of an implant material have less impact on the biocompatibility and biodegradability and should primarily be designed to match the mechanical properties of the tissue they are replacing, the surface of an implant material deserves outmost attention. Biodegradable polymers are currently investigated as temporary implants, porous structures for tissue engineering and for pharmacological applications such as delivery

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vehicles for drugs and genes. Depending on the application, size and shape of biodegradable materials vary significantly, ranging from nanometer-sized drug carrying micelles to centimeter-sized bone screws and plates. Biodegradable polymers for macroscale applications can be fabricated into solid three-dimensional objects, spun into filaments and fabricated into meshes and prepared as porous foams. Polyesters have gained the widest application in biomedical materials because of their susceptibility to hydrolysis, acceptable degradation rate and non-toxic degradation products. Poly(amino acids) are equally, or possibly more biocompatible if comprised of natural amino acids, but their degradation rates are considerably slower due to the higher hydrolytic stability of the amide (peptide) bond, - C(O) – NH -.

POLYESTERS AS BIOMEDICAL MATERIALS As indicated before, polyesters are the most suitable materials for orthopedic and skeletal applications, because of their degradation rate, which matches the bone healing rate and their biocompatible degradation products. Polyesters are thermoplastic materials, meaning they are linear polymers that soften when heated. These materials can be processed from the melt, and do not require the use of potentially harmful solvents for their processing. A multitude of monomers is available, resulting in a large number of polyesters with varying physical, chemical and mechanical properties.

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1. Poly(lactic acid), PLA Lactic acid is the starting material for the production of poly(lactic acid) and can be produced chemically and biologically. The biological approach is based upon the fermentation of carbohydrates by engineered Lactobacilli strains and is more cost-effective than the chemical approach; therefore, it is now prevalent [6]. Lactic acid is an optical active molecule with one stereo-center, indicated by *, see Figure 9. Different strains of Lactobacilli produce either predominantly the D or L-form of lactic acid. [7]. Lactobacilli amylophilus, L. bavaricus, L. casei, L. maltarmicus and L. salivarius produce the L-isomer, L. delbrueckii, L. jensenii and L. acidophilus produce either the D-isomer or a racemic mixture of the two isomers. Dimerization of lactic acid yields lactide, which is the starting material for the production of poly(lactic acid), see Figure 10. The dimerization product is a chiral molecule with two optically active carbon atoms. The chiral carbon atoms can be either in D- or Lconfiguration or the lactide exists as DL- or meso lactide. L-lactic acid is the naturally occurring isomer and is also produced in muscle during intense activity under oxygen limiting conditions, catalyzed by lactate dehydrogenase.

O HO

CH* C

OH

CH3 Figure 9. Structure of lactic acid, note the stereocenter (*).

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Figure 10. Ring-opening polymerization of lactide yields poly(lactic acid).

O

CH2 C

O C CH2

O

O

O Ring opening Polymerization

O

CH2 C n

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Figure 11. Ring-opening polymerization of glycolide yields poly(glycolic acid).

Figure 12. Ring-opening polymerization of dioxanone yields poly(dioxanone), (top) and ring-opening polymerization of ε-caprolactone yields poly(caprolactone), (bottom).

Lactide is polymerized in the melt (melting point of meso lactide: 126-127 °C, melting point of D- and L-lactide: 97°C) by ring-opening polymerization with 0.2% stannous octanoate as initiator at a reaction temperature between 170 – 190°C. The polymerization reactions follow a rather complicated coordination/insertion mechanism [8]. Polymerization of L-lactide leads to the formation of poly(L-lactide), PLLA, a semi-crystalline polymer with a degree of crystallinity of approximately 40%. As described above, the degree of crystallinity depends on the processing conditions and can be adjusted. PLLA has a Tg of 53 64 °C and a melting point of 175°C [9]. The ring-opening polymerization of the racemic meso-lactide yields the amorphous poly(D,L-lactide), PDLLA, with a Tg of 50 – 57 °C, and a melting point of approximately 180°C [9].

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A new azeotropic dehydration condensation has been developed in the 1990’ies and allows for the direct polymerization of lactic acid, also yielding molecular weights above 100,000 Da. [10-13]. This technique yields high molecular weights but has the disadvantage that considerable amounts of catalyst are carried over into the product as impurities. The presence of minute amounts of these tin catalysts can be problematic in biomedical applications, causing catalyst toxicity, unwanted degradation, uncontrolled and nonreproducible hydrolysis thus requiring extensive purification procedures. Due to their different stereochemistry, PLLA and PDLLA have different physical characteristics and a different degradation behavior, thus resulting in different applications. The semi-crystalline PLLA has a high modulus of approximately 4.8 GPa, good tensile strength and a low extension and is therefore considered to be a useful material for loadbearing applications. Phantom Soft Thread Soft Tissue Fixation Screw®, Phantom Suture Anchor® (DePuy), Full Thread Bio Interference Screw® (Arthrex), BioScrew®, BioAnchor®, Meniscal Stinger® (Linvatec) and Clearfix Meniscal Dart® (Innovasive Devices) are PLLA-based orthopedic products [14]. The amorphous PDLLA has a much lower modulus (~ 1.9 GPA) than PLLA. Since it is amorphous it degrades faster than PLLA. Combining the lower strength with an accelerated degradation makes it a preferred material for drug delivery systems, scaffolds and degradable sutures when copolymerized with PGA.

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2. Poly(glycolic acid), PGA Poly(glycolic acid), PGA, is a highly crystalline polymer (crystallinity > 50%) that crystallizes readily due to the absence of side groups, see Figure 11, PGA has a Tg of 36 - 45 °C and a melting point of approximately 230°C [9]. It has high mechanical strength with a modulus of about 12.5 GPa. Based on its structure PGA is chemically closely related to PLA, compare Figure 10, but lacks the asymmetric carbon atom. The absence of the methyl side groups leads to the above-mentioned high crysallinity. PGA is considered the first synthetic biodegradable polymer for biomedical application and is sold as synthetic sutures under the trade name Dexon® since its approval by the FDA in 1969. Monofilaments are formed by melt-extrusion and braided into sutures.

3. Copolymers of PGA and PLA, Poly(lactide-co-glycolide), PLGA PGA provides a high modulus and a rather rapid rate of degradation due to the polymer’s high hydrophilicity. PLLA on the other hand is characterized by a lower modulus and also a lower rate of degradation. One way to tailor the properties to match a specific application is by copolymerizing lactide with glycolide. Poly(lactide-co-glycolide)s with varying ratios of glycolide to lactide have been developed commercially, using both, PLLA and PDLA. Poly(L-lactide-co-glycolide) copolymers form amorphous polymers in the composition range between 25 and 75%. Compared to the homopolymers, PLGA copolymers are much more hydrolytically labile. 50/50 mixtures are very hydrolytically instable and degrade within 1 to 2 months, 75/25 poly(DL-lactide-co-glycolide) degrades within 4 to 5 months and the 85/15 mixture within 5 to 6 months [15]. Due to their easier processing PGLA copolymers have

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been developed for a variety of biomedical applications, including sutures [16], screws and plates [17], meshes for skin replacement (Dermagraft®), tissue engineering scaffolds due to their good cell adhesion [18, 19] and drug-delivery vehicles, such as micro- and nano-spheres, microcapsules and nanofibers [20, 21].

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4. Bacterial Polyesters Poly(β-hydroxybutyrate), see Figure 8a is a highly crystalline (crystallinity > 80%) polymer with a Tg of 4°C and a melting point of 175°C [9]. The molecular weight is dependent on the bacterial strain and varies between 100,000 and more than 1 Mio Da. Due to their enzymatic synthesis all bacterial polyesters are isotactic with the chiral carbon atoms in R-configuration. The resulting hydrolysis product, R-β-hydroxy butyric acid, is a mammalian metabolite that is present in low concentrations in blood. Because of its brittleness the polymer is rarely used in its homopolymer form, but is rather produced as a random copolymer; poly(β-hydroxybutyrate-co-β-hydroxy valerate), PHBV, Figure 8b. The physical and mechanical properties of the copolymer can be tailored according to application needs. Increasing the amount of hydroxyvalerate leads to reduced Tg‘s and melting points. Flexibility and toughness increase as impact strength increases and Young’s modulus decreases with increasing amounts of hydroxyvalerate [22]. Increasing amounts of hydroxyvalerate reduce the crystallinity, and the degradation rates for these copolymers increase. While PHB and PHBV are not well-suited materials for immediate and extended blood contact [23], they appear to be very beneficial in bone related treatments. Doyle [24] studied PHB composites reinforced with hydroxyapatite and found that bone is rapidly formed and becomes highly organized as new bone tissue. There was no chronic inflammation and also no evidence of PHB degradation was found over a period of 12 months. As such, bacterial polyesters may be candidates for long-term implants. Yagmurlu [25] reported a positive effect on treating osteomyelitis with antibiotics administered by PHBV drug delivery vehicles. The advantageous influence of PHB and PHBV in orthopedic applications is believed to be based on the material’s piezoelectric properties, which promote osteoconductivity and are claimed to induce bone reformation in load-bearing sites [26]. Chen’s research focuses on the use of bacterial polyesters as biomedical materials [27, 28].

5. Other Polyesters The hydrolytic instability of the ester bond has made polyesters a prime candidate for biomedical implants. PLLA, PGA and copolymers of PGA and PLLA or PDLA have advanced the furthest with an extensive list of commercial products. Using similar ringopening polymerization techniques as discussed for PLA and PGA, Dioxanone and εcaprolactone have been polymerized yielding polymers that have received attention in the biomedical field, see Figure 12. Polydioxanone has been investigated for orthopedic applications as fixation screws, and also as monofilament sutures. It is a highly crystalline

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polymer (no side groups) with a rather low modulus of ~ 1.5 GPa and a low Tg of -10 - 0°C [16].

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Figure 13. Schematic depiction of a poly(urethane), with R typically being either a polyester or polyether.

Poly-ε-caprolactone, PCL, is also obtained by ring-opening polymerization and is characterized by an even lower Tg than polydioxanone, -60°C, which is the result of the polymer chain having no side groups and a comparatively large number of methylene (-CH2-) groups, in other words with the extension of the methylene segments the physical properties approximate those of poly(ethylene) [-(-CH2-CH2-)n-]. The ester functionality guarantees however the hydrolytic degradability, even though the degradation is rather slow with 2 to 3 years. PCL is highly biocompatible, its physical properties, low tensile strength of 23 MPa and extremely high elongation [9] make it less of a candidate for orthopedic application and more a candidate for long-term drug delivery devices [29] and tissue engineering scaffolds, for instance as blend with hyaluronic acid as potential meniscus replacement [30]. The latest developments in degradable polymers for biomedical applications seem to shift from polyesters to polycarbonates. Kohn and coworkers have developed tyrosine-derived polycarbonates [31]. Copolymers of the slow degrading poly(desaminotyrosyl-tyrosine ethyl ester carbonate) and the fast degrading poly(desaminotyrosyl-tyrosine ethyl ester carbonateco-desaminotyrosyl-tyrosine) have been fabricated into fibers by electrospinning and used in tissue engineering [32].

6. Polyurethanes Recently, polyurethanes, Tg ~ -40°C, have gained some consideration as biomedical materials. The polyurethane linkages (also called carbamate) see Figure 13 can be considered as a hybrid of an ester- and an amide-bond. Polyurethanes are produced by a steppolymerization process: A polymeric polyester- or polyether diol or diamine (Mw ~ 2000), which constitutes the soft segment, reacts with a diisocyanate to yield the diisocyantae macromer. If a polymeric diamine is used as soft segment a polyurea macromer will be formed. Nevertheless, manufacturers typically refer to the final product also as polyurethanes. Aromatic diisocyantaes, which are typically used in polyurethane syntheses of structural materials or flexible and rigid foams, have also been used for the synthesis of biomedical materials, but they are rather toxic. Thus, diisocyanobutane and hexamethylene diisocyanate have been developed specifically for polyurethanes in biomedical applications. The thus formed diisocyanate macromers are subsequently linked by reactions with chain extenders, which are small molecule diols or diamines. The chain extenders together with the diisocyanate groups form the “hard segments” in the final polyurethane materials. Depending

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on whether the chain extender is either a diol or a diamine, another urethane linkage (reaction of an isocyanate and a hydroxyl group) or a urea linkage (reaction of an isocyanate with an amine group) is formed. Based on the complexity of the polyurethane chemistry, see Table 1, a multitude of polyurethanes can be produced owing to the variety in starting materials. Hence, mechanical properties can be tailored to application specifications. The situation is even more complex, in that, isocyanate groups can induce crosslinking based on the formation of allophanate groups or due to trimerization of the isocyante, the chemistry of which will however, exceed the scope of this chapter. The biodegradability of polyurethanes is determined by the nature of the soft segment. Only poly(ester urethane)s are biodegradable; the urethane and urea linkage are hydrolytically stable and do not degrade. The polyester segment in these polymers typically consists of wellestablished biodegradable polyesters, such as poly(lactide-co-glycolide) or polycaprolactone. Poly(ether-urethane)s have been investigated as biostable implant materials, for instance as artificial blood vessels and artificial blood vessel coatings. The flexibility of the material makes it ideal for tubing, gaskets and catheters. Degradation of polyurethanes has however, been observed and it was shown to be due to mechanical failures. Polyurethanes are prone to stress cracking. Additional annealing processes can alleviate this problem [33]. Oxidative enzymes present in blood have been determined to induce and catalyze the crack formation on polyurethane surfaces [34].

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BIODEGRADATION The degradation of a material is determined by its chemical and physical structure. In order to be hydrolytically degradable, the material must contain hydrolytically labile linkages. The ester linkage in polyesters is susceptible to react with water, thus progressively cleaving the main chain. Once the cleavage products, which are terminated by carboxyl and hydroxyl groups, have reached a sufficiently small molecular weight, they become water-soluble and get secreted. Hydrolysis is pH-dependent and is catalyzed by bases as well as acids. In the human body the pH is however fixed at the almost neutral value of 7.4, thus catalysis, as it is common in ester chemistry does not apply to biomedical implants. Esterases, responsible for the degradation of fats, which are glycerol esters of fatty acids, do not participate in the degradation of macromolecular esters, i.e. polyesters. Polyamides, polycarbonates, polyureas and polyurethanes have hydrolytically sensitive linkages but degrade very slowly. The slow degradation rate combined with their physical properties make those materials biostable, and they are used in long-term, permanent implants. Polyanhydrides and poly(ortho esters) on the other hand are very hydrolytically instable and degrade within the timescale of hours, they are used as drug delivery systems, but cannot be used as implant materials.

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Table 1. Summary of reactions that yield polyurethanes Starting material

Diisocyanate + poly(diol) 2 OCN – R – NCO + HO-XXX-OH XXX = poly(ether diol) or poly(ester diol)

Macromer obtained from the reaction of the starting materials.

OCN–R–NH-C(O)–O-XXX-OC(O)–NH-R-NCO Poly(ester urethane) or poly(ether urethane) macromer, depending on the nature of polymeric segment XXX The urethane linkage is underlined.

Diisocyanate + poly(diamine) 2 OCN – R – NCO + NH2-XXX-NH2

OCN–R–NH-C(O)-NH-XXXNH-C(O)–NH-R-NCO

XXX = poly(ether diamine) or poly(ester diamine)

Poly(ester urea) or poly(ether urea) macromer, depending on the nature of polymeric segment XXX. The urea linkage is underlined.

Reaction of macromer with chain extender yields the polymer. Macromer + HO – R’ – OH or NH2 –R” – NH2 Poly(urethane) -(-R’-O-C(O)-NH-R-NH-C(O)-OXXX-O-C(O)-NH-R-NH-C(O)-O-)nPoly(urea urethane) -(-R”-NH-C(O)-NH-R-NH-C(O)-OXXX-O-C(O)-NH-R-NH-C(O)-NH-)n-

Poly(urea urethane) -(-R’-O-C(O)-NH-R-NH-C(O)-NHXXX-NH-C(O)-NH-R-NH-C(O)-O-)n Poly(urea) -(-R”-NH-C(O)-NH-R-NH-C(O)-NHXXX-NH-C(O)-NH-R-NH-C(O)-NH)n-

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All polyurethane syntheses consist of a two-step process: first, the synthesis of a diisocyanate macromer from a diiosocyante and a polydiol or polydiamine and second, the reaction of the diisocyantae macromers with chain extenders. The latter are small molecules, either diamines or diols.

The presence of ester bonds linking the repeat units of a polymer backbone is not yet sufficient to make a material biodegradable. Poly(ethylene terephthalate), PET, is also a polyester, but it does not degrade. It is mainly used as textile fibers and soda bottles. When considering biodegradability or hydrolytic stability the chemistry of the entire molecule must be considered. In PET ester bonds link ethyl terephthalate units, which are highly hydrophobic and induce a high level of crystallinity. The aromatic nature of the terephthalic acid prevents interaction of the polymer with water on a molecular level, that is, the water molecules are unable to enter into the material and actually perform a nucleophilic attack on the ester bond as illustrated in Figure 4. Hence the presence of a hydrolytically labile bond is no guarantee that the material undergoes hydrolysis, only aliphatic polyesters undergo biodegradation. The impact of the hydrophilicity on the degradation rate is well demonstrated when comparing the degradation rates for PLA and PGA. One additional methyl group in the side chain of PLA, is the only difference between the two polymers, see Figures 10 and 11. This methyl group makes PLA more hydrophobic, which results in a slower degradation. Middleton and Tipton have provided an excellent summary of the application of PLA and PGA in orthopedic devices [35], and show that orthopedic implants made from semi-

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crystalline PLLA, Figure 10, take more than two years to degrade, while the more hydrophilic PGA, Figure 11, degrades within 6 to 12 months. As alluded to earlier the crystal structure of polymers must also be considered. Highly crystalline polymers, in which the polymer chains are organized in lamellae withstand hydrolysis much better than amorphous materials. Water molecules cannot penetrate readily the lamellae, and degradation proceeds solely from the surface of the crystalline regions. There is no organized arrangement of polymer chains in amorphous polymers and chains exist as random coils, where water molecules can penetrate easily into this loose, disorganized structure, starting hydrolysis reactions at several ester bonds simultaneously. Using again PLA as example, it can be shown that changing the crystallinity by going from the semi-crystalline PLLA to the racemic PDLLA reduces the degradation time from more than 24 months to 12 to 16 months [35]. This difference in degradation time is the sole result of the change of a physical parameter (crystallinity), the chemistry has not been altered. The physical structure of a polymer can be determined by the processing conditions. Polymers processed out of the melt, as it is typical for polyesters, form crystallites when allowed to undergo a slow cooling process, thus yielding a high degree of crystallinity. If this same polymer-melt is quenched, the polymer chains are frozen in the random coil configuration, thus yielding a completely amorphous material. The combination of a chemically hydrophobic material with a high degree of crystallinity, as it is the case in PET, yields a material that does not undergo hydrolysis despite the fact that it is a polyester. Moreover, the size and shape, surface chemistry and surface energy, lubricity and also the molecular weight determine the rate of depolymerization. Special emphasis should always be placed on the implant surface as this forms the interface with the surrounding tissue [5]. Finally the mode of degradation needs to be considered. Implant materials undergo either surface or bulk degradation. Materials that undergo surface modification show pit-formation. Electron-microscopic studies of the surfaces of these materials show in the initial stage of the degradation the formation of indentations or pits on the surface. Degradation has proceeded preferably in the amorphous regions, degrading all the amorphous material and leaving only crystallites behind. Measuring the crystallinity of the material at this point shows actually a higher degree of crystallinity than that of the starting material. This is an apparent increase in crystallinity and actually the result of the removal of the amorphous region. Materials that undergo bulk erosion degrade simultaneously all over the cross section of the material [36]. Bacterial polyesters and polycarbonates undergo surface erosion, most other polyesters undergo bulk degradation.

REFERENCES [1] [2] [3] [4]

Staudinger, H. “Polymerization” Ber. dt. Chem. Gesellsch. 1920, 53(6) 1073-1085 Chemical and Engineering News 2005, 83(28), 74 Anderson, A.J., Dawes, E.A. “Occurrence, Metabolism, Metabolic Role, and Industrial Uses of Bacterial Polyhydroxyalkanoates” Microbiol. Rev. 1990, 54, 450 Madison, L.L., Huisman, G.W. “Metabolic Engineering of poly(3-hydroxyalkanoates): From DNA to Plastic” Microbiol. Molec. Biol. Rev. 1999, 63, 21

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The Molecular Structure of Degradable Polymers [5]

Castner, D.G., Ratner, B.D. “Biomedical Surface Science: Foundations to Frontiers” Surface Sci. 2002, 500, 28-60

[6] [7]

www.natureworksllc.com

[8]

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[11] [12] [13] [14] [15] [16] [17]

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[18]

[19]

[20] [21] [22] [23]

Hartmann, M.H. in: “High Molecular Weight Polylactic Acid Polymers” Biopolymers from Renewable Resources (ed. D.L. Kaplan), Springer Verlag, Berlin, Heidelberg, New York, 1998 Du, Y.J., Lemstra, P.J., Nijenhuis, A.J., van Aert, H.A.M., Bastiaansen, C. “ABA Type Copolymers of Lactide with Poly(ethylene glycol). Kinetic, Mechanistic, and Model Study” Macromolecules, 1995, 28, 2124 Polymer Data Handbook, Oxford University Press, New York, Oxford 1999 Ajioka, M., Enomoto, K., Suzuki, K., Yamaguchi, A. “The basic prpoperties of poly(lactic acid) produced by the direct condensation polymerization of lactic acid” J. Polym. Environm. 1995, 3(4) 225 Ohta, M., Obuchi, S. Yoshida, Y. 1995 US Patent 5,444,143 Ichikawa, F., Kobayashi, M., Ohta, M., Yoshida, Y., Obuchi, S., Itoh, H. 1995 US Patent 5,440,008 Enomoto, K., Ajioka, M., Yamaguchi, A. 1994, US Patent 5,310,865 Nair, L.S. Laurencin, C.T. “Biodegradable Polymers as Biomaterials” Progr. Polym. Sci, 2007, 32(8-9), 762 Middleton, J.C., Tipton, A.J. Synthetic Biodegradable Polymers as Medical Devices. www.devicelink.com/mpb/archive/98/03/002.html Barber, A.F., Boothby, M.H., Richards, DP, “New sutures and suture anchors in sports medicine” Sports Med. Arthrosc. 2006, 14(3) 177 Tiainen, J., Veiranto, M., Suokas, E., Tormala, P., Waris, T., Ninkoviv, M., Ashammakhi, N. “Bioabsorbable ciprofloxacin-containing and plain self reinforced poly(lacitide-polyglycolide 80/20 screws: pullout strength properties in human cadaver parietal bones” J. Craniofac. Surg. 2002, 13, 427 Nair, L.S., Laurencin, C.T. “Polymers as biomaterials for tissue engineering and controlled drug delivery” (ed.: D. Kaplan, K. Lee) Springer Verlag Review Series, Adv. Biochem. Eng. Biotechnol. 2006, 47 Laurencin, C.T., Ambrosio, A.M.A., Attawia, M.A., Ko, F.K., Borden, M.D. “In Vitro Cell Adhesion and Proliferation on Novel Bioresorbable Matrices for Use in Bone Regeneration Applications” in Polymers from Renewable Resources – Biopolyesters and Biocatalysis (ed.: C. Scholz, R.A. Gross) ACS Symposium Series 764, Washington 2000 Shive, M.S., Anderson, J.M. “Biodegradation and biocompatibility of PLA and PLGA” Adv. Drug Deliv. Rev. 1997, 28(1) 5 Jain, R.A. “The manufacturing techniques of various drug loaded biodegradable poly(lactide-co-glycolide) (PLGA) devices” Biomaterials 2000, 21(23) 2475 Holmes, P.A. in: Developments in Crystalline Polymers, (ed.: D.C. Basset), Elsevier, New York 1988, 2, p. 1 van der Giessen, W.J., Lincoff, A.M., Schwartz, R.S., van Beusekom, H.M., Serruys, P.W., Holmes, D.R., Jr., Ellis, S.G., Topol, E.J. “Marked inflammatory sequelae to implantion of biodegradable and nonbiodegradable polymers in porcine coronary arteries” Circulation, 1996, 94, 1690

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[24] Doyle, C., Tanner, E.T., Bonfield, W. “In vitro and in vivo evaluation of polyhydroxybutyrate and of polyhydroxybutyrate reinforced with hydroxyapatite” Biomaterials 1991, 12, 841 [25] Yagmurlu, M.F., Korkusuz, F. Gürsel, I. Korkusuz, P., Örs, Ü., Hasirci, V. “Sulbactam-cefoperazone polyhydroxybutyrate-co-hydroxyvalerate (PHBV) local antibiotic delivery system: In vivo effectiveness and biocompatibility in the treatment of implant-related experimental osteomyelitis” Biomed. Mater. Res. 1999, 46, 494 [26] Pouton, C.W., Akhtar, S. “Biosynthetic polyhydroxyalkanoates and their potential in drug delivery” Adv. Drug Del. Rev. 1996, 18, 133 [27] Chen, G-Q., Wu, Q. “The application of polyhydroxyalkanoates as tissue engineering materials” Biomaterials 2005, 26(33) 6565 [28] Deng, Y., Lin, X-S., Zheng, Z., Deng, J-G., Chen, J-C., Ma, H., Chen, G-Q. “Poly(hydroxybutyrate-co-hydroxyhexanoate) promoted production of extracellular matrix of articular cartilage chondrocytes in vitro” Biomaterials, 2003, 24, 4273 [29] Sinha, V.R., Bansal, K., Kaushik, K., Kumria, R., Trehan, A., “Poly-e-caprolactone microspheres and nanospheres: an overview” Int. J. Pharm. 2004, 278, 1 [30] Chiari, C., Koller, U., Dorotka, R., Eder, C., Plasenzotti, R., Lang, S., Ambrosio, L., Tognana, E., Kon, E., Salter, D., Nehrer, S. “A tissue engineering approach to meniscus regeneration in a sheep model” Osteoarthritis and Cartilage 2006, 14, 1056 [31] www.medicaldevice-network.com/features168.html [32] Meechaisue, C., Dubin, R., Supaphol, P., Hoven, V.P., Kohn, J. “Electrospun mat of tyrosine-derived polycarbonate fibers for potential use as tissue scaffolding material” J. Biomater. Sci. Polym. Ed.. 2006, 17(9), 1039 [33] Wintermantel, E. and Ho, S.W. “Medizintechnik mit biokompatiblen Werkstoffen und Verfahren” , 3. Auflage, Chapter 11. Polymere, Springer Verlag, Berlin, Heidelberg, New York, 2002 [34] Sutherland, K., Mahoney, J.R., Coury, A.J., Eaton, J.W., “Degradation of biomaterials by phagocyte-derived oxidants” J. Clin. Invest. 1993, 92, 2360 [35] Middleton, J.C., Tipton, A.J. “Synthetic biodegradable polymers as orthopedic devices” Biomaterials 2000, 21, 2335 [36] Vert, M. “Aliphatic Polyesters: Great Degradable Polymers That Cannot Do Everything” Biomacromolec. 2005, 6, 538

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In: Degradable Polymers for Skeletal Implants Editors: P.I.J.M. Wuisman and T. H. Smit

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Chapter 2

TIME-DEPENDENT FAILURE IN LOAD-BEARING POLYMERS. A POTENTIAL HAZARD IN STRUCTURAL APPLICATIONS OF POLYLACTIDES Leon E. Govaert1, Tom A.P. Engels1, Serge H.M. Söntjens1, and Theo H. Smit2 1

Eindhoven University of Technology, Mechanical Engineering, Polymer Technology, The Netherlands 2 VU Medical Centre, Amsterdam and Skeletal Tissue Engineering Group Amsterdam, The Netherlands

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ABSTRACT Polylactides are commonly praised for their excellent mechanical properties (e.g. a high modulus and yield strength). In combination with their bioresorbability and biocompatibility, they are considered prime candidates for application in load-bearing biomedical implants. Unfortunately, however, their long-term performance under static load is far from impressive. In a previous in vivo study on degradable polylactide spinal cages in a goat model it was observed that, although short-term mechanical and real-time degradation experiments predicted otherwise, the implants failed prematurely under the specified loads. In this chapter we demonstrate that this premature failure is attributed to the time-dependent character of the material used. The phenomenon is common to all polymers, and finds its origin in stressactivated segmental molecular mobility leading to a steady rate of plastic flow. The main conclusion is that knowledge of the instantaneous strength of a polymeric material is insufficient to predict its long-term performance.

INTRODUCTION When skeletal structures or tissues fail due to trauma or disease, additional support is required to take over their mechanical function. For example, in spinal diseases leading to degeneration, instability and/or severe deformations, fusion of one or more spinal segments

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Leon E. Govaert, Tom A.P. Engels, Serge H.M. Söntjens et al.

may be indicated. Devices used for this purpose not only should maintain or restore the spinal anatomy, but also create the proper mechanical environment for bony fusion. As bones and implants must resist considerable loads, metal is a popular material for this purpose. Although metal implants are routinely used and quite successful, there also are some drawbacks. In the first place: metals (and non-degradable polymers alike) are permanent materials and as such remain susceptible to long-term complications like wear [1], migration [2], and late foreign body reactions [3;4]. For this reason, some surgeons recommend to remove all metallic implants used for fixation [5]. Worldwide, the removal of the metallic hardware varies from a routine procedure in all patients (Germany, Australia), to selective removal only from patients with symptoms (e.g. Netherlands) [6]. In the USA, retrieval surgeries for the spine were reported in 25-40% of the patients [7-9]. A second disadvantage of metal implants is that they eclipse the fusion zone on radiological imaging. It is in fact impossible to determine whether healing has been achieved within a metal implant or not. In case of persistent back pain, this is a complicating factor to determine surgical success or failure. Finally, metal implants cause stress shielding over the fusion area, resulting in delayed unions. Obviously, these are undesired properties for fixation implants aiming at healing or fusion. Skeletal fixation devices essentially have a temporary function: once healing is achieved, removal is desired both from the clinical and biomechanical point of view. Indeed, living bone itself is the optimal mechanical support, because living bone is a self-optimizing tissue through mechanical adaptations [10-12]. Furthermore, with permanent implants there always will remain a possible risk for long-term complications. These considerations motivated the development of degradable polymer implants [13-16], which have evident advantages over metal devices: their stiffness is comparable to that of bone; they do not interfere with radiography, computer tomography, or magnetic resonance imaging [17]; and they degrade over time and thus eliminate the necessity of retrieval surgeries. In addition, the healing process may be stimulated by the successive loss of their mechanical properties, thereby gradually increasing the loads on the healing tissues. This concept of degradable polymers for skeletal tissue regeneration has amply been described, but its practical implementation remains challenging, in particular for load-bearing implants as used in trauma or spine. Although degradable polymers are interesting materials for surgical implants, there are some caveats. First, polymer degradation can cause a severe host tissue responses: late complications like osteolytic reactions have been reported with the use of different polyester implants [18-24]. Degradation and intensity of the inflammatory response are influenced by implant related factors (polymer type, purity, crystallinity, design, processing techniques), and environment related factors (implantation site, vascularization, micro-motion, dynamic loading) [25-27]. New implants thus should be carefully evaluated at their specific anatomic location before clinical introduction. Mechanical strength is a second concern of degradable polymers: the skeleton -in particular the spine and long bones- is subject to relatively large amplitudes of dynamic loading. Polymers not only have limited strength as compared to metals, but they also appear to degrade faster under such conditions [25;28]. Early loss of mechanical strength of the implants destabilizes the spinal segment, thus causing non-unions and clinical failure. Another observation in pre-clinical studies is the plastic deformation of implants made of 70/30 PLDLLA [29]. Cages which apparently had sufficient strength for bearing spinal loads during at least eight months appeared to have been broken and deformed after only three

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months (Figure 1). It appeared in additional studies, that the mechanical strength of 70/30 PLDLLA was lower for lower loading rates, higher temperature, and higher humidity [30]. Thus, 70/30 PLDLLA appears to show strong time-and load-dependent behaviour which is typical for glassy polymers [31].

Figure 1. A 70/30 PLDLLA cage after three and six months follow-up in a goat spine. Histology shows micro-cracks already after three months of implantation. Micro-MRI confirms that micro-cracks are formed and also shows some plastic deformation of the cage. After six months, severe deformation of the cage is typically seen, along with failed fusion [29].

This chapter addresses this typical mechanical behaviour of glassy polymers. First, the phenomenology of delayed failure in glassy polymers is discussed. The problem is in fact quite well-known in non-degradable polymers like polycarbonate. Subsequently, the timedependent behaviour of three amorphous polylactides is quantified in a series of short- and long-term loading experiments. It will be shown that the time-dependency of glassy polylactides is so strong, that early failure under static loading conditions can be explained.

DELAYED FAILURE IN POLYMERS Phenomenology Let us first illustrate the problem at hand. In Figure 2a the stress-strain curve of polycarbonate (PC), a ductile glassy polymer, is shown, measured at a constant strain rate of

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Leon E. Govaert, Tom A.P. Engels, Serge H.M. Söntjens et al.

10-2 s-1 in uniaxial extension. Initially the material displays a linear region, where the stress increases proportionally with strain. At higher stresses the response becomes non-linear and subsequently reaches a maximum; the so-called yield stress. This maximum marks the onset of plastic deformation, and soon after the material displays necking, a strain-localization phenomenon [32]. In this process a localized plastic deformation zone is formed that subsequently propagates along the entire length of the specimen whereas the applied force a

b

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Figure 2. Deformation behavior of polycarbonate in uniaxial extension. a) Tensile test at a constant strain-rate of 10-2 s-1. b) Creep test under a constant load of 55 MPa.

remains almost constant. Upon further straining the material breaks at a stress level that depends strongly on molecular weight [33,34]. Please note, from a mechanical point of view, the moment of neck initiation is where the material is regarded to fail as this is the point where it loses its mechanical integrity. We are now confronted with the question whether the information in the stress-strain curve in Figure 2a (short-term), can be used to determine the maximum load level that can be sustained without failure over the lifetime of our product: the design strength. In the case of construction metals, this is usually the yield strength, as static loads below this level will only induce macroscopic elastic deformations. In other words: for metals any static load below the yield strength can be endured almost indefinitely1. Unfortunately, the situation is more complicated in case of polymers. To illustrate this we perform an experiment were we load an identical sample, with the same strain rate, to a stress value of 55 MPa, about 15% below the yield stress. Subsequently this load is kept constant over time. The material’s response to this excitation is shown in Figure 2b. Under the applied static stress the polymer displays a time-dependent mechanical response. The deformation of the sample is observed to increase gradually in time, and after a short while it deforms at a constant rate of strain. At longer loading times the deformation rate increases and the material subsequently fails after being loaded for only 9 hrs. The observed mode of failure is very similar to that observed in the short-term tensile test; necking. For this reason the phenomenon is sometimes referred to as delayed yielding [35,36]. The time-scale on which polycarbonate fails, proves to depend strongly on the level of the applied load. This is illustrated in Figure 3a, which gives the creep response of 1

This does not hold for cyclic loading where fatigue failure may occur.

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polycarbonate for three different stress levels. The higher the stress level, the faster the material deforms and fails. The stress-dependence of the time-to-failure is shown in Figure 3b. A semi-logarithmic relation is observed where a decrease in stress of approximately 3 MPa leads to an increase in lifespan by an order of ten. This brings us to the remarkable insight that, in the case of polymers, it is not the question whether the material will fail under static load, but rather when it will fail under the specified load. a

b

Figure 3. a) Deformation of polycarbonate under various constant stress levels. The markers indicate the point where necking occurs (ductile failure). b) Stress dependence of the time-to-failure, taken from (a).

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Origin Let us now examine the physical background of the observed time-dependent failure process. Amorphous polymers consist of long, covalently-bonded molecules that are randomly distributed throughout the material. Each molecule has the ability to change its spatial conformation by rotation over covalent bonds that form the back-bone of the chain, and in its equilibrium state a random coil will be the most probable conformation. The rate at which a chain can change its conformation depends strongly on temperature. At high temperature conformational changes are fast and the chains can move freely with applied deformation (rubberlike behavior). At low temperature (below the glass transition temperature), chain mobility decreases drastically and the material virtually “vitrifies” [37] (becomes glass-like). Here it is essential to realize that changes in chain conformation are still feasible! However, due to the extremely low mobility of the chains, these conformational amendments are generally not observable within practical experimental time scales. This situation changes drastically upon the application of stress. Similar to temperature, applied stress enhances main-chain mobility and, as a result, the effects of changes in chain conformations become visible. To demonstrate the profound influence of applied stress and temperature on chain mobility we take a closer look at the deformation behavior of a polymer glass around the yield point. In Figure 4a the stress-strain curves of PC are presented for 2 temperatures at various strain rates. The yield stresses determined from these experiments (maxima of the

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curves), are plotted in Figure 4b as a function of the applied strain rate. Typically, the yield stress is observed to increase with increasing strain rate and decreasing temperature [38]. In the initial stage of loading, where the stress is still low, chain mobility will be negligible, and the modulus is determined by the intermolecular interactions between individual chains. With increasing stress, the chain mobility increases and changes in chain conformation gradually start to contribute to the deformation of the material. This contribution progressively increases, until it reaches a stress level where the plastic strain rate resulting from chain mobility exactly matches the one experimentally applied; the yield stress. In other words, applied stress induces a state of enhanced molecular mobility that stimulates a dynamic rearrangement of molecular segments, resulting in a steady rate of plastic flow. The magnitude of this plastic flow rate depends on the applied stress and temperature. For further illustration we consider the deformation of PC under static stress; the creep curves in Figure 3a. From this figure we determine the evolution of strain rate as a function of strain, a so-called Sherby-Dorn plot [39] (Figure 5a).

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a

b

Figure 4. a) Stress-strain curves for polycarbonate at various strain rates (10-4, 10-3 and 10-2 s-1) and temperatures. b) Yield stress as a function of strain rate for various temperatures. a

b

Figure 5. a) Evolution of strain rate during creep at various loads (derived from the data in Figure 2a). b) Stress as a function of strain rate. Tensile experiments (open symbols), creep experiments (closed symbols).

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b

Figure 6. Intrinsic deformation behavior of PC in uniaxial compression (a) at a constant compressive strain rate of 10-2 s-1, (b) under a constant true compressive stress of 64 MPa.

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In this plot we can observe that, at each load, the strain rate initially decreases (primary creep) until it reaches a steady state of flow, where the strain rate remains more or less constant (secondary creep). It was first demonstrated by Bauwens-Crowet et al. [40] that the steady state obtained in static loading is identical to that obtained at the yield stress in a constant strain rate experiment. This is demonstrated in Figure 5b, that presents the steady state values of stress and strain rate obtained from tensile tests at constant strain rate and creep tests under static load. Both yield exactly the same curve. In summary: we now established that applied stress induces a state of enhanced molecular mobility in polymer glasses that results in a steady rate of plastic flow. The question that remains is: how does this lead to strain localization and subsequent failure?

Intrinsic Deformation Behavior To understand the reason for strain localization, we have to study the stress-strain response in an experimental set-up in which a sample can deform homogeneously up to large plastic deformations. Examples of such techniques are uniaxial compression tests [41,42] or videocontrolled tensile tests [43], and the stress-strain curves thus obtained are generally referred to as the intrinsic stress-strain response. An illustrative example is presented in Figure 6a, which shows the intrinsic response of PC in an uniaxial compression test at a constant rate of strain. In contrast to the behavior in uniaxial extension, the sample does not display necking but deforms homogeneously over the entire strain range covered. After the yield point, the intrinsic stress-strain response of polymer glasses displays two characteristic phenomena: strain softening, the initial decrease of true stress with strain, and strain hardening, the subsequent upswing of the true stress-strain curve (Figure 6a). Strain hardening is generally interpreted as the result of a stress contribution of the orienting molecular network [41,42,44]. A similar picture is observed for the response to a constant true compressive stress (Figure 6b). Comparable to figure 1b, the static load induces a steady rate of plastic flow. Plastic deformation accumulates steadily, until, at a critical level of plastic strain, strain softening sets in and the deformation rapidly increases until stabilized by strain hardening [45].

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b

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Figure 7. Intrinsic deformation behavior of PC in uniaxial compression (a) influence of applied strain rate, (b) influence of physical aging.

Regarding the intrinsic mechanical response of glassy polymers, there are two timedependencies that need to be considered [31]. The first one we already encountered: it is the time-dependence of the mechanical properties itself. The influence of strain rate on the intrinsic behavior of polycarbonate is presented in Figure 7a. With increasing strain rate the yield stress increases similarly to the observations in Figure 3a,b. Moreover, an identical effect is observed in the post-yield region, where the curves shift upwards by the same amount as the yield stress. The second time-dependency that is of importance is the influence of the age of the material. Polymer glasses are generally not in thermodynamic equilibrium and, as a result, display a persistent drive towards it (physical aging), leading to a gradual change in mechanical properties over time [46,47]. This process can be accelerated by storing the material at an elevated temperature below its glass transition temperature; a heat-treatment called annealing. In Figure 7b, the intrinsic response of an annealed sample of PC is compared to that of a rapidly cooled sample (quenched). It is clear that physical aging results in an increase of both modulus and yield stress, but upon plastic deformation the differences between the curves disappear and eventually they fully coincide at a strain of approximately 0.2. Apparently all influence of thermal history is erased at that strain and both samples are transformed to a similar, mechanically ‘rejuvenated’ state [31,48]. From Figure 7b it is obvious that an increase of yield stress, due to a thermal treatment, will directly imply an increase in strain softening. The influence of molecular weight on the intrinsic response is usually small and negligible [31], which makes thermal history the key factor in influencing the intrinsic properties of a specific polymer glass. A change in thermal history will also reflect in the long-term failure behavior of polymer glasses [45]. This is demonstrated for PC in Figure 8a,b. Figure 8a presents the yield stress vs. strain rate in uniaxial extension for quenched and annealed PC. As a result of annealing the yield stress increases, but the kinetics (slope of the curve) remain unchanged. As can be witnessed in Figure 8b, the increase in yield stress is accompanied by an improvement in the life-time under constant stress. Depending on the effectiveness of the treatment, the improvement may be by orders of magnitude [45]. The second point of interest is that the slope of applied stress versus the logarithm of time to failure is the same for both thermodynamic states. Moreover,

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the absolute value of this slope is virtually identical to that observed for the yield stress versus the logarithm of strain rate [40,45] (Figure 8a). With respect to the mode of failure, it is in particular the post-yield characteristics of the polymer, i.e. strain softening and strain hardening, that play a determining role [48,49]. In the vicinity of stress concentrations, strain softening will inevitably lead to the formation of localized plastic deformation zones (DZ’s). The initial size and evolution of the DZ is determined by a subtle interplay between the amount of strain softening and the amount of strain hardening. If the strain hardening is sufficiently strong, the DZ’s are stabilized and expand in a controlled fashion to the bulk of the material. Typical examples of such ductile behavior are shear band formation and necking. In the case of insufficient strain hardening, on the other hand, the material will be inclined to deform plastically by crazing, extremely localized zones of plastic deformation that act as a precursor for cracks and thus induce a brittle failure mode [48,49]. To illustrate this we compare the mechanical responses of PC and polystyrene (PS). Most striking here is the difference in macroscopic failure behavior of PC, which shows a ductile failure mode (necking) and PS, which fails in a fully brittle fashion (Figure 9a).

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a

b

Figure 8. a) Yield stress vs. strain rate in uniaxial extension for annealed and quenched PC. b) Time-tofailure vs. applied stress for annealed and quenched PC.

a

b

Figure 9. a) Mode of failure of PC (ductile) and PS (brittle). b) Comparison of the intrinsic deformation behavior of PC and PS. Degradable Polymers for Skeletal Implants, Nova Science Publishers, Incorporated, 2009. ProQuest Ebook Central,

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This can be rationalized by the differences in the intrinsic stress-strain curves of PC and PS presented in Figure 9b. In uniaxial compression, both polymers behave ductile and can be deformed to high compressive strains. Polystyrene exhibits pronounced strain softening and only a weak contribution of the strain hardening. In uniaxial extension, strain localizations cannot be stabilized and evolve almost without limits. As a result this extreme strain localization leads to the initiation of crazes, and, ultimately, macroscopic failure. Polycarbonate, on the other hand, displays only a moderate amount of strain softening and a much stronger contribution of the strain hardening. Localized plastic deformation zones, induced by strain softening, are now stabilized and spread out to other regions in the material. As a result a larger volume participates in the deformation and shear yielding and stable necking are observed. Nevertheless, also polycarbonate will initiate crazes if a more severe localization is introduced by changing the geometry of the test, e.g. by adding a notch [34,50]. As we explained above, the amount of strain softening can be altered by thermal treatments, like annealing. Even small changes in yield stress can have major consequences for the macroscopic deformation behavior. For instance, a subtle increase in strain softening induced by annealing, leads to severe localization of strain and brittle fracture in lowmolecular weight polycarbonate [45, 49]. On the other hand, by removing strain softening through mechanical pre-conditioning [51], PS becomes ductile and can be deformed in uniaxial extension up to strains of 30%. These experiments clearly indicate the dominant role of strain softening in localization and failure of glassy polymers.

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Predictability of Failure So far, we have established that applied stress induces a state of enhanced molecular mobility in polymer glasses that stimulates a dynamic rearrangement of molecular segments, resulting in a steady rate of plastic flow. The deformation can propagate at this steady pace until strain softening sets in, accelerating the rate of deformation and initiating a localized plastic deformation zone; catastrophic failure occurs. A successful prediction of time-dependent failure under static and dynamic excitations can be obtained by application of a constitutive model able to capture the intrinsic deformation characteristics of glassy polymers [45,52] (Figure 7a, b). However, this method requires a substantial investigation of the mechanical performance of the polymer glass. If we limit ourselves to a single loading geometry (uniaxial compression or tension) the approach is rather straight forward. In the physical picture described above, a logical first step towards lifetime prediction is the prediction of the plastic flow rate as a function of applied stress and temperature ε&pl (σ , T ) . In most cases the kinetics can be captured excellently by a simple Eyring flow rule [53]. Subsequently, it is hypothesized that strain softening always sets in at the same critical plastic strain εcr. The lifetime under load tfail can then be estimated by:

t fail =

ε cr ε& pl (σ , T )

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(1)

Time-Dependent Failure in Load-Bearing Polymers a

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b

Figure 10. a) Compressive stress-strain curves of PLLA, PDLLA, and PLDLLA at a strain rate of 10-3s-1. b) Compressive stress-strain curves of PLDLLA at strain rates of 10-4, 10-3 and 10-2 s-1; influence of applied strain rate.

In this straightforward approach the kinetics of failure under static stress are directly linked, and therefore identical, to the kinetics of plastic flow. This is exactly what is observed in practice (see Figure 8a, b), and consequently the approach proved very successful in predicting time-dependent failure of PC and PVC in long-term static loading [54,55]. Equivalent expressions were also successfully applied to predict failure in static and dynamic loading conditions for other polymers [56-58].

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Delayed Failure of Polylactides Materials In this study we set out to capture the deformation behavior of amorphous polylactides. In the field of biomedical resorbable implants used as structural fixation components in the spine, mostly amorphous PLDLLA is used [59-64]. This choice is related to the fact that degradation times of semi-crystalline PLA are longer than needed and that unfavorable inflammatory responses can occur when highly crystalline debris is left after degradation of the amorphous phase [23,65]. Since we use observations on amorphous PLDLLA resorbable spinal cages as a starting point [29,30], this investigation will be restricted to amorphous PLA’s. From the available polylactides, we chose three different materials: a homochiral poly(Llactide) (PLLA) with a glass transition temperature of 58°C, a racemic co-polymer poly(D,Llactide) (PDLLA) with a glass transition temperature of 42°C, and a 70:30 copolymer of Llactide and D,L-lactide (PLDLLA) with a glass transition temperature of 52°C. All materials were studied in the fully amorphous state. For this purpose PLLA was rapidly quenched from the melt to avoid crystallization. The other two are amorphous by nature.

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Mechanical Evaluation The intrinsic mechanical responses of all three materials, measured in uniaxial compression at a constant strain rate, are presented in Figure 10a. The differences between the three materials appear only minor. The main difference is the value of the yield stress; i.e. PDLLA (78 MPa), PLDLLA (71 MPa), and PLLA (81 MPa). Please note, these data were obtained on dry samples, in wet conditions the yield stress will be substantially lower [30]. a

b

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Figure 11. a) Yield stress vs. strain rate in uniaxial compression for PLLA, PLDLLA and PDLLA. b) Time-to-failure vs. applied stress in uniaxial compression for PLLA, PLDLLA and PDLLA.

Influences of annealing treatments proved minimal, which could be related to the fact that, close to Tg, glassy polymers can obtain the equilibrium state with the result that the treatment will induce no further changes [66,67]. The intrinsic responses display a very pronounced strain softening, more than a factor of 2 for PLLA (even more than PS, see Figure 9b), and a rather weak strain hardening. This intrinsic response is in full accordance with the fact that amorphous PLA’s are very brittle in tension or bending, similar to PS fracturing through a crazing mechanism [68,69]. Of course, this also makes the material extremely sensitive for stress concentrations, e.g. the stress-concentration induced by the thread root of a screw will act as a crack initiation site [70]. By plasticization, blending, co-polymerization and/or rubber toughening the toughness of the material can be improve [71-75], but also orientation of the polymer chains, e.g. by cold- and hot drawing, is a route to improvement [69,76]. The strain rate-dependence of PLDLLA is presented in Figure 10b. Remarkable is the rather strong strain-rate dependence of the yield stress; approx. 13 MPa per decade of strain rate. It is also clear that the strain rate dependence in the strain hardening region is much lower; approx. 4 MPa/decade. As a result the amount of strain softening has a quite strong strain-rate dependence which will make the material more brittle at higher strain rates. The strain-rate dependence of the yield stress is presented in figure 11a for all 3 PLA’s. It can be observed that the deformation kinetics (slope of the curve) are identical for all three materials. The curves of the individual materials appear shifted vertically with respect to one another, with PLLA on the top position and PLDLLA at the bottom. The vertical shift between these two extremes is approximately 11 MPa. The lifetime under constant stress is presented in Figure 11b. The picture here compares well to that of the strain rate dependence

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b

Figure 12. a) Yield stress vs. strain rate in uniaxial compression for PLLA and PC. b) Time-to-failure vs. applied stress in uniaxial compresson for PLLA and PC.

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of the yield stress. The three materials show identical kinetics, but appear shifted vertically with respect to one another. Also in this case, the absolute slope of applied stress versus the logarithm of time to failure (Figure 11b) is virtually identical (13 MPa/decade) to that observed for the yield stress versus the logarithm of strain rate (Figure 11a).

Long-Term Performance of PLA’s Vs other Glassy Polymers From a qualitative point of view, the deformation behaviors of the three PLA’s are similar to that of other glassy polymers. From a quantitative point of view, however, there are some major differences. As an illustration, Figure 12 compares the strain-rate dependence of the yield stress and the stress-dependence of the time to failure of PLLA with that of PC. For a standard tensile test (strain rate 10-3 s-1) the yield stress of PLLA is approximately 100 MPa (Figure 12a), considerably higher than that of PC (65 MPa). However, the most striking difference is that in the slopes of yield stress versus the logarithm of strain rate. For an increase in strain rate by a factor of ten the yield stress increases 13 MPa for PLLA, but only 3 MPa for PC. Apparently the plastic flow rate is much more stress-dependent in PC than in PLLA. This difference can also be observed in the slopes of applied stress versus time to failure (Figure 12b), where the same kinetics apply. This implies that an decrease in applied stress by 13 MPa leads to an increase in lifetime by a factor of 10 for PLLA, whereas the same reduction in stress would result in an increase of over a factor of 10.000 for PC. Despite the excellent short term properties of PLLA, its lifetime under a static stress of 50 MPa (50% of the short term strength) is only little more than a single day. In comparison, the lifetime of PC under the same load of 50 MPa (76% of the short term strength) is over 3 months! This is actually the expected lifetime of PLLA at a static load of only 25% of its short term strength. From a quantitative point of view it is quite clear that the long-term performance of PLLA is worrying. Taking into account that the example given here relates to its performance under dry conditions, at room temperature, the picture becomes even more disturbing.

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b

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Figure 13. a) The maximum load of the PLDLLA cage as a function of loading speed (displacement rate), temperature, and humidity. b) Time-to-failure for dry cages at 37ºC. The dotted line is a prediction for the time-to-failure of the wetted cages, based on a decrease by 0.5 kN as determined in Figure 13a.

To illustrate this, Figure 13 presents some mechanical data on the PLDLLA spinal cage which displayed premature failure in a goat study [29]. Figure 13a shows the dependence of cage strength as a function of loading speed, temperature and humidity [30]. The dependence on loading speed presents a similar picture to the data presented in Figure 12a. In general, the cage appears to be stronger at higher loading velocity, lower temperature, and under dry conditions. An increase in loading velocity by a factor 10 increases the cage strength by 0.86 kN. Taking into account the cross-sectional area of the spinal cage (70 mm2), this would imply an increase of yield stress of 12.3 MPa/decade, which compares well to the slope observed in Figures 11a and 12a. The cage strength, determined according to ASTM D695 at a loading rate of 1.3 mm.min-1, is 7.1 kN. An increase in temperature from 23ºC (room temperature) to 37ºC (body temperature) substantially decreases the cage strength by 1.5 kN. An extra decrease in strength of 0.5 kN is observed when the cage is tested after being soaked in water for 24 hrs. This influence of temperature and humidity is equal for all loading velocities, i.e. the slope of the curve remains unchanged. Figure 13b presents the time-to-failure for dry cages at 37ºC under various compressive forces below its instantaneous strength. Also here the general observation corresponds with that presented in Figures 11b and 12b. Time-to-failure was generally short and strongly dependent on the applied load. The long term performance of wet cages was estimated here by shifting the fit through the “dry” data by 0.5 kN, as determined from the compression experiments in Figure 13a. Under a compression load of 50% of the short-term strength (ASTM D695), wet cages are expected to collapse after less than one hour of loading. At 25% of the cage strength the lifetime is approximately one month.

Influence of Molecular Degradation In the case of degradable polymers, like the PLA’s studied here, the molecular weight steadily decreases over the course of time. It seems justified to question the possible influence of degradation on the failure kinetics. An important indication for a relation between

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molecular weight and mechanical performance is the reduction in cage strength that is observed during in-vitro degradation of ETO and E-beam sterilized PLDLLA spinal cages [77]. Similar results were found in a recent study on the influence of molecular weight on the mechanical performance of PDLLA [78]. Nevertheless, combined modeling of molecular degradation and time-dependent failure seems to indicate only a very minor acceleration of failure at very low stress levels. At higher stresses time-dependent failure will occur before there is sufficient degradation to influence flow kinetics [78].

CONCLUDING REMARKS

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We have established that glassy polymers can best be regarded as highly viscous fluids. Upon application of stress a state of enhanced molecular mobility is induced, that stimulates a dynamic rearrangement of molecular segments, resulting in a steady rate of plastic flow. This deformation can propagate at this steady pace until strain softening sets in, accelerating the rate of deformation initiating a localized plastic deformation zone; failure occurs. This implies that under static load the implant will always fail, the main question being when. This specific type of failure mechanism clearly requires a new approach to the design of load-bearing polymer constructs. It is clearly demonstrated in Figure 13, that, even if we consider only a single loading geometry, it is difficult to speak of the strength of an implant. With predictive models becoming available, the long-term performance of a degradable construct can in principle be estimated in early stages of design. Since mechanical testing protocols for interbody devices, like ASTM-F2077, currently do not consider long-term failure under static loading, it seems most appropriate to reconsider the standardized testing protocols for load-bearing constructs when these are made of polymers like amorphous PLA’s.

REFERENCES [1] [2]

[3]

[4] [5] [6]

Togawa D, Bauer TW, Brantigan JW et al. Bone graft incorporation in radiographically successful human intervertebral body fusion cages. Spine 2001;26:2744-50. Taneichi H, Suda K, Kajino T et al. Unilateral transforaminal lumbar interbody fusion and bilateral anterior-column fixation with two Brantigan I/F cages per level: clinical outcomes during a minimum 2-year follow-up period. J Neurosurg.Spine 2006;4:198205. Togawa D, Bauer TW, Lieberman IH et al. Lumbar intervertebral body fusion cages: histological evaluation of clinically failed cages retrieved from humans. J.Bone Joint Surg.Am. 2004;86-A:70-9. Ohlin A, Karlsson M, Duppe H et al. Complications after transpedicular stabilization of the spine. A survivorship analysis of 163 cases. Spine 1994;19:2774-9. Mueller M, Allgoewer M, Schneider R et al. Manual of internal fixation. Techniques recommended by the AO Group. 2nd ed. New York: Springer, 1979. Chapman M, Woo S. Principles of Fracture Healing. In: Chapman M, Madison M, eds. Operative Orthopaedics. Philadelphia: JB Lippincott Company, 1988:115-23.

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36 [7]

[8] [9]

[10] [11] [12] [13]

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[17]

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Leon E. Govaert, Tom A.P. Engels, Serge H.M. Söntjens et al. Brantigan JW, Steffee AD, Lewis ML et al. Lumbar interbody fusion using the Brantigan I/F cage for PLIF and the VSP pedicle screw system: two year results of a FDA IDE clinical trial. In: Husson J, LeHuec J, eds. In Intersomatique du Rachis Lumbaire. Montpellier: Sauramps Medical, 1996. Ray CD. Threaded fusion cages for lumbar interbody fusions. An economic comparison with 360 degrees fusions. Spine1997;22:681-5. Bjarke CF, Stender HE, Laursen M et al. Long-term functional outcome of pedicle screw instrumentation as a support for posterolateral spinal fusion: randomized clinical study with a 5-year follow-up. Spine 2002;27:1269-77. Smit TH, Odgaard A, Schneider E. Structure and function of vertebral trabecular bone. Spine 1997;22:2823-33. Huiskes R, Ruimerman R, van Lenthe GH et al. Effects of mechanical forces on maintenance and adaptation of form in trabecular bone. Nature 2000;405:704-6. Smit TH, Muller R, van DM et al. Changes in bone architecture during spinal fusion: three years follow-up and the role of cage stiffness. Spine 2003;28:1802-8. van Dijk M, Smit TH, Burger EH et al. Bioabsorbable poly-L-lactic acid cages for lumbar interbody fusion: three-year follow-up radiographic, histologic, and histomorphometric analysis in goats. Spine 2002;27:2706-14. Wuisman PI, van DM, Smit TH. Resorbable cages for spinal fusion: an experimental goat model. J.Neurosurg. 2002;97:433-9. van Dijk M, Tunc DC, Smit TH et al. In vitro and in vivo degradation of bioabsorbable PLLA spinal fusion cages. J.Biomed.Mater.Res. 2002;63:752-9. van Dijk M, Smit TH, Arnoe MF et al. The use of poly-L-lactic acid in lumbar interbody cages: design and biomechanical evaluation in vitro. Eur.Spine J. 2003;12:34-40. Cordewener FW, Bos RR, Rozema FR et al. Poly(L-lactide) implants for repair of human orbital floor defects: clinical and magnetic resonance imaging evaluation of long-term results. J.Oral Maxillofac.Surg. 1996;54:9-13. Rokkanen PU, Bostman O, Hirvensalo E et al. Bioabsorbable fixation in orthopaedic surgery and traumatology. Biomaterials 2000;21:2607-13. Partio EK, Bostman O, Hirvensalo E et al. Self-reinforced absorbable screws in the fixation of displaced ankle fractures: a prospective clinical study of 152 patients. J Orthop.Trauma 1992;6:209-15. Bergsma EJ, Rozema FR, Bos RR et al. Foreign body reactions to resorbable poly(Llactide) bone plates and screws used for the fixation of unstable zygomatic fractures. J Oral Maxillofac Surg 1993;51:666-70. Bostman OM. Osteoarthritis of the ankle after foreign-body reaction to absorbable pins and screws: a three- to nine-year follow-up study. J Bone Joint Surg.Br. 1998;80:333-8. Bostman OM, Pihlajamaki HK. Late foreign-body reaction to an intraosseous bioabsorbable polylactic acid screw. A case report. J Bone Joint Surg.Am. 1998;80:1791-4. Bergsma JE, de Bruijn WC, Rozema FR et al. Late degradation tissue response to poly(L-lactide) bone plates and screws. Biomaterials 1995;16:25-31. Bostman OM, Pihlajamaki HK. Adverse tissue reactions to bioabsorbable fixation devices. Clin.Orthop.Relat Res. 2000;216-27.

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37

[25] Middleton JC, Tipton AJ. Synthetic biodegradable polymers as orthopedic devices. Biomaterials 2000;21:2335-46. [26] Bostman O, Pihlajamaki H. Clinical biocompatibility of biodegradable orthopaedic implants for internal fixation: a review. Biomaterials 2000;21:2615-21. [27] Vert M. Poly(lactic acid)s. In: Wnek GE, Bowlin GL, eds. Encyclopedia of Biomaterials and Biomedical Engineering. New York: Marcel Dekker, inc., 2004:125464. [28] Athanasiou KA, Agrawal CM, Barber FA et al. Orthopaedic applications for PLA-PGA biodegradable polymers. Arthroscopy 1998;14:726-37. [29] Krijnen MR, Mullender MG, Smit TH et al. Radiographic, histologic, and chemical evaluation of bioresorbable 70/30 poly-L-lactide-CO-D, L-lactide interbody fusion cages in a goat model. Spine 2006;31:1559-67. [30] Smit TH, Engels TAP, Wuisman PI et al. Time-dependent mechanical strength of 70/30 Poly(L, DL-lactide): shedding light on the premature failure of degradable spinal cages. Spine 2008;33:14-8. [31] Klompen ETJ, Engels TAP, Govaert LE, Meijer HEH, Modeling of the postyield response of glassy polymers: influence of thermomechanical history. Macromolecules 2005;38:6997-7008. [32] Vincent PI, Necking and cold drawing, Polymer 1960:1:7-19. [33] Flory PJ, Tensile strength in relation to molecular weight of high polymers, J. Amer. Chem. Soc. 1945:67:2048-50. [34] Vincent PI, The tough-brittle transition in thermoplastics, Polymer 1960:1:425-44. [35] Matz DJ, Guldemond WG, Cooper SL, Delayed yielding in glassy polymers, J. Polym. Sci., Polym. Phys. Ed. 1972:10:1917-30. [36] Narisawa I, Ishikawa M, Ogawa H, Delayed yielding of polycarbonate under constant load, J. Polym. Sci., Polym. Phys. Ed. 1978:16:1459-70. [37] McKenna GB, Glass formation and glassy behavior, In Comprehensive Polymer Science, Vol. 2:Polymer Properties; Booth C, Price C, Eds., Pergamon: Oxford, 1989; Ch. 10, pp 311–362. [38] Bauwens-Crowet C, Bauwens J-C, Homès G, Tensile yield-stress behavior of glassy polymers, J. Polym. Sci., A-2 1969:7:735-42. [39] Sherby OD, Dorn JE, Anelastic creep of PMMA, J. Mech. Phys. Solids 1954:6:145-62. [40] Bauwens-Crowet C, Ots J-M, Bauwens J-C, The strain-rate and temperature dependence of yield of polycarbonate in tension, tensile creep and impact tests. J. Mater. Sci. Let. 1974:9:1197-1201. [41] Arruda EM, Boyce MC, Evolution of plastic anisotropy in amorphous polymers during finite straining, Int. J. Plast. 1993:9:697-720. [42] van Melick HGH, Govaert LE, Meijer HEH, On the origin of strain hardening in glassy polymers, Polymer 2003:44:2493-2502. [43] G’Sell C, Hiver JM, Daoun A, Souahi A, Video-controlled tensile testing of polymers and metals beyond the necking point, J. Mater. Sci. 1992:27:5031-39. [44] Tervoort TA, Govaert LE, Strain-hardening behavior of polycarbonate in the glassy state, J. Rheol. 2000:44:1263-77. [45] Klompen ETJ, Engels TAP, van Breemen LCA, Schreurs PJG, Govaert LE, Meijer HEH, Quantitative prediction of long-term failure of polycarbonate. Macromolecules 2005;38:7009-17.

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Leon E. Govaert, Tom A.P. Engels, Serge H.M. Söntjens et al.

[46] Struik LCE, Physical aging in amorphous polymers and other materials. Amsterdam: Elsevier, 1978. [47] Hutchinson JM, Physical aging of polymers, Prog. Polym. Sci. 1995:20:703-60. [48] Meijer HEH, Govaert LE, Mechanical performance of polymer systems: the relation between structure and properties, Prog. Polym. Sci. 2005:30:915-938. [49] van Melick HGH, Govaert LE, Meijer HEH, Localisation phenomena in glassy polymers: influence of thermal and mechanical history, Polymer 2003:44:3579-91. [50] Fraser RAW, Ward IM, The impact fracture behaviour of notched specimens of polycarbonate, J. Mater. Sci. 1977:12:459-68. [51] Govaert LE, van Melick HGH, Meijer HEH, Temporary toughening of polystyrene through mechanical pre-conditioning, Polymer 2001:42:1271-4. [52] Janssen RPM, de Kanter D, Govaert LE, Meijer HEH, Fatigue life predictions for glassy polymers: a constitutive approach, Macromolecules 2008:41:2520-30. [53] Eyring H, Viscosity, plasticity, and diffusion as examples of absolute reaction rates, J. Chem. Phys. 1936:4:283-91. [54] Visser HA, Bor TC, Wolters M., Engels TAP, Govaert LE, Life-time assessment of load bearing polymer glasses: Application to polymer pipes under internal, Macromol. Mater. Eng., submitted. [55] Visser HA, Bor TC, Wolters M., Engels TAP, Govaert LE, A new engineering approach to predict the hydrostatic strength of uPVC pipes, Proc. PPS07 EA, Göthenburg, Sweden, 2007. [56] Coleman BD, Application of the theory of absolute reaction rates to the creep failure of polymeric filaments, J. Polym. Sci. 1956:20:447-55. [57] Coleman BD, Knox AG, The interpretation of creep failure in textile fibers as a rate process, Text. Res. J. 1957:27:393-9. [58] Janssen RPM, Govaert LE, Meijer HEH, An analytical method to predict fatigue life of thermoplastics in uniaxial loading: sensitivity to wave type, frequency and stress amplitude, Macromolecules 2008:41:2531-40. [59] Lowe TG, Coe JD. Bioresorbable polymer implants in the unilateral transforaminal lumbar interbody fusion procedure. Orthopedics 2002; 25: S1179-S1183. [60] Lowe TG, Tahernia AD. Unilateral transforaminal posterior lumbar interbody fusion. Clin Orthop 2002; 394: 64-72. [61] Austin RC, Branch CL, Jr., Alexander JT. Novel bioabsorbable interbody fusion spacer-assisted fusion for correction of spinal deformity. Neurosurg Focus 2003; 14: E11 . [62] Kuklo TR, Rosner MK, Polly DW, Jr. Computerized tomography evaluation of a resorbable implant after transforaminal lumbar interbody fusion. Neurosurg Focus 2004;16: E10. [63] Couture DE, Branch CL, Jr. Posterior lumbar interbody fusion with bioabsorbable spacers and local autograft in a series of 27 patients. Neurosurg Focus 2004; 16: E8. [64] Coe JD, Vaccaro AR, Instrumented transforaminal lumbar interbody fusion with bioresorbable polymer implants and iliac crest autograft. Spine 2005; 30: S76-S83. [65] Park MC, Tibone JE, False magnetic resonance imaging persistence of a biodegradable anterior cruciate ligament interference screw with chronic inflammation after 4 years in vivo. Arthroscopy 2006; 22: 911e1-911-e4.

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[66] Engels TAP, Govaert LE, Peters GWM, Meijer HEH, Processing-induced properties in glassy polymers: application of structural relaxation to yield stress development. J. Polym. Sci. B 2006:44:1212-25. [67] Bauwens-Crowet C, Bauwens J-C, Annealing of polycarbonate below the glass transition, Polymer 1982:23:1599-1604. [68] Grijpma DW, Pennings AJ. (Co)polymers of L-lactice,2 Mechanical properties. Macromol. Chem. Phys. 1994: 195: 1649-1663. [69] Grijpma DW, Altpeter H, Bevis MJ, Feijen J. Improvement of the mechanical properties of poly(D,L-lactide) by orientation. Polym. Int. 2002; 51: 845-851. [70] Brkaric M, Baker KC, Israel R, Harding T, Montgomery DM, Herkowitz HN, Early failure of bioabsorbable anterior cervical fusion plates, J. Spinal Disord. Tech. 2007:20:248-54. [71] Jacobsen S, Fritz HG. Plasticizing polylactide – The effect of different plasticizers on the mechanical properties. Polym. Eng. Sci. 1999; 39: 1303-1310. [72] Martin O, Avérous L. Poly(lactic acid): plasticization and properties of biodegradable multiphase systems. Polymer 2001; 42: 6209-6219. [73] Shibata M, Teramoto N, Inoue Y. Mechanical properties , morphologies, and crystallization behavior of plasticized poly(L-lactide)/poly(butylene succinate-co-Llactate) blends. Polymer 2007; 48: 2768-2777. [74] Li Y, Shimizu H. Toughening of polylactide by melt blending with a biodegradable poly(ether)urethane elastomer. Macromol. Biosci. 2007; 7: 921-928. [75] Grijpma DW, Hofslot DA van, Supèr H, Nijenhuis AJ, Pennings AJ. Rubber toughening of poly(lactide) by blending and block copolymerization. Polym. Eng. Sci. 1994; 34:1674-1684. [76] Grijpma DW, Penning JP, Pennings AJ. Chain entanglement, mechanical properties and drawability of poly(lactide). Colloid. Polym. Sci. 1994; 272: 1068-1081. [77] Smit TH, Thomas KA, Hoogendoorn RJW, Strijkers GJ, Helder MN, Wuisman PIJM, Sterilization and strength of 70-30 polylactise cages, Spine 2007:32:742-7. [78] Söntjens SHM, Engels TAP, Smit TH, Govaert LE, Influence of molecular weight on the deformation and failure kinetics of PDLLA constructs, Biomaterials, in preparation.

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In: Degradable Polymers for Skeletal Implants Editors: P.I.J.M. Wuisman and T. H. Smit

ISBN 978-1- 60692-426-6 © 2009 Nova Science Publishers, Inc.

Chapter 3

POLY(LACTIDE)S AND THEIR COPOLYMERS: PHYSICAL PROPERTIES AND HYDROLYTIC DEGRADATION Hideto Tsuji* Department of Ecological Engineering, Faculty of Engineering Toyohashi University of Technology Tempaku-cho, Toyohashi, Aichi 441-8580, Japan

ABSTRACT

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Poly(lactide), i.e., poly(lactic acid) (PLA) is a biodegradable polyester produced from renewable resources. Due to its high mechanical performance and very low toxicity, PLA is utilized for biomedical and pharmaceutical applications. PLA is susceptible to hydrolytic degradation and yield its monomer, lactic acid in the human body, and then the formed lactic acid is metabolized to finally give CO2 and H2O. In biomedical and pharmaceutical applications, the hydrolytic degradation mechanism and rate are crucial factors for determining its performance and should be accurately adjusted depending on the applications and purposes. Also, consistency in the mechanical properties of organ and artificial material is crucial for biomedical applications to avoid damage to organs. Furthermore, the surface hydrophilicity of a material is closely related to the cell adhesion property and, therefore, should be controlled to elevate the cell affinity for a material or adhesion to it. This review deals with the methods for manipulating the mechanical and also the surface properties and hydrolytic degradation rate of PLA-based materials.

*

.

E-mail: [email protected]

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INTRODUCTION Poly(lactide), i.e., poly(lactic acid) (PLA) is biodegradable (bioabsorbable) in the human body and has very low toxicity to the human body [1-14]. Because of these reasons, PLA has been intensively studied for more than forty years in terms of scientific interest and a wide variety of applications. Also, PLA has high mechanical performance comparable to that of representative commercial polymers such as polystyrene and poly(ethylene terephthalate) (PET). Because of its biodegradability and very low toxicity, in addition to its high mechanical performance, PLA is utilized as biomedical materials for skeletal implants and tissue regeneration and as matrices for drug delivery systems. Typical biodegradable polyesters are summarized in Table 1 [6]. A variety of biodegradable polyesters, including PLA and copolymers, can be utilized for medical and pharmaceutical purposes. Actual and possible applications of biodegradable polymers for medical and pharmaceutical purposes are shown in Table 2 [6]. Skeletal implants or biomaterials for bone tissue regeneration are effective applications of biodegradable polyesters [14,15]. Currently, for fixation of fractured bones in orthopedic and oral surgery metals are normally used in the form of plates, pins, screws, and wires. However, they should be removed after the healing of fractured bones by further surgery. It would be very beneficial to patients if these fixation devices could be fabricated using biodegradable polymers because they don't require re-operation [6].

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Table 1. Classification of Biodegradable Polyesters [6] Polyesters Poly(α-hydroxylacid)s

Chemical structure -(O-CHR-CO)n-

Examples R: H Poly(glycolide) (PGA) R: CH3 Poly(lactide) (PLA)

Poly(3-hydroxyalkanoate)s

-(O-CHR-CH2-CO)n-

R: CH3 Poly(3-hydroxybutyrate) (PHB) R: CH3, C2H5 Poly(3-hydroxybutyrate-co-3hydroxyvalerate) [P(HB-HV)]

-[O-(CH2)m-CO] Xm=3-5

m=3 Poly(γ-butyrolactone) m=4 Poly(δ-valerolactone) m=5 Poly(ε-caprolactone) (PCL)

-[O-(CH2)m-O-CO(CH2)n-CO] x-

m=2, n=2 Poly(ethylene succinate) (PES) m=4, n=2 Poly(butylene succinate) (PBS) m=4, n=2,4 Poly(butylene succinate / adipate) (PBSA)

Miscellaneous Poly(ω-hydroxyalkanoate)s

Poly(alkylene dicarboxylate)

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Table 2. Medical Applications of Bioabsorbable Polymers [6] Function Bonding

Purpose Suturing Fixation Adhesion

Examples Vascular and intestinal anastomosis Fractured bone fixation Surgical adhesion

Closure

Covering Occlusion

Wound cover, Local hemostasis Vascular embolization

Separation

Isolation Contact inhibition

Organ protection Adhesion prevention

Scaffold

Cellular proliferation Tissue guide

Skin reconstruction, Blood vessel reconstruction Nurve reunion

Capsulation

Controlled drug delivery

Sustained drug release

In biomedical and pharmaceutical applications, the hydrolytic degradation mechanism and rate of a material, in addition to its physical properties, are crucial for determining its performance. This review deals with the physical properties and degradation of PLA-based materials having different molecular and highly ordered structures, fillers, material shape, and surface structure, and summarizes the effects of these material factors. However, before discussing the physical properties and hydrolytic degradation, the synthesis of PLA-based materials is briefly summarized (it is stated in detail in Chapter 6).

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1. SYNTHESIS PLAs are synthesized by two methods; polycondensation of lactic acids and ringopening-polymerization (ROP) of lactides, as illustrated in Figure 1 [13]. Figure 2 shows the structure of lactic acid, lactide (the cyclic dimer of lactic acid), and comonomers. Here, Llactide (LLA), D-lactide (DLA), and meso-lactide (meso-LA) are composed of two L-lactic acid (L-lactyl) units, of two D-lactic acid (D-lactyl) units, and an L-lactic acid unit and a Dlactic acid unit, respectively [11]. Racemic lactide or DL-lactide (DLLA) is a 1:1 physical mixture or a 1:1 racemic compound of LLA and DLA, and has a much higher melting temperature (Tm)=124°C compared with those of LLA or DLA (95-99°C). Utilizing various initiators, catalysts, and procedures, PLAs having different tactic structures of isomers can be synthesized (Figure 1), whereas copolymerization of lactic acid or lactide with comonomers such as glycolide and ε-caprolactone (Figure 2) will yield a variety of copolymers (Figure 3) from block-type to random-type.

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Figure 1. Synthesis and molecular structures of poly(L-lactide), i.e., poly(L-lactic acid) (PLLA) [(a) and (b)], poly(D-lactide), i.e., poly(D-lactic acid) (PDLA) [(c) and (d)], and stereo-block isotactic PLA [(e) and (f)] [13].

Also, ROP of LA lactide with multi-functional coinitiators such as glycerol, pentaerythritol, and poly(vinyl alcohol) gives branched polymers and graft copolymers [11]. These PLAs and copolymers belong to the family of aliphatic polyesters and, therefore, their ester groups are hydrolytically degraded in the presence of water according the following reaction (1): -COO- + H2O → -COOH + HO-

(1)

In biomedical and pharmaceutical applications, "bioabsorbability (biodegradability)" is functionality. In these applications, the degradation rate as well as the physical properties of PLA-based materials should be accurately manipulated. For example, the degradation rate of scaffold materials is required to be adjusted to that of tissue regeneration, while the degradation rate of matrices for drug delivery systems should be selected to prolong the effectiveness of drugs.

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Poly(lactide)s and their Copolymers

45

Figure 2. Structure and melting temperature (Tm) of L- and D-lactic acids, lactides (LAs), and their comonomers [11].

2. MECHANICAL PROPERTIES Consistency in the mechanical properties of organ and artificial material is crucial for biomedical applications to avoid damage to organs. The physical properties of the PLLA, PDLLA, and PLA stereocomplex are given in Table 3 [11], together with those of poly[(R)-3hydroxybutyrate] (R-PHB), poly(ε-caprolactone) (PCL), and poly(glycolide) (PGA). The physical properties, including mechanical properties, of polymeric materials depend on their molecular characteristics, highly ordered structures, fillers, and material morphology.

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Figure 3. Structures of LA homopolymer and linear copolymers utilized in medical applications.

Table 3. Physical properties for some biodegradable aliphatic polyesters [11]

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Tm T

a)

0 m

(°C)

b)

c)

(°C)

PLLA

PDLLA

syn-PLA

PCL

151

PLA stereocomplex 220-230

170-190

-

205, 212, 215

-

R-PHB

PGA

60

180

225-230

-

279

71,79

188, 197

-

50-65

50-60

34

65-72

-60

5

40

93,135, 142, 203 1.25-1.29 19-20.5, 22.7

-

-

142, 146

142

146

180-207

1.27 21.1

-

-

1.06-1.13 20.8

1.177-1.260 20.6

1.50-1.69 -

-155±1

0

-

-

0

+44e)

-

WVTRf) (g/m2/day) σBh) (GPa)

82-172 0.12-2.3i)

-

0.8i)

177 0.1-0.8i)

13g) 0.18-0.20i)

0.08-1i)

Ek) (GPa) εBl) (%)

7-10i) 12-26i)

0.040.05j) 1.5-1.9j) 5-10j)

-

8.6i) 30i)

20-120i)

5-6i) 50-70i)

4-14i) 30-40i)

Tg

(°C)

ΔHm0

d)

(J/g)

Density (g/cm3) Solubility parameter (δp) (25°C) [(J/cm3) 0.5] [a] 58925 in chloroform (deg.dm-1.g1.cm3)

a)

Melting temperature; b) Equlibrium melting temperature; c) Glass transtion temperature; d) Enthalpy of melting for 100% crystallinity; e) 300nm, 23°C; f) Water vapor transmission rate at 25°C; g) P(HB-HV) (94/6); h) Tensile strength; i) Oriented fiber; j) Non-oriented film; k) Young's modulus; l) Elongation-atbreak.

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Changes in the thermal and mechanical properties of PLAs upon hydrolytic degradation are connected with the structural changes in the crystalline and amorphous regions. The mechanical properties of PLA-based materials are controllable by varying the material parameters, such as their molecular characteristics, highly ordered structures, fillers, and material morphology.

2.1. Effect of Molecular Characteristics Molecular weight is an important parameter in determining mechanical properties, as given by the following equation:

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P = P0 - K / Mn

(2)

where P is the physical property of a polymeric material, P0 is the P of a polymer with an infinite number-average molecular weight (Mn), and K is a constant. We previously revealed that as-cast PLLA films have non-zero tensile strength (σB) below 1/Mn of 2.5x10-5 g-1 mol, or above Mn of 4.0x104 g mol-1, and that σB increases with 1/Mn [6,16], according to Eq. (2) (Figure 4). Eling et al. [17] showed similar σB dependence of PLLA fibers on 1/viscosityaverage molecular weight (Mv). Perego et al. [18] reported that flexural strength reaches a plateau at Mv above 35,000 and 55,000 for PDLLA and amorphous-made PLLA, respectively. Ultimate σB and Young's modulus (E) of fibers from LA copolymers, such as poly(L-lactideco-ε-caprolactone) [P(LLA-CL)] and poly(L-lactide-co-D-lactide) [P(LLA-DLA)], are smaller than those of a homopolymer like PLLA, while the elongation-at-break (εB) of P(LLA-DLA) fiber becomes higher than that of PLLA fiber when they are melt-spun and thermally drawn (Table 4) [19]. On the other hand, Grijpma et al. [20] showed that the impact strength of PLLA can be increased by cross-linking.

2.2. Effect of Highly Ordered Structures The mechanical properties of PLLA vary depending on their highly ordered structure such as crystallinity (Xc) and crystalline thickness (Lc) or T m [21]. An increase in Xc increases the σB and E of PLLA but decreases the εB. A decrease in the σB of PLLA films prepared at high crystallization temperature (Tc) may be ascribed to the formation of large-size spherulites and crystallites in addition to their high Xc. These results indicate that the mechanical properties of PLLA can be controlled to some extent by altering their highly ordered structures. Similar to other polymers, the values of the σB and E of PLLA fibers increase but the εB value decreases with an increase in degree of molecular orientation [17,2225]. Leenslag and Pennings prepared PLLA fiber having σB=2.1GPa and E=16GPa by hotdrawing of a dry-spun fiber [26].

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Figure 4. Tensile strength (σB) of as-cast PLLA and PDLA as a function of Mn-1 [6].

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Table 4. Ultimate mechanical properties and heats of fusion of hot-drawn fibers from various LLA copolymers [19] Preparation method Dry spinning

Melt-spinning

a)

Sample

PLLA P(LLA-CL) (80/20) P(LLA-DLA) (85/15) PLLA P(LLA-CL) (90/10) P(LLA-CL) (80/20) P(LLA-DLA) (85/15)

2.8x10 3.6x105

(MPa) 2300 1050

Ultimate E (GPa) 16 7.3

Ultimate ΔHm a) σB (Jg-1) (%) 22 25 15

6.0x105

950

9.2

21

19

2.8x105 3.8x105

530 400

9.3 8.2

26 23

57 45

3.6x105

350

5.6

29

32

6.0x105

185

5.0

50

0

Mv (g mol-1) 5

Ultimate

σB

Enthalpy of melting.

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Okuzaki et al. prepared PLLA fiber with σB=275 MPa and E=9.1GPa by the zonedrawing method from a low-molecular-weight PLLA (Mv=13,100 g mol-1) [27]. These results prove that molecular orientation as well as molecular weight are important parameters which influence the mechanical properties.

2.3. Effect of Polymer Blending In contrast to the mechanical properties of miscible polymer blends, those of phaseseparated polymer blends are discontinuous at a polymer mixing ratio where inversion of the continuous and dispersed phases takes place. Even when phase separation occurs in the blends, the σB, yield stress (σY), and E of the blends of glassy PLLA (or PDLLA) with rubbery poly(ethylene oxide) (PEO) [28] or PCL [29-31] can be widely varied by altering the polymer mixing ratio (Figure 5). In Figure 5, the weight fraction of PLLA (XPLLA) is defined by the following equation: XPLLA= WPLLA / (WPCl + WPLLA)

(3)

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where WPLLA and WPCL are the weights of PLLA and PCL in the blends. The impact strength of PLLA is reported to become higher by the addition of rubbery biodegradable polymers such as PCL [32]. The stereocomplexation between PLLA and PDLA is said to enhance the tensile properties of blends compared to those of the non-blended PLLA or PDLA [16]. Probably, the microstructure formed by gelation or the increased number of spherulites per unit mass enhance the tensile properties of blends. The increased tensile properties of blends may be also caused by dense chain packing in the amorphous region due to a strong interaction between the L- and D-unit sequences, as evidenced by the increased Tg of blends.

2.4. Effects of Material Morphology The methods used for the preparation of porous PLA-based materials are (1) the removal of the additive or solvent from melt-molded, solution-cast, gelled [3,33-36], and frozen mixtures [37] of PLA-based materials with additives or solvents and from the phase-separated mixtures of PLA / solvent / non-solvent systems [38-40], and (2) the batchfoaming technique (pressure quench) using supercritical CO2 as the blowing agent [41]. Pore formation is effective for lowering E and elevating the εB of biodegradable polymers. The reported examples are concerned with the compression modulus of poly(lactide-co-glycolide) [P(LAGA)] (50:50) [35] and poly(hydroxybutyrate-co-hydroxyvaler ate) [P(HB-HV)] [42], and the tensile modulus of PCL [43]. Figure 6 shows porous PLLA materials prepared from water extraction of PEO from PLLA/PEO blends [36]. The combination of blending and pore formation can yield PLA-based materials having a wide range of E values. For example, E of PLLA/PCL blends was controllable in the range of 0.1-1.4 GPa by altering the blending ratio and pore formation, as shown in Figure 5 [31].

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Figure 5. Young's modulus of non-porous and porous blends of PLLA and PCL as a function of weight fraction of PLLA (XPLLA) in blends [31].

Figure 6. Porous PLLA prepared by water extraction of PEO from PLLA/PEO blends. PLLA/PEO(w/w)=80/20 (A), 60/40(B) [36].

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Recently, electrospining of PLA and copolymers [44], and the PLA stereocomplex [45] have been successfully utilized to fabricate nanofibers or nanomats for biomedical applications. The typical morphology and strain-stress curves are shown in Figures 7 and 8 [46]. Similar to the aforementioned pore formation, the formation of nanomats is effective to lower E and elevate the εB of PLA-based materials. The fiber and nanomat morphology are affected by the applied voltage, the effluent rate and concentration of the solution, the type of solvent, and so on. The mechanical properties of obtained nanomats can be controlled by drawing and annealing [46].

Figure 7. SEM images of P(LLA-GA) (mol/mol=10/90) membranes electrospun under 2 kV/cm electric field strength, at a feed rate of 100 μL/min and different concentrations: (A) 7.5 wt%; (B) 10 wt%; and (C) 15 wt% in HFIP. [Reprinted from Polymer, vol. 44, Zong, X., Ran, S., Fang, D., Hsiao, B.S., and Benjamin Chu, B. Control of Structure, Morphology and Property in Electrospun Poly(glycolide-colactide) Non-Woven Membranes via Post-Draw Treatments, pp.4959-4967, Copyright (2003), with permission from Elsevier.] [46].

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Figure 8. Stress–strain curves for P(LLA-GA) (mol/mol=10/90) electrospun membranes prepared under the 2 kV/cm electric field, at a feed rate of 100 μL/min and different concentrations of 7.5, 10 and 15 wt%, respectively. [Reprinted from Polymer, vol. 44, Zong, X., Ran, S., Fang, D., Hsiao, B.S., and Benjamin Chu, B. Control of Structure, Morphology and Property in Electrospun Poly(glycolide-colactide) Non-Woven Membranes via Post-Draw Treatments, pp.4959-4967, Copyright (2003), with permission from Elsevier.] [46].

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2. SURFACE PROPERTIES Well-known methods for altering the surface hydrophilicity of polymers include plasma treatment, coating, and surface grafting [47]. However, for hydrolyzable biodegradable polyesters such as PLA, PGA, and PCL, alkaline treatment can be used. There are two advantages to utilize the surface treatments of PLA-based materials, such as coating and alkaline treatment. That is, with surface treatments one can manipulate hydrophilicity without altering the bulk physical properties of the materials, such as the mechanical properties, and the treatments are applicable on the spot after the production of the materials. Among the surface treatments, coating with a hydrophilic polymer and alkaline treatment are effective methods for enhancing hydrophilicity. Alkaline hydrolytic degradation proceeds only on the surface of PLA-based materials. Therefore, the ester groups on the surface are hydrolyzed to give hydrophilic hydroxyl and carboxyl groups, resulting in increased surface hydrophilicity (Figure 9) [48]. Actually, Tsuji and Ishida indicated that hydrophilic PVA coating [49] and alkaline treatment [50,51] are effective in enhancing the hydrophilicity of PLA-based materials (Figure 10) and that hydrophilicity is controllable by varying the PVA or alkaline solution concentration and treatment time [49,51]. A polymer surface with a contact angle of around 70 degrees exhibits the highest cell adhesion [52]. Therefore, adjusting the hydrophilicity of PLA-based materials by alkaline treatment will lead to higher cell affinity for the material or adhesion to it.

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Figure 9. Mechanism of alkaline treatment of PLA-based material [48].

Figure 10. Advancing contact angle (θa) of crystallized and amorphous PLLA films (PLLA-C and PLLA-A films, respectively) before treatment and after PVA coating (in 1 g dL-1 of PVA solution at 25°C for 24 hours) and alkaline treatment (in 0.01 N NaOH solution at 37°C for 48 hours) [49,51].

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3. HYDROLYTIC DEGRADATION Numerous factors affect the degradation mechanism and rate of PLA-based materials. The material factors are summarized in Table 5. When catalytic molecules or substances such as enzymes and alkalis are present in the degradation media or environment, the degradation of PLA-based materials proceeds via a surface erosion mechanism, as schematically illustrated in Figure 11(A) [11,53]. Here, the increase in brightness of the material means a molecular weight decrease. In this surface erosion mechanism, catalytic molecules or ions act only on the surface of materials but will not diffuse into the material so the material is eroded from the surface and the core part of the material remains unchanged. In contrast, in the absence or at a very small concentration of catalytic molecules and ions, as in vivo or in a phosphate-buffered solution, the degradation of PLA materials takes place via a bulk erosion mechanism [Figure 11(B)]. This degradation mechanism takes place when PLA-based materials are implanted in the human body. The hydrolytic degradation mechanism depends on the thickness of biodegradable materials. As shown in Table 6, the critical thickness above which the degradation mechanism changes from bulk erosion to surface erosion depends on the molecular structure of bioabsorbable (or hydrolyzable) polymers [54]. Also, in the case of PLA, when the PLA materials is thicker than 2 mm, the hydrolysis-forming oligomers and monomers with a high catalytic effect are entrapped and accumulated in the core part of the materials [55]. This will result in accelerated hydrolytic degradation in the core part (coreaccelerated bulk erosion), as shown in Figure 11(C). On the other hand, crystallized PLA materials contain crystalline and amorphous regions, as depicted in Figure 12 [11,53]. The chains in amorphous regions in crystallized PLA materials are more susceptible to hydrolytic degradation than those in crystalline regions, leaving the chains in crystalline regions intact (Figure 12). The remaining crystalline regions, called "crystalline residues", have the structure of "extended chain crystallites". Such crystalline residues have a very low hydrolytic degradation rate compared to that of amorphous regions and, therefore, are expected to remain for a long period. During the course of hydrolytic degradation, the thickness of the crystalline regions is not fixed but is known to increase. In addition to the thickening of the crystalline regions, new crystalline regions are formed, resulting in an increase in crystallinity (Xc) in vitro and in vivo [11,53-57], as shown in Figure 13 [57]. Such crystallization is due to enhanced chain mobility by a decrease in molecular weight, a reduced entanglement effect, and the presence of water as a plasticizer. Table 5. Material factors which affect degradation behavior and rate of PLA-based materials. Molecular Structures Molecular weight and distribution Tacticity (Optical purity) and distribution Comonomer structure, content, and distribution Terminal groups Branching Crosslinks

Highly-Ordered Structures Crystallinity Crystalline thickness Spherulitic size and morphology Orientation Hybridization (blends and composites)

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Material morphology Material shape and dimension Porosity and pore size

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Figure 11. Hydrolytic degradation mechanisms of bulky PLA materials [11,53].

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Table 6. Critical thickness (Lcritical) of biodegradable polyesters where hydrolytic degradation mechanism changes from bulk erosion to surface erosion [54]. Polymer Poly(anhydride) Poly(ketal) Poly(ortho ester) Poly(acetal) Poly(α-hydroxycarboxylic acid) Poly(ε-hydroxycarboxylic acid) Poly(amide)

Molecular structure / Example (-R-CO-O-CO-)n (-O-CR1R2-O-R3-) n [-O-CR1(OR2)-O-R3-] n (-O-CHR1-O-R2-) n [-O-CH(CH3)-CO-]n Poly(lactic acid), Poly(lactide) [-O-(CH2) 5-CO-]n Poly(ε-caprolactone) (-NH-R-CO -) n

Lcritical 75 μm 0.4 mm 0.6 mm 2.4 cm 7.4 cm 1.3 cm 13.4 m

Figure 14 shows the changes in the weight-average molecular weight (Mn) and peak molecular weight (Mt) of crystalline residues during hydrolytic degradation in a phosphatebuffered solution at 37°C [58]. These crystalline residues were prepared by accelerated hydrolytic degradation of crystallized PLLA in a phosphate-buffered solution at 97°C for 40 hours. Assuming linear decreases of Mn and Mt during the hydrolysis period of from 192 to 512 days, the estimated average hydrolysis rates [using equation (4)] for changes in Mn and Mt in that period were 5.31 and 5.01 g mol-1 day-1, respectively.

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Figure 12. Schematic representation of structures of crystallized PLA material before and after hydrolytic degradation, or of the formation of crystalline residues (or extended chain crystallites) [53].

Figure 13. Xc change of PLLA films having different initial Xc during hydrolytic degradation in phosphate-buffered solution at pH 7.4 and 37°C [57]. Degradable Polymers for Skeletal Implants, Nova Science Publishers, Incorporated, 2009. ProQuest Ebook Central,

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Figure 14. Mn and Mt of PLLA crystalline residues during hydrolytic degradation in phosphate-buffered solution at pH 7.4 and 37°C [58].

Mn(t2)=Mn(t1)-k (t2 - t1)

(4)

Using these two lines for Mn and Mt in the period of from 192 to 512 days and assuming that the hydrolysis rates were constant irrespective of the molecular weights of the crystalline residues for hydrolysis periods exceeding 192 days, the total periods required for the crystalline residues having initial Mn=10,200 and Mt=13,900 to be degraded into half of an Llactide unit with molecular weight of 72.1 g mol-1, were 1.82x103 and 2.25x103 days, respectively. This finding means that the crystalline residues (extended-chain crystallites) can remain for a long period, such as ca. 2x103 days (ca. 5.5 years) after PLLA loses its functions as a biomaterial. The Arrhenius plots of the k values obtained at different degradation temperatures (50-97°C) yield the ΔEh value of PLLA crystalline residues [59]. The obtained -1

-1

value of 18.0 kcal mol (75.2 kJ mol ) in the temperature range of 50-97°C is higher than -1

12.2 kcal mol (50.9 kJ mol-1) for PLLA hydrolytic degradation in the melt (250-350°C) [60], revealing that the PLLA chains in the crystalline residues are, in terms of hydrolysis, much more degradation-resistant than those in an amorphous state or in the melt. Among the factors shown in Table 5, the representative factors which determine the degradation of PLA-based materials in vivo are molecular weight, structure and content of comonomer unit, crystallinity, orientation, blending, material thickness and porosity.

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Normally, the incorporation of hydrophilic monomer units or polymers accelerates the hydrolytic degradation. On the other hand, increasing molecular weight, crystallinity, or the degree of orientation reduces the hydrolytic degradation rate. The indexes for hydrolytic degradation or biodegradation are tabulated in Table 7. The indexes which can be used for tracing the hydrolytic degradation of materials depend on the erosion mechanism. For example, gravimetry is effective to monitor surface erosion because a significant weight loss is observed at an early stage of degradation. However, gravimetry is ineffective in the case of bulk erosion because the weight loss occurs at a late stage of degradation when a large decrease in molecular weight takes place and, thereby, water-soluble oligomers and monomers are formed and elute from the materials. In contrast, molecular weight change is most effective to trace a bulk erosion, but is ineffective in the case of surface erosion. Table 7. Indexes for degradation and their tracing methods. Indexes Weight remaining Molecular weight and distribution Physical properties Material morphology

Tracing methods Gravimetry Gel permeation chromatography (GPC), i.e., size exclusion chromatography (SEC) Thermal properties-Differential scanning calorimetry (DSC), Mechanical properties-Tensile testing, and so on. Microscopy [scanning electron microscopy (SEM), transmission electron microscopy (TEM), polarized optical microscopy]

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5.1. Effects of Molecular Structures Among the effects of molecular structure, those of the monomer unit structure and the molecular weight of a polymer are crucial. In the sections 5.1.-5.4., the hydrolytic degradation media have neutral pH and the temperature is 37°C, unless otherwise specified. The incorporation of comonomer units such as glycolide (GA) and ε-caprolactone (CL) accelerates the hydrolytic degradation of PLA-based copolymers, irrespective of the hydrophilicity of the comonomer. This is in marked contrast to alkaline degradation, wherein the hydrophilicity of the comonomer has a dominant effect on hydrolytic degradation [61]. Figure 15 shows the Mn and residual tensile strength changes of amorphous PLLA, P(LLADLA)(77:23), P(LLA-GA)(81:19), and P(LLA-CL)(82:18) with respect to hydrolytic degradation in a phosphate-buffered solution [62]. Here, to remove the effects of highly ordered structure, all the initial specimens were made amorphous. The hydrolytic degradation rate constant (k) values of the initially amorphous PLLA were evaluated from the changes in number-average molecular weight (Mn) according to the following equation: ln Mn(t2)=ln Mn(t1)-k (t2 - t1)

(5)

where Mn(t2) and Mn(t1) are the Mn values at the hydrolytic degradation times of t2 and t1, respectively. The k values were estimated from the Mn change in Figure 15(a) using equation (3). This equation should be used when no significant weight loss was observed. For the estimation of the k values, the Mn data for the period of 0-24 weeks were used, except for

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those for the period of 0-16 weeks in the case of P(LLA-GA), where the slope of log Mn changed at 16 weeks. The estimated k values for PLLA, P(LLA-DLA), P(LLA-GA), and P(LLA-CL) were 1.11x10–3, 2.39x10–3, 1.61x10–2, and 1.81x10-2 day–1. The k values of PLLA and P(LLA-DLA) are comparable to 2.59 x10–3 day–1 (0-52 weeks) reported for an amorphous-made PLLA (Mw = 1.2x106 g mol–1, Mw/Mn = 2.6) [62,63], and that of P(LLAGA) is very similar to 1.91x10–2 day–1 reported for a P(LLA-GA)(75:25) scaffold (Mn = 1.81x 105 g mol–1, Mw/Mn = 1.79) [64]. The estimated k value for P(LLA-CL) is comparable to 3.16x10-2 day-1 reported for P(LLA-CL)(60:40) (Mn = 7.4x104 g mol–1, Mw/Mn = 2.0) [65].

Figure 15. Mn and residual tensile strength changes of amorphous PLLA, poly(L-lactide-co-D-lactide) [P(LLA-DLA)] (77:23), poly(L-lactide-co-glycolide) [P(LLA-GA)] (81:19), and poly(L-lactide-co- caprolactone) [P(LLA-CL)] (82:18) with respect to hydrolytic degradation in phosphate-buffered solution at pH 7.4 and 37°C [62,63].

The rather higher k value reported for P(LLA-CL)(60:40) may be due to the higher fraction of CL units and thickness of the copolymer specimens compared to those in the present study. The k values were higher for P(LLA-GA) and P(LLA-CL) than those for PLLA and P(LLA-DLA). The higher hydrolytic degradation rates of P(LLA-GA) and P(LLA-CL) are partly attributable to their higher chain mobility. This is evidenced by the fact that the Tg values of P(LLA-GA) and P(LLA-CL) (56 and 27ºC, respectively) were lower than those of PLLA and P(LLA-DLA) (60 and 58ºC, respectively) [62,63]. Such higher chain mobility of P(LLA-GA) and (PLLA-CL) facilitates faster water supply and, thereby, enhances hydrolytic degradation. Furthermore, in the case of P(LLA-GA), the higher hydrophilicity and lower steric hindrance of GA units compared with those of the L-lactide (LLA) units should have enhanced the water diffusion and concentration of P(LLA-GA), resulting in rapid hydrolytic degradation. Incorporation of D-lactide (DLA) units in PLLA chains is known to increase the hydrolytic degradation rate [66,67]. We have recently found that even the incorporation of a very small amount of DLA units (1.2%) in PLLA chains significantly elevates the hydrolytic degradation rate [68].

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Although molecular weight has a very small effect on the hydrolytic degradation rate of PLA-based materials with molecular weights exceeding 8x104 g mol-1 [68], the effect becomes significant for molecular weights below 8x104 g mol-1 [69]. The enhanced hydrolytic degradation with decreasing molecular weight at the range below 8x104 g mol-1 can be explained by four factors: (1) the elevated molecular mobility, (2) the increased density (number per unit mass) of hydrophilic terminal carboxyl and hydroxyl groups, (3) the increased density of catalytic terminal carboxyl groups, (4) the higher probability of the formation of water-soluble oligomers and monomers upon hydrolytic degradation. Factors (1) and (2) increase the water diffusion rate and content, enhancing the hydrolytic degradation. Factors (1)-(4) should be dramatic for PLA at a molecular weight lower than 1x104 g mol-1.

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5.2. Effects of Highly Ordered Structures Figure 16(C) shows the molecular weight distribution change of crystallized PLLA hydrolytically degraded in a phosphate-buffered solution. As depicted in Figure 12, the chains in the amorphous regions are selectively degraded and removed, and thereby the crystalline residues (extended chain crystallites) are formed. The peak at around 1x104 g mol-1 is ascribed to one fold of the thickness of the PLLA crystalline residues. Among the effects of highly ordered structures, that of crystallinity is reported to have a crucial influence on the hydrolytic degradation of PLLA in neutral media [11]. Figure 17 shows the Mn change of PLLA having different crystallinity (Xc) values in a phosphate-buffered solution [57]. Surprisingly, the elevated crystallinity of PLLA enhances its hydrolytic degradation [57,7072]. This is in marked contrast with the results for proteinase K-catalyzed enzymatic degradation [73] and alkaline degradation [74,75], wherein the degradation rate is higher for PLLA materials having a lower crystallinity. The result here can be explained as follows. Upon the crystallization of PLLA, the hydrophilic terminal groups (-OH and -COOH) and the catalytic terminal group (-COOH) are excluded from the crystalline regions and are condensed in the amorphous region between the crystalline regions, as depicted in Figure 18 [53,57]. The open and closed circles in the figure refer to hydroxyl and carboxyl groups, respectively, or vice versa. The high densities (numbers per unit mass) of terminal groups will cause loose chain packing in the amorphous region between the crystalline regions compared to those in the completely amorphous specimen. Such loose chain packing and high densities of the hydrophilic terminal groups enhance the diffusion of water molecules and increase the water content. The elevated water supply rate and content and the increased catalytic effect by the high density of the carboxylic group will synergize to remarkably accelerate hydrolytic degradation of crystallized PLLA. In contrast to the enzymatic degradation of PLLA [76,77], orientation has a relatively small effect on hydrolytic degradation in a neutral medium [78]. The initial crystalline thickness affects the hydrolytic degradation behavior of PLLA materials especially at a late stage. Figure 19 shows the initial Tm before hydrolytic degradation and peak molecular weight (Mt) after hydrolytic degradation [79]. This shows that the molecular weight of crystalline residues depends on the initial crystalline thickness, while the molecular weight of crystalline residues formed by hydrolytic degradation of hydrolysis-forming crystalline regions depends on the media and other conditions. The molecular weight of Peak I (the lowest) (Mn,s) can be converted to the lamella thickness (Lc) of PLLA after hydrolytic degradation using the following equation assuming the PLLA

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chains take 103 helices in the α-form unit cell having a dimension of c=2.78 (or 2.88) nm (fiber axis) [80,81]: Lc (nm) = 0.278 (or 0.288) x Mn,s / 72.1

(6)

where 72.1 is the mass per mole of the lactyl unit (half of the lactide unit). The following equations show the relationship between final Tm and Lc: Tm (K) = 471 [1-1.59/Lc (nm)] (Degradation in phosphate-buffered solution)

(7)

Tm (K) = 472 [1-1.46/Lc (nm)] (Degradation in the presence of proteinase K)

(8)

Here, the c values of 2.78 nm and 2.88 nm, respectively, were used to calculate the Lc values in equations (5) and (6). The Thomson-Gibbs expression between Tm and Lc is given by the following equation [82]: Tm =Tm0 (1-2σ/Δh0ρcLc)

(9)

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where σ, Δh0, and ρc are the specific fold surface free energy, heat of fusion (per unit mass), and crystal density, respectively. The comparison between (7) or (8) and (9) gives the Tm0 values of 198 and 199°C as the Tm0 values, which are slightly lower but almost in agreement with those obtained from the Hoffman-Weeks procedure for the melt-crystallized PLLA specimens by Kalb and Pennings (215°C) [83] and by ourselves (212°C) [21]. This agreement on Tm0 estimated by different methods also confirms our assumption that Mn,s is the molecular weight of one fold of the PLLA chain in the crystalline regions.

Figure 16. Molecular weight distribution curves of crystallized PLLA film hydrolytically degraded in (A) enzymatic, (B) alkaline, and (C) phosphate-buffered solutions [11]. Degradable Polymers for Skeletal Implants, Nova Science Publishers, Incorporated, 2009. ProQuest Ebook Central,

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Figure 17. Mn and residual tensile strength changes in PLLA films having different initial crystallinity (Xc) values with respect to hydrolytic degradation in phosphate-buffered solution at pH 7.4 and 37°C [57].

Figure 18. Schematic representation of completely amorphous and crystallized PLLA [53,57].

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Figure 19. Initial Tm before hydrolytic degradation and peak molecular weight (Mt) after hydrolytic degradation [79].

5.3. Effects of Blending and Additives Five or more factors for a second polymer or an additive are anticipated to influence the hydrolytic degradation rate and behavior of PLA-based materials in vivo and in a neutral media; (1) miscibility and dispersibility, (2) hydrophilicity, (3) acidity or basicity, (4) molecular weight, (5) sizes and morphology of the polymer domains (if the second polymer is immiscible with the first polymer) or additive. For example, the addition of hydrophilic polymers such as PVA [84,85] and PEO [86], basic additives such as thioridazine [87], or catalytic additives such as lauric acid [88] and lactide [70], is known to enhance the hydrolytic degradation of PLLA in a neutral medium. Normally, the hydrolytic degradation rate constant (k) values, estimated using equation (5), are in the range of 2.2-3.4x10-3 day-1 (012 months) [57,89]. The addition of PVA and lauric acid respectively increased the k value to 9.4x10-3day-1 [85] and 0.03 day-1 [88]. A peculiar example is the addition of PDLA to PLLA. In the blends, the interaction between a PLLA chain and a PDLA chain is much higher than that between PLLA chains or PDLA chains, resulting in a stereocomplex formation [90].

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Such strong interaction disturbs the diffusion of water into the material and lowers the hydrolytic degradation rate. For example, as-cast pure PLLA and PDLA had k values of 1.41x10-3 and 2.14x10-3day-1, respectively, whereas the k value of their 1:1 blend was 7.3x104 day-1 [90]. That is, the 1:1 blend has a k value one order lower than those of nonblended specimens. In summary, utilizing blending and additives, PLA-based materials having k values of 7.3x10-4 to 3x10-2 day-1 can be fabricated. However, it is surprising that the addition of hydrophobic biodegradable polyesters such as PCL can accelerate the hydrolytic degradation rate of PLLA in phase-separated blends. Figure 20 shows the Mn and residual tensile strength changes in PLLA/PCL blends having different initial XPLLA values with respect to hydrolytic degradation in a phosphate-buffered solution [91]. The k values obtained using equation (5) were 2.80, 3.07, 5.11, 7.13, and 5.2x10-3 day-1, for XPLLA=0, 0.25, 0.5, 0.75, and 1, respectively. As-cast PCL having XPLLA=0 (initial Xc=67%) is susceptible to hydrolytic degradation when compared to that of as-cast PLLA having XPLLA=1 (initial Xc=51%), although PCL had a much lower initial molecular weight than that of PLLA. The blends with XPLLA=0.5 and 0.75 had significantly higher k values than those calculated from the k values of PCL and PLLA. The probable explanation for this result is as follows. PCL and PLLA are partially miscible and the PCL molecules incorporated in the PLLA-rich phase elevated the hydrolytic degradation rate of PLLA due to its higher catalytic terminal carboxyl groups originating from its lower molecular weight. Moreover, PLLA has a higher hydrophilicity because of its higher number per unit mass of hydrophilic ester groups and, therefore, the PLLA molecules incorporated in the PCL-rich phase should have an increased water-supply rate and content of the PCL-rich phase, resulting in a rapid hydrolytic degradation of PCL. 106

175

(a) Mn

(b) Residual σ

B

125

B

Residual σ (%)

105

Mn

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150

4

10

XPLLA=0

0

4

8

75 50

0.25 0.5 0.75 1 103

100

25

12

16

20

Hydrolytic degradation time (months)

24

0

0

4

8

12

16

20

24

Hydrolytic degradation time (months)

Figure 20. Mn and residual tensile strength changes in PLLA/PCL blend films having different initial XPLLA values with respect to hydrolytic degradation in phosphate-buffered solution at pH 7.4 and 37°C [91].

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5.4. Effects of Material Morphology Both thinning materials and pore formation decrease the average length required for the hydrolysis-forming water-soluble oligomers and monomers to diffuse outside and for water molecules to diffuse inside the material. Therefore, such shape changes will increase the transfer rates of lactic acid oligomers and monomers and water molecules. The increased transfer rates of lactic acid oligomers and monomers will decrease their catalytic effects for hydrolytic degradation, whereas the elevated transfer rate of water will elevate the hydrolytic degradation rate. The reported result that porous PLLA had a lower hydrolytic degradation rate than that of nonporous PLLA [92] indicated that the former effect prevailed over the latter one.

CONCLUSION The mechanical properties and the hydrolytic degradation rate of PLA-based materials can be manipulated by varying the molecular and highly ordered structures, fillers, and material morphology. The surface property (cell affinity or adhesion) is controllable by alkaline surface treatment, coating, and surface grafting. For biomedical applications of PLAbased materials, including skeletal applications, the mechanical and surface properties and hydrolytic degradation rate should be highly adjusted for each purpose.

REFERENCES

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[6] [7] [8] [9]

Kharas G.B., Sanchez-Riera F., and Severson D.K. (1994). Polymer of lactic acids. In Mobley D.P. (Ed.). Plastics from microbes (pp. 93-137). New York: Hanser Publishers. Doi Y., and Fukuda K. (Eds.) (1994). Biodegradable plastics and polymers. Amsterdam: Elsevier. Coombes A.G.A. and Meikle M.C. (1994). Bioabsorbable synthetic polymers as replacements for bone graft, Clin. Mater. 17, 35-67. Vert M., Schwarch G., and Coudane J.J. (1995). Present and future of PLA polymers. J. Macromol. Sci.: Pure Appl. Chem. 1995; A32: 787-796. Hartmann, M.H. (1998). High molecular weight polylactic acid polymers. In Kaplan, D.L., (Eds.), Biopolymers from renewable resources (pp. 367-411). Berlin (Germany): Springer; 1998. Ikada, Y. and Tsuji, H. (2000). Biodegradable polyesters for medical and ecological applications. Macromol. Rapid Commun. 21, 117-132. Garlotta, D. (2001). A literature review of poly(lactic acid). J. Polym. Environ., 9, 6384. Albertsson, A.-C. (Ed.) (2002). Degradable aliphatic polyesters (Advances in polymer science, vol.157). Berlin (Germany): Springer. Södergård, A. and Stolt, M. (2002). Properties of lactic acid based polymers and their correlation with composition. Prog. Polym. Sci. 27, 1123-1163.

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[10] Scott, G. (Ed.) (2002). Biodegradable polymers: Principles and applications (2nd edition). Dordrecht (The Netherlands): Kluwer Academic Publishers; 2002. [11] Tsuji, H. (2002). Poly(lactide). In Doi, Y. and Steinbüchel, A. (Eds.), Polyesters III (Biopolymers, vol. 4) (pp. 129-177). Weinheim (Germany): Wiley-VCH. [12] Auras, R., Harte, B., and Selke, S. (2004). An overview of polylactides as packaging materials. Macromol. Biosci., 4, 835-864. [13] Tsuji, H. (2005). Poly(lactide) Stereocomplexes: Formation, structure, properties, degradation, and applications. Macromol. Biosci., 2005, 5, 569-597. [14] Salgado, A.J., Coutinho, O.P., and Reis, R.L. (2004). Bone tissue engineering: State of the art and future trends, Macromol. Biosci., 4, 743-765. [15] Wuisman, P.I.J.M. and Smit, T.H. (2006). Bioresorbable polymers: Heading for a new generation of spinal cages, Eur. Spine J., 15, 133-148. [16] Tsuji, H. and Ikada, Y. (1999). Stereocomplex formation between enantiomeric poly(lactic acid)s. 11. Mechanical properties and morphology of solution-cast films, Polymer, 40, 6699-6708. [17] Eling, B., Gogolewski, S., and Pennings, A.J., (1982). Biodegradable materials of poly(L-lactic acid): 1. Melt-spun and solution spun fibres, Polymer, 23, 1587-1593. [18] Perego, G., Cella, G. D., and Bastioli, C. (1996). Effects of molecular weight and crystallinity on poly(lactic acid) mechanical properties, J. Appl. Polym. Sci., 59, 37-43. [19] Penning, J.P., Dijkstra, H., and Pennings, A.J. (1993). Preparation and properties of absorbable fibers from L-lactide copolymers, Polymer, 34, 942-951. [20] Grijpma, D.W., Kroeze, E., Nijenhuis, A.J., and Pennings, A.J. (1993). Poly(L-lactide) crosslinked with spiro-bis-dimethylene-carbonate, Polymer, 34, 1496-1503. [21] Tsuji, H. and Ikada, Y. (1995). Properties and morphologies of poly(L-lactide): 1. Annealing condition effects on properties and morphologies of Poly(L-lactide), Polymer, 36, 2709-2716. [22] Hyon, S.-H., Jamshidi, K., and Ikada, Y. (1984). Melt spinning of poly-L-lactide and hydrolysis of the fiber in vitro. In: Polymers as biomaterials, Shalaby, S.W., Hoffman, A.S., Ratner, B.D., Horbett, T.A. (Eds.). New York: Prenum Press, pp.51-65. [23] Horacek, I. and Kalísek, V. (1994). Polylactide. I. Continuous dry spinning-hot drawing preparation of fibers, J. Appl. Polym. Sci., 54, 1751-1757. [24] Horacek, I. and Kalísek, V. (1994). Polylactide. II. Discontinuous dry spinning-hot drawning preparation of fibers, J. Appl. Polym. Sci., 54, 1759-1765. [25] Fambri, L., Pegoretti, A., Fenner, R., Incardona, S.D., and Migliaresi, C. (1997). Biodegradable fibres of poly(L-lactic acid) produced by melt-spinning, Polymer, 38, 79-85. [26] Leenslag., J.W. and Pennings, A.J. (1987). High-strength poly(L-lactide) fibres by a dry-spinning/Hot-drawing process, Polymer, 28, 1695-1702. [27] Okuzaki, H., Kubota, I., and Kunugi, T. (1999). Mechanical properties and structure of the zone -drawn poly(L-lactic acid) fibers, J. Polym. Sci., Part B: Polym. Phys., 37, 991-996. [28] Sheth, M., Kumar, R.A., Davé, V., Gross, R.A., and McCarthy, S.P. (1997). Biodegradable polymer blends from poly(lactic acid) and poly(ethylene glycol), J. Appl. Polym. Sci., 66, 1495-1505.

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[60] Tsuji, H., Daimon, H., and Fujie, K. (2003). A new strategy for recycling and preparation of poly(L-lactic acid): Hydrolysis in the melt. Biomacromolecules, 4, 835840. [61] Tsuji, H. and Tezuka, Y. (2005). Alkaline and enzymatic degradation of L-lactide copolymers: 1. Amorphous-made films of L-lactide copolymers with D-lactide, glycolide, ε-caprolactone, Macromol. Biosci., 5, 135-148. [62] Saha, S.K. and Tsuji, H. (2006). Hydrolytic degradation of amorphous-made films of Llactide copolymers with glycolide and D-lactide, Macromol. Mater. Eng., 291, 357-368. [63] Saha, S.K. and Tsuji, H. (2006). Effects of rapid crystallization on hydrolytic degradation and mechanical properties of poly(L-lactide-co-e-caprolactone), React. Funct. Polym., 66, 1362-1372. [64] Wu, L. and Ding, J. (2004). In vitro degradation of three-dimensional porous poly(D,Llactide-co-glycolide) scaffolds for tissue engineering. Biomaterials, 25, 5821-5830. [65] Malin, M., Hiljanen-Vainio, M., Karjalainen, T., and Seppälä, J. (1996). Biodegradable lactone copolymers: II. Hydrolytic study of ε-caprolactone and lactide copolymers. J. Appl. Polym. Sci., 59, 1289-1298. [66] Vert, M., Chabot, F., Leray, P., and Christel, P. (1981). Stereoregular bioresorbable polyesters for orthopaedic surgery. Makromol. Chem. Suppl., 5, 30-41. [67] Tsuji, H. (2002). Autocatalytic hydrolysis of amorphous-made polylactides: Effects of L-lactide content, tacticity, and enantiomeric polymer blending, Polymer, 43, 17891796. [68] Saha, S.K. and Tsuji, H. (2006). Effects of molecular weight and small amounts of Dlactide units on hydrolytic degradation of poly(L-lactic acid)s. Polym. Degrad. Stab., 91, 1665-1673. [69] Hyon, S.-H., Jamshidi, K., and Ikada, Y. (1998). Effects of residual monomer on the degradation of DL-lactide polymer. Polym. Int., 46, 196-202 [70] Nakamura, T., Hitomi, S., Watanabe, S., Shimizu, Y., Jamshidi, K., Hyon, S.-H., and Ikada, Y. (1989). Bioabsorption of polylactides with different molecular properties, J. Biomed. Mater. Res., 23, 1115-1130. [71] Li, S., Garreau, H. and Vert, M. (1990). 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. [72] Duek, E.A.R., Zavaglia, C.A.C., and Belangero, W.D. (1999). In vitro study of poly(lactic acid) pin degradation. Polymer, 40, 6465-6473. [73] Tsuji, H. and Miyauchi, S. (2001). Poly(L-lactide): 7. Enzymatic hydrolysis of free and restricted amorphous regions in poly(L-lactide) films with different cystallinities and a fixed crystalline thickness. Polymer, 42, 4463-4467. [74] Cam, D., Hyon, S.-H., and Ikada, Y. (1995). Degradation of high molecular weight poly(L-lactide) in alkaline medium. Biomaterials, 16, 833-843. [75] Tsuji, H. and Ikada, Y. (1998). Properties and morphology of poly(L-lactide): II. Hydrolysis in alkaine solution, J. Polym, Sci. :Part A: Polym. Chem., 36, 59-66. [76] Cai, H. Dave, V. Gross, R.A., and McCarthy, S.P. (1996). Effects of physical aging, crystallinity, and orientation on the enzymatic degradation of poly(lactic acid). J. Polym. Sci.: Part B: Polym. Phys., 34, 2701-2708.

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[77] Tsuji, H., Ogiwara, M., Saha, S.K., and Sakaki, T. (2006). Enzymatic, alkaline, and autocatalytic degradation of poly(L-lactic acid): Effects of biaxial orientation, Biomacromolecules, 7, 380-387. [78] Hyon, S.-H., Jamshidi, K., and Ikada, Y. (1984). Melt spinning of poly-L-lactide and hydrolysis of the fiber in vitro. In Shalaby, S.W., Hoffman, A.S., Ratner, B.D., and Horbett, T.A. (Eds.), Polymers as Biomaterials (pp. 51-65). NewYork: Prenum Press. [79] Tsuji, H., Ikarashi, K., and Fukuda N. (2004). Poly(L-lactide): XII. Formation, growth, and morphology of crystalline residues as extended-chain crystallites through hydrolysis of poly(L-lactide) films in phosphate-buffered solution. Polym. Degrad. Stab., 84, 515-523. [80] Hoogsteen, W., Postema, A.R., Pennings, A.J., ten Brinke, G., and Zugenmaier, P. (1990). Crystal structure, donformation, and morphology of solution-spun poly(Llactide) fibers, Macromolecules, 23, 634-642. [81] Okihara, T., Tsuji, M., Kawaguchi, A., Katayama, K., Tsuji, H., Hyon, S.-H., and Ikada, Y. (1991). Crystal structure of stereocomplex of poly(L-lactide) and poly(Dlactide), J. Macromol. Sci.-Phys. B30, 119-140. [82] For example, Gedde, U.W. (1995). Polymer physics (pp.131-198). London: Chapman and Hall. [83] Kalb, B. and Pennings, A.J. (1980). General crystallization behaviour of poly(L-lactic acid), Polymer, 21, 607-612. [84] Pitt, C.G., Cha, Y., Shah, S.S., and Zhu, K.J. (1992). Blends of PVA and PGLA: Control of the permeability and degradability of hydrogels by blending, J. Controlled Release, 19, 189-200. [85] Tsuji, H. and Muramatsu, H. (2001). Blends of aliphatic polyesters: 5. Non-enzymatic and enzymatic hydrolysis of blends from hydrophobic poly(L-lactide) and hydrophilic poly(vinyl alcohol), Polym. Degrad. Stab., 71, 403-413. [86] Nijenhuis, A.J., Colstee, E., Grijpma, D.W., and Pennings, A.J. (1996). High molecular weight poly(L-lactide) and poly(ethylene oxide) blends: Thermal characterization and physical properties, Polymer, 37, 5849-5857. [87] Kishida, A., Yoshioka, S., Takeda, Y., and Uchiyama, M. (1989). Formulation-assisted biodegradable polymer matrices. Chem. Pharm. Bull., 37, 1954-1956. [88] Renouf-Glauser, A.C., Rose, J., Farrar, D., and Cameron, R.E. (2005). A degradation study of PLLA containing lauric acid. Biomaterials, 26, 2415-2422. [89] Tsuji, H. and Ikada, Y. (2000). Properties and morphology of poly(L-lactide): 4. Effects of structural parameters on long-term in vitro hydrolysis of poly(L-lactide) in phosphate-buffered solution. Polym. Degrad. Stab., 67, 179-189. [90] Tsuji, H. (2000). In vitro hydrolysis of blends from enantiomeric poly(lactide)s: 1. Well-stereocomplexed blend and non-blended films. Polymer, 41, 3621-3630. [91] Tsuji, H. and Ikada Y. (1998). Blends of aliphatic polyesters. II. Hydrolysis of solutioncast blends from poly(L-lactide) and poly(ε-caprolactone) in phosphate-buffered solution, J. Appl. Polym. Sci., 67, 405-415. [92] Lam, K.H., Nieuwenhuis, P., Molenaar, I., Esselbrugge, H., Feijen, J., Dijkstra, P.J., and Shakenraad, J.M. (1994). Biodegradation of porous versus non-porous poly(Llactic acid) dilms. J. Mater. Sci. Mater. Med., 5, 181-189.

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Chapter 4

COMPOSITES BASED ON DEGRADABLE POLYMERS K. E. Tanner* Departments of Civil and of Mechanical Engineering, University of Glasgow, Glasgow, Scotland, UK

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ABSTRACT While polymers have been used in medical applications since the 19th Century, it was only towards the end of 20th Century that degradable polymers were first used in the manufacture of biomedical composites. In the vast majority of cases ceramic and glass fillers have been added to increase both the bioactivity, that is to encourage integration of the material into body tissue, and the mechanical properties, in particular the stiffness, so that materials can be produced which have similar mechanical properties to bone. These materials have been subjected to mechanical, in vitro and in vivo testing, however, currently only polymer-polymer composites have been reported to be used clinically. The addition of the fillers has generally increased the positive biological response to the degradable polymers and the choice of filler has allowed the control of the changes in pH around the polymers, thus increasing the biological acceptability compared to the nonfilled polymer, and allowed control of the degradation rate. Additionally particulate filled polymer composites have been used to manufacture scaffolds for potential bone tissue engineering scaffolds, encouraging the ingrowth of cells into the scaffold.

INTRODUCTION A composite is composed of two or more phases combined so that the two phases can be differentiated at the microscopic or sub-microscopic scale. This definition therefore excludes metal alloys, such as stainless steel where the carbon, chromium, nickel and molybdenum atoms are interspersed throughout the iron atoms or polymer blends where the individual molecules are intermingled. In an alloy it is the individual atoms that interact to give the *

Author’s address: Departments of Civil and of Mechanical Engineering, James Watt (South) Building University of Glasgow, Glasgow, G12 8QQ, UK, Telephone +44 141 330 3733, Fax +44 141 330 4343, e-mail: [email protected].

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properties of the alloy, while in a polymer blend the individual polymer chains interact. In a composite the individual phases interact as section of bulk materials, albeit extremely small sections of bulk. The rationale for composite production is to optimise one or more of the material properties of the individual phases and to overcome any deficiencies of these phases. In many cases the properties to be optimised are mechanical, for example in a Formula 1 racing car the aim is to manufacture a composite cockpit at a minimum mass with limited consideration of costs, but which will allow the driver to walk away from a head-on crash at over 200km per hour. However, depending on the application, other properties can be optimised. In composites for biomedical applications the biological properties are commonly one of the properties of interest. Composites can be based all four of the major classes of materials: metals, ceramics, glasses and polymers and may be based on between one and four of these classes of materials. Thus one can have glass-in-polymer, ceramic-in-metal, ceramic-in-glass or polymer-polymer composites. Composites normally consist of one continuous phase, called the matrix, with a second phase, called the filler, distributed in the matrix as particles, fibres or fabric. These two phases have defined properties and functions. The major mechanical function of the matrix is to hold the filler in place. Polymers generally are tough, ductile, have low modulus, low wear resistance, high creep and poor thermal stability, while ceramic and glass based fillers are brittle, stiff, have good wear resistance, low creep and high thermal stability. Thus ceramic or glass in polymer composites can (depending on formulation) have reasonable strength, toughness, creep resistance, wear resistance and medium stiffness. In a composite no property can be “better” than that of any individual phase but, for example, by combining the stiff, brittle and bioactive ceramic, hydroxyapatite, with low modulus, ductile, but biologically inert, polyethylene, the resultant composite has medium mechanical stiffness, thus approaching the properties of cortical bone, reasonable toughness and yet is still bioactive [1]. Four main groups of degradable polymers have been used in medical applications and are described in greater detail in other chapters of this book. However, briefly the groups are in the order of the time they started to be used in biomaterial composites, polyhydroxybutyrate and its co-polymers, polyglycolic and polylactic acid, chitosan based polymers and hydrogels. The fillers used in the degradable polymers are three principle groups. The two rigid fillers are calcium phosphate based ceramics, principally hydroxyapatite (HA – Ca10(PO4)6(OH)2) and tricalcium phosphate (TCP – Ca3(PO4)2), as well as mixtures of these two ceramics plus various biologically active glasses such as Bioglass® and Wollastonite based glass-ceramics. The third type of filler is drawn fibres of a biodegradable polymer where the fibres drawing aligns the polymer chains and thus increases the stiffness and strength of the polymer, as has been described by Joukainen et al. [2] from the group in Tampere, Finland.

MATRIX MATERIALS In the late 1980s polyhydroxybutyrate (PHB) and its co-polymer polyhydroxyvalyrate (PHV) became available as “medical grade” and various groups started to use it both as a

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non-filled polymer and as the matrix phase of composites. Composites based on PHB and PHV were developed with fillers including hydroxyapatite [3] and phosphate based glasses [4]. However, in the mid 1990s “medical grade” PHB and PHV become unobtainable for about 10 years, but more recently various PHB/PHV based composites have been developed. PHB and PHV have been used as they break down to form hydroxybutyric acid which can be processed by the human body as it occurs naturally in blood. PHV co-polymer is used to increase the strain to failure, for example Bergmann and Owen [5] found that the addition of 27mol% of PHV increased the strain to failure of the unfilled polymer from 1.8% strain to 9.7% strain, however, this level of co-polymer reduced the modulus from 3.2GPa to 0.64GPa. At the same time as PHB/PHV ceased to be available in medical grade, polyglycolic acid (PGA) sutures become available and then polylactic acid (PLA) and more recently chitosan has been used as the matrix of composites, along with various polymer blends.

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MECHANICAL PROPERTIES OF COMPOSITES In composites the major factors that control the mechanical properties of the composite are the filler, the matrix, the amount of filler and the interface developed between the filler and the matrix, which therefore controls the load transfer between filler and matrix. In considering the effect of filler addition it is the volume content of filler that is the most relevant, however the easiest way to measure the amount of filler added is by the weight content and conversion from weight content to volume content requires calculation using on the densities of the two phases. In the case of ceramic filled polymers the ceramic density normally is a factor of 3 to 6 higher than the polymer density so the difference in filler content between say 20wt% and 20vol% is substantial. We can easily consider the two most extreme cases, that is the stress throughout the composite is constant or the strain throughout the composite is constant. Considering first when the stress in the filler and matrix is constant. Here the simplest model is layers of filler and layers of matrix aligned perpendicular to the applied load, as shown in figure 4.1. The stress is F/a in all layers and the strain in each layer is σ/E. The deformation of each layer is σ/E × t (layer thickness). The constant stress model gives:

1 ⎛ vf v ⎞ = ⎜⎜ + m ⎟⎟ Ec ⎝ E f Em ⎠

eqn (4.1)

where Ec is the modulus of the composite, Ef and Em are the moduli of the filler and the matrix respectively and vf and vm are the volume fraction of the filler and the matrix respectively.

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Figure 4.1. Uniform stress model composite.

Figure 4.2. Uniform strain model composite.

For the uniform strain model the strains in the filler and matrix are the same and the simplest model is again sheets of filler and sheets of matrix in parallel, but now parallel with the load direction (figure 4.2). The strain is Flayer/tlayerElayer in each sheet and the force in each layer is tlayerElayer × ε while the constant strain model gives: Ec = vfEf + vmEm.

eqn (4.2)

These models are also known as the Reuss and Voigt models respectively. They are the simplest of the composite models, but all other models and experimental data give intermediate values to these two (figure 4.3). The earliest of these models was by Einstein [6], but the Einstein model only works at low filler content. A good review of theoretically and experimentally produced models is given by Bigg [7].

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Ef

75

Constant Strain or Voigt Model

Constant Stress or Reuss Model

Em 0

20

40

60

80

100

Volume % filler

Figure 4.3. Comparison the constant stress and constant strain models for composites, all composites fit between these two boundaries.

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IN VITRO AND IN VIVO TESTING OF COMPOSITES In vitro testing can be divided into three types, all of which are important when dealing with degradable composites. The first is degradation of the material due to exposure to a model physiological environment to follow the change in mechanical or other physical properties. The second is soaking the material in simulated body fluid, SBF, a solution developed by Kokubo et al. [8]. This solution is virtually chemically identical to blood plasma (table 4.1), but without cells or other biological entities. If a bone bioactive material is placed in this solution apatite crystals are deposited on the surface and the rate of deposition and thickness of the apatite layer developed is indicative of the bioactivity of the material. The third type of in vitro testing is cell culture to assess the likely biological response to the material, both as initially produced and after degradation. The first and third types are applicable to all degradable materials, but soaking in SBF is only applicable to the testing of potentially bone bioactive materials. Table 4.1 Comparison of the ion levels in Blood Plasma and SBF-K9 (from Kokubo et al. [8])

Blood Plasma SBF

Na+ 142.0 142.0

K+ 5.0 5.0

Ion Concentrations (mM) Ca2+ Mg2+ HCO32.5 1.5 27.0 2.5 1.5 4.2

Cl103.0 147.8

HPO421.0 1.0

SO40.5 0.5

In the early days of composite biomaterials production in vitro testing had not developed to the level of being able to be used for the analysis of the biological response to biomaterials and thus all testing was performed in vivo. Now in vivo testing normally follows extensive in vitro cell culture testing prior to clinical trials.

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MANUFACTURE OF COMPOSITES PHB/PHV BASED COMPOSITES Doyle et al. [3] compounded HA with polyhydroxybuterate (PHB) and then injection moulded plaques. The modulus of the unfilled PHB was 4.0GPa and increased to 11GPa with the addition of 40vol% HA, however the tensile strength dropped from 37MPa to 22MPa and the strain to failure dropped from over 2.5% to under 0.5%. Knowles and colleagues [4, 9-11] used PHB with 7mol% PHV and dry blended in four phosphate based glasses with different solubility rates at 20, 30 or 40wt% prior to injection moulding. Even after 500 days they found no reduction in the dry mass of the specimens of non-filled PHBV, but that the addition of the glasses produced mass loss within the first 100-200 days depending on the glass formulation. Furthermore they showed that the composites were piezoelectric, with higher piezoelectric signals from the more highly filled composites. Bergmann and Owen [5] room temperature mixed at 0 to 40wt% HA filler into PHB/PHV and then used compression moulding to produce plaques, so probably did not obtain a good distribution of the filler in the matrix. Their results showed that increasing the HA content did not increase the modulus as indicated by theory and decreased the strength and strain at failure as would be expected, thus showing the poor interface and probable particle-to-particle contact which occurred. Other more recent producers of PHB/PHV based composites include Coskun et al. [12] who used PHB either as a homopolymer or with 8mol% or 22mol% PHV and then reinforced these polymers with 5 or 15 wt% short rods of HA which were 2-4µm across and 20-30µm long. These materials were blended and the specimens were injection moulded and mechanically tested. As expected increasing the HA content increased the flexural and Young’s moduli and decreased the deformation at failure while increasing the PHV content decreased the moduli and increased the ductility. The ultimate strengths for the homopolymer and the 22% PHV increased with increasing HA content, but for the 8% PHV the maximum strength was obtained with 5wt% HA. It is interesting to note from SEMs of freeze fractured specimens that the rods shape of the HA have been retained and that the rods are approximately aligned with the injection direction (figure 4.4). The authors considered these materials to have mechanical properties which approached the required values for use in cortical bone with maximum flexural modulus of 5.39GPa and Young’s modulus of 2.21GPa obtained for PHB with 15wt% HA and maximum flexural strength of 78.2MPa obtained for the same material and maximum tensile strength of 23.4MPa for PHB/8%PHV with 5wt% HA, although these values are below those of cortical bone [13]. Chen et al. [14] used dynamic mechanical analysis (DMA) to measure the moduli of PHB12%PHV reinforced with 0, 3.2, 9.1 and 14.2vol% nanoHA, the composite was produced by solution casting with ultrasonication used to distribute the filler particles. Scanning electron microscopy (SEM) showed that the nanoHA particles were well distributed. DMA showed that the storage modulus increased from 3.5 to 7.0GPa and the loss modulus also increased with increasing HA content and that, depending on filler content and test frequency, the glass transition temperature varied between 12 and 26˚C, thus always below body temperature. They soaked the 14.2vol% HA composite in simulated body fluid (SBF [8]) for 30 and 50 days and saw the deposition of an apatite layer, indicating the potential bioactivity of the composite.

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Figure 4.4. Scanning electron micrograph of PHB8%PHV co-polymer reinforced with 15wt% hydroxyapatite rods (reproduced from Ciskun et al. [12]).

Misra et al. [15] reinforced PHB with 5 or 20wt% Bioglass®45S5 using solvent casting. Surprisingly, the addition of the Bioglass® powder reduced the tensile modulus although in dynamic mechanical analysis (DMA) the storage and loss moduli both increased. The authors suggest that the difference may be due to de-wetting of the polymer/glass interface or agglomeration of the filler particles leading to premature failure of the polymer/glass interface at the higher stress levels used in the tensile testing compared to the lower stress levels used in DMA testing. Soaking in SBF indicated that the composites would be bioactive. Wang et al. [16] used PHBHHx a co-polymer of polyhydroxybuterate with 12% poly hydroxyhexanoate (PHHx) and reinforced it with 10% HA. The addition of the PHHx copolymer substantially decreased the stiffness and strength, although it increased the biological response to the polymer. However, while the addition of HA to PHB increased the cellular response the addition of HA to PHBHHx reduced the cellular response.

POLYGLYCOLIC ACID (PGA) AND POLYLACTIC ACID (PLA) BASED COMPOSITES Shikinami and Okuno [17] synthesised their own hydroxyapatite particles and then mixed them into poly-L-lactide before processing the resultant composite via extrusion and forging before finally machining into devices (figure 4.5). Increasing the HA content increased the moduli up to the maximum filler level of 50wt% (32.0vol%) HA, but the highest bending, tensile and impact strengths were obtained for 30wt% (16.7vol%) composite. The bending strength gradually decreased with time of soaking in buffered saline and there was no much difference in the strengths after 52 weeks soaking all having dropped from 200-250MPa

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down to approximately 120MPa. Thus unlike most of the other composites discussed in this chapter the strengths were above those of cortical bone, while the stiffness was within the lower bounds of reported values for cortical bone [13].

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Figure 4.5 Screws, washers, a pin, and plates made of forged u-HA/PLLA composites (from Shikinami and Okuno [17]).

Bleach and colleagues [18-20] manufactured triphasic composites of drawn PLA96 fibres in a PLDLA (70L:30DL) matrix which was itself reinforced with hydroxyapatite or tricalcium phosphate (TCP). The drawn PLA fibres were used to increase the stiffness and strength of the composite and produce a flexural stiffness of 6GPa without the additional HA reinforcement and over 7GPa with the calcium phosphates, which were only about 12wt% in the overall composite. The flexural strengths were between 65 and 80MPa and for the PLA only composites started to fall after 8 weeks soaking in SBF while with the calcium phosphates reinforced plates the mechanical properties started to drop after 12 weeks. The presence of the calcium phosphates also increased the bioactivity of the materials when assessed by soaking in SBF. Rich et al. [21] investigated the effect of particle size on the properties of P(CL/DLLA) filled with between 40 and 70wt% of a bioglass. They used DMA to measure the mechanical properties and found that using particles below 45µm compared to 90-315µm lead to higher stiffness as did increasing the filler content upto 60wt%. However, the smaller size filler lead to increased water absorption and degradation rates (figure 4.6). Wright-Charlesworth et al. [22] produced 0 to 40wt% HA in PLLA composites by compounding-extrusion followed by injection moulding. The addition of HA, as expected, increased the stiffness, but also decreased the strength of the PLLA with less drawing of the polymer matrix between the particles. Furthermore after soaking in PBS the increasing the HA content the decreased the degradation rate. As with most such degradable composites the strength dropped faster than the stiffness for all compositions.

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Figure 4.6 Scanning electron micrographs of the fracture surfaces of composite containing 60wt% of the bioactive glass (particle size 6 months) and D,L-PLA (degradation time: < 6 months) cylinders were implanted in the femur of the rabbits. It was found that while D,L-PLA degraded faster than L-PLA, significant differences in bone growth, repair, and remodeling were not observed [47].

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4.2. Enzymatic Degradation: As in chemical hydrolysis, enzymatic degradation of PLA devices contributes to the loss in molecular mass and mechanical strength. Enzymatic degradation occurs when the surfacebinding domain of enzyme adsorbs onto the polymer and initiates hydrolysis of the ester bond [48]. It has been reported that PLA-degrading enzymes selectively cleave the α-ester bond of the L-isomer of PLA. The mechanism of enzymatic degradation may be explained by the fact that PLA-degrading enzymes also degrade silk fibroin. Owing to the structural similarity of the L-lactic acid unit of PLA and L-alanine of silk fibroin, the enzyme recognizes the L-lactic acid of PLA as L-alanine resulting in enzymatic cleavage of the ester bond. Among various classes of enzymes that have been identified as capable of PLA hydrolysis, proteases, depolymerases and lipases have been reviewed [48-50]. A summary of the most important conclusions are given below. Proteinase K was the first enzyme reported to hydrolyse PLA and has since been extensively used to investigate enzymatic degradation of a variety of PLA devices [48]. After adsorption to the polymer surface followed by enzymatic degradation, proteinase K strongly moves to the adjacent surface of the polymer, initiating hydrolysis. The interaction between the enzyme and PLA is irreversible [51]. Among the different strains of proteinase K, acid and neutral proteases significantly degrade L-PLA. In addition to proteinase K, other serine proteases such as, trypsin, elastase, and subtilisin degrade L-PLA. In general, lipases cleave low molecular weight and racemic PLA only although at high temperature and alkaline pH, commercially-available Lipase PL completely degraded PLA [52]. Although a number of microbes degrade PLA, the isolation, purification, and characterization of the enzymes that degrade PLA from these organisms has not been extensively investigated. Recently, L-PLA depolymerase, a protease enzyme, has been isolated and purified from Amycolatopsis sp and demonstrated optimal activity at a pH of 6.0 and between 37 - 45oC [53]. PLA is also degraded by several commercially-available proteases that are mainly alkaline proteases derived from Bacillus sp. Several enzymes isolated from yeast have also demonstrated significant PLA-degrading activity. Recent experiments proved that enzymatic hydrolysis of PLA occurs both by exo- and endo-chain scission. Similar to chemical hydrolysis, the important factors that affect the enzymatic hydrolysis of PLA are the molecular weight, composition, crystallinity and, purity [54]. An increase in molecular weight of L-PLA increases the rate of enzymatic hydrolysis.

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The effect of temperature (> Tg) on the rate of enzymatic hydrolysis of PLA has not been investigated. While enzymatic degradation of gelatin-coated PLA nanoparticles was minimal in gastric fluid, they were rapidly hydrolyzed into lactate in pancreatic fluid, emphasizing the role of pH on degradation [55]. Proteases are stereospecific and hydrolyze L-PLA and D,L-PLA but not D-PLA. Stereocomplexes formed between L- and D-PLA resist enzymatic hydrolysis because the enzyme is less able to penetrate into the matrix [56]. The rate and extent of enzymatic hydrolysis decrease as the crystallinity and melting temperature of PLA increase. Enzymatic hydrolysis in PLA crystals preferentially occurs at the disordered chain-packing regions without affecting the folded chains. The rate of enzymatic hydrolysis is also affected by the tacticity of PLA chains. In the amorphous regions, proteinase K-catalyzed hydrolysis occurs at the free ends of the polymer chain. Atomic Force Microscopy identified two distinct amorphous regions in PLA where enzymatic hydrolysis was faster in the free amorphous region compared to the restricted amorphous region [57]. Adsorption of water molecules by PLA films increases the molecular mobility of the polymer chains in the amorphous region, enhancing Proteinase Kinduced enzymatic hydrolysis [58]. In blends of PLA with other polymers, the rate of enzymatic degradation is determined by the ratio of crystalline and amorphous regions [54].

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4.3. Microbial Degradation: Although a number of PLA-degrading bacteria have been identified, they are primarily distributed in the soil and therefore are irrelevant to human application. Microbial degradation of PLA has been extensively reviewed elsewhere [48, 50]. However, to study the mechanism of microbial degradation several enzymes have been isolated and purified. Most of the microbial strains that degrade PLA also cleave silk fibroin suggesting ester bond cleavage. Degradation of PLA by the fungus, Tritirachium album was significantly enhanced by gelatin. In addition to gelatin, several amino acids, peptides, and poly (L-amino acids) such as, silk fibroin, elastin, keratin, and collagen enhance the microbial degradation of PLA.

5. SURFACE MODIFICATION OF PLA Although PLA is the material of choice for several tissue engineering applications, the degradation time of the polymer needs to be tailored for individual needs. PLA can be structurally modified in bulk or on its surface using other polymers to prepare composites and blends with well-defined degradation times. For example, PLA can be stabilized against degradation by grafting other polymers in high density or by introducing terminal groups such as methyl groups on the side chains. The bulk and surface modifications of PLA have been reviewed [59]. Briefly, methods to modify PLA include copolymerization of lactide with other lactone monomers such as glycolic acid and ε-caprolactone. The composite polymer obtained will have a slower rate of degradation, reduced crystallinity, and lower mechanical strength compared to PLA. Multi-block co-polymers of PLA and PEG with increased hydrophilicity may be synthesized. The degradation profile of PLA-PEG-ε-caprolactone

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polymer blends can be tailored by altering the co-polymer ratio. Various other polymer blends including PLA dendrimers, PLA-dextran, and PLA-chitosan have been prepared [59]. Similarly, modification of the surface of PLA by proteins such as fibronectin, vitronectin, collagen, and gelatin enhances the adhesion of the cells to the surface of the polymer scaffolds. Additionally, alkali hydrolysis treatment of PLA enables immobilization of peptides/ligands onto the PLA scaffold. Extensive research has been performed on the degradation of these polymer blends. However, the mechanism of degradation is similar to that of PLA.

CONCLUSION

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The review focuses on the degradation of PLA, its mechanism, degradation of PLA scaffolds, risks of autocatalysis, and methods of controlling the extent of autocatalysis. PLA is an excellent biocompatible material for implants. The degradation of PLA is influenced by a number of factors that can be classified as polymer properties (method of polymerization, crystallinity, purity, tacticity), environmental factors (temperature, ionic strength, buffer concentration, pH), and device–related factors (geometry, method of preparation). A general scheme for the mechanism of degradation of the polymer has been outlined. The rate of degradation of PLA influences the mechanical integrity of the implant, and ultimately the biological response such as cell adhesion and bone formation. In addition, surface and bulk modification of PLA alters the degradation of the polymer, resulting in tailored biomaterials for specific applications. Although significant literature exists on the degradation mechanism of these polymer blends, this review exclusively focuses on the degradation of PLA.

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[44] Leenslag JW, Pennings AJ, Bos RR, Rozema FR, Boering G. Resorbable materials of poly(L-lactide). VII. In vivo and in vitro degradation. Biomaterials 1987;8(4):311-4. [45] Pistner H, Stallforth H, Gutwald R, Muhling J, Reuther J, Michel C. Poly(L-lactide): a long-term degradation study in vivo. Part II: Physico-mechanical behaviour of implants. Biomaterials 1994;15(6):439-50. [46] Tschakaloff A, Losken HW, von Oepen R, Michaeli W, Moritz O, Mooney MP, et al. Degradation kinetics of biodegradable DL-polylactic acid biodegradable implants depending on the site of implantation. Int. J. Oral. Maxillofac. Surg. 1994;23(6 Pt 2):443-5. [47] Merolli A, Gabbi C, Cacchioli A, Ragionieri L, Caruso L, Giannotta L, et al. Bone response to polymers based on poly-lactic acid and having different degradation times. J. Mater. Sci. Mater. Med. 2001;12(9):775-8. [48] Tokiwa Y, Calabia BP. Biodegradability and biodegradation of poly(lactide). Appl. Microbiol. Biotechnol. 2006;72(2):244-51. [49] Williams DF. Enzymic hydrolysis of polylactic acid. Eng. Med. 1981;10:5-7. [50] Tokiwa Y, Jarerat A. Biodegradation of poly(L-lactide). Biotechnol. Lett. 2004;26(10):771-7. [51] Yamashita K, Kikkawa Y, Kurokawa K, Doi Y. Enzymatic degradation of poly(Llactide) film by proteinase K: quartz crystal microbalance and atomic force microscopy study. Biomacromolecules 2005;6(2):850-7. [52] Hoshino A, Isono Y. Degradation of aliphatic polyester films by commercially available lipases with special reference to rapid and complete degradation of poly(Llactide) film by lipase PL derived from Alcaligenes sp. Biodegradation 2002;13(2):1417. [53] Pranamuda H, Tokiwa Y, Tanaka H. Polylactide Degradation by an Amycolatopsis sp. Appl. Environ. Microbiol. 1997;63(4):1637-1640. [54] Tsuji H, Miyauchi S. Enzymatic hydrolysis of poly(lactide)s: effects of molecular weight, L-lactide content, and enantiomeric and diastereoisomeric polymer blending. Biomacromolecules 2001;2(2):597-604. [55] Landry FB, Bazile DV, Spenlehauer G, Veillard M, Kreuter J. Degradation of poly(D,L-lactic acid) nanoparticles coated with albumin in model digestive fluids (USP XXII). Biomaterials 1996;17(7):715-23. [56] Lee WK, Iwata T, Gardella JA, Jr. Hydrolytic behavior of enantiomeric poly(lactide) mixed monolayer films at the air/water interface: stereocomplexation effects. Langmuir 2005;21(24):11180-4. [57] Kikkawa Y, Abe H, Iwata T, Inoue Y, Doi Y. Crystallization, stability, and enzymatic degradation of poly(L-lactide) thin film. Biomacromolecules 2002;3(2):350-6. [58] Kikkawa Y, Fujita M, Abe H, Doi Y. Effect of water on the surface molecular mobility of poly(lactide) thin films: an atomic force microscopy study. Biomacromolecules 2004;5(4):1187-93. [59] Wang S, Cui W, Bei J. Bulk and surface modifications of polylactide. Anal. Bioanal. Chem. 2005;381(3):547-56.

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In: Degradable Polymers for Skeletal Implants Editors: P.I.J.M. Wuisman and T. H. Smit

ISBN 978-1- 60692-426-6 © 2009 Nova Science Publishers, Inc.

Chapter 9

BIOLOGICAL TESTING OF DEGRADABLE POLYMERS IN VIVO O.M. Böstman∗ Dept. of Orthopaedics and Trauma Surgery, Helsinki University Hospital, Finland

ABSTRACT

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The advantages of degradable skeletal implants over metallic ones include avoidance of hardware removal procedures and absence of the long-term problems associated with retained metallic fixation devices such as corrosion and stress-protection induced osteopenia. However, with modern surgical technique and first-class metallic implants removal operations are unusual, and corrosion seldom constitutes a problem. Instead, degradable implants may offer other advantages. They are radiolucent and do not interfere with magnetic resonance imaging either. Pre-clinical testing of new surgical appliances is essential to rule out major biological hazards. Firstly, the implants must be free of toxic, immunological, cancerogenic and teratogenic risks. Then the mechanical properties have to be evaluated. A long degradation time of several years may pose its own difficulties as the life span of only few test animals exceed five years. Moreover, findings in other species may not be directly applicable to humans. Consequently, carefully planned clinical trials must be regarded as part of the biological testing of the degradable devices, like any other novel surgical implants. In the experimental and clinical studies performed, the only biological adverse effect of importance so far encountered is the occurrence of inflammatory foreign-body reactions to the degrading polymers, especially those with a short degradation time. The mechanical reliability and strength retention of the implants seems not to be associated with significant problems. Histomorphometric studies have shown that polyglycolide disappears from the tissues within six months, whereas large amounts of high-molecularweight polylactide may retain within cancellous bone still five years after implantation. Consequently, research and development is being focused on implants made of polymers having a medium-long duration time such as stereoisomeric copolymers of polylactide.

∗

Address: Dr. O. Böstman, Tiirasaarent. 24, FIN-00200 Helsinki, Finland, E-mail: [email protected].

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O.M. Böstman To improve the understanding of all the factors influencing the behaviour of degradable materials in the tissues systematic studies are required.

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INTRODUCTION When a fracture, osteotomy, arthrodesis, osteochondral defect or ligamentous injury necessitates the use of internal skeletal fixation or supporting devices, the implants are needless and may be harmful as soon as healing of the tissues has taken place. In contrast to metallic devices, degradable implants leave no hardware retained in the tissues. Even if modern metallic appliances do not often require removal in adult patients, they are associated with some other well-known disadvantages, such as excessive rigidity resulting in stressprotection osteopenia, and corrosion. The psychological advantages of avoiding implant removal procedures would seem to be of particular value in children. These facts constitute the incitement behind the efforts made during the past three decades to develop degradable implants for skeletal fixation procedures [1-10]. In this context, the terms absorbable, resorbable, degradable and biodegradable are used interchangeably. Avoiding the need of considering implant removal procedures and the long-term problems of retained hardware are not the only advantages of degradable skeletal implants. Biodegradability would be a desirable feature of a fixation pin or screw, when the implant is to be inserted through articular surfaces, as persistent metallic devices may disturb joint function. Such an intra-articular fixation of displaced osteochondral fractures is sometimes necessary, especially in the knee and elbow joints [11-13]. Degradable scaffolds can also be useful in attempting to restore osteochondral defects [14]. In endoprosthetic and spinal surgery it would be of clinical importance if at least one of the auxiliary implant components would be totally radiolucent [15-17]. Many organic macromolecular compounds are degradable and resorbable in living tissues [18], but rather few possess the chemical and physical properties required to be processed into internal surgical fixation devices. The most commonly used are polyglycolic acid (PGA), polylactic acid (PLA) and their derivatives. PLA is used in several stereoisomeric forms such as poly-levo-lactide (PLLA) and poly-dextro-levo-lactide (PDLLA). Because of their thermal behaviour, several of the synthetic biodegradable polymers are difficult to shape into complex designs such as screws and plates. To give an example, polylactide in general is relatively brittle and rigid in room temperature. Moreover, the demands made by many clinical applications on the initial mechanical strength of the devices set their limitations to the geometric shaping of the implants. Initially, the degradable implants tested and studied were prepared by casting the polymers into sheets or films which permitted basic investigations on the biological behavior of the compounds in bone tissue but were not suitable for fracture fixation. Fabrication of implants subsequently was accomplished by melt molding and extrusion into pins and rods. The production of screws and small plates became possible later, but the mechanical strength of the first generation screws was modest. The initial strength and the strength retention characteristics of rods and screws were improved by the introduction of a fibre-reinforced composite texture in which the polymer matrix is reinforced with suture fibres of the same material [19].

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The assortment of absorbable internal fixation implants from different manufacturers that is commercially available for orthopaedic applications today includes cylindrical pins, rods, screws, small plates, plugs, staples, anchors, tacks, cords and cages. The pins measure from 1.0 to 4.5 mm in diameter and up to 70 mm in length. The screws have a core diameter from 2.0 to 4.5 mm. Also partially threaded lag screws are available. The screw profiles are best suited for fixation of cancellous bone fragments. Small plates may be used in craniofacial surgery and other small-fragment fixations. The staples, anchors, arrows, tacks and 2.0-mmcords are mainly intended to fix or reconstruct ligaments, meniscal tissue and joint capsules. Degradable cages have been developed for spinal surgery. Also degradable synthetic carriers of bone morphogenetic protein and antimicrobial agents as well as chondrocyte-seeded implants for cartilage repair and scaffolds for the regeneration of nerves and vessels are currently subject of research and experimentation. All surgical implants, metallic as well as degradable, must be free of certain biological hazards (Table 1) throughout the degradation and depolymerisation process. The degradation time of different degradable polymers varies greatly, from a couple of months to five years or more. A question of importance for both the fast- and slow-degrading implants is, whether the normal tissue elements will be restored within the implant track. The quality of the tissues replacing the internal fixation devices after degradation is clinically relevant especially when multiple screws and other space-occupying intraosseous implants are used. Thorough pre-clinical in vivo evaluation of skeletal degradable devices is essential [20]. However, long-term experimental studies in order to investigate the ultimate biodegradation of implants made of slow-degrading polymers are difficult to design because of the fairly short life span of most laboratory animals. Follow-up times of four to five years are the upper limit for most of the commonly used animals. All the potential hazards associated with surgical innovations are not necessarily recognisable in animal experiments and not even during the first years after the clinical introduction of a new method but first several years later. Therefore the clinical introduction of degradable implants in new applications must, to some extent, be regarded as being part of the biological testing of the polymers in vivo. Indeed, the only way to proceed after animal experiments is often a judicious clinical trial. Table 1. Potential biological hazards of degradable implants Toxic Immunological Infectious (contamination) Inlammatory Metabolic (interference) Cancerogenic Teratogenic

DEGRADATION BEHAVIOUR IN VIVO The degradation behaviour of these implants cannot be directly and systemically studied in humans. The species used in animal experiments include rats, rabbits, goats and sheep. As for the

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study design, simple implantation has been the most common method, but also bone defects and osteotomies mimicking fractures have been created (Figure 1). Most studies have been concerned with cancellous or membranous bone only. The evaluation methods include physical macroscopic examination, conventional radiography, microradiography, qualitative histology, histomorphometry, oxytetracycline fluorescence measurement, transmission electron microscopy and molecular-weight analysis.

Figure 1. A microradiograph showing a transcondylar osteotomy mimicking a cancellous bone fracture in the distal rabbit femur that has been fixed with a degradable screw (asterisk). (Previously unpublished.)

Already in the 1970s and early 1980s it was demonstrated by several investigators that implants made of PGA, PLA and their copolymers are completely absorbable within cancellous bone tissue. The degradation of these polymers occurs mainly by hydrolytic scission and to a lesser extent through non-specific enzymatic action, the main route of final elimination being respiration. The differences in the final metabolism of the polymers are relatively small but the depolymerisation rates vary. High-molecular-weight PLLA has the longest degradation time, with a half-life clearly exceeding one year, while the degradation

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rates of PDLLA, PGA, PGA:PLA copolymers and polydioxanone are more rapid [21-25]. A myriad of implant-related factors influence and modify the degradation behaviour of skeletal degradable devices. These include molecular weight, inherent viscosity, crystallinity, thermal history, geometry (weight to surface area ratio) and porosity of the implant, presence of residual impurities and oligomers in the polymer, and the fabrication and sterilisation processes used. Implants from different sources often differ considerably from each other in their physicochemical characteristics. Understandably, the manufacturers are not often able or willing to supply all these data of their implants, and it is not easy to find a report in the open literature having all these details recorded. This renders a comparison of the findings reported in different studies difficult. A porous thin sheet depolymerises in general more rapidly than a dense compact block. Yet in bulky implants made of PLA a process of autocatalysis may under some circumstances induce rapid degradation in the core area of the device. The complexity of the depolymerisation behaviour of biodegradable screws was elucidated in a study, in which a major difference in the hydrolysis rate was found between PLA screws obtained using two different initiator systems, zinc or tin derivatives, in the polymerisation of lactide monomers [26]. In addition, the histological findings of that study seemed to confirm the hypothesis of a heterogeneous degradation mechanism of PLLA [27]. Degradation does not, however, imply immediate absorption of an implant. In an early experimental study using pellets of carbon-14 and tritium labeled PGA implanted in rat tibiae, approximately 70 % of the implant material remained at the site after three months [21]. As seen in experimental studies by light microscopy, PGA totally disappears from the cancellous bone of the distal rabbit femur within 36 weeks [22]. In one study, high-strength PLLA rods implanted in the medullary cavity of rabbit femur had resorbed completely by 62 months but only partially at 42 months after implantation [28]. In another study, PLLA was detected five years after fixation of a mandibular osteotomy with a PLLA plate [29]. Consequently, so far it has not been possible to determine the rate of degradation of highmolecular-weight PLLA implants precisely. The same applies to the degree of the tissue restoration accompanying the vanishing of the polymer. In a long-term 4.5-year follow-up study on PLLA screw implanted in the cancellous bone of distal rabbit femur, the tissue replacement was consistently poor, indicating that normal bony restoration would not take place [30]. In that study, as much as one half of the PLLA material implanted was left within the screw track after a follow-up time of 54 months (Figure 2). As for the spatial pattern of the degradation process, an interesting finding was the occurrence of a secondary new-bone front within the implant cavity around the core area of the PLLA screw. Such a walling-off compact newbone front possibly lengthens the degradation and absorption processes further by interfering with the centripetal tissue replacement within the implant cavity. High-molecular-weight PLLA implant used for experimental orbital floor defect in one goat was not completely degraded in 1.5 years [31]. It was concluded that full resorption would take 3.5 years. In another study, after subcutaneous implantation of high-molecular-weight PLLA plates and screws, complete degradation of the implants had not taken place within 2.8 years in the only rat alive in the study at that time [32]. Furthermore, the tissue milieu probably plays an important role, an implant being more efficiently depolymerised by richly vascularised cancellous bone than by dense avascular connective tissue. Thus an estimate of the ultimate degradation time of an individual implant in a clinical situation can only be given in rough terms. In animal experiments, the polymer is gradually seen to become invaded by connective tissue as the degradation proceeds. The degree

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of final restoration of the original tissue architecture within the implant track varies greatly of reasons not yet fully understood. After insertion of an absorbable pin or screw made of PGA through an articular surface, the articular cartilage and the subchondral bony architecture may show complete restitution (Figure 3), but occasionally the implant is replaced by loose connective tissue only [33].

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Figure 2. A coronal plane photomicrograph showing the restoration of subchondral bone at the intraarticular entrance (arrow) of a PGA screw into the distal rabbit femur 36 weeks after the implantation. Stain: Masson-Goldner trichrome. (Previously unpublished.)

Figure 3. A transverse plane photomicrograph showing the abundant presence of PLLA within the implant cavity (asterisk) 54 months after the implantation of a PLLA screw into the distal rabbit femur. The cavity is walled-off by a new-bone front (arrow). Stain: Masson-Goldner trichrome. (Previously unpublished.)

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A concentric tissue invasion from the periphery towards the centre of a degradable implant seems to be a constant finding in the specimens. No migration of PLLA particles outside the implant track into the neighbouring tissues has been reported, but it has been observed for PGA during its degradation [34]. The hydrophobicity of PLLA due to its methyl groups could be an explanation, PGA being a more hydrophilic compound. A transmission electron microscopy study showed that PLLA degrades by disintegrating into polygonal particles, 10 x 20 μm in size, digested by mononuclear phagocytic cells [35]. The life span of test animals sets its limitations to studies on the tissue behaviour of slowdegrading polymers such as PLLA. Although the fracture fixation devices made of PLLA by different manufacturers constitute a heterogeneous group of implants, it can be stated that the ultimate degradation time of high-molecular-weight PLLA within bone tissue in general is so far known only approximately. It has been hypothesised that PLLA particles or macrophages with PLLA particles, could migrate to the lymph nodal tissue from the implantation site , but this possibility has only seldom been taken into consideration in the study designs. In one systematic study on rabbits, no birefringent material could be found in the lymph nodes three years after intraosseous implantation of PLLA screws [35]. This supports the concept of a local degradation to molecular level and absorption via blood rather than through the lymphatic system. Only sporadic data on the degradation of implants made of high-molecular-weight PLLA in humans is available. Macroscopic remnants of devices made of PLLA in clinical use have been found to be present in the ankle after four years [36] and in the maxillo-facial region still after five years [37]. Attempts have been made to assess the degradation indirectly by magnetic resonance imaging [38] but such a method is able to supply only approximate information.

TISSUE RESPONSE AND RESTORATION At the time writing, degradable skeletal implants have not been found to be associated with toxic, teratogenic or cancerogenic hazards. The tumorigenicity of PLLA in rodents seems not to differ from that of polyethylene [39]. However, there are other concerns (Table 2). The results of biocompatibility studies in test animals, however, cannot be directly extrapolated to humans. The only method-specific adverse reaction so far recorded for these implants is a local inflammatory foreign-body reaction. Such reactions have been recorded in both experimental studies and clinical trials. There are several animal studies reporting signs of an inflammatory foreign body reaction [31,32,34,40,41]. In clinical series, it has occurred in 2% to 25% of the patients depending on the type of implant used and the anatomical site operated on [42-46]. As for the histological characteristics of the tissue response, it seems to be a non-allergic, as such non-specific foreign-body reaction composed mainly of multinucleate giant cells and a few inflammatory cell elements. The relatively sudden accumulation of liquid degradation products of the polymer within a cavity surrounded by unyielding trabecular bone makes great demands on the clearing capacity of the tissues. If this capacity is exceeded, the polymeric debris is retained, and an inflammatory foreign-body reaction ensues.

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O.M. Böstman Table 2. Potential concerns regarding the tissue response to degradable implants

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Non-specific local irritation Inflammatory foreign-body reaction Dissemination of polymeric debris in the lymphatic system Osteoinhibition, osteolysis Excessive osteostimulation Degeneration of articular cartilage Induction of osteoarthritis Device retention due to protracted degradation Poor tissue replacement

The foreign-body reaction obviously is an inevitable tissue response to the polymers, while the inter-individual variation of the intensity of the reaction is dependent on factors that influence the debris-clearing capacity of the bone section concerned. In an in vitro study, it was found that one of the possible factors responsible for the adverse tissue effects, the decrease of the pH values at degrading alpha-hydroxy-polyester implants, can be offset by incorporation of basic calcium carbonate within the implants [47]. The tissue response may also be modulated by antibodies against macrophage-activating cytokines within the polymer [48]. Besides different physicochemical characteristics of different polymers, the implantation site also has an effect on the biocompatibility and possible adverse tissue responses. Intraosseously placed PLLA screws have elicited less tissue reactions than PLLA plates placed on the bone surfaces [46]. In the distal rabbit femur, around an intra-articular entrance of PLLA pins and screws, the pre-existing cartilage showed degenerative changes within a 400 μm wide zone as measured from the tissue-implant boundary [49]. The width of this zone of detrimental influence, 400 μm, was nearly 20 % of the radius of the 4.5-mm-diameter implant. The screw-implanted specimens showed a more active osteoid formation at the tissue-implant boundary than the pin-implanted ones. This was probably due to the larger surface-area of the screws, approximately 1.7 times that of the pin. As a whole, the new-bone response to PLLA at the tissue-implant interface was less conspicuous than that previously reported for PGA implants under similar experimental conditions [33]. The foreign-body reactions are sometimes accompanied by osteolytic lesions, well-known from experimental studies [34,39] as well as clinical studies [50-53]. The follow-up times of the vast majority of the experimental and clinical studies have been too short to demonstrate the ultimate behaviour and regression potential or fate of the osteolysis. In one long-term clinical study [54] ankle fracture patients with extensive osteolytic reactions to PGA screws were prone to develop osteoarthritis later on. This indicates that the inflammatory foreign-body reactions may not always be of transient nature and free of permanent adverse consequences. Implants made of PLA have proved to be biologically more inert than those made of PGA to elicit inflammatory reactions, which already in the early 1990s resulted in a shift from PGA to PLA as the principal polymer utilised in the development and production of the degradable devices. However, it is obvious that such reactions do not occur until the degradation has reached its final phase. Obviously, considering the ultimate biocompatibility of high-molecular-weight PLLA, little can be concluded from studies having follow-up times shorter than four years. Experimental laboratory in vivo studies with longer follow-up times

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are seldom feasible in larger scale. Clinical trials might be a more practical way to obtain relevant data. Indeed, a few clinical studies have already demonstrated that foreign-body reactions of clinical significance to PLLA implants can occur four years or more after an internal fixation procedure of a fracture [55-58]. The quality and quantity of the reparative tissue that replaces a degrading skeletal implant is of importance in particular when the entrance of the device in the bone is intra-articular and when large implants are used. As the restoration pattern of articular cartilage after implantation of fast-degrading polyglycolide implants was studied, proper new subchondral bone formation and restoration of the trabecular bone architecture at the level of the original chondro-osseous junction seemed to be a prerequisite for the regeneration of normal or nearnormal articular cartilage [33]. Consequently, well-differentiated hyaline cartilage restoration after intra-articular insertion of PLLA implants cannot be expected to take place until the degradation and tissue replacement of the polymer have progressed far enough. With highmolecular-weight PLLA this can take six or seven years. Meanwhile, the articular-surface defect may have resulted in degenerative changes in the vicinity of the implant channel. As already was discussed earlier in this chapter, long-term experimental findings do not support the idea that the degradation of the implants would be followed by proper restoration and normalisation of the tissues. However, the size of the implants in relation to the dimensions of the recipient bone section probably play a role in the regeneration potential. During the early years of bioabsorbable fracture fixation, it was speculated that these implants might possess some osteostimulatory potential and promote bone healing. Certain experiments have suggested an osteoconductive potential of absorbable polyester implants. Osseous defects in rat tibiae filled with a 50:50 PGA:PLA copolymer implant displayed an accelerated rate of healing compared to empty control defects [59]. Similar results were obtained when the copolymer implant was combined with allogenic demineralised freezedried bone powder [60]. In more recent studies, no evidence for an osteostimulatory effect of PGA screws could be found as compared with the tissue restoration within the empty tapped drill-holes. On the contrary, PGA seemed to be actually osteoinhibitory, the amount of trabecular bone being less in the PGA implanted specimens [30]. The scarce osteoid formation fraction in the 36-month specimens of that study indicates that the tissue restoration process already had ceased at that time. In another study on rabbit calvarial bone defects, PGA membranes were found to retard osteogenesis [61]. The intensity of the osteolytic foreign-body reactions to these implants and the degree of restoration of the tissues seem to be influenced by some factors so far unidentified. Consequently, the risk of a severe reaction and detrimental late tissue effects cannot be predicted in an individual patient.

MECHANICAL STRENGTH The mechanical properties of absorbable implants must be discussed in terms of initial strength, strength retention during degradation and elasticity. The failure behaviour of absorbable polymers is influenced by their glass-transition temperature, at which the polymer becomes brittle and rigid. This ranges from 58 degrees Celsius for PLLA to -16 degrees for polyparadioxanone. The initial strength is influenced more by the manufacturing technique of

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the implant than by the polymer utilised. Simple melt-moulding or extrusion of synthetic biodegradable polymers into pins result in weaker implants than certain special manufacturing techniques. A special fibre-reinforcing technique using high pressures and temperatures gives high-strength composite implants with fibres embedded within a matrix of the same biodegradable polymer [19]. The initial bending strength of a 4.5-mm-diameter PLLA pin manufactured by a fibre-reinforcing technique has been recorded to be 245 MPa [62]. For pins of PGA the initial strength is even higher, 350 MPa [63], close to that of stainless steel. On the whole, however, the mechanical properties of absorbable materials are very different from those of stainless steel. Indeed, a direct comparison seems meaningless, since the absorbable implants have not been developed to mimic the metallic ones. The strength of a polyester implant declines before a macroscopic degradation commences. When placed in the subcutaneous tissue of rabbit, PGA rods with an initial bending strength of 350 MPa lost more than 50 % of their bending strength within two weeks [63]. PLLA rods with an initial bending strength of 210 MPa lost 50 % of their bending strength within twelve weeks [64]. To delay the rapid degradation and loss of strength of PGA implants the effect of various coatings acting as hydrolysis barriers has been investigated. The strength retention of an absorbable implant is determined by its degradation rate. As already was discussed in this chapter, the degradation rate is influenced by micro- and macrostructural properties of the implant as well as by environmental factors. In living tissues, absorbable fixation devices lose most of their mechanical strength long before the implant loses its gross appearance and the actual decomposition starts. The fixation properties of pins, rods, screws and plates of alpha-hydroxy polyesters have been studied in animal experiments under conditions simulating fractures of cancellous or cortical bone [65-67], and epiphyseal separation in immature animals [68]. With the exception of one study [68] in which plates and screws of PLA failed to stabilise an osteotomy of a canine radial shaft, all these investigations demonstrated that the polyester implants have adequately fixed the bony fragments. In human cadaver studies satisfactory fixation properties have been reported with 2.0 millimeter PGA rods in osteotomies of the distal radius [69]. In another study on the distal radius, satisfactory fixation was reported with 2.7-mm-diameter PLA rods [70]. In a bovine experimental model, a 6.3-mm-diameter PLLA screw was found to be as good as a conventional metal screw to fix a bone-patellar tendon-bone graft for the anterior cruciate ligament [71]. It has to be emphasised that cadaver testing may not give a reliable picture of the potential of an implant in a certain clinical situation. Certain mechanical properties of biodegradable implants may be expressed in terms of Young's elasticity modulus. The elasticity modulus of absorbable polymers is much less than that of stainless steel. Instead, it is close to that of cortical bone and only slightly higher than that of cancellous bone. The ultimate mode of failure of an implant can be ductile or brittle. From a clinical point of view, a ductile mode of failure is desirable, since a brittle breakage of an implant before union of a fracture usually results in immediate redisplacement between the fragments. Due to the ductile mode of failure of fibre-reinforced degradable implants, they lose their shear strength more slowly than their bending strength. Therefore angulation instead of lateral displacement is likely to occur between fixed fracture fragments in case of moderate overloading. Nevertheless, complete redisplacements are seen when the forces acting on the

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fracture clearly exceed the mechanical capability of the implants. This may occur due to the fracture type as in supracondylar fractures of the humerus [72], or due to decreased compliance of the patient because of alcoholism or other mental disorder [73,74]. With adequate patient co-operation, the mechanical reliability of the degradable fixation devices is not a problem e.g. in common injuries such as malleolar fractures and syndesmotic disruptions of the ankle [6,75].

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CONCLUSION With an extensive experience accumulated during more than 20 years, it seems unlikely that PGA, PLA or other degradable polymers closely related to them would possess biological hazards so far unknown. Occasional inflammatory foreign-body reactions with local osteolytic lesions and poor tissue restoration are the only significant drawbacks. Experimental studies have been unable to disclose the exact pathogenesis of the adverse tissue responses. Certain precautions in the clinical use of bioabsorbable devices are necessary. Large implants in proportion to the bone section concerned should be avoided to minimise the risk of subsequent bone failures. Secondly, it is questionable whether devices made of high-molecular-weight PLLA with a degradation time of more than five years can be regarded as bioabsorbable from a clinical point of view. A very long degradation time is a disadvantage in clinical situations, where intra-articular breakage or subcutaneous prominence of an implant may cause problems or discomfort. Devices made of polymers with medium-long degradation time such as stereoisomeric copolymers of polylactide (poly-DLlactic acid) may not be burdened by the problems associated with the fast- and slowdegrading ones [76]. The preliminary results are promising [12,77], with the exception of one clinical study that had to be stopped due to loss of bone around a degradable poly-L/DL-lactic acid anchor for shoulder stabilisation [53]. More clinical experience of these implants is required. Since there may exist some species-related differences between rabbits and humans in the tissue restoration potential, also more data on the tissue replacement in humans should be gathered by retrieving proper tissue specimens whenever feasible. Biodegradable polymers with a short degradation time seem to be prone to elicit inflammatory foreign-body reactions. This has resulted in a decrease of the use of implants made of polyglycolide. Instead, polylactide in its various stereoisomeric forms has become the most popular polymer utilised in the manufacturing of biodegradable internal fixation devices. Since clinical studies on human patients concerned with implants made of PLLA are fully dependent on the availability of biopsy specimens from possible re-operations performed due to complications, long-term controlled in vivo studies in humans are impossible to perform. A very long degradation time of more than four years, although advantageous for the biocompatibility of the material, can be a problem in some clinical applications. Awareness of the relative biostability of PLLA is important when clinical applications are considered. The question of the biocompatibility of these materials is of general importance if absorbable alpha-hydroxy polyesters are increasingly being used also within other fields of orthopaedic surgery than internal fracture fixation. Indeed, it seems that the use of degradable implants in the stabilisation of ligamentous structures in shoulder and knee surgery [78,79] already outnumbers the fracture fixation procedures. Spinal cages for intervertebral fusion

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may be one of the important new clinical applications [16,17]. The true risks of the introduction of surgical innovations will be revealed only with increasing clinical experience of the method concerned. Animal experiments do not always give valid information of all the biological aspects relevant in humans. Prompt reporting of unexpected complications emerging in association with surgical innovations should be given high priority in the scientific literature. The international orthopaedic community must have the right to receive reliable information also of the adverse effects, not only the benefits of enthusiastically marketed novelties. To improve our understanding of all the critical factors influencing the behaviour of degradable materials in the tissues, systematic studies are required. Such an approach is, however, much more difficult in in vivo studies, whether they are experimental or clinical, than in vitro.

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[14] Nagura, I; Fujioka, H; Kokubu, T; Makino, T; Sumi, Y; Kurosaka, M. Repair of osteochondral defects with a new porous synthetic polymer scaffold. J. Bone Joint. Surg. 2007, 89-B:258-264. [15] Thanner, J; Kärrholm, J; Malchau, H; Wallinder, L; Herberts, P. Migration of press-fit cups fixed with poly-L-lactic acid or titanium screws: A randomized study using radiostereometry. J. Orthop. Res. 1996, 14:895-900. [16] Wuisman, PIJM; Smit, TH. Bioresorbable polymers:heading for a new generation of spinal cages. Europ. Spine J. 2006, 2:133-148. [17] Smit, TH; Thomas, KA; Hoogendorn, RJW; Strijkers, GJ; Helder, MN; Wuisman, PIJM. Sterilization and strength of 70/30 polylactide cages: e-beam versus ethylene oxide. Spine 2007, 7:742-747. [18] Williams, DF. Biodegradation of surgical polymers. J Mater Sci 1982, 17:1233-1246. [19] Törmälä, P. Biodegradable self-reinforced composite materials; manufacturing, structure and mechanical properties. Clin. Mater. 1992, 10:29-34. [20] An, YH; Woolf, SK; Friedman, RJ. Pre-clinical in vivo evaluation of orthopaedic bioabsorbable devices. Biomaterials 2000, 21:2635-2652. [21] Miller, RA; Brady, JM; Cutright, DE. Degradation rates of oral resorbable implants (polylactates and polyglycolates): rate modification with changes in PLA/PGA copolymer ratios. J. Biomed. Mater. Res. 1977, 11:711-719. [22] Böstman, O; Päivärinta, U; Partio, E; Vasenius, J; Manninen, M; Rokkanen, P. Degradation and tissue replacement of an absorbable polyglycolide screw in the fixation of rabbit femoral osteotomies. J. Bone Joint. Surg. [Am] 1992, 74-A:10211031. [23] Pistner, H; Gutwald, R; Ordung, R; Reuther, J; Mühling, J. Poly(L-lactide): a long-term degradation syidy in vivo. I. Biological results. Biomaterials 1993, 14:671-677. [24] Yamamuro, T; Matsusue, Y; Uchida, A; Shimida, K; Shimozaki, E; Kitaoka, K. Bioabsorbable osteosynthetic implants of ultra high strength poly-L-lactide. Int Orthop. (SICOT) 1994, 18:332-340. [25] Weir, NA; Buchanan, FJ; Orr, JF; Farrar, DF; Boyd, A. Processing, annealing and sterilisation of poly-L-lactide. Biomaterials 2004, 25:3939-3949. [26] Schwach G, Vert M. In vitro and in vivo degradation of lacticacid –based interference screws used in cruciate ligament reconstruction. Int .J. Biol. Macromol. 1999, 25:283291 [27] Vert, M; Li, S; Garreau, H. New insights on the degradation of bioresorbable polymeric devices based on lactic and glycolic acids. Clin. Mater. 1992, 10:3-8. [28] Matsusue, Y; Hanafusa, S; Yamamuro, T; Shikinami, Y; Ikada, Y. Tissue reaction to bioabsorbable ultra-high-strength poly-L-lactide rod. Clin. Orthop. 1995, 317:246-253. [29] Suuronen, R; Pohjonen, T; Hietanen, J; Lindqvist, C. A 5-year in vitro and in vivo study of the biodegradation of poly-lactide plates. J. Oral Maxillofacial Surg. 1998, 56:604-614. [30] Böstman, OM; Laitinen, OM; Tynninen, O; Salminen, ST; Pihlajamäki, HK. Tissue restoration after resorption of polyglycolide and poly-laevo-lactic acid screws. J. Bone Joint Surg. 2005, 87-B:1575-1580. [31] Rozema, FR; Bos, RRM; Pennings, AJ; Jansen, HWB. Poly(L-lactide) implants in repair of defects of the orbital floor: an animal study. J. Oral Maxillofac. Surg. 1990, 48:1305-1309.

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[32] Bos, RRM; Rozema, FR; Boering, G; Nijenhuis, AJ; Pennings, AJ; Verwey, AB; Nieuwenhuis, P; Jansen, HWB. Degradation of and tissue reaction to biodegradable poly(L-lactide) for use as internal fixation of fractures: a study in rats. Biomaterials 1991, 12:32-36. [33] Böstman, O; Päivärinta, U. Restoration of tissue components after insertion of absorbable fracture fixation devices of polyglycolide through the articular surface: an experimental study in the distal rabbit femur. J. Orthop. Res. 1994, 12:403-411. [34] Böstman, O; Päivärinta, U; Manninen, M; Rokkanen, P. Polymeric debris from absorbable polyglycolide screws and pins: intraosseous migration studied in rabbits. Acta Orthop. Scand. 1992, 63:555-559. [35] Laitinen, O; Pihlajamäki, H; Sukura, A; Böstman O. Transmission electron microscopic visualization of the degradation and phagocytosis of a poly-L-lactide screw in cancellous bone: A long-term experimental study. J. Biomed. Mater. Res. 2002, 61:3339. [36] Böstman, O; Pihlajamäki, H; Partio, EK; Rokkanen, P. Clinical biocompatibility and degradation of polylevolactide screws in the ankle. Clin. Orthop. 1995, 320:101-109. [37] Bergsma, EJ; de Bruijn, WC; Rozema, FR; Bos, RRM; Boering, G. Late degradation tissue response to poly(L-lactide) bone plates and screws. Biomaterials 1995,16: 25-31. [38] Pihlajamäki, H; Karjalainen, P; Aronen, H; Böstman, O. MR imaging of biodegradable poly-levolactide osteosynthesis devices in the ankle. J. Orthop. Trauma 1997, 11:559564. [39] Nakamura, T; Shimizu, Y; Okumura, N; Matsui, T; Hyon, SH; Shimamoto, T. Tumorigenicity of poly-L-lactide (PLLA) plates compared with medical-grade polyethylene. J. Biom. Mater Res. 1994, 28:17-25. [40] Weiler, A; Helling, HJ; Kirch, U; Zirbes, TK; Rehm, KE. Foreign-body reaction and the course of osteolysis after polyglycolide implants for fracture fixation. Experimental study in sheep. J. Bone Joint Surg. 1996, 78-B:369-376. [41] Pihlajamäki, H; Salminen, S; Laitinen, O; Tynninen, O; Böstman, O. Tissue response to polyglycolide, polydioxanone, polylevolactide, and metallic pins in cancellous bone: an experimental study on rabbits. J. Orthop. Res. 2006, 24:1597-1606. [42] Böstman, O; Hirvensalo, E; Mäkinen, J; Rokkanen P. Foreign-body reactions to fracture fixation implants of biodegradable synthetic polymers. J. Bone Joint Surg. 1990, 72-B:592-596. [43] Hoffmann, R; Krettek, C; Hetkämper, A; Haas, N; Tscherne, H. Osteosynthese distaler Radiusfrakturen mit biodegradablen Frakturstiften. Zweijahresergebnisse. Unfallchirurg 1992, 95:99-105 [44] Bergsma, EJ; Rozema, FR; Bos, RRM; de Bruijn, WC. Foreign body reactions to resorbable poly(L-lactide) bone plates and screws used for the fixation of unstable zygomatic fractures. J. Oral Maxillofac. Surg. 1993, 51:666-670. [45] Miketa, JP; Prigoff, MM. Foreign body reactions to absorbable implant fixation of osteotomies. J. Foot Ankle Surg. 1994, 33:623-627. [46] Böstman, OM; Pihlajamäki, HK. Adverse tissue reactions to bioabsorbable fixation devices. Clin. Orthop. 2000, 371:216-227. [47] Agrawal, CM; Athanasiou, KA. Technique to control pH in vicinity of biodegrading PLA-PGA implants. J. Biomed. Mater Res. 1997, 38:105-114.

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[48] Khouw, IM; van Wachem, PB; de Leij, LF; van Luyn, MJ. Inhibition of the tissue reaction to a biodegradable biomaterial by monoclonal antibodies to IFN-gamma. J. Biomed. Mater. Res. 1998, 41: 202-210. [49] Böstman, OM; Viljanen, J; Salminen, S; Pihlajamäki, H. Response of articular cartilage and subchondral bone to internal fixation devices made of poly-L-lactide: a histomorphometric and microradiographic study on rabbits. Biomaterials 2000, 21: 2553-2560. [50] Böstman, OM. Osteolytic changes accompanying degradation of absorbable fracture fixation implants. J. Bone Joint. Surg. 1991,73-B: 679-682. [51] Fraser, RK; Cole, WG. Osteolysis after biodegradable pin fixation of fractures in children. J. Bone Joint Surg. 1992,74-B: 929-930. [52] Rovinsky, D; Nissen, TP; Otsuka, NY. Osteolytic reaction to polylevolactic acid fracture fixation. Orthopedics 2001,24: 177-179. [53] Müller, M; Kääb, MJ; Villiger, C; Holzach, P. Osteolysis after open shoulderstabilization using a new bioresorbable bone anchor: A prospective, non-randomized clinical trial. Injury 2002, 33SB: 30-36. [54] Böstman, OM. Osteoarthritis of the ankle after foreign-body reaction to absorbable pins and screws. A three- to nine-year follow-up study. J. Bone Joint. Surg. 1998, 80-B: 333-338. [55] Böstman, OM; Pihlajamäki, H. Late foreign-body reaction to an intraosseous bioabsorbable polylactic acid screw. A case report. J. Bone Joint Surg 1998, 80-A: 1791-1794. [56] Takizawa, T; Akizuki, S; Horiuchi, H; Yasukawa, Y. Foreign body gonitis caused by a broken poly-L-lactic acid screw. Arthroscopy 1998,14: 329-330. [57] Yoshino, N; Takai, S; Watanabe, Y; Kamata, K; Hirasawa, Y. Delayed aseptic swelling after fixation of talar neck fracture with a biodegradable poly-L-lactide rod: Case report. Foot Ankle Internat 1998,19: 634-637. [58] Mosier-Laclair, S; Pike, H; Pomeroy, G. Intraosseous bioabsorbable poly-L-lactic acid screw presenting as a late foreign-body reaction: A case report. Foot Ankle Internat 2001, 22: 247-251. [59] Hollinger, JO. Preliminary report on the osteogenic potential of a biodegradable copolymer of polylactide (PLA) and polyglycolide (PGA). J. Biomed Mater. Res. 1983, 17: 71-82. [60] Schmitz, JP; Hollinger, JO. A preliminary study of the osteogenic potential of a biodegradable alloplastic-osteoinductive alloimplant. Clin. Orthop. 1988, 237: 245-255. [61] Askar, I; Gultan, SM; Erden, E; Yormuk, E. Effects of polyglycolic acid bioabsorbable membrane and oxidised cellulose on the osteogenesis in bone defects: an experimental study. Acta Chir. Plast 2003, 45: 131-138. [62] Manninen, MJ; Pohjonen, T. Intramedullary nailing of cortical bone osteotomies in rabbits with self-reinforced poly-L-lactide rods manufactured by the fibrillation method. Biomaterials 1993, 14: 305-312. [63] Vasenius, J; Vainionpää, S; Vihtonen, K; Mero, M; Mikkola, J; Rokkanen, P; Törmälä, P. Biodegradable self-reinforced polyglycolide (SR-PGA) composite rods coated with slowly biodegradable polymers for fracture fixation; strength and strength retention in vitro and in vivo. Clin. Mater. 1989, 4: 307-317.

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[64] Majola, A; Vainionpää, S; Mikkola, HM; Törmälä, P; Rokkanen, P. Absorbable selfreinforced polylactide composite rods for fracture fixation: strength and strength retention in the bone and subcutaneous tissues of rabbits. J. Mater. Sci. Mater. Med. 1992, 3: 4-7. [65] Claes, L; Burri, C; Kiefer, H; Mutschler, W. Resorbierbare Implantate zur Refixierung von osteochondralen Fragmenten in Gelenkflächen. Aktuel Traumat 1986, 16: 74-77 [66] Jukkala-Partio, K; Pohjonen, T; Laitinen, O; Partio, EK; Vasenius, J; Toivonen, T; Kinnunen, J; Törmälä, P; Rokkanen, P. Biodegradation and strength retention of polyL-lactide screws in vivo. An experimental long-term study in sheep. Ann. Chir. Gynaecol. 2001, 90: 219-224. [67] Eitenmüller, J; Gerlach, KL; Schmickal, T; Muhr, G. Semirigide Plattenosteosynthesen unter verwendung absorbierbare Polymere als temporäre Implantate. II Tierexperimentelle Untersuchungen. Chirurg 1987, 58: 831-839. [68] Mäkelä, EA. Healing of epiphyseal fracture fixed with a biodegradable polydioxanone implant or metallic pins. An experimental study on growing rabbits. Clin. Mater. 1988, 3: 61-71. [69] Hoffmann, R; Krettek, C; Haas, N; Tscherne, H. Die distale Radiusfraktur. Frakturstabilisierung mit biodegradablen Osteosynthese-Stiften (Biofix). Experimentelle Untersuchungen und erste klinische Erfahrungen. Unfallchirurgie 1989, 92: 430-434. [70] Rader, CP; Räuber, C; Rehm, KE; Koebke, J. Internal fixation of the distal radius. A comparative, experimental study. Arch. Orthop. Trauma 1995, 114: 340-343. [71] Kousa, P; Järvinen, TLN; Pohjonen, T; Kannus, P; Kotikoski, M; Järvinen, M. Fixation strength of a biodegradable screw in anterior cruciate ligament reconstruction. J. Bone Joint Surg. 1995, 77-B: 901-905. [72] Böstman, O; Mäkelä, EA; Södergård, J; Hirvensalo, E; Törmälä, P; Rokkanen, P. Absorbable polyglycolide pins in internal fixation of fractures in children. J. Pediat. Orthop. 1993, 13: 242-245. [73] Kankare, J; Hirvensalo, E; Rokkanen, P. Malleolar fractures in alcoholics treated with biodegradable internal fixation: 6/16 reoperations in a randomized study. Acta Orthop. Scand. 1995, 66: 524-528. [74] Pelto-Vasenius, K; Hirvensalo, E; Vasenius, J; Partio, EK; Böstman, O; Rokkanen, P. Redisplacement after ankle osteosynthesis with absorbable implants. Arch. Orthop. Trauma Surg .1998, 117:159-162. [75] Hovis, WD; Kaiser, BW; Watson, JT; Bucholz, RW. Treatment of syndesmotic disruptions of the ankle with bioabsorbable screw fixation. J. Bone Joint Surg. 2002, 84-A: 26-31. [76] Gogolewski, S; Mainil-Varlet, P. Effect of thermal treatment on sterility, molecular and mechanical properties of various polylactides. 2. Poly(L/D-lactide) and poly(L/DLlactide). Biomaterials 1997, 18: 251-255. [77] Tiainen, J; Soini, Y; Törmälä, P; Waris, T; Ashammakhi, N. Self-reinforced polylactide/polyglycolide 80/20 screws take more than 1(1/2) years to resorb in rabbit cranial bone. J. Biomed. Mat. Res. 2004, 70B: 49-55. [78] Freedman, KB; Smith, AP; Romeo, AA; Cole, BJ; Bach, BR Jr. Open Bankart repair versus arthroscopic repair with transglenoid sutures or bioabsorbable tacks for recurrent

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anterior instability of the shoulder: a meta-analysis. Am. J. Sports Med. 2004, 32:15201527. [79] Harvey, A; Thomas, NP; Amis, AA. Fixation of the graft in reconstruction of the anterior cruciate ligament. J. Bone Joint. Surg. 2005, 87-B: 593-603.

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PART II:

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CLINICAL APPLICATIONS OF DEGRADABLE IMPLANTS

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In: Degradable Polymers for Skeletal Implants Editors: P.I.J.M. Wuisman and T. H. Smit

ISBN 978-1- 60692-426-6 © 2009 Nova Science Publishers, Inc

Chapter 10

FRACTURE REPAIR WITH BIO-RESORBABLE IMPLANTS C.J. van Manen and M. van der Elst Department of Surgery and Traumatology, Reinier de Graaf Gasthuis, Delft, The Netherlands

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ABSTRACT Since the 1950’s surgical procedures for fracture repair have increased in absolute numbers. Compared to conservative therapy it has many advantages but definitely also some disadvantages. Advantages are the possibility of anatomical reduction, rigid fixation and early mobilisation of the limb and joints. Disadvantages consist of introduction of a foreign body, infection risk, anaesthetic risks and injury of the soft tissues. After consolidation osteosynthesis material loses its purpose and often needs to be removed during a secondary operation. Fractures at different locations in the body ask for different types of osteosynthesis material. At many locations it is possible to replace the ‘standard’ metal implant for a bio-resorbable implant. Research has shown some benefits of bio-resorbable implants over metal implants because it obviates the need for secondary operations for removal of the implant, with comparable union rates and functional results. Complicating factors like adverse tissue reactions and material failure has resulted in scepticism on the use of bio-resorbable implants. Fundamental research needs to be performed to optimise implants and to improve the strength of the implants, without augmenting the resorption time. At the same time the incidence of adverse tissue reaction needs to be decreased. More prospective randomised clinical trials need to be performed to prove the beneficiary effects of degradable implants in fracture treatment.

INTRODUCTION Fractures of the skeletal system are a common injury after trauma. Since ancient times the treatment of fractures consist of immobilisation of the injured limb. Through hieroglyphs ancient Egyptians showed us the art of immobilisation by means of a cast manufactured with

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twigs and mud, a technique used throughout history and evolving into the cast of paris-plaster first mentioned by Dutch army-surgeon Matthijsen (+/-1850). The technique is still used for initial or definite immobilisation in emergency wards worldwide and will probably remain in use because of the efficiency, low costs and ease of use. Since the beginning of the 20th century, following improvements in surgical and anaesthetic methods and the introduction of antibiotics, surgeons have developed new surgical techniques. Cerclage wiring, percutaneous pinning and intramedulary rods were developed, later followed by external fixators, special screws and plates, now commonly used in open reduction and internal fixation (ORIF) of fractures. Different fractures at different locations in the body ask for a wide range of surgical techniques and materials. Since the 1950’s there has been an increase in absolute numbers of surgical procedures in fracture treatment. The advantages of surgical interventions are the possibility of anatomical reduction, rigid fixation giving stability and early mobilisation of the limb. Disadvantages are the risk of infection after incurring a surgical wound combined with an introduced foreign body, anaesthetic risks and injury of the soft tissues. After consolidation the implant loses its purpose but still might give rise to problems. K-wires, percutaneously introduced, are often removed during secondary surgery. The plates used in ORIF are known for giving complaints of pain over the fracture site, long after complete consolidation of a fracture. Soft tissues, like tendons are known to be irritated by sharp edges of the plates and screws. Pain and tendon problems are complications that sometimes can be solved by re-operating the injured limb, removing the implant and reconstructing the torn tendon. In the 1970’s bio-resorbable implants were first introduced for musculo-skeletal injuries and fractures. Bio-resorbable implants can provide anatomical reduction, sufficient stability and early mobilisation of the limb. Because of bio-degradation the implant completely disappears from the fracture site, gradually transferring the mechanical load, hereby leaving healthy and strong bone without the need for a secondary operation for removal of the implant or reconstruction of soft tissues. During the operation the implant can be trimmed to a custom fit, to provide a perfect implant for each patient. CT-scans show less image distortion and MRI imaging in different parts of the body is possible, because there is no metal implant in the patient’s body. Known disadvantages of bio-resorbable implants are tissue reactions, limited mechanical properties, initial costs and the requirements of storage because of decay over time. During introduction the radiolucency is also a disadvantage for it is impossible to see the direction of the implant during operation, unless a screw is cannulated and a metallic guide wire is inserted. The most common complication in bio-resorbable implants is a sterile tissue reaction: Bostman, in a review of 2528 implants, reported an incidence of 4% of significant sterile local tissue reactions. [Bostman 2000]. Such a fluid accumulation contains polymer degradation products and might result in formation of a sinus (often wrongfully called a fistula). The reason is probably a higher rate of degradation than the rate of absorption (elimination). This reaction was more common in the period when predominantly homopolymeric devices such as PGA and PLLA implants were used. Nowadays most implants are made of co-polymers and the risk of a sterile local tissue reaction should be minimal. [Ashammakhi 2003] The reaction does not interfere with fracture healing and can often be treated conservatively with use of non-steroidal anti-inflammatory drugs alone. Sometimes the swelling needs to be aspirated with a fine needle. In some cases a formal

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debridement will be necessary in the operating theatre, diminishing the advantage over metal implants in regarding to secondary operations. In a meta-analysis of randomised controlled trials comparing bio-resorbable fixation devices with metal devices in different kinds of musculoskeletal injuries no significant difference between the groups was found with respect to functional outcome, infections and complications. In some groups of patients treated with bio-resorbable implants the reoperation rates were lower. In 11 out of 943 patients (1,2%) in this meta-analysis, who were treated with bio-resorbable implants, sterile sinusitis, giant cell reaction or fluid accumulation was reported. [Jainandunsing, 2005]

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CLINICAL USE The clinical use of bio-resorbable implants for fractures in the literature was analysed. In our clinic the use of bio-resorbable implants is limited to the fracture repair of fractures of the distal radius, fractures of the olecranon, fractures of the hand, malleolair fractures of the ankle and certain types of epiphysiolysis. In Scandinavia, where research into bio-resorbables has achieved a high level of evidence, fractures at many more sites have been treated with bioresorbable implants and extensive research has been done. This research has demonstrated that the following fractures can be treated with bio-resorbable implants: glenoidal rim fractures, fractures of the proximal humerus, fractures of the lateral humeral condyle, fractures of the medial condyle of the humerus, fractures of the olecranon, fractures of the radial head, fractures of the distal radius, fractures of the hand, fractures of the femoral head and neck, fractures of the femoral condyles, fractures of the patella, fractures of the tibial condyles, malleolar fractures, fractures of the talus, fractures of the calcaneus, and fractures of the metatarsal bones and phalanges of the toes. [Rokkanen 1998] Craniomaxillofacial (CMF) surgery is not mentioned in the review by Rokkanen, but also has an extensive number of uses for these implants. Bio-resorbable tacks are used on a large scale in the repair of ligaments in the knee as well.

MATERIALS FOR BIO-RESORBABLE IMPLANTS Since 1984 bio-resorbable implants have been used for the treatment of fractures. Bioresorbable polymers (repeated sequences of monomers) are based on polyesters that can be broken down when immersed in a hydrolytic surrounding. In vivo the process is catalysed by enzymes. Decades of research shows that polymers made from the polyglycolic acid group (PGA) and polylactic acid group (PLA) are the strongest of all bio-resorbable polymers. PLA has two isomers, the L isomer (PLLA) and the D isomer (PLGA). The L isomer is a crystalline polymer with a longer degradation time than the D isomer. By changing the ratios between L and D isomer the degradation rate of an implant can be varied. Initially most implants were entirely made from PLLA but recently more PLGA has been used together with the PLLA. A third generation of implants has become available now that contains reinforcing elements in the material that have come from the material itself, so-called Self Reinforced (SR) implants, like SR-PLLA. The implants molecular weight, degree of

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orientation, crystallinity, mass, solidity and chain regularity decrease the rate of degradation, while highly vascularised implantation sites, load acting on the implant, and the presence of additives accelerate the rate of degradation. [Gogolewski 2000]

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DISTAL RADIUS Wrist fractures are the most common of all fractures in man [Court-Brown 2006]. Because of man’s increasing age and the subsequent increase of the incidence of osteoporosis, wrist fractures are predicted to remain among the fractures with the highest incidence. Besides a higher incidence in an older population, there is also a rising incidence among a younger, healthier population because of changing recreational pastimes with increased risk of fractures of the wrist, like in-line skating and snowboarding. The primary treatment of a distal radius fracture will be reduction and fixation after which reduction is maintained in a plaster cast. General consensus is that after reduction of a wrist fracture in a young active patient, any inter-fragmentary step or gap of more than 2 mm, any radial shortening of more than 2 mm, any loss of the normal palmar angulation of more than 10° and/or radial inclination of 5° demands surgical treatment. Although no specific superior operative treatment is proven in any meta-analysis [Handoll 2003], many studies have shown the positive effects of open reduction and internal fixation (ORIF) with dorsal and volar plates on the fracture healing of unstable distal radius fractures. [Ring 1997, Campbell 2000]. The main disadvantage of ORIF with a metal plate has been the extensor tenosynovitis because of irritation by the metal implant. Attempts have been undertaken to develop low-profile, contourable implants to minimalise irritation, but still between 12% and 23% of patients require a removal of the implant after fracture consolidation. Bio-resorbable plates seem to address many of these problems: an easy contourable plate, easy to trim, low-profile, which provides angular stability. In case of tendon irritation a second operation will be obsolete because of resorption from the fracture site after fracture healing. The gradual loss of mass and rigidity allows for healing of the bone without stress shielding. The surgical technique is similar to the operation with standard metal plates, except that the plate can be moulded using a heatpack. When the perfect fit for this specific radius has been produced by moulding and trimming, the plate is fixed to the bone with bioresorbable screws. The screw and plate can be cauterised to provide angular stability. In a recent series of 32 patients randomly assigned to metal implants (N=13) and bioresorbable implants (N=19), results showed comparable function and pain after a follow-up of 52 weeks, but also an equal number of re-operations due to re-operations in the experimental group for material failure (and replacement with a metal implant) or sterile sinuses/swelling. The results were disappointing, because a decrease in number of reoperations was hypothesized. One future advice resulting from this trial could be to ensure a longer period of immobilisation by means of a softcast, in order to compensate for loss of rigidity of the implant over time, before complete consolidation has taken place. In a different series of 26 distal radius fractures treated with the same Lactosorb ®dorsal plate (a compound of 82% L polylactic acid and 18% polyglycolic acid), only two foreign body reactions were found, one resulting in re-operation. In this study the plate was combined with Biobon ®, a

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bio-resorbable bone substitution made from calcium phosphate. At 6-8 weeks 5 out of 26 fractures did show a redislocation. Therefore the researchers concluded that in a fracture with a metaphyseal gap of more than 7 mm a treatment with a metal implant is preferable because the consolidation of the fracture will take longer than the period of complete rigidity of the bio-resorbable implant. [Gangopadhyay, 2006]

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OLECRANON Since the 1970’s the treatment of choice for simple intra-articular fractures of the olecranon is the Tension Band Wire. This technique uses K-wires for fixation of the fracture and adds a cerclage-wire (called the Zuggertung) which is wrapped around the K-wires in an eight-shaped form. In this setup the musculus Triceps Brachii is used during extension and applies a dynamic force perpendicular to the fracture line, creating compression. Worldwide the TBW is applied using metal wires. The advantages are anatomical reposition of the intraarticular fracture, reducing the risk of arthrosis, and quick return to normal use of the elbow, reducing stiffness. [Rowland, 1992] Disadvantages are of a general surgical nature and based on the introduction of a foreign body such as the formation of a sinus in the bone or irritation of soft tissues. The K-wires are introduced through the proximal side of the olecranon and pass through the M. Triceps Brachii and the Bursa olecrani. The most distal part will always remain as a processus. The dynamic force of the muscle might eventually dislocate the K-wires and cause them to migrate. Bursitis Olecrani is frequent because the olecranon is a pressure point when resting the arm on a surface. Another problem mentioned here is failure of the material such as breakage of the cerclage-wire. All these problems might result in a secondary operation to remove the materials. This happens in the majority of the Zuggertung-osteosynthesis of the olecranon. Using bio-resorbable K-wires and a PDS loop for the cerclage-wire some of the complications are diminished. The two most often experienced problems, a painful proximal olecranon because of the K-wires and breakage of the cerclage wire, do not need a secondary operation, because of natural resorption from the fracture site. [Juutilainen, 1995] In this study 25 patients were randomly assigned to treatment with bio-resorbable (N=15) and metal (N=10) implants, with complete consolidation in all fractures and equal functional results in both groups. The technique using bio-resorbable K-wires is equal to the one using metal ones. The initial operation is easier, because of the ease of tying a knot in a PDS loop compared to a metal cerclage wire.

HAND/CARPAL FRACTURES The treatment of unstable fractures or complex injuries of the hand frequently demands for internal fixation. The demanding nature of the function of the hand requires anatomic restoration and early return to mobilisation. Metal screws and plate implants are common and over time comparable shaped bio-resorbable implants, with comparable mechanical properties have been developed.

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Most literature on bio-resorbable implants in fractures of the hand are either case reports or reports on bio-mechanical properties. [Waris 2002,2003,2004, Kujala 2004, Fitoussi 1998, Prevel 1996, Maruyama 1996] One randomised trial compared Herbert screws with polyglycolide screws in non-union and delayed union in fractures of the carpal scaphoid. 34 patients were included and the rate of union was 64% in the polyglycolide group and 60% in the Herbert screw group. The functional outcome was better in the Herbert screw group than in the polyglycolide group. A transient, local, nonbacterial tissue reaction occurred in 5 out of 20 (25%) patients in the polyglycolide group. Two Herbert screws were removed because the screw penetrated into the radial cartilage. The researchers concluded that the complication rate was relatively high in both methods, and that fractures of the scaphoid should not be treated with bio-resorbable implants. [Pelto-Vasenius 1995] In a review on bio-resorbable implants in hand trauma the author claims that the failure of treatment with different kinds of bio-resorbable screws might not have the material itself to blaim, but might be due to limitations of industry production: compared to metallic fixation there are fewer different kinds of screws. This claim is illustrated by the fact that no bio-resorbable screws exist with variable pitch threads, which is necessary for good compression of the fracture. A second limitation is that no headless screw implants, which would allow for complete subchondral placement, exist. Finally, there are fewer length and diameter options available in all implants. [Hughes 2006] The techniques are similar in bio-resorbable K-wires compared to metal ones, except for the pre-drilling of the holes, before inserting the K-wires.

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LATERAL/MEDIAL MALLEOLUS FRACTURE, RUPTURE OF SYNDESMOSIS Ankle fractures are common after inversion and direct trauma. The treatment is often based on the classification according to Weber. Weber A fractures are treated conservatively and Weber B/C fractures demand a surgical treatment when dislocation greater than 2 mm is present and/or the posterior tibial fracture fragment is larger than one-quarter of the joint surface. Occasionally a fracture is treated with screws and plates but many ankle fractures can be treated by screws alone. In 1994 the first scientific proof was given that replacing metal screws with bio-resorbable screws was safe and effective for this specific application [Bucholz, 1994]. Treatment with bio-resorbables was later shown to be efficient in both malleolar fractures [Michelson 2003] and ruptures of the syndesmosis. [Hovis 1997,2002, Sauer 2004, Sinisaari 2002, Thordarson 2001, Kaukonen 2005] Joukainen even stated the superiority of treatment with bio-resorbables by setting up a trial comparing two different kinds of bio-resorbable implants. [Joukainen 2007] The beneficiary effect of the bioresorbable syndesmotic screws might result from the fact that both metallic and bioresorbable screws inhibit full bearing of weight for eight weeks, while after this period the metal screws are removed with additional immobilisation. The technique of handling the screws is comparable for bio-resorbable and metal screws, especially when the bio-resorbable screw is cannulated with a guide wire.

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EPIPHYSEAL FRACTURES A special kind of fracture is the epiphyseal fracture in children. This occurs when the growth plate becomes separated from the bone at the epiphysis, often in a hyperextension or hyperflexion trauma. Classification is performed by the Salter and Harris classification. All types of epiphysiolysis carry a risk of shortening or angulation when the physis closes, either completely or partial. In case of surgical treatment, by transphyseal pinning, the implant needs to be removed to guarantee normal growth after consolidation. Two studies describing transphyseal fixation of the lateral and humeral epicondyles with bio-resorbable pins reported good results, because no secondary operations proved necessary and there were no growth disturbances. The omission of a second operation is of course a great psychological advantage for little children. [Bostman 1989, Makela 1992] A study was performed by Partio on treatment of distal femoral epiphyseal fractures. They conclude that treatment with bioresorbable pins offers an attractive alternative for treatment of this injury, especially compared to traction followed by a plaster cast. 7 out of 9 patients did not have any growth problems and could return to their previous level of sports, but they report that special care needs to be taken to achieve exact anatomical reduction. [Partio, 1997] These studies were not randomised nor controlled and only consisted of small series.

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CONCLUSION Bio-resorbable implants have been in clinical use since 1970’s and good results have been described in some implants. Simple, non-weight bearing fractures can be safely treated with bioresorbable implants, but still mechanical failure is a big concern in bio-resorbable devices. Another consideration to be made on the mechanical properties is that stable fixation of fractures allows early motion. Because load acting actually increases the rate of degradation of bio-resorbable implants, it is unclear if these implants are adequate to allow rapid mobilisation. The use of longer lasting and stronger materials may actually negate some of their theoretical advantages, because resorption rates will extend over too long a period. A second problem that needs to be overcome is that of adverse tissue reactions. Even though fracture healing is not influenced by the sterile inflammation surrounding the implant, a secondary treatment for a complication is sometimes necessary. Many theoretical advantages have been described for bio-resorbable implants, but in clinical use they have not all been proven. Some disadvantages have been shown, but in the end they still result in similar re-operation rates. Little scientific research has been produced comparing the rates of complications and the financial burden of these complications. The cost/benefit ratio of either implant needs to be investigated. This can only be done in a clinical trial comparing the costs of initial implantation, absence of work due to immobilisation, re-operation rates and the costs of secondary procedures and absence of work when a complication actually occurs. This way clinical indications can be set for different types of fractures in which bio-resorbable implants are advantageous to metal implants or vice versa. The authors conclude that fundamental research is needed in order to develop new materials to be able to control the rate of resorption, hereby developing implants with

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predictable mechanical properties. Research into resorption time and immunological responses should be performed to minimise the incidence of adverse tissue reactions. Design of new implants has to be intensified to widen the arsenal of implants. Only after extensive, additional research, bio-resorbable implants for fracture treatment can be used on a larger clinical scale.

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Figure 1. The reunite plate by biomet during surgery.

Figure 2. X-ray of wrist, three months after trauma. Degradable Polymers for Skeletal Implants, Nova Science Publishers, Incorporated, 2009. ProQuest Ebook Central,

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Figure 3. X-ray of elbow, pre-operative and 5 months after surgery.

Figure 4. X-ray of lateral malleolus fracture at trauma and after 7 months.

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Figure 5. Fixation of lateral malleolus with bio-resorbable implants.

Figure 6. X-ray of syndesmosis after fixation using a bio-resorbable screw.

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Figure 7. Fixation of syndesmosis using a cannulated screw.

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REFERENCES Ashammakhi N, Suuronen R, Tiainen J, Tormala P, Waris T. Spotlight on naturally absorbable osteofixation devices. J. Craniofac. Surg. 2003 Mar;14(2):247-59. Review. Bostman O, Makela EA, Tormala P, Rokkanen P. Transphyseal fracture fixation using biodegradable pins. J. Bone Joint. Surg. Br. 1989 Aug;71(4):706-7. Bostman OM, Pihlajamaki HK. Adverse tissue reactions to bioabsorbable fixation devices.Clin. Orthop. Relat Res. 2000 Feb;(371):216-27. Review. Bucholz RW, Henry S, Henley MB. Fixation with bioabsorbable screws for the treatment of fracture of the ankle. J. Bone Jt. Surg. 1994;76-A:319-24 Fitoussi F, Lu W, Ip WY, Chow SP. Biomechanical properties of absorbable implants in finger fractures. J. Hand Surg. [Br]. 1998;23:79-83. Gogolewski S. Bioresorbable polymers in trauma and bone surgery. Injury. 2000 Dec;31 Suppl 4:28-32. Review. Hovis WD, Bucholz RW. Polyglycolide bioabsorbable screws in the treatment of ankle fractures. Foot Ankle Int 1997;18:128–31. Hovis WD, Kaiser BW, Watson JT, Bucholz RW. Treatment of syndesmotic disruptions of the ankle with bioabsorbable screw fixation. J. Bone Joint. Surg. Am. 2002;84:26–31. Hughes TB. Bioabsorbable Implants in the Treatment of Hand Fractures: An Update.Clin. Orthop. Relat Res. 2006 Apr;445:169-174. Jainandunsing JS, van der Elst M, van der Werken CC. Bioresorbable fixation devices for musculoskeletal injuries in adults. Cochrane Database Syst. Rev. 2005 Apr 18;(2):CD004324. Review.

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Joukainen A, Partio EK, Waris P, Joukainen J, Kroger H, Tormala P, Rokkanen P. Bioabsorbable screw fixation for the treatment of ankle fractures. J. Orthop. Sci. 2007 Jan;12(1):28-34. Juutilainen T, Patiala H, Rokkanen P, Tormala P. Biodegradable wire fixation in olecranon and patella fractures combined with biodegradable screws or plugs and compared with metallic fixation. Arch. Orthop. Trauma Surg. 1995;114(6):319-23. Kaukonen JP, Lamberg T, Korkala O, Pajarinen J. Fixation of syndesmotic ruptures in 38 patients with a malleolar fracture: a randomized study comparing a metallic and a bioabsorbable screw. J. Orthop. Trauma. 2005 Jul;19(6):392-5. Kujala S, Raatikainen T, Kaarela O, Ashammakhi N, Ryhanen J. Successful treatment of scaphoid fractures and nonunions using bioabsorbable screws: report of six cases. J. Hand Surg. 2004;29:68–73. Makela EA, Bostman O, Kekomaki M, Sodergard J, Vainio J, Tormala P, Rokkanen P. Biodegradable fixation of distal humeral physeal fractures. Clin. Orthop. Relat. Res. 1992 Oct;(283):237-43. Maruyama T, Saha S, Mongiano DO, Mudge K. Metacarpal fracture fixation with absorbable polyglycolide rods and stainless steel Kwires: a biomechanical comparison. J. Biomed. Mater. Res. 1996;33: 9–12. Michelson JD. Ankle fractures resulting from rotational injuries. J. Am. Acad. Orthop. Surg. 2003;11:403–12. Partio EK, Tuompo P, Hirvensalo E, Bostman O, Rokkanen P. Totally absorbable fixation in the treatment of fractures of the distal femoral epiphyses. A prospective clinical study. Arch. Orthop. Trauma Surg. 1997;116(4):213-6. Pelto-Vasenius K, Hirvensalo E, Böstman O, et al. Fixation of scaphoid delayed union and non-union with absorbable polyglycolide pin or Herbert screw. Consolidation and functional results. Arch. Orthop. Trauma Surg. 1995;114:347–351 Prevel CD, Eppley BL, Ge J, Winkler MM, Katona TR, D’Alessio K, Sarver D. A comparative biomechanical analysis of resorbable rigid fixation versus titanium rigid fixation of metacarpal fractures. Ann Plast Surg. 1996;37:377–385. Rokkanen PU. Bioabsorbable fixation devices in orthopaedics and traumatology. Ann Chir. Gyn. 1998;87:13}20. Rokkanen PU, Bostman O, Hirvensalo E, Makela EA, Partio EK, Patiala H, Vainionpaa SI, Vihtonen K, Tormala P. Bioabsorbable fixation in orthopaedic surgery and traumatology. Biomaterials. 2000;21:2607–2613. Rowland SA, Burkhart SS. Tension band wiring of olecranon fractures. A modification of the AO technique. Clin. Orthop. Relat Res. 1992 Apr;(277):238-42. Sauer ST, Marymont J, Mizel MS. What’s new in foot and ankle surgery? J. Bone Joint Surg. Am. 2004;86:878–86. Sinisaari IP, Lüthje PM, Mikkonen RM. Ruptured tibio-fibular syndesmosis: comparison study of metallic to bioabsorbable fixation. Foot Ankle Int 2002;23:744–8. Thordarson DB, Samuelson M, Shepherd LE, Merkle PF, Lee J. Bioabsorbable versus stainless steel screw fixation of the syndesmosis in pronation-lateral rotation ankle fractures: a prospective randomized trial. Foot Ankle Int 2001;22:335–8. Waris E, Ashammakhi N, Happonen H, Raatikainen T, Kaarela O, Tormala P, Santavirta S, Konttinen YT. Bioabsorbable miniplating versus metallic fixation for metacarpal fractures. Clin. Orthop. Relat Res. 2003;410:310–319.

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Waris E, Ashammakhi N, Raatikainen T, Tormala P, Santavirta S, Konttinen YT. Selfreinforced bioabsorbable versus metallic fixation systems for metacarpal and phalangeal fractures: a biomechanical study. J. Hand Surg. 2002;27:902–909. Waris E, Ninkovic M, Harpf C, Ashammakhi N. Self-reinforced bioabsorbable miniplates for skeletal fixation in complex hand injury: three case reports. J. Hand Surg. 2004;29:452– 457.

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In: Degradable Polymers for Skeletal Implants Editors: P.I.J.M. Wuisman and T. H. Smit

ISBN 978-1- 60692-426-6 © 2009 Nova Science Publishers, Inc

Chapter 11

FIBULA REGENERATION AFTER VASCULARIZED FIBULAR GRAFT HARVESTING Arthur de Gast*,1, Hay A.H. Winters2 and Paul I.J.M. Wuisman1 1

Department of Orthopedic Surgery and Department of Plastic and Reconstructive Surgery, VU University Medical Center, Amsterdam, The Netherlands 2

1. ABSTRACT

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Vascularized and non-vascularized fibular graft transfer for skeletal reconstruction purposes has been shown to be an effective treatment strategy for various congenital and acquired skeletal deformities, including previously failed arthrodesis, and reconstruction of large segmental bone defects due to tumor, infection and trauma[1-13]. When a fibular graft is harvested leaving the periosteal sleeve and the vascular supply intact at the donor site, complete regeneration of the fibula can be expected particularly in children under the age of 15 years. This regeneration potential is however lost after harvesting a vascularized fibular graft. Most reports focus on the outcome of the reconstructed skeletal area, with little attention given to the clinical problems resulting from donor site morbidity. Consequently, little data is available on strategies for donor site reconstruction and enhancement of fibula regeneration after vascularized fibular graft harvesting. Some clinical and biomechanical studies have shown serious donor site morbidity after harvesting vascularized as well as non-vascularized fibular grafts.[14-25] Clinical symptoms at the donor site ranges from mild subjective complaints of ache and discomfort to severe cases of ankle instability and valgus deformity as well as significant plantar flexion weakness with resultant gait abnormality.[26-30] To prevent the above mentioned complaints and complications, enhancement or promotion of regeneration of the fibula, particularly after vascularized graft harvesting, is desirable. Successful fibula regeneration using beta-tricalcium phosphate has been described after reconstruction of a none vascularized harvesting procedure[31], however, to our knowledge there are no reports on successful fibula regeneration of a resected vascularized fibula defect. *

Author data: Arthur de Gast, MD PhD, staff orthopedic surgeon, Department of Orthopedic Surgery, VU University Medical Center, De Boelelaan 1117, 1007 MB Amsterdam , The Netherlands, Tel: +31 20 444 2987, Fax: +31 20 444 2357, E-mail: [email protected].

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This chapter reports on our prospective study in which a fibula defect created upon a vascularized graft harvest was reconstructed using a bioresorbable porous PLDLLA mesh (70/30 copolymer of poly[L-lactide-co-D,L-lactide], Macropore Biosurgery, San Diego, CA), shaped to a tube, and gentle impacted with bicalciumphosphate granules (BCP Bicalphos, Medtronic Sofamore Daneck, Memphis, Ten, USA, 60% hydroxy apatite and 40% tricalciumphospate, 100% porosity, pore diameter 400- 600 µm, and pore interconnectivity 100-150 µm).

Keywords: Fibula, graft, bone regeneration, resorbable, polylactic-acid, calcium phosphate

2. PATIENTS AND SURGICAL TECHNIQUE 2.1. Patients Between February 2002 and June 2003 a reconstruction of the harvested vascularized fibula defect was performed in 13 patients, after approval was obtained from the Institutional Review Board of the VU University Medical Center. The surgical indications for which the vascularized fibular grafts were harvested included segmental defect reconstruction after bone/soft tissue tumor resection (6), augmentation of mechanical stability in chronic osteomyelitis (3), and enhancement of fusion in progressive de novo scoliosis (2), kyphosis (1), and pseudarthrosis (1) (Table 1). Patients included 5 men and 8 women with a mean age of 42 years (range 8 to 71 years) at surgery. The average clinical follow up period was 45 months (range 36 to 54 months).

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2.2. Surgical Technique 2.2.1. Surgical Dissection The surgical intervention was performed with the patient in a supine or lateral decubitus position depending on the site of the primary lesion. The vascularized fibula was harvested separately by a second team. The fibula head and the lateral malleolus were marked. A straight line or garland skin incision was made from several centimeters below the fibular head extending distally until 5 cm proximal to the lateral malleolus. The muscular fascia was exposed and the posterior crural septum between the peroneal and soleus muscles was identified. The peroneal muscles were stripped at their origin from the anterolateral part of the fibula, leaving the periosteum intact. The crural septum was divided. The dissection was continued anteriorly detaching the extensor digitorum longus, extensor hallucis longus and dividing the interosseous membrane. Posterior to the crural septum, the peroneal artery and veins were identified. After osteotomies of the fibula at the appropriate levels, the vessels were cut distally and dissection proceeded towards proximal. Proximally, dissection and ligation of the peroneal vessels were performed at the junction of the posterior tibial artery and veins and subsequently the vascularized graft was removed, further prepared and transplanted.

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Table 1. Patient demographics and clinical data Patient number

Age (years/ sex)

Diagnosis

1

54/ f

2

51/ f

3

59/ m

4

71/ m

5*

8/ m

6*

12/ m

7

46/ m

8

58/ f

9

18/ f

10

66/ m

11*

14/ m

12

30/ m

13

49/ f

Lumbar De Novo Scoliosis Lumbar De Novo Scoliosis Thoracic Kyphosis TBC T1-T3 Chondro-sarcoma pelvis Osteo-myelites femur Ewings sarcoma femur Pancoast tumor ingrowth T2-T4 Oropha-ryngeal carcinoma Fibrous Dysplasia prox humerus Pseudar-throsis distal femur Spondyli-tis TBC L4 Osteomye-litis tibia Glandular carcinoma

Length of fibular Resection (cm) 21

Duration Follow up (months)

Sheet +/BCP

54

21

Continuity of distal side with the fibula

Continuity of proximal side with the fibula

No BCP

New bone formation (no/partial/ complete) No

No

No

51

No BCP

No

No

No

19

52

No BCP

No

No

No

20

48

BCP

Complete

Yes

No

13

46

BCP

Complete

Yes

Yes

21.5

46

BCP

Complete

Yes

No

22

44

BCP

Complete

Yes

Yes

19

58

BCP

Partial

No

Yes

20

36

BCP

Complete

No

Yes

27

40

BCP

Complete

No

No

18

38

BCP

Complete

Yes

Yes

19

39

BCP

Complete

Yes

Yes

19

42

BCP

Partial

No

Yes

* Patients under age of 15 years not developing valgus deformation of ankle joint .

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Only the diaphysis of the fibula was harvested and care was taken to retain appropriate lengths of the fibula distally (minimum length 8 to 10 cm) and proximally (minimum length 3 to 6 cm distal of the fibular head).

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2.2.2. Reconstruction of the Fibula at the Donor Site To reconstruct the fibular defect, a bioresorbable PLDLLA mesh (70/30 copolymer of poly[L-lactide-co-D,L-lactide], Macropore Biosurgery, San Diego, CA, USA), sized 20 by 20 cm by 0.8 mm, was bathed in sterile water at a temperature of 40 to 50 degrees Celsius, subsequently cut to the appropriate size and molded around a 8 mm rod. Upon cooling a PLDLLA tube was obtained with the appropriate shape and size (Figure 1). The molded tube was tested in the recipient bed and should overlap both remaining fibula ends for about 1 cm (Figure 2). In 10 patients the tube was first gently impacted with BCP granules (Sofamore Daneck, Memphis, Ohio, USA) (Figure 3) and, depending on the length of the tube, 50 - 75 ml autologous blood was injected in the scaffold prior to definitive insertion, thus creating a coherent cloth. During insertion, care was taken to have contact between the distal and proximal fibular remnants with the BCP (Figure 4). The reconstruct was covered with surrounding muscles followed by skin closure over a suction drain. In 3 patients the reconstruction was performed with an empty PLDLLA tube alone. 2.2.3. Postoperative Protocol Postoperatively, all patients were allowed immediate full weight bearing. All patients were seen at the outpatient clinic and examined both clinically and radiographically at regular intervals (6 weeks, 3, 6, 12, 18, 24 months and then yearly). At each time point clinical symptoms such as pain, ankle instability and limitations in normal walking or sports activities were enquired and documented. Physical examination which was always in comparison with the contra-lateral side included range of motion of ankle joint and toes, as well as clinical signs of valgus axis deviation and plantar flexion weakness of the ankle. In addition, anteroposterior (AP) and lateral radiographs of the donor leg were taken at every examination time. Furthermore, a six-monthly weight bearing AP-radiographs of both ankle joints were performed in all children to assure further normal growth and development of the distal tibial and fibular growth plates and epiphyses.

3. RESULTS Demographic data of the patients and the treatment have been summarized in Table 1. The mean length of fibula harvested was 20 cm (range 13 to 27 cm). In all 3 patients with an empty PLDLLA tube reconstruction, only few isolated bone islets were observed scattered along the tube trajectory. However, a gradual callus formation replacing the BCP granules was seen in 8 of 10 patients having a PLDLLA – BCP construct (Figure 5D-F). In these patients the BCP was gradually replaced by newly formed bone as observed in the first 24 to 36 months postoperatively and sometimes beyond. Two adult patients showed partial resorption of the granules with only proximal continuity between the partially regenerated fibula and the native fibula.

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Figure 1. PLDLLA sheet moulded to a tube.

Figure 2. Probe insertion of the tube: the tube envelops the proximal fibula end of about 1cm.

Figure 3. PLDLLA tube containing the BCP granules.

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Figure 4. In situ placement of the PLDLLA – BCP construct.

Figure 5A . Case 6, male, 12 years old, T1 weighted MRI scan with contrast showing an Ewing Sarcoma of the right proximal femur with substantial soft tissue invasion. Degradable Polymers for Skeletal Implants, Nova Science Publishers, Incorporated, 2009. ProQuest Ebook Central,

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Figure 5B. Radiograph of the proximal femur one month after surgery showing a reconstruction of the proximal femur diaphysis by a vascularized fibular graft protected by an allograft shell and stabilized by a hookplate osteosynthesis.

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Figure 5C. Radiograph of the proximal femur 46 months after surgery showing the complete intergration and remodeling of the allograft/fibular graft reconstruction. Degradable Polymers for Skeletal Implants, Nova Science Publishers, Incorporated, 2009. ProQuest Ebook Central,

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Figure 5D. Radiograph of the right lower leg one month after surgery showing the complete fibula defect reconstruction by the PLDLLA – BCP reconstruct. Note, only the BCP granules are visible and there is no remodeling seen. Normal open distal fibula/tibia epiphysis.

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Figure 5E. Radiograph of the right lower leg 24 months after surgery showing the intact PLDLLA – BCP reconstruct with remodeling of the BCP granules from proximal to distal. Unchanged, open distal fibula/tibia epiphysis. Degradable Polymers for Skeletal Implants, Nova Science Publishers, Incorporated, 2009. ProQuest Ebook Central,

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Figure 5F. Radiograph of the right lower leg 46 months after surgery showing an intact and almost complete remodeled neo fibula with intact fibula – neo fibula anchorages. The distal fibula/tibia epiphysis is almost closed and there is no valgus deformation.

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Figure 5G. Anteroposterior and posteroanterior Technetium 99 bone scans of the lower extremities showing a slightly enhanced uptake at the side of the neo fibula pointing to both vital bone and some enhanced bone remodeling of the neo fibula.

Complete fusion at both proximal and distal fibula-construct junction was seen in 2 of the 6 other adult patients whereas in 3 other adult patients failure of fusion was seen distally (twice) and proximally (once). One adult patient showed failure of fusion at both ends of the construct to the fibula. In the 3 children, complete, gradual replacement and intact continuity between the regenerated fibula and the native fibula parts was observed (Figure 5D-F). There were no differences in time to remodeling between the children and adults. In all cancer patients (cases 4,6,7,8,13) by whom postoperatively total body bone scans were performed, tracer-uptake of the neo fibula was found underscoring vital bone formation of the regenerated fibula (Figure 5F). There were no postoperative wound complications; neither were there any complications related to the implant. Postoperatively, all patients started with full weight-bearing as soon as tolerated. At the time of last follow-up 1 patient (case 10, Table 1) having a PLDLLA – BCP reconstruction complained of pain and swelling at the distal tube – fibula junction. Physical

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examination showed no palpable swelling, however, radiography showed the absence of union between the remodeled BCP and the distal fibula remnant. Another patient (case 3, Table 1), having a PLDLLA reconstruction alone, complained of hypoesthesia of the lower lateral part of the calf and dorsum of the foot. None of the patients had signs or symptoms of ankle instability nor did any patient show limitations in range of motion of the ankle and toes. Gait pattern was normal in all patients with no patient complaining of walking impairment. All patients that participated in sporting activities prior to surgery have resumed such activities in full, including jumping and jogging. On physical examination all patients had a stable ankle joint and normal ankle and toe functions comparable to the opposite site. None of the children experiences limitations in the activities of daily life and sports. On examination there were normal symmetrical ankle joint and toe functions. Clinically and radiographically there were patients with a valgus deformity of the ankle joint nor were there any signs of growth disturbances of the distal tibia/fibula growth plate and epiphysis.

4. DISCUSSION 4.1. Advantages of Free Vascularized Bone Graft

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The advantages of using a free vascularized bone graft to reconstruct osseous bone defects have been well described in the literature. Vascularized bone grafts are less prone to resorption, exhibit superior mechanical properties, fuse and incorporate rapidly at the anchorage sites, and are more resistant to infection and the adverse effects of radiation therapy.[19,32,33] Specially, in spinal reconstructive surgery we favor the use of vascularized fibular graft because of its superior mechanical properties, besides, the graft is easy to fit to the type, size, shape and location of the osseous defect in different anatomical areas. The graft provides up to 30 cm of bone in the adult male.[34]

4.2. Adverse Effects of Harvesting Fibular Grafts It is generally accepted that harvesting fibular graft has a low donor site morbidity[23], however, serious complaints and complications such as weakness in plantar flexion of the ankle and rotation of the leg, ankle pain, ankle instability and the development of a progressive valgus deformity of the ankle in children have been reported.[15,27,28,30,35] Considering these complaints and complications, one should consider methods to regenerate a neo fibula at the site of the harvested vascularized fibula especially in children and in those individuals performing heavy physical activity. After non vascularized fibular graft harvesting, fibula regeneration can occur in children under the age of 15 when the periosteal sleeve is left intact.[36,37] However, spontaneous fibula regeneration very rarely occurs after non vascularized harvesting above the age of 15 and never occurs after vascularized fibular graft harvesting irrespective of age.[30,37] To overcome potential post resection donor site morbidity, several surgical methods have been developed intended on local stabilization of the distal tibiofibular complex. Fragniére et al described two methods of preventing valgus deformity of the ankle after vascularized

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fibular graft harvesting in children.[27] In one group a tibial cortical autograft stabilized with a Kirschner wire was used to reconstruct the fibular defect. The outcome of this reconstruction which showed similar results compared to a historical control group (no reconstruction), however, was worse when compared to a group treated with a tibiofibular syndesmotic screw with respect to the development of ankle valgus deformity/instability. Another method for prevention of progressive ankle valgus deformation is a tibiofibular metaphyseal synostosis (the Langenskiöld procedure).[38] Evaluation of this procedure showed that it can delay but not prevent the ankle valgus.

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4.3. Regeneration of the Fibula Diaphysis Recently, a novel application of beta-TCP blocks in regeneration of the fibula in patients undergoing non-vascularized fibular graft harvesting have been reported.[31] In 12 of 14 reconstructed fibula defects with beta-TCP newly formed trabecular bone was seen at an average of 9.3 months after surgery. One of the two failed cases had a vascularized graft harvesting procedure. Although regeneration was seen, continuity between the regenerated fibula and fibular ends was relatively uncommon, especially in adults at the distal neo fibula native fibula junction. Our clinical and radiographic results are quite comparable: we achieved a high percentage of fibula regeneration even in the less promising patient group undergoing a vascularized fibular graft harvesting procedure. Similar to Arai et al.31 we had a consistent success rate in children compared to adults with respect to BCP absorption, bone remodeling and bone bridging at the junction sites. Furthermore, no patient developed a valgus deformity. Our case series showed a higher tendency towards fusion between the host fibula and the PLDLA-BCP construct proximally compared to the distal end. However, in our series the fibula regeneration was a much slower process: we observed a more gradual transformation of the BCP granules to a mixture of bone and granules. The remodeling process occurred earlier proximally. The observed differences in remodeling between this study and that of Arai et al. may be explained by the differences in the harvesting procedure (vascularized versus non-vascularized) and the materials used (beta-TCP versus PLDLLA mesh containing BCP). Significant bone regeneration of a critical size tricortical iliac crest defect using a similar bioresorbable protective PLDLLA mesh has been reported in an animal study.[39] However, this concept failed in the three patients having a tubular PLDLLA mesh reconstruction only and therefore this method is not recommended. Although no previous report exists about successful fibula regeneration using our technique, two previous clinical papers have reported favorable results when using the same PLDLLA mesh and osteoconductive materials to reconstruct iliac crest defects.[40,41] We confirmed these favorable results in an animal study performed at our institute comparing the outcome of reconstruction of a standardized critical iliac crest defect (unpublished data). Three methods were compared (no reconstruction vs. PLDLLA mesh vs. PLDLLA mesh and BCP). At every time point (3, 6, and 12 months) a significant larger area of diffuse bone formation was observed in the PLDLLA – BCP group. However, substitution of the BCP granules by trabecular bone did not start before 6 months and was more pronounced at 12 months. At this last time point increased resorption of the BCP by multinucleate giant cells

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(MNGCs) was observed which provided the necessary space for bone tissue development. Although the final result of the remodeling process could not be observed (the study ended at 12 months), the findings fit well with the observations of the current study: the remodeling/resorption of the BCP granules in bone is a slow process and could be observed up to 24 to 36 months and occasionally beyond. Furthermore, the remodeling process occurred more pronounced and earlier proximally, probably due to a more muscular embedding of the construct. We assume that this environment facilitates both development of in growing vessels into the graft and better migration of cells responsible for the bone remodeling process. We assume that earlier trabecular bone formation could be established by increasing the resorption rate by e.g. using a beta - TCP instead of BCP or by using BCP granules with a different pore size, microstructure and/or a lower HA content.

4.4. Conclusions and Future Developments In conclusion, the data presented in this chapter show the first clinical case series demonstrating that the fibula can regenerate after harvesting a vascularized fibular graft, using a bioresorbable protective sheet in combination with a BCP, both in adults and in children. Although the number of children in this case series is little, it is important to note that no patient developed a valgus ankle deformity postoperatively. The surgical technique is save, simple, reliable and does not significantly lengthen the duration of surgery. The concept also makes it feasible to perform different studies using other osteoconductive/osteoinductive materials, and adding growth factors and/or mesenchymal stem cells to the construct.

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ACKNOWLEDGEMENTS The following persons have contributed in various ways to this chapter: Rick M.A. Paul MD PhD, Timothy U. Jiya MD, Hay A.H. Winters MD and Matthijs R. Krijnen.

REFERENCES [1]

[2] [3] [4] [5]

Asazuma, T., Yamagishi, M., Nemoto, K., Amako, M., Osada, M., Fujikawa, K. Spinal fusion using a vascularized fibular bone graft for a patient with cervical kyphosis due to neurofibromatosis. J. Spinal Disord. 10: 537-540, 1997. Cybulski, G. R. Vascularized fibular grafts for vertebral body replacement. J. Neurosurg. 72: 519-520, 1990. Doi, K., Kawai, S., Sumiura, S., Sakai, K. Anterior cervical fusion using the free vascularized fibular graft. Spine 13: 1239-1244, 1988. Esses, S. I., Natout, N., Kip, P. Posterior interbody arthrodesis with a fibular strut graft in spondylolisthesis. J. Bone Joint Surg. Am. 77: 172-176, 1995. Hardes, J., Gosheger, G., Halm, H., Winkelmann, W., Liljenqvist, U. Three-level en bloc spondylectomy for desmoplastic fibroma of the thoracic spine: a case report. Spine 28: E169-E172, 2003.

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[15] [16] [17] [18] [19] [20] [21]

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Arthur de Gast, Hay A.H. Winters and Paul I.J.M. Wuisman Hubbard, L. F., Herndon, J. H., Buonanno, A. R. Free vascularized fibula transfer for stabilization of the thoracolumbar spine. A case report. Spine 10: 891-893, 1985. Kim, C. W., Abrams, R., Lee, G., Hoyt, D., Garfin, S. R. Use of vascularized fibular grafts as a salvage procedure for previously failed spinal arthrodesis. Spine 26: 21712175, 2001. Krishnan, K. G., Muller, A. Ventral cervical fusion at multiple levels using free vascularized double-islanded fibula - a technical report and review of the relevant literature. Eur. Spine J. 11: 176-182, 2002. Macdonald, R. L., Fehlings, M. G., Tator, C. H. et al. Multilevel anterior cervical corpectomy and fibular allograft fusion for cervical myelopathy. J. Neurosurg. 86: 990997, 1997. Meyers, A. M., Noonan, K. J., Mih, A. D., Idler, R. Salvage reconstruction with vascularized fibular strut graft fusion using posterior approach in the treatment of severe spondylolisthesis. Spine 26: 1820-1824, 2001. Nijland, E. A., van den Berg, M. P., Wuisman, P. I., van Royen, B. J., Winters, H. A., van Ouwerkerk, W. J. Correction of a dystrophic cervicothoracic spine deformity in Recklinghausen's disease. Clin. Orthop. 349: 149-155, 1998. Noorda, R. J., Wuisman, P. I., Fidler, M. W., Lips, P. T., Winters, H. A. Severe progressive osteoporotic spine deformity with cardiopulmonary impairment in a young patient. A case report. Spine 24: 489-492, 1999. Zarzycki, D., Rymarczyk, A., Bakalarek, B., Kalicinski, M., Winiarski, A. Surgical treatment of congenital vertebral displacement Type A in the sagittal plane only: a retrospective study involving eleven cases. Spine 27: 72-77, 2002. Anthony, J. P., Rawnsley, J. D., Benhaim, P., Ritter, E. F., Sadowsky, S. H., Singer, M. I. Donor leg morbidity and function after fibula free flap mandible reconstruction. Plast. Reconstr. Surg. 96: 146-152, 1995. Babhulkar, S. S., Pande, K. C., Babhulkar, S. Ankle instability after fibular resection. J. Bone Joint Surg. Br. 77: 258-261, 1995. Babovic, S., Johnson, C. H., Finical, S. J. Free fibula donor-site morbidity: the Mayo experience with 100 consecutive harvests. J. Reconstr. Microsurg. 16: 107-110, 2000. Ganel, A., Yaffe, B. Ankle instability of the donor site following removal of vascularized fibula bone graft. Ann. Plast. Surg. 24: 7-9, 1990. Lee, E. H., Goh, J. C., Helm, R., Pho, R. W. Donor site morbidity following resection of the fibula. J. Bone Joint Surg. Br. 72: 129-131, 1990. Minami, A., Kasashima, T., Iwasaki, N., Kato, H., Kaneda, K. Vascularised fibular grafts. An experience of 102 patients. J. Bone Joint Surg. Br. 82: 1022-1025, 2000. Shpitzer, T., Neligan, P., Boyd, B., Gullane, P., Gur, E., Freeman, J. Leg morbidity and function following fibular free flap harvest. Ann. Plast. Surg. 38: 460-464, 1997. Takakura, Y., Yajima, H., Tanaka, Y., Komeda, T., Tamai, S. Treatment of extrinsic flexion deformity of the toes associated with previous removal of a vascularized fibular graft. J. Bone Joint Surg. Am. 82: 58-61, 2000. Tang, C. L., Mahoney, J. L., McKee, M. D., Richards, R. R., Waddell, J. P., Louie, B. Donor site morbidity following vascularized fibular grafting. Microsurgery 18: 383386, 1998. Vail, T. P., Urbaniak, J. R. Donor-site morbidity with use of vascularized autogenous fibular grafts. J. Bone Joint Surg. Am. 78: 204-211, 1996.

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[24] Youdas, J. W., Wood, M. B., Cahalan, T. D., Chao, E. Y. A quantitative analysis of donor site morbidity after vascularized fibula transfer. J. Orthop. Res. 6: 621-629, 1988. [25] Zimmermann, C. E., Borner, B. I., Hasse, A., Sieg, P. Donor site morbidity after microvascular fibula transfer. Clin. Oral Investig. 5: 214-219, 2001. [26] Pacelli, L. L., Gillard, J., McLoughlin, S. W., Buehler, M. J. A biomechanical analysis of donor-site ankle instability following free fibular graft harvest. J. Bone Joint Surg. Am. 85-A: 597-603, 2003. [27] Fragniere, B., Wicart, P., Mascard, E., Dubousset, J. Prevention of ankle valgus after vascularized fibular grafts in children. Clin. Orthop. 245-251, 2003. [28] Hsu, L. C., Yau, A. C., O'Brien, J. P., Hodgson, A. R. Valgus deformity of the ankle resulting from fibular resection for a graft in subtalar fusion in children. J. Bone Joint Surg. Am. 54: 585-594, 1972. [29] Paluska, D. J., Blount, W. P. Ankle valgus after the Grice subtalar stabilization: the late evaluation of a personal series with a modified technic. Clin. Orthop. 59:137-46.: 137146, 1968. [30] Weiland, A. J., Weiss, A. P., Moore, J. R., Tolo, V. T. Vascularized fibular grafts in the treatment of congenital pseudarthrosis of the tibia. J. Bone Joint Surg. Am. 72: 654-662, 1990. [31] Arai, E., Nakashima, H., Tsukushi, S. et al. Regenerating the fibula with beta-tricalcium phosphate minimizes morbidity after fibula resection. Clin. Orthop. Relat Res. 233-237, 2005. [32] Goldberg, V. M., Shaffer, J. W., Field, G., Davy, D. T. Biology of vascularized bone grafts. Orthop. Clin. North Am. 18: 197-205, 1987. [33] Gonzalez, d. P., Bartolome, d., V, Grana, G. L., Villanova, J. F. Free vascularized fibular grafts have a high union rate in atrophic nonunions. Clin. Orthop. Relat Res. 3845, 2004. [34] Moran, C. G., Wood, M. B. Vascularized bone autografts. Orthop. Rev. 22: 187-197, 1993. [35] Anderson, A. F., Green, N. E. Residual functional deficit after partial fibulectomy for bone graft. Clin. Orthop. Relat Res. 137-140, 1991. [36] Edelman, R., Barbacci, D. Fibular regeneration. J Foot Surg. 31: 368-371, 1992. [37] Bettin, D., Bohm, H., Clatworthy, M., Zurakowski, D., Link, T. M. Regeneration of the donor side after autogenous fibula transplantation in 53 patients: evaluation by dual xray absorptiometry. Acta Orthop. Scand. 74: 332-336, 2003. [38] Kanaya, K., Wada, T., Kura, H., Yamashita, T., Usui, M., Ishii, S. Valgus deformity of the ankle following harvesting of a vascularized fibular graft in children. J. Reconstr. Microsurg. 18: 91-96, 2002. [39] Cornwall, G. B., Thomas, K. A., Turner, A. S., Wheeler, D. L., Taylor, W. R. Use of a resorbable sheet in iliac crest reconstruction in a sheep model. Orthopedics. 25: s1167s1171, 2002. [40] Epstein, N. E., Hollingsworth, R. Does donor site reconstruction following anterior cervical surgery diminish postoperative pain? J Spinal Disord. Tech. 16: 20-26, 2003. [41] Wang, M. Y., Levi, A. D., Shah, S., Green, B. A. Polylactic acid mesh reconstruction of the anterior iliac crest after bone harvesting reduces early postoperative pain after anterior cervical fusion surgery. Neurosurgery. 51: 413-416, 2002.

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In: Degradable Polymers for Skeletal Implants Editors: P.I.J.M. Wuisman and T. H. Smit

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Chapter 12

DEGRADABLE POLYMERS IN CRANIOMAXILLOFACIAL SURGERY U. Eckelt Department for Oral and Maxillofacial Surgery, University Hospital Carl Gustav Carus, Dresden, Germany

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ABSTRACT Biodegradable osteosynthetic materials are a component of clinical application in cranio-maxillofacial surgery. The usage ranges from the cranium to the midface to the mandible. Nowadays, semi-crystalline polylactide (PLLA), amorphous polylactide (PDLLA), polyglycolide (PLGA) and their copolymers [P(L/Ld)LA and PGA] are most frequently applied. Besides the biocompability, the material stability has to precisely correspond to the area of application. Due to the higher mechanical stress of the mandible, more stable materials are necessary. Especially in the condylar process region, the need of more stable material makes an application difficult. The most indications for resorbable material are made in craniofacial surgery and midfacial fractures.

INTRODUCTION In cranio-maxillofacial surgery, titanium plates and screws have been the golden standard for a stable fixation and were examined in numerous studies for more than 30 years [14]. Disadvantages of this method are the possibility of corrosion, titanium incorporation into the tissue, thermal paresthesia, future difficulties in radiologic diagnostics, palpation of the material beneath the skin and the inward migration of the osteosynthetic material in cases of infantile cranio-facial interventions. Further, the possible loosening of screws and inflammation represent an indication for the removal of non-resorbable osteosynthetic material. These disadvantages have been the reason for research regarding the application of biodegradable osteosynthetic materials in the cranio-maxillofacial field for a long time. In 1971, Kulkarni et al. [10] reported about the use of biodegradable material for the fixation of

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craniofacial fractures for the first time. Since that time, various publications evaluated different materials. Semi-crystalline polylactides (PLLA), amorphous polylactides (PDLLA), polyglycolides (PLGA) and their copolymers [P(L/Ld)LA and PGA] are currently most frequently applied [5]. Besides the biocompability, it is especially the stability of biodegradable osteosynthetic materials which is of high significance in the craniomaxillofacial field. In general, two main indications of application are currently differed in cranio-maxillofacial surgery. This is the craniofacial and the midfacial area – where osteosynthesis is performed between rather stable fragments– and the mandibular region, which consists of osseous fragments that are directly exposed to mastication load and movement. In conclusion, biodegradable osteosynthetic materials used in the mandibular area have to withstand higher stability criteria. A further feature of cranio-maxillofacial surgery is that the material is situated directly beneath the skin and, therefore, has to have a volume as thin as possible to avoid patient disturbance by a possible palpation through the skin. This may be a lasting problem for many years depending to the duration of resorption of the material. These general criteria always have to be in mind while applying biodegradable materials and the special information of the manufacturer regarding the material has to be taken into account in the medical indication of application. Mechanical properties, biocompability and the degradation time considerably vary with the different systems. This data depends on the fact whether polymers or copolymers are applied and on the degree of crystallization as well as the molecular weight and on the way of production of the material, e. g. by reinforced techniques (SR).

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SONICWELD®-SYSTEM A disadvantage of biodegradable osteosynthetic materials is the general necessity of thread-cutting after creating a drill hole. This procedure does not only imply a longer time of surgery but also might cause a re-dislocation of the already positioned fragments in many cases. Nevertheless, an absolutely accurate and rectangular insertion of the screws is imperative. While inserting the screws, the screw head is prone to fracture before being completely adjusted. This possible fracture depends on the stability of the applied biodegradable material. In such a case, a new drill hole and thread-cutting will be necessary. Due to this problem, self-tapping resorbable screws have been developed, e. g. Tacker™ system (Inion Ltd. Tampore, Finland) [6]. A completely new method for the fixation of biodegradable osteosynthetic materials is the ultrasound-aided insertion of bioresorbable pins whereupon a melting of the pins into the osteosynthetic material and the retention of the melted resorbable pin material within the cancellous bone occur [17] (Figure 1). These both features cause a very high stability of the osteosynthesis compared with a screw fixation. After the adaptation of the osteosynthetic material at the relocated fracture or the fragments of craniosynostosis, a hole will be drilled and a suitable pin will be inserted in this hole by ultrasound. The SonicWeld® system is based upon the usage of Resorb X® material which consists of 50 % L-lactide and 50 % Dlactide (PDLLA). The pins have a diameter of 2.1 mm and are available in lengths of 4 to 9 mm. Plates or mesh material are available in strengths of 0.3 mm, 0.6 mm and 1.0 mm.

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Animal research and stability tests showed [18] that this osteosynthesis has a higher stability compared to osteosynthesis with resorbable screws and plates. Histological examination showed that the thermal energy caused by the insertion of the pin does not lead to a negative impairment of the bone tissue, surrounding the pin [15]. Besides, there were no signs of adverse tissue reactions.

CRANIOFACIAL SURGERY

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Craniosynostoses are an important area within craniofacial malformations. They result from a premature ossification of the sutures. In craniosynostosis the direction of growth of the cranium vertical to the affected suture is delayed. At the same time, an exaggerated growth occurs towards the affected suture.

Figure 1. Principle of the SonicWeld® system with anchorage of the biodegradable material in cancellous bone.

This causes a disproportion between the volume of the cranium capsule and the growing brain whereby the intracranial pressure increases. In most cases the synostosis is not only limited to the neurocranium but impairs the development of the viscerocranium. The pathogenesis of this disease has been unclear to this day. Moderate craniosynostoses can be clinically less noticeable and do not lead to an increased intracranial pressure. More distinctive cases can imply brain atrophy and cerebral disorders. Especially the optical nerve can be affected so that its pressure damage may result in vision impairment up to blindness. An important clinical symptom is the exophthalmos. Besides, headache is one of the most common neurologic symptoms. In infants, the increasing intracranial pressure is expressed by agitation, sleep disorders, frequent crying, vomitus, drinking and growing disorders. Cerebral seizures are also possible. Besides a neuropediatric examination, the radiological diagnostic

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investigation is essential. Thus the localisation of the synostosis and the evaluation of the extent of the cranial malformation is possible. Besides computer tomography, magnetic resonance imaging is necessary for radiological diagnostics. Due to the impact of neurocranial malformation on the viscerocranium, a stomatognathic examination is extraordinarily important. Kephalometric x-ray images have to be included into the examination. According to basic publications from Tessier [21], the frontoorbital advancement has proved to be of high value in the treatment of craniosynostoses. During this surgery, the cranium will be re-shaped into the configuration it would have under non-pathological growth conditions. In doing so, the intracranial volume will be increased at the same time. Today, the surgical method is largely standardized. The osteotomy lines run along the sutures (Figure 2). In doing so, the growth inhibition is disrupted until reossification so that the development of the cranium can be guided in a more favourable direction.

Figure 2. Frontoorbital advancement.

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The result of the surgical intervention can only be ensured by a stable fixation. Whereas osteosyntheses by wire osteosutures were favoured in the beginning of the treatment of craniosynostoses, titanium plates were preferred later on due to their higher stability [16]. As the appositional growth of the cranium causes a relative movement of the osteosynthetic material in a centripetal direction, the applied non-resorbable osteosynthetic material has to be removed 3 to 6 months postoperatively [20]. Otherwise, a removal would not be possible any more. In order to avoid this disadvantage, the application of biodegradable materials increased. In many cases of craniosynostoses surgery, multiple screws have to be placed for the fixation of the osteosynthetic material. Therefore the disadvantage of time consuming thread cutting is of high relevancy. Another disadvantage is the possible fracture of the screw head before the screw has been completely placed. In such a case, the advantages of the SonicWeld® system are especially noticeable [1]. The easier handling enormously facilitates the connection of the first segments (Figure 3).

Figure 3. Trigonocephalus intraoperatively – fixation by biodegradable material and pin fixation (SonicWeld®).

Experimental examinations showed a reduction of the time of osteosynthesis of 50 % compared to the application of bioresorbable plate and screw systems [18]. In clinical usage, a combination of titanium materials and biodegradable materials is possible. This is especially interesting in cases where extensive movements of parts of the calvarium are necessary. In such a case, a fixation by biodegradable material is advantageous in the frontal area which is difficult to be reached. The stabilisation in the region of the coronal suture can be performed by titanium plates. These plates can be easily removed during a potential scar correction in the cranial area without another exposure of the sensitive frontal area.

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MIDFACIAL FRACTURES

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Nowadays, the fixation by titanium micro and mini plates is the preferred treatment for the fixation of midfacial fractures which are divided into central, lateral and centrolateral midfacial fractures. However, the osteosynthetic material is quite often palpable in the supraorbital and infraorbital area. That causes the patients to ask for a removal of the osteosynthetic material. In contrast to mandibular fractures, the midfacial osseous parts are not actively movable during the mastication process. Especially in this aesthetically sensitive region, it is essential to use biodegradable osteosynthetic materials with a low dimensioned volume. On the one hand, the placement of osteosynthetic material depends on the type of fracture; on the other hand, the application will be generally performed according to the stabilisation of the power trajectories in the midface which are located in the lateral nasal area, in the area of the zygomatic alveolar crista, in the area of the frontal zygomatic suture and in the infraorbital area (Figure 4).

Figure 4. Stabilisation of mid face fractures by biodegradable material.

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Figure 5. Fixation at the frontal zygomatic suture – insertion of a resorbable pin.

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The surgical approach will be executed identically to the usage of titanium material, using a supraorbital (Figure 5) or upper eyelid approach, lower eyelid approach and oral approach in the area of the upper vestibule. The coronary approach can also be used for complex fractures in the frontal face area. A preauricular approach can be necessary for the treatment of zygomatic fractures. Heidemann and Gerlach [7] noticed an ordinary wound and fracture healing in a study of 50 patients and an observation period of 6 to 30 months postoperatively. No second surgery was necessary in any case.

FRACTURES OF THE ORBITAL FLOOR Fractures of the orbital floor are encountered in combination with central or centrolateral midfacial fractures but also as isolated fractures which is referred to as a blowout fracture. The blowout fracture causes a displacement of orbital content towards the maxillary sinus. Fractures of the medial orbital boundary towards the ethmoidal cells are also possible. A displacement of orbital tissue towards the ethmoidal cells is to be observed. Due to the shift of the orbital content, an enophthalmos with the clinical signs of double image occurrence arises. Therefore, a surgical intervention is absolutely necessary. Autogenous bone is often recommended for the reconstruction. The bone must be harvested from the area of the cranial calvarium, from the mandibular ramus or from the iliac crest. Whereas the transplantation of autogenous bone always implicates a donor site defect, the application of alloplastic material unrestricted in availability. Biodegradable material has been used for a long time [2]. There are a lot of references regarding the usage of polydioxanone foils. These foils are inserted by an infraorbital, subciliar or trans-conjunctival approach after the relocation of soft tissue from the maxillary sinus or from the ethmoidal cells. Disadvantages are the rigidity of the material and the secondary anterior displacement especially in the region of the orbital floor. In general, the usage of biodegradable material is indicated for defect sizes up to 2 x 2 cm in the

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orbital floor area. New resorbable mesh materials can be easily adapted after heating and which fixed to the infraorbital edge by pin or screw synthesis. This prevents secondary displacement [9]. The biocompability of the applied material is of a high significance in the orbital area of reconstruction.

MANDIBLE Fractures of the Mandibular Body

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Osteosynthesis of fractures of the mandibular body requires a high degree of osteosynthetic material stability. This has to be taken into account while choosing the material. The stability of different biodegradable materials in fractures of the mandibular body could be demonstrated in several studies [3, 12]. The surgical procedure is comparable to the procedure with titanium mini plates. According to the recommendations of Champy [4], the osteosynthetic material is placed at the mandible by an intraoral approach (Figure 6). The usage of two plates is necessary in the frontal up to the canine area in order to neutralize torsion power whereas only one plate is necessary in the molar and mandibular angle region (Figure 7). Different authors also recommend the usage of two plates for the area of the mandibular angle for reasons of stability [11]. A study regarding the INION® system [13] showed an early exposure of the biodegradable osteosynthetic material in the area of the mandibular angle in cases where plates crossed the oblique line.

Figure 6. Blowout fracture on the right side.

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Figure 7. PDS foil at the orbital floor.

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Hence, the fixation of the resorbable plates at the exterior of the mandible in the area of the mandibular angle is essentially recommended. In this area, they are covered with thicker soft tissue and an early exposure is less likely. However, the screws have to be inserted transbuccally or by a cranked wrench key. Altogether, the osteosynthetic materials currently available for mandibular applications could not be generally established due to their difficult handling. Further it has to be mentioned that the material costs are higher compared to titanium. Apart of the avoidance of a secondary surgery, the more difficult handling is a disadvantage in mandibular applications.

Fractures of the Mandibular Condylar Process In principle, the fractures of the mandibular condylar process are difficult to be reached due to the facial nerve. These fractures are particularly interesting for biodegradable materials because a second surgical approach to remove the osteosynthesis material represents a higher risk of damage to the facial nerve due to scar formation [8]. As the approach to this narrow region offers very little space, the more difficult handling of voluminous biodegradable materials is a further disadvantage. Landes et al. [11] showed a sufficient stability for condylar process fractures. However the condylar process fracture often proves difficult to be repositioned and fixed. It is therefore occasionally prefixed by a titanium mini plate. This plate is removed after the fixation of the first resorbable plate (Figure 8). Rasse et al. [19] demonstrated the sufficient stability under usage of two Resorb-X® mini plates in animal research experiments. In this case, one of the plates was fixed at the retral mandibular edge and the second plate was placed near the semilunar incisure. In order to enable an extended usage of biodegradable materials in the field of fractures of the condylar process, more gracile bioresorbable plate systems with a higher stability have to be developed. This has not been the case up to now.

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Figure 8. Positioning of resorbable plates in mandibular fractures according to Champy.

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CONCLUSION Biodegradable osteosynthetic materials are a component of clinical application in craniomaxillofacial surgery. The usage ranges from the cranium to the midface to the mandible. Nowadays, semi-crystalline polylactide (PLLA), amorphous polylactide (PDLLA), polyglycolide (PLGA) and their copolymers [P(L/Ld)LA and PGA] are most frequently applied. Besides the biocompability, the material stability has to precisely correspond to the area of application. Due to the higher mechanical stress of the mandible, more stable materials are necessary. Problems with handling of biodegradable materials in terms of thread cutting and screw head fracture while placing the screws are solved by the SonicWeld® system. Here a bioresorbable pin is melted into the drill hole by ultrasound and anchors in the cancellous bone. It melts into the osteosynthetic material. In the area of the mandibular condylar process, the current materials are too voluminous as there is only a narrow and unstable surgical approach in order to enable an immediate function.

REFERENCES [1]

[2]

Eckelt, U; Nitsche, M; Müller, A; Pilling, E; Pinzer, T; Roesner, D. Ultrasound aided pinfixation of biodegradable osteosynthetic materials in cranioplasty for infants with craniosynostosis. J. CranioMaxillofac. Surg., 2007; 35, 218-221 Enislidis, G. Treatment of Orbital fractures: The case for treatment with resorbable materials. J. Oral Maxillofac. Surg., 2004, 62: 869-872

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Degradable Polymers in Cranio-Maxillofacial Surgery [3]

[4] [5] [6] [7]

[8]

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[10] [11]

[12]

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[13]

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[17]

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Chacon, GE; Dillard, F; Clelland, N; Rashid, R. Comparison of strains produced by titanium and poly D, L-lactide acid plating systems to in vitro forces. J. Oral. Maxillofac. Surg. 2005, 968-972 Champy, M; Lodde, JP; Jaeger, JH, Wilk, A. Osteosynthéses mandibularies selon la technique de Michelet. I-Bases biomécaniques. Rev. Stomatol., 1976, 77 :569-576 Gerlach, KL. Resorbierbare Polymere als Osteosynthesematerialien. Mund Kiefer Gesichtschir. 2000; 4: 91-102 Haers, PE. Keeping oral and maxillofacial surgeons informed. Int. J. Oral Maxillofac. Surg. 2005; 34: 589 Heidemann, W ; Gerlach, KL. Anwendung eines resorbierbaren Osteosynthesesystems aus Poly(D, L)laktid in der Mund-, Kiefer- und Gesichtschirurgie. Dtsch. Zahnärztl. Z, 2002, 57: 50-53 Hlawitschka, M ; Eckelt, U. Assessment of patients treated for intracapsular fractures of the mandibular condyle by closed techniques. J. Oral. Maxillofac. Surg., 2002, 60: 784791 Kontio, R; Ruuttila, P; Lindroos, L; Suuronen, R; Salo, A; Lindqvist, C; Virtanen, I; Konttinen, YT. Biodegradable polydioxanone and poly(L/D)lactide implants: an experimental study on peri-implant tissue response. Int. J. Oral. Maxillofac. Surg., 2005, 34: 766-776 Kulkarni, RK; Moore, EG; Hegyeli, AF; et al. Biodegradable poly(lactid acid) polymers. J. Biomed. Mater. Res., 1971, 5: 169-181 Landes, C; Ballon, A. Indications and Limitations in resorbable P (L70/30DL)LA osteosyntheses of displaced mandibular fractures in 4.5-Year follow-up. Plast. Reconstr. Surg., 2006, 117: 577-589 Laughlin, R; Block, M; Wilk, R; Malloy, R.; Kent, J. Resorbable plates for the fixation of mandibular fractures: a prospective study. J. Oral Maxillofac. Surg., 2007, 65: 89-96 Leonhardt, H; Demmrich, A; Loukota, R; Müller, A; Mai, R; Eckelt, U. Inion®– versus Titanium-osteosynthesis system – a prospective investigation of mandibular fracture treatment. Br. J. Oral Maxillofac. Surg., in press Luhr, HG. Entwicklung der modernen Osteosynthese. Mund Kiefer Gesichtschir, 2000, 4: 84-90 Mai, R; Lauer, G.; Pilling, E; Jung, R; Leonhardt, H; Proff, P et al. Bone welding – histological evaluation in the jaw. Ann. Anat 2007; 189: 350-355 Mühling, J; Reuther, J; Sörensen N. Operative Behandlung cranio-facialer Fehlbildungen (Operative methods in craniofacial malformations). Kinderarzt, 1984; 15: 1022-1023 Pilling, E; Mai, R; Theissig, F; Stadlinger, B; Loukota, R; Eckelt, U. An experimental in vivo analysis of the resorption to ultrasound activated pins (Sonic weld®) and standard biodegradable screws (Resorb-X®) in sheep. Br. J. Oral Maxillofac. Surg,. 2007, 45: 447-450 Pilling, E; Meissner, H; Jung, R; Koch, R; Loukota, R; Mai, R; Reitemeier, B; Richter, G; Stadlinger, B; Stelnicki, E; Eckelt, U. An experimental study of the biomechanical stability of ultrasound-activated pinned (SonicWeld Rx + Resorb-X ) and screwed fixed (Resorb-X ) resorbable materials for osteosynthesis in the treatment of simulated craniosynostosis in sheep. Br. J. Oral Maxillofac. Surg., 2007, 45: 451-456

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[19] Rasse, M; Moser, D; Zahl, C; et al. Osteosynthesis of condylar neck fractures with resorbable ply(DL)lactide plates and screws. An animal experiment. Br. J. Oral Maxillofac. Surg., 2007, 45: 35-40 [20] Stelnicki, EJ; Hoffman, W. Intracranial migration of microplates versus wires in neonatal pigs after frontal advancement. J. Craniofac. Surg., 1998; 9, 60-64 [21] Tessier P. Osteotomies totales de la face: syndrome de Crouzon, syndrome d’Apert ; oxycéphalies, scapocéphalies, turricéphalies. Ann. Chir. Plast., 1967 ; 2 : 273

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In: Degradable Polymers for Skeletal Implants Editors: P.I.J.M. Wuisman and T. H. Smit

ISBN 978-1- 60692-426-6 © 2009 Nova Science Publishers, Inc.

Chapter 13

ABSORBABLE MATERIALS IN SHOULDER SURGERY Lennart Magnusson1, Jüri Kartus*,2 and Lars Ejerhed2 1

2

Västerås Ortopedpraktik, Västerås, Sweden, NÄL/Uddevalla Hospital, Trollhättan/Uddevalla, Sweden

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ABSTRACT The used of absorbable implants in shoulder surgery is extensive. Both absorbable and non-absorbable implants appear to render good clinical results. However, there are reports of both early and late bony reactions, such as the formation of cysts at the site of implantation, especially after the use of polylactic- or polylevolactic acid polymer implants. Long-term radiographic studies have revealed that the formation of cysts is not as frequent after using polyglyconic acid co-polymer implants. In spite of good early clinical results, it is possible that the surgeon will have to handle inferior bone tissue in the anterior gleniod if revision surgery is required. Due to the frequent formation of radiographically visible cysts, even after seven to eight years, the authors do not recommend the use of polylactic- or polylevolactic acid polymer implants today. In the future, major efforts should be made to develop new polymers with osteoinductive/osteoconductive properties. However, the introduction of new polymers should also include long-term radiographic and clinical assessments in order to detect early and late failures that could be related to the design of the implants or their chemical composition.

BACKGROUND The performance of shoulder reconstruction due to both rotator cuff lesions and posttraumatic instability has become less technically demanding after the introduction of suture anchors. The use of suture anchors and absorbable devices has made procedures such as the Bankart procedure easier, without any negative effect on the results in terms of shoulder stability. [12,26,27] Different types of suture anchor or fixation device have been widely used *

Contact: [email protected].

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in soft tissue repair in the shoulder [1,15,25,30] and both open and arthroscopic procedures can be performed using similar devices. There are some obvious advantages to using bioabsorbable devices that make them a potentially attractive alternative to their non-absorbable counterparts. They do not generate artefacts on MRI and, if they break or become loose, they are not as hazardous to the joint cartilage as metallic implants.[32] The potential for restoring bone tissue after resorption is another advantage that has been described in conjunction with the use of bioabsorbable anchors. However, their biocompatibility is not yet fully understood.

DIFFERENT IMPLANT MATERIALS AND COMPOSITIONS Suture anchors or fixation devices generally consist of polyglyconic acid (PGA), polylactic acid (PLA), poly-L-lactic acid (PLLA) or various co-polymers of these macromolecules and are regarded as bioresorbable/absorbable/biodegradable. [11] Suture anchors made of other materials such as non-absorbable plastic or metal have also been used. PGA and its co-polymers were the first synthetic suture materials (e.g. the Dexon® sutures) and they have been commercially available since the early 1970s or thereabouts. Following the commercial and clinical success of these suture materials, they were also introduced as anchor- or tack-like implants primarily designed for use in the shoulder (Figure 1).

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RADIOGRAPHIC FINDINGS AT THE SITE OF IMPLANTATION Osteolytic changes after using absorbable implants have been reported in oral surgery, fracture surgery and after Bankart reconstruction in the shoulder. [2,3,4,9,10,20,21] When used in oral surgery and for fracture stabilisation, the changes have not appeared to influence the clinical or radiographic healing in the medium and long term. [2,3,4,6,8,24] There are also studies that have reported no lytic or resorptive bone changes following the implantation of absorbable materials. [1,30] Böstman [6] has stated that bicortical perforation during the drilling manoeuvre allows for more complete healing of the drill holes containing absorbable materials. The reason for this might be improved drainage of the absorbable material; in other words, there are two ways out for the debris after bicortical perforation Figure 2 A and B). Even the use of non-absorbable anchors in the shoulder region has led to clearly visible cystic osteolytic changes, as reported by Ejerhed et al. [10] in a non-randomised study comparing plastic polyacetal (Delrin®) implants and PGA co-polymer implants with an identical design. These authors speculated that the cause of the formation of cysts in the case of the absorbable anchors was osmotic action, while it was a foreign-body reaction for the non-absorbable ones. To our knowledge, the use of non-absorbable and absorbable suture anchors in the shoulder region has been compared in a few randomised studies. Warme et al. [30] reported that the drill holes containing non-absorbable anchors had a sclerotic rim of 1 mm and a diameter that was near the anchor size, while the drill holes containing absorbable anchors appeared to have been replaced by bone after one to two years. One weakness in their study is, however, that radiographs were only available for 10/20 patients i.e., 50% drop-outs, which makes it difficult to draw any definite conclusions.

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Figure 1. Example of two different designs of PGA co-polymer fixation devices primarily designed for use in the shoulder. Copyright Jüri Kartus

Figure 2 A. If an implant is retained within the bone tissue, there is no way out for the debris. This allows for both foreign-body reactions and an osmotic process and can result in the formation of cysts. Copyright Catarina Kartus

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Figure 2 B. If, however, there is an opportunity for the debris to drain from the absorbable material, there is less likelihood that cysts will form. Drainage can occur either through bicortical perforation or, in some cases, because the repair is not as tight as the surgeon might wish. Copyright Catarina Kartus

In spite of this, their clinical results were good. In line with the results of Ejerhed et al. [10] and Warme et al., [30] Tan et al. [29] reported that metallic and absorbable anchors both render good clinical results with a failure rate in terms of stability of 9%. However, the minimum follow-up was only 18 months and no radiographs were taken.

RADIOGRAPHIC CLASSIFICATION OF DRILL HOLES For more than 15 years, our research group, which is affiliated to Göteborg University in Sweden, has been focusing special attention on the radiographic appearance of the drill holes in the shoulder region after the implantation of both absorbable and non-absorbable materials. Before the start of the studies, a special evaluation system was developed to make it possible radiographically to classify drill holes and osteolytic changes in conjunction with the drill holes in the shoulder region. This classification system was first published in 1998 (Table 1). [17] The intra-rater reproducibility of the system has been tested and has rendered a good kappa value of approximately 0.8. [19] Using this classification system, radiographic examinations have shown that the drill holes used for implanting PGA co-polymer implants heal in most patients after a period of approximately two years, at least if the implants are placed partly outside the bone tissue. [17,18]

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Table 1. Radiographic evaluation of the drill holes Rating 0 I II III

Description Invisible Hardly visible Visible Visible with cystic changes

A special evaluation system to assess the drill holes after implanting absorbable or nonabsorbable materials was developed by Kartus et al. [17]

EARLY PROBLEMS AFTER IMPLANTING PGA CO-POLYMER IN THE SHOULDER Clinically, it has been reported in the literature that these implants might break or cause synovitis in the short term. Karlsson et al. [15] reported the early breakage of the PGA copolymer tack implant in 1/66 patients, while Burkart et al. [7] reported the corresponding findings, together with massive aseptic synovitis in a case report involving four patients. All the patients in both studies healed uneventfully after arthroscopic intervention. Kartus et al. reported that 2/72 [18] patients had pain in the shoulder and a short period of increased body temperature within three weeks after the implantation of PGA co-polymer implants. For both patients, the reaction subsided without intervention within three days. It appears that the inflammatory reaction after the implantation of PGA co-polymer implants is mild and rare,

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without any negative effect on the clinical outcome.

DOES THE MODE OF IMPLANTATION OF THE ABSORBABLE MATERIAL MATTER? There appears to be a difference in the development of the drill holes depending on whether the PGA co-polymer implants are placed fully intra-osseously or whether the implant is placed partly outside the bone tissue. In the first case, there is probably a greater risk of cysts developing due to an osmotic process and, in the second case, there is greater potential for the drainage of debris as shown in Figure 2 A and Figure 3. It has been shown in several studies that, if a PGA co-polymer implant is placed partly outside the bone tissue, the near complete healing of the drill holes appears to take place in the long term. [16,17,18] In a study by Kartus et al., [17] 56% of the patients had clearly visible or cystic drill holes two years after the full intra-osseous implantation of a PGA co-polymer implant, while the corresponding findings were only made in 23% if the implant was placed partly outside the bone tissue. Magnusson et al. [21] subsequently re-examined the patients who had their implants placed fully intra-osseously eight years after the index operation and found that 33% still had cystic or clearly visible drill holes.

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In the study by Magnusson et al., [21] 16% of the patients still had radiographically visible cystic reactions in the bone eight years after open Bankart reconstruction using PGA co-polymer implants. The reason for this is not fully understood, but, if revision surgery is needed, managing bone of inferior quality might be a problem as a result of using absorbable implants.

DIFFERENCES BETWEEN FAST AND SLOW ABSORBING MATERIALS

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Clinically, there appears to be a conflict of interest between the desire to use an implant that provides safe, long-term fixation to enable the repaired tissue to heal and an implant that resolves quickly without causing any local or systemic damage. PLA and PLLA polymer implants do not resorb as quickly as PGA co-polymer implants. [5]

Figure 3 A and B. If the implant is placed partly outside bone tissue and is also cannulated, there appears to be less risk of cystic changes developing. Copyright Catarina Kartus

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This appears to be beneficial, in terms of both a longer period for the tissues to heal before they break and possibly less synovitis. However, if the period before these implants lose their strength and resorb is very long, they might possibly cause degenerative changes in the glenohumeral joint, as has been shown after loose or incorrectly placed metal implants in the anterior shoulder region. [32] In a randomised study by Magnusson et al., [20] two similarly designed but chemically different absorbable implants consisting of either a synthetic co-polymer of PGA or a polymer of self-reinforced PLLA were compared (Figure 4). In spite of different degradation times, both types of absorbable implant appeared to have sufficient strength during the early period until capsular labrum healing occurred, since the authors found no early re-dislocations during the rehabilitation period in either group. Furthermore, no clinical signs of early synovitis or any increase in C-reactive protein response were found in either group until six weeks postoperatively (unpublished results, Magnusson et al.).

Figure 4. In their randomised study, Magnusson et al. [20] used either self-reinforced PLLA polymer or PGA co-polymer implants. Copyright Elsevier

The degradation of both these polymers is known to take place through hydrolysis and enzymatic action and the PGA co-polymer degrades more rapidly than the PLLA polymer does. [5, 31] After two years, tacks of PLLA had caused larger radiographically visible drill holes including cystic changes when compared with the PGA co-polymer tacks (Figure 5 A and B). The same patients subsequently underwent X-ray examinations seven to eight years after the index procedure. At that point, 100% of the drill holes after implanting the PGA copolymer had healed, while more than 50% of the patients in whom PLLA implants had been used had cystic or clearly visible drill holes (unpublished material, Magnusson et al.).

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change to 6 months instead of From 6 mån

change to 2 years instead of From 2 år Figure 5 A and B. Example of radiographs obtained at six months and two years in a female patient who underwent an arthroscopic Bankart procedure using self-reinforced PLLA implants. At two years, the formation of cystic changes at the site of implantation can still be seen. Copyright Elsevier

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In the above two-year study by Magnusson et al., [20] the authors were not able to identify any correlation between the development of radiographically visible degenerative changes and osteolytic reactions in the bone. Furthermore, the failure rate in terms of stability was approximately 15% in both groups, which is well on a par with the long-term results after open Bankart reconstruction. [22] However, it might be the case that, during a longer followup, there would be a correlation between the development of degenerative changes and the presence of osteolytic changes in the shoulder region. In the literature, there are other reports of cystic formations in the shoulder region after the implantation of absorbable materials. In a case report, Spoliti [28] found osteolytic changes in the superior pole of the glenoid 10 months after SLAP lesion repair using a single PLLA suture anchor. The author speculated that a decrease in the pH in the surrounding tissue led to a foreign-body reaction and osteolysis. Correspondingly, Müller el al. [23] found osteolysis in the anterior glenoid in 7/15 patients four months after the implantation of a poly-L/DL-lactic 70/30 implant during open shoulder stabilisation. After two to five years, three of these patients still displayed an unchanged radiographic situation at the site of implantation.

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CLINICAL IMPLICATIONS Considering the fact that absorbable implants of PLLA are frequently used in current shoulder surgery, these results are of great interest. If 10-20% of patients might require revision surgery within a five-year period after the index operation, the quality of the bone tissue in the anterior glenoid appears to be of major interest. [16] So, even though the early clinical results after arthroscopic Bankart procedures using absorbable implants are good, as reported in some studies, [13,14,26] it is important to consider these findings of clearly visible and cystic drill holes in the long term. Due to the long-term radiographic findings after the implantation of PLLA polymer implants in the shoulder region, we have now discontinued the use of these implants in favour of either PGA co-polymer or metal ones.

THE FUTURE In the future, a major effort should be made to develop new materials which allow full bony in-growth at the site of implantation. Recently, steps towards this have been taken through the introduction of new osteoconductive materials on the market. It is also important to make long-term clinical and radiographic evaluations of all the new biocompatible materials that are introduced on the market in order to evaluate both the early and late clinical and radiographic findings after the use of absorbable implants in the shoulder region.

ACKNOWLEDGEMENTS The authors thank Catarina Kartus for the illustrations.

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[10]

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[16]

Barber FA, Snyder SJ, Abrams JS, Fanelli GC, Savoie FH, 3rd. Arthroscopic Bankart reconstruction with a bioabsorbable anchor. J. Shoulder Elbow Surg. 2003;12:535-538. Bergsma EJ, Rozema FR, Bos RR, de Bruijn WC. Foreign body reactions to resorbable poly(L-lactide) bone plates and screws used for the fixation of unstable zygomatic fractures. J. Oral Maxillofac. Surg. 1993;51:666-670. Bergsma JE, de Bruijn WC, Rozema FR, Bos RR, Boering G. Late degradation tissue response to poly(L-lactide) bone plates and screws. Biomaterials. 1995;16:25-31. Bostman O, Hirvensalo E, Makinen J, Rokkanen P. Foreign-body reactions to fracture fixation implants of biodegradable synthetic polymers. J. Bone Joint Surg. [Br]. 1990;72:592-596. Bostman OM. Absorbable implants for the fixation of fractures. J Bone Joint Surg [Am]. 1991;73:148-153. Bostman OM. Osteolytic changes accompanying degradation of absorbable fracture fixation implants. J. Bone Joint. Surg. [Br]. 1991;73:679-682. Burkart A, Imhoff AB, Roscher E. Foreign-body reaction to the bioabsorbable suretac device. Arthroscopy. 2000;16:91-95. Cheung LK, Chow LK, Chiu WK. A randomized controlled trial of resorbable versus titanium fixation for orthognathic surgery. Oral Surg. Oral Med. Oral Pathol. Oral Radiol. Endod. 2004;98:386-397. Ejerhed L, Kartus J, Funck E, Kohler K, Sernert N, Karlsson J. Absorbable implants for open shoulder stabilization: a clinical and serial radiographic evaluation. J. Shoulder Elbow Surg. 2000;9:93-98. Ejerhed L, Kartus J, Funck E, Kohler K, Sernert N, Karlsson J. A clinical and radiographic comparison of absorbable and non-absorbable suture anchors in open shoulder stabilisation. Knee Surg. Sports Traumatol. Arthrosc. 2000;8:349-355. Gunja NJ, Athanasiou KA. Biodegradable materials in arthroscopy. Sports Med Arthrosc. 2006;14:112-119. Karlsson J, Jarvholm U, Sward L, Lansing O. Repair of Bankart lesions with a suture anchor in recurrent dislocation of the shoulder. Scand. J. Med. Sci. Sports. 1995;5:170174. Karlsson J, Kartus J, Brandsson S, Magnusson L, Lundin O, Eriksson BI. Comparison of arthroscopic one-incision and two-incision techniques for reconstruction of the anterior cruciate ligament. Scand. J. Med. Sci. Sports. 1999;9:233-238. Karlsson J, Kartus J, Ejerhed L, Gunnarsson AC, Lundin O, Sward L. Bioabsorbable tacks for arthroscopic treatment of recurrent anterior shoulder dislocation. Scand. J. Med Sci Sports. 1998;8:411-415. Karlsson J, Magnusson L, Ejerhed L, Hultenheim I, Lundin O, Kartus J. Comparison of open and arthroscopic stabilization for recurrent shoulder dislocation in patients with a Bankart lesion. Am. J. Sports Med. 2001;29:538-542. Kartus C, Kartus J, Matis N, Forstner R, Resch H. Long-term independent evaluation after arthroscopic extra-articular Bankart repair with absorbable tacks. A clinical and radiographic study with a seven to ten-year follow-up. J. Bone Joint Surg. Am. 2007;89:1442-1448.

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[17] Kartus J, Ejerhed L, Funck E, Kohler K, Sernert N, Karlsson J. Arthroscopic and open shoulder stabilization using absorbable implants. A clinical and radiographic comparison of two methods. Knee Surg Sports Traumatol. Arthrosc. 1998;6:181-188. [18] Kartus J, Kartus C, Povacz P, Forstner R, Ejerhed L, Resch H. Unbiased evaluation of the arthroscopic extra-articular technique for Bankart repair: a clinical and radiographic study with a 2- to 5-year follow-up. Knee Surg. Sports Traumatol. Arthrosc. 2001;9:109-115. [19] Magnusson L. Posttraumatic recurrent anterior shoulder instability.Aspects of surgical techniques, implants and diagnostic methods. Thesis, Göteborg University. 2005. [20] Magnusson L, Ejerhed L, Rostgard-Christensen L, Sernert N, Eriksson R, Karlsson J, et al. A prospective, randomized, clinical and radiographic study after arthroscopic Bankart reconstruction using 2 different types of absorbable tacks. Arthroscopy. 2006;22:143-151. [21] Magnusson L, Ejerhed L, Rostgard L, Sernert N, Kartus J. Absorbable implants for open shoulder stabilization. A 7-8-year clinical and radiographic follow-up. Knee Surg. Sports Traumatol Arthrosc. 2006;14:182-188. [22] Magnusson L, Kartus J, Ejerhed L, Hultenheim I, Sernert N, Karlsson J. Revisiting the Open Bankart Experience: A Four- to Nine-Year Follow-up. Am.. J. Sports Med. 2002;30:778-782. [23] Muller M, Kaab MJ, Villiger C, Holzach P. Osteolysis after open shoulder stabilization using a new bio-resorbable bone anchor: a prospective, non-randomized clinical trial. Injury. 2002;33 Suppl 2:B30-36. [24] Pelto-Vasenius K, Hirvensalo E, Bostman O, Rokkanen P. Fixation of scaphoid delayed union and non-union with absorbable polyglycolide pin or Herbert screw. Consolidation and functional results. Arch. Orthop. Trauma Surg. 1995;114:347-351. [25] Potzl W, Witt KA, Hackenberg L, Marquardt B, Steinbeck J. Results of suture anchor repair of anteroinferior shoulder instability: a prospective clinical study of 85 shoulders. J. Shoulder Elbow Surg. 2003;12:322-326. [26] Resch H, Povacz P, Wambacher M, Sperner G, Golser K. Arthroscopic extra-articular Bankart repair for the treatment of recurrent anterior shoulder dislocation. Arthroscopy. 1997;13:188-200. [27] Richmond JC, Donaldson WR, Fu F, Harner CD. Modification of the Bankart reconstruction with a suture anchor. Report of a new technique. Am. J. Sports Med. 1991;19:343-346. [28] Spoliti M. Glenoid osteolysis after arthroscopic labrum repair with a bioabsorbable suture anchor. Acta Orthop. Belg. 2007;73:107-110. [29] Tan CK, Guisasola I, Machani B, Kemp G, Sinopidis C, Brownson P, et al. Arthroscopic stabilization of the shoulder: a prospective randomized study of absorbable versus nonabsorbable suture anchors. Arthroscopy. 2006;22:716-720. [30] Warme WJ, Arciero RA, Savoie FH, 3rd, Uhorchak JM, Walton M. Nonabsorbable versus absorbable suture anchors for open Bankart repair. A prospective, randomized comparison. Am. J. Sports Med. 1999;27:742-746.

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[31] Vert M, Christel P, Chabot F, Leray J. Bioabsorbable plastic materials for bone surgery. In: Hastings GW, Ducheyne P, editors. Macromolecular Biomaterials. Boca Raton: CRC Press, 1984:119-142. [32] Zuckerman JD, Matsen FAd. Complications about the glenohumeral joint related to the use of screws and staples. J. Bone Joint Surg. [Am]. 1984;66:175-180.

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Chapter 14

DEGRADABLE POLYMERS IN HAND SURGERY Abigail R. Hamilton1, Chaitanya S. Mudgal *2 and Jesse B. Jupiter2 1

Harvard Combined Orthopaedic Surgery Residency Program Massachusetts General Hospital, Boston, MA, USA 2 Massachusetts General Hospital, Department of Orthopaedic Surgery Harvard Medical School, Yawkey Center, Boston, MA, USA

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ABSTRACT Implants used in the hand and wrist have to satisfy certain criteria. They should be small, biomechanically strong enough to withstand loading in the hand and wrist and should have a low profile. In addition to these features, it is extremely useful, if they are inert enough to avoid causing local soft tissue or bony reaction and therefore do not require additional procedures for removal. Bioabsorbable implants fulfill all these criteria, and although their use in the hand and wrist is still in its infancy, early data suggest that the rates of success and complications associated with their use, are comparable to that associated with their metal counterparts. As costs associated with their production and use reduce, and more data regarding their efficacy become available, it appears that universal acceptance will follow.

INTRODUCTION Over the past few decades, there has been a shift in the treatment of hand fractures from non-operative immobilization techniques to a more surgically oriented approach. An important component of this shift has been the production of mini- and micro-plating systems and implantation devices specific to the complex needs of the hand. The interplay of intricate *

Address for correspondence:Chaitanya S. Mudgal, MD, MS(Orth.), M.Ch(Orth.) Orthopaedic Hand Service, Massachusetts General Hospital,Instructor, Orthopaedic Surgery, Harvard Medical School, Yawkey Center, Suite 2100,55 Fruit Street, Boston, MA 02114., Telephone: 617-643-3945, Fax: 617-724-8532, Email: [email protected].

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ligamentous, capsular and tendinous structures of the hand make development of these implants challenging since implant strength needs to be balanced with a low profile and small size. The first mini and micro-plating systems were initially developed for use in craniomaxillofacial surgery and were then transferred to use in surgical procedures on the hand. The use of bio-absorbable compounds has followed this same progression with the most extensive use documented and studied within the field of craniomaxillofacial surgery. Long-term follow-up of over ten years of use is available mainly within craniomaxillofacial literature [5] and remains largely unavailable within hand surgery due to its relatively new implementation. The clinical results available are mainly composed of small cohorts of case series or retrospective reviews. Large cohorts of patients have been reviewed retrospectively but include fractures from all anatomic areas and are not specific to hand surgery [8, 33]. Suture materials were the first bio-absorbable devices used within surgery but since the late 1980s the use of other bio-absorbable devices including pins, rods, suture anchors, plates and screws in procedures specific to the hand has been increasing. Interest in bio-absorbable compounds within hand surgery arises from its potential ability to have strength to maintain anatomic reduction during osseous healing and then degrade when no longer required therefore eliminating the potential need for future hardware removal. Another potential benefit is the radiolucency of bio-absorbable implants allowing for ease in radiographic follow-up in the assessment of osseous healing. The implants also do not interfere with computed tomography (CT) or magnetic resonance (MR) imaging causing the scatter effect typically seen with metallic implants. Rigid metallic devices are known to be associated with osteopenia deep to the plate that may lead to re-fracture; this osteopenia and its effects may also be decreased by the less rigid fixation of bio-absorbable devices. The new technology is not without its own complications, most notably local irritation of soft tissues from foreign body reactions resulting in local swelling, effusions and sinus tract formation [8, 39]. Another limitation to widespread use is the cost and restricted availability of implants compared with metallic implants. Due to the decreased strength of bioabsorbable implants, they are not designed to be either self-drilling or self-tapping [39]. Previous experience and practice is strongly recommended in order to achieve the best clinical outcomes, as their use may be more technically demanding [39]. Accurate fracture reduction is essential as interfragmentary compression with miniscrews is not possible and bioabsorbable pins are less forgiving and harder to remove than Kirschner wires requiring redrilling in some cases. The proper rationale must be applied in each surgical case to determine if a bioabsorbable implant may be of benefit to a patient. However, with further materials science research, industry production of implants as well as ongoing basic science and clinical research, bio-absorbable implants may offer new possibilities in the armamentarium of treatment options for hand surgeons.

BIO-ABSORBABLE COMPOUNDS As a full review of bio-absorbable compounds is available in Chapter X, this will be a brief overview of bio-absorbable compounds available for use in surgery specific to the hand.

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Of the numerous bio-absorbable polymers available for medical use, in-depth study of osteofixation devices has focused on high molecular weight polyhydroxyacids formed by ring-opening polymerization of cyclic diesters, polyglycolide [PGA], polylactide [PLA] and their copolymers. PLA is the cyclic dimer of lactic acid existing as two optical isomers, D and L. PLA is seen in implants as polylevolactic acid (PLLA) or various ratios of copolymers of polylevo- and polydextrolactic acid (P(L/DL)LA). The strength characteristics of these polymers have been further enhanced with the development of self-reinforcing (SR-devices). These compounds have been used in a multitude of procedures in surgical treatment of hand disorders given their biocompatibility, inherent mechanical properties and the ability to adjust degradation rates by varying their composition. Biodegradation: Degradation of bio-absorbable polymers occurs mainly via hydrolysis. However, in vivo testing has also implicated enzymatic degradation processes to a lesser extent [43]. In vivo, PLA and PGA are degraded to lactic acid and glycine which produce carbon dioxide and water via the citric acid cycle with byproducts excreted mainly via respiration. The acidic byproducts produced in the initial stages (lactic acid and glycine) further act to catalyze the degradation over time as well as to create an acidic environment relative to surrounding tissues, an occurrence which has been theorized to produce an environment adverse to bacterial ingrowth [2, 5]. Final elimination of polymeric debris is thought to be a function of macrophages and giant cells. Degradation rates vary, with PLLA having the longest degradation time due to the crystallinity and hydrophobicity of the compound from methyl groups, while PGA degrades faster due to the hydrophilic nature of the compound. As with any compound, the chemical make-up, the porosity, surface to volume ratio and sterilization method affect degradation time. Specific to the bio-absorbable polymers used are the size of the polymer, geometric isomerization, monomer concentration and crystallinity. Work by Vasenius et al. shows that in vivo degradation is faster than in vitro and is also increased based on host factors such as vascularity of surrounding tissues, mechanical stress and implantation site [43]. In this study, degradation was noted to be faster in well vascularized cancellous bone than in subcutaneous tissue. Cellular enzymes have been described to enhance degradation of polymers and a vascular environment rich in cellular elements may explain the faster rate of degradation in bone [42]. Loss of strength occurs prior to macroscopically detectable degradation of implants [33]. Clinically, the absorption time for SR-PGA is 6-12 months [34]. Pure PLLA make take 5 years or more to fully degrade and P(L/DL)LA implants between 2-3 years [6]. Mechanical properties: A multitude of factors affect the mechanical properties of bioabsorbable implants. The most important being their biochemical structure and composition, the technique used to produce the implant, and sterilization technique. The homopolymer of the L isomer of PLA (PLLA) is a semicrystalline polymer which exhibits high tensile strength and low elongation. This results in a high modulus that makes them more suitable for load-bearing applications required in orthopedic fixation and sutures. These polymers exhibit a glass-transition temperature. Above this temperature the material is soft and malleable. In non-SR devices, heating above this transition temperature is required for plate molding. PLLA is about 37% crystalline with a glass-transition temperature of 60— 65°C. P(L/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 organized crystalline structure. This effectively decreases the modulus as the percentage of D-isomer presence

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increases. PGA is an amorphous material with a glass-transition temperature of approximately 36°C. Clinical studies reveal that in vivo, strength of SR-PLLA implants decreases overtime and becomes equivalent to the strength of cancellous bone in 36 weeks [24]. Complete mechanical strength of SR-P(L/DL)LA 70/30 occurs in approximately four months.

AVAILABLE IMPLANT DEVICES Implant production using bio-absorbable polymers is performed via multiple techniques. Production via melt-extrusion, injection molding and compression molding have all been reported in early production. Implants produced via these techniques are usually brittle and flexible and are unsuitable for areas of high stress. In 1992 Tormala introduced selfreinforcing as a manufacturing technique revolutionizing the use of bio-absorbable implants in osteosynthesis. This involves formation of a composite polymer structure with reinforcing units of the same chemical structure laid in parallel with a binding matrix. This technique produces high-strength implants that can be molded at room temperature with routine plate bending techniques as opposed to non-reinforced plates that require heating above their glass transition temperature prior to contouring [37, 39]. This self-reinforcing technique also allows for sterilization with γ-irradiation without significant loss of strength of the implant.

CLINICAL APPLICATIONS

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Distal Radius In recent history there has been a large influx of new implant designs available commercially for the treatment of distal radius fractures. While the role of internal fixation has increased in the treatment of unstable fractures with this influx, it is not without its own complications. Flexor and extensor tendon irritation have been reported at a rate of over 10 % [4, 11, 12, 19, 32] and tendon ruptures have also been reported. Recent introduction of a bioabsorbable dorsal distal radius plate offers a potential solution to the need for implant removal because of persistent tendon irritation which has been reported at a rate of 12-23% [11, 12, 19, 32]. Gangopadhyay et al. reported results of treatment of 26 unstable, intraarticular dorsally displaced fractures of the distal radius with a bioabsorbable dorsal distal radius plate (ReUnite) and calcium phosphate (Biobon) bone substitute. The Reunite plate is a copolymer of PLLA and PGA (82%/18%) with a height of 2.5 mm available in small and large sizes. The implant is heated above the glass transition temperature with a portable heat plate to produce malleability for contouring. Five patients lost the reduction achieved at time of surgery, one at 6 weeks and four between 6 and 12 weeks. On re-operation, in the patient who lost reduction at 6 weeks the plate was found to be broken and was revised with a metal plate and iliac crest bone graft. These authors noted severe dorsal comminution in all patients who lost operative reduction and concluded that use of this plating system should not be recommended for fracture fixation, if a metaphyseal gap of greater than 7 mm is noted

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following open fracture reduction. Other complications included rupture of the long extensor of the thumb and long finger, two cases of extensor tenosynovitis that resolved with 2 weeks of anti-inflammatory medication and splinting and two cases of inflammatory reaction to plate resorption between 8 and 11 weeks. In the cases of the foreign body reactions, one patient was treated with aspiration and one required a return to the operating room for a formal debridement. Both of these patients had an excellent functional result at final review [15]. Further study of bio-absorbable implants in treatment of distal radius fractures will help determine their potential role. Patient perception in the United Kingdom is promising for this research. One hundred consecutive adults with distal radius fractures were interviewed and given detailed information regarding bioabsorbable plating vs. metal plating techniques with 95% reporting that they appreciated the bioabsorbable feature, 91% felt that potential need for removal was the most negative aspect of metal plating techniques and 80% stated they would be willing to participate in a randomized controlled trial to compare the use of the two. The most common concern reported by 29% of patients regarding the bioabsorbable implant was regarding the strength [25].

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WRIST ARTHRODESIS Wrist arthrodesis is considered the gold-standard of treatment options for painful degenerative joint disease of the wrist. Multiple fixation techniques have been used to effectively stabilize the wrist while fusion occurs, including pins, staples as well as metal plate and screw fixation [18, 27, 31]. Arthrodesis procedures are valuable in both osteoarthritis and rheumatoid disease of the wrist. Rates of fusion throughout the literature reach 96-100% among the various techniques; however complication rates up to 23% have been reported [31]. Due to the relatively common occurrence of secondary procedures for removal of metal pins, plates or screws for migration or discomfort, bio-absorbable implant use in wrist arthrodesis could offer a significant benefit in this area [31]. The most commonly performed surgical technique implementing bio-absorbable implant was initially described by Juutilainen and Patiala in 1995. This technique involves implantation of a SR-PLLA rod (3.2 x 50 mm) in pre-drilled 3.2 mm holes in the distal radius and capitate with the resected ulnar head being utilized as bone graft. Juutilainen and Patiala (1995) described fusion procedures in multiple joints (18 wrist, 18 hand, 6 talocrural, 11 subtalar-calcaneoucuboid-talonavicular joint) in 47 patients with rheumatoid arthritis between 1989 and 1994. They obtained a 100% fusion rate in the wrist and hand with two reported non-unions of the talocrural joint. Long term results were reviewed by the same group in 21 fusions performed on 18 patients between 1991-1996 with a mean follow-up of 5.4 years with one patient developing a non-union and one patient continued to have severe intermittent pain despite clinical and radiographic evidence of union. In review, complications including infection, median nerve compression, sinus tract formation or aseptic swelling did not occur[46]. Voutilainen et al. (2002) further reported results on 24 wrist arthrodeses in 18 patients with rheumatoid arthritis performed between 1997-2000 using the same technique and obtained fusion in 21 out of 24 cases. One non-union which required re-operation was noted

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to have rod migration into the medullary canal of the proximal radius and two non-unions did not require further treatment [45]. Satisfactory pain relief was noted in 22 of 24 patients mirroring results reported with metal plates and screws [27]. Patients were followed for an average of 1.7 years and no reports of foreign body reaction were noted. Other than nonunion, there were no other complications. These initial results are encouraging, and suggest that use of bio-absorbable implants in wrist arthrodesis is as effective at providing stability during fusion as metal implants. They do however, eliminate the morbidity associated with secondary removal of hardware. There is still limited data on fusion in non-rheumatoid patients and long-term follow-up is limited in multiple centers given the new introduction of bio-absorbable implants.

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SCAPHOID Treatment of delayed or non-union of scaphoid fractures using various bioabsorbable implants has been reported in literature. Use of quinine dye-coloured SR-PGA pins coated with PDS (2.0 mm diameter) was compared to Herbert Screws (4.0mm diameter) in 34 patients with delayed union or non-union of the scaphoid [28]. Of the 24 patients available to long-term follow-up, (14 in the SR-PGA group and 10 in the Herbert screw group) high complication rates were noted in patients in both groups. Transient sinus formation was reported in 25% of the SR-PGA pin group which was attributed to the color dye and poor vascularity of the scaphoid bone leading to fewer phagocytic cells to absorb the degradation products. Two patients who had Herbert screws placed required removal of the implant due to radiocarpal penetration of the screw. Union rate was reported to be 64% in the SR-PGA group and 60% in the Herbert screw group. However, functional outcome was better in the Herbert screw group. Yamamuro et al. (1994) reviewed six cases of scaphoid non-union treated with PLLA devices, and noted that union was obtained in all six cases. In their review, there were no noted foreign body reactions [47]. A review of treatment of 12 patients with non-union of the scaphoid with either a SR-PLLA lag screw (6 patients) or two SR-PLLA pins (6 patients), reported a 100% union rate with average time to union of 4.5 months [1]. Four patients had clear drainage from the wound post-operatively that resolved without intervention. These authors noted increased cost to be the major drawback of use of bioabsorbable implants in the treatment of scaphoid nonunion. Kujala et al. (2004) reported results using non-cannulated 2.0-2.7mm SR-PLLA 70/30 screws in the treatment of 6 patients with scaphoid fractures (3 patients) and non-unions (3 patients). Union was obtained in 5 of 6 patients with one persistent non-union. The authors concluded that bioabsorbable SR-PLLA screws might offer a viable alternative in the treatment of scaphoid non-unions and that the development of cannulated screws over a radio-opaque guide wire may allow more accurate intra-operative positioning of these screws which are radiolucent [21].

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CARPUS Trapeziometacarpal (TMC) arthritis: Symptomatic arthritis of the TMC joint often requires surgical intervention. There are a variety of procedures performed including but not limited to ligament reconstruction, osteotomy of the first metacarpal, silicone elastomer arthroplasty, trapezial excision with tissue interposition with or without ligament reconstruction, and total joint arthroplasty. A biodegradable T-shaped TMC device made of a degradable polycaprolactone-based polyurethaneurea (Artelon; Artimplant AB, Sweden) has been used in clinical pilot studies in Sweden [26]. The Artelon device serves two purposes: (1)to resurface the distal part of the trapezium and (2)to stabilize the TMC joint by augmentation of the joint capsule without subsequent shortening of the thumb that may cause decreased pinch strength [17, 36]. The complete hydrolysis of the spacer takes approximately six years [16]. When the hydrolysis is completed, part of the degraded material (approximately 50% of the initial weight) remains incorporated at the site of implantation [26]. Fifteen patients with radiographically verified, isolated stage 3 TMC osteoarthritis as described by Eaton and Glickel [13], were included in an open, controlled prospective study [26]. Five patients received the Artelon TMC Spacer anchored to bone with osteosutures, five patient were treated with tendon arthroplasty using the long abductor of the thumb (APL), and five patients received the Artelon TMC Spacer anchored with titanium screws. Implantation of the Artelon TMC Spacer was performed after removal of the distal 2 mm of the trapezium. A transient inflammatory reaction with moderate local swelling and tenderness at 2 weeks post-operatively that resolved without intervention, was noted in two patients who received the spacer. No other post-surgical complications were reported. Outcomes at three years were equivalent in terms of pain relief, ability to flatten the hand and range of motion among the three cohorts. The pinch strength measured by tripod pinch and key pinch (lateral pinch) was significantly higher in the spacer group compared with the APL group. Histologic specimens taken from one patient at 6 months (during removal of a titanium screw owing to discomfort) showed bone in contact with Artelon fibers without intervening structures as well as soft tissue in-growth into the woven structure at the periphery. There were no chronic inflammatory cells or any evidence of foreign-body response to the spacer. These initial results are promising and further clinical assessment of this device is ongoing.

METACARPAL / PHALANGEAL INJURIES Fracture Fixation Treatment of metacarpal and phalangeal fractures requires a balance between providing mechanical stability to allow for early mobilization during post-operative rehabilitation while causing the least disruption to surrounding soft tissues. The deformity in metacarpal fractures is typically apex dorsal while proximal phalangeal fractures typically deform in an apex volar pattern due to the deforming effects of the extrinsic and intrinsic musculature of the hand. In fractures of either bone there are variable degrees of rotational or shortening forces depending

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on the specific fracture pattern [7]. Any form of fracture fixation must be sufficiently strong to counteract these forces. Kirschner wire (K wire) fixation is considered by many to be a highly reproducible and useful method of operative treatment of fractures of tubular bones of the hand given the limited amount of soft tissue exposure required for placement. This treatment method is not without complications. Unless these are left protruding outside the skin, they require removal which may increase risk to soft tissues. On the other hand, when they are cut long to increase the ease of removal they can cause skin ulceration and be associated with pin tract infection. The inability to cut them flush to the bone may also interfere with free gliding of tendons inhibiting early range of motion rehabilitation protocols. The use of bio-absorbable implants in the treatment of metacarpal and phalangeal fractures has many potential benefits. They do not require a later procedure for removal. They can be cut flush to the bone to decrease soft tissue irritation and may decrease the rate of infection as they do not leave the skin open to potential pin site infections. They are also radiolucent which aids in radiographic assessment of healing. While not in widespread use, initial studies are promising in that bio-absorbables may offer another treatment choice in the management of hand fractures. In a cadaveric biomechanical study Fitoussi et al. (1998) compared the mechanical rigidity of K-wire to bio-absorbable pin fixation in proximal phalangeal fractures. Transverse and oblique osteotomy models were tested. Transverse fractures were treated with either two cross pins or one oblique pin with wire loop. Oblique fractures were treated with either two cross pins, two or three oblique pins. Metallic and bio-absorbable implants were found to be mechanically comparable under apex volar bending and compression forces. Bio-absorbable pins failed under torsional loads significantly earlier than the metal pins [14]. In a combined biomechanical and clinical study on extra-articular fractures in the hand 1.5 mm PGA rods (Biofix: Bioscience Ltd, Tampere Finland) were found to have 73% reduction in bending strength when implanted into bone and 64.4% reduction after implantation into subcutaneous tissue [22]. At three weeks all implants extracted were fragmented. The clinical trial consisted of 30 patients randomized to treatment either with 1.5mm or 2mm PGA rods or K-wire with wire loop added for further stability given the bio-absorbable implant was known to degrade to zero strength at three weeks. Patients were followed to 24 weeks and bony union was noted in all patients except one in the study group that became displaced at five weeks requiring reoperation attributed to faulty wiring technique and one patient in the control group that required re-operation for loss of reduction following K-wire migration. Both patients requiring re-operation went onto fracture healing by eight weeks. All patients except the two requiring re-operation returned to work ten to sixteen weeks after injury and no allergic reactions were noted in the study group. With the development of self-reinforced bio-absorbable implants, the strength of fixation is increased and maintained for longer periods than the PGA pins tested in initial studies. The biomechanical results obtained with the use of non-reinforced bio-absorbable plates have also been less promising [10]. Bio-absorbable self-reinforced miniplates and screws have been developed for applications in craniomaxillofacial surgery and have recently been adapted for use in hand surgery. Waris et al. (2002) performed a biomechanical study using a pig metacarpal osteotomy model investigating the fixation stabilities of self-reinforced bioabsorbable SR-PLLA pins, and SR-P(L/DL)LA 70/30 screws and miniplates compared with those of standard metallic

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fixation. In this study, two interfragmentary 2.0 mm SR-P(L/DL)LA 70/30 screws provided similar fixation rigidity to two 1.5mm K-wires. SR-P(L/DL)LA 70/30 screws and 1.5mm SRPLLA pins also showed similar rigidities in palmar and dorsal apex loading tests; however the screws were noted to have significantly increased rigidity when tested in lateral apex bending and in torsion. The SR-P(L/DL)LA screws did not show any statistically significant difference from titanium lag screws under bending or torsional stresses. The use of a single interfragmentary bio-absorbable screw provided low rotational rigidity as the bio-absorbable miniplate implant screws can not be applied as a lag screw barring interfragmentary compression. SR-P(L/DL)LA and titanium plating showed the highest mean values of failure torque and did not differ statistically from each other but the titanium plating had statistically significant greater rigidity [40]. While initial biochemical studies present data to suggest that newer self-reinforced bioabsorbable implants offer mechanical stability only slightly less than that of metallic implants [14, 23, 30, 38, 40] there are very few clinical studies available assessing the in-vivo behavior of these devices. In a case report of treatment with self-reinforced miniplates in complex hand injuries of three patients, all three patients went onto union without signs of plate failure [41]. There were no clinical signs of adverse reactions including sterile abscess formation, local irritation or swelling indicative of foreign body reactions to the implants. Arata et al. (2003) reported a series of 26 cases of digital replantation using an intramedullary PLLA rod for the fixation of diaphyseal fractures. One case of transient bone resorption was observed but all patients went on to bony union and there were no reports of infection [3]. The ability to cut bio-absorbable implants directly at a bony surface presents an obvious potential benefit in the treatment of peri- or intra-articular fractures. Pelto-Vasenius et al.. (1996) reported clinical results in the treatment of 13 metacarpal and phalangeal fractures, 12 of which were intra-articular using SR-PLLA and SR-PGA pins of 1.1, 1.5 and 2.0mm diameter. All fractures went on to bony union with report of only one minor re-displacement in a comminuted phalangeal fracture. Results of treatment specific to Bennett’s fracture using similar methodology by the same surgeon were poor with only two of five patients having satisfactory results [29]. The treatment of hand fractures with bio-absorbable implants is still in its infancy, and as the experience of surgeons grows and industry follows with further development of devices, larger scale randomized controlled clinical studies are required to better understand the potential benefits and pitfalls of this new technology.

Soft Tissue Injury The use of bio-absorbable suture anchors in the treatment of soft tissue injuries is commonly used in arthroscopic procedures of both the knee and shoulder. PLA/PGA composite anchors used in rotator cuff repairs have been shown to have comparable strength to metal anchors. Bio-absorbable suture anchors are used less frequently for soft tissue repairs in hand surgery. Their most notable use is in the repair of ulnar collateral ligaments of the thumb and repair of the scapholunate ligament. Vihtonen et al. reported excellent results in the treatment of 70 patients with total rupture of the ulnar collateral ligament of the first metacarpophalangeal joint of the thumb (skier’s / gamekeeper’s thumb) with SR-PLLA minitacks. Sixty-nine out of 70 patients healed without complication [44].

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Abigail R. Hamilton, Chaitanya S. Mudgal and Jesse B. Jupiter

COMPLICATIONS

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Immune Response As bio-absorbable implants undergo degradation via hydrolysis, patients treated with PGA implants were noted to have foreign body reactions at a rate close to 5% [9]. Of 2528 patients treated with bio-absorbable pins, rods, bolts, and screws made of PGA or PLA from 1985-1995 at the University Hospital, Helsinki, 108 developed a clinically significant local inflammatory, sterile tissue reaction. The incidences were 5.3% (107 of 2037) with PGA implants and 0.2% (one of 491) with a PLA implant. This study excluded painless transient minor swelling which is also a frequently reported phenomenon [26, 33, 40]. The incidence varied among anatomic sites of implantation, from 1.8% in fractures of the radial head to 25% of scaphoid nonunions. The overall volume of the implant did not affect the rate of incidence of soft tissue reaction; however the geometry of the implant did. Screws and serrated bolts had a statistically significant higher incidence of adverse tissue responses than pins and rods. In theory the larger surface area in contact with the host tissue increased the cellular reactivity and allowed for areas along the screw threads where an incubation process for an adverse tissue reaction could occur. Radiographs obtained at the time of presentation showed evidence of osteolysis along the implant tracks in 62 patients (57.4%). The implants manufactured between 1985 and 1988 included an additive green aromatic quinone dye within the polymer as a stain which increased the risk of a reaction (p100 kDa) has a Tm that ranges between 170 to 180 ºC and a Tg that ranges between 58-65 ºC.[24, 25] PLLA exhibits high tensile (11.482.7 MPa) and flexural (45-145 MPa) strength suggesting that it could be suitable for loadbearing applications.[26, 29-32] PLLA implants have been shown to retain their mechanical properties for at least six months in vivo, however in some cases complete resorption in the joint site can take up to five years.[30, 33, 34] Polycaprolactone (PCL): PCL is a semi-crystalline hydrophobic aliphatic polyester studied for long term drug delivery systems and implants. The melting point and glass transition temperatures of PCL are unusually low, approximately 58 to 63 ºC and -60 to 65 ºC, respectively.[24, 35] The tensile strength and elongation of PCL has been shown to vary between 21-35 MPa and 300-500%, respectively.[24] PCL’s slow degradation rate (> 2 years) makes it an excellent candidate for long-term implantable systems, when compared to PGA or PLLA implants.[36] Polydioxanone (PDS): PDS is a crystalline polymer (55% crystallinity) with a melting point of approximately 115 ºC and a glass transition temperature of -10 ºC; thus, the polymer is usually processed at the lowest possible temperature to prevent depolymerization back to monomer.[37] In vitro studies have shown that tensile strengths of PDS sutures can range from 296-358 MPa, with a percent elongation of 43.7-58.1%.[38] An experimental study of PDS implants in rabbits showed that the degradation time in vivo ranged from 16 to 25 wks.[39, 40] Poly(trimethylene carbonate) (PTMC): PTMC is a soft, rubbery polymer that, until recently, had not been considered for biomedical applications owing to its tackiness and weak mechanical properties. However, recent studies have shown that amorphous PTMC of high MW (~500 kDa and molecular number > 200,000) exhibits rubber-like properties with a Young’s modulus of 6 MPa and a tensile strength of 12-16 MPa.[41, 42] The Tm and Tg of PTMC can vary from 30 to 50 ºC and -17 to -19 ºC, respectively depending on the fabrication technique.[41] Degradation occurs mostly enzymatically and to a lesser extent by hydrolysis since autocatalysis, a phenomenon observed in polyesters like PLLA, is absent in PTMC.[43] In vivo studies in rabbit tibia and femur have shown that high MW PTMC rods (~450 kDa) degrade rapidly losing about 60% of their mass by the 8th week.[43] Table 1. Some FDA approved polymers and their mechanical contributions in a copolymer system Polymer PLLA PDLA PGA TMC

Major properties Hydrophobic Slow degradation Disrupts crystallinity Hydrophilic Quick degradation Subzero glass transition temperature Rubbery at room temperature

Contribution to implants Provides strength Provides flexibility Lowers stiffness Improves degradation rate Enhances malleability and toughness

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In addition to the above mentioned polymers, numerous other materials have been developed and used experimentally in recent years, including poly(ortho-esters),[44] poly(3hydroxybutyrate),[45] and poly(3-hydroxybutyrate-co-3-hydroxyvalerate).[46, 47]

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CHALLENGES IN THE DEVELOPMENT OF DEGRADABLE POLYMERS For a polymer to function as an efficient skeletal implant, it must exhibit several properties, biocompatibility notwithstanding. A polymeric implant should be able to withstand the harsh loading environments in vivo early in the healing process. At the same time, the implant must provide a suitable anchorage site to cells toward integration with the surrounding native tissue. The implant should also be expected to gradually degrade over time while transferring the load to healing tissues in the surrounding area. These properties are largely influenced by the chemical stability of the polymer backbone and the polymer processing technique which play a large role in determining the polymers’ hydrophilicity, crystallinity, melt and glass-transition temperatures, and MW.[48-50] In addition, the presence of residual monomers or additives can also significantly affect the properties of the polymer.[49] A polymer scientist working with biodegradable materials must evaluate each of these variables and determine the appropriate properties for a specific load-bearing skeletal implant. The stability of the polymer backbone is an important factor to consider when choosing a particular polymeric biomaterial for a load-bearing application. Common degradable polymers with ester, anhydride, orthoester and amide functional groups have hydrolytically unstable linkages in their backbone, and exhibit time-dependent properties resulting in viscoelastic behavior. Polymer degradation is accelerated by greater hydrophilicity in the backbone or end groups and greater reactivity among hydrolytic groups in the backbone.[51, 52] Thus, the biomaterial of choice must maintain adequate mechanical properties over time in an aqueous environment; i.e., not degrade too rapidly or slowly, while healing occurs. The mechanical properties of a polymer implant may be manipulated by the use of different processing techniques. Some of the current techniques to fabricate polymers include extrusion, injection molding, compression molding, gel casting, solution casting, solvent casting particulate leaching, electrospinning and self-reinforcement.[31, 53, 54] The choice of technique depends largely on the desired properties and morphology of the final product. For instance, the MW of a polymer might be controlled by varying the temperature at which the extrusion process takes place.[55] Nano-fibrous polymers can be created using electrospinning to enhance cell migration and adhesion on a scaffold.[56] The strength of a polymer can be increased by self-reinforcement, where polymer fibers made from the same material as the desired matrix are sintered at high pressures and temperatures.[54, 57, 58] For instance, the flexural strength of PGA can be increased from 100 MPa to 400 MPa upon selfreinforcement.[59] Thus, determining a polymer with the appropriate chemical and mechanical properties is crucial to the survival and success of an implant in vivo.

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CO-POLYMERS In addition to the techniques mentioned above to manipulate the mechanical properties of individual polymers, an alternate approach is to chemically synthesize two or more different monomers to form a co-polymer. By varying the individual amounts of each polymer, the mechanical and degradation properties of the copolymer can be controlled, thereby allowing for greater flexibility in choosing an appropriate implant material for a particular application (see Table 1). A general description of some of the salient copolymers is given below. Poly-DL-lactic acid (PDLLA): PDLLA is an amorphous copolymer with a random distribution of the two enantiomers of PLA (PDLA and PLLA). A combination of PLLA and PDLA serves to disrupt the crystallinity of PLLA and accelerate the degradation process. Thus, PDLLA has a lower tensile strength (27.6-41.4 MPa), higher elongation (3-10%), and quicker degradation time (12-16 months) than PLLA, making it a more attractive option in a load-bearing system requiring shorter fixation times. [24, 60, 61] Polylactic acid–polyglycolic acid (PLLA-PGA): The combination of two distinct polymers, PLLA and PGA offers to extend the range of mechanical properties not provided individually by either homopolymer. Experiments have shown that a non-linear relationship exists between the copolymer composition and the mechanical and degradation properties of the individual homopolymers. For instance, a copolymer implant of 50% PGA and 50% PLLA degrades faster than either homopolymer. Copolymers of PLLA with 25 to 70% PGA are amorphous due to the disruption of the regularity of the polymer chain by the other monomer.[62, 63] A copolymer of 90% PGA and 10% PLLA, developed as an absorbable suture material, degrades within three to four months but has a slightly longer strengthretention time.[64] Poly-L-lactic acid-hydroxyapatite (PLLA-HA): HA, an inorganic component of bone, has been shown to facilitate and promote bone formation.[65] HA is highly brittle and exhibits low fracture toughness upon mechanical loading. To overcome this issue, researchers have combined HA and degradable polymers (e.g., PLLA) using techniques such as electrospinning to produce more robust scaffolds that can aid in bone regeneration.[65] The crystallinity of PLLA-HA can be increased by increasing the HA content in the co-polymer. Accordingly, the mechanical properties of the PLLA/HA composites vary with HA content with increased levels of HA resulting in increased bending modulus from 2 to 7.4 GPa and strength from approximately 3 to 114 MPa.[66] Poly-L-lactide-β–tricalcium phosphate (PLLA-TCP): β-TCP is a ceramic with osteoconductive properties that degrades via hydrolysis into phosphate and calcium ions.[67] Like HA, TCP is brittle and exhibits low fracture toughness. Thus, the addition of β-TCP to PLLA creates a mechanically superior osteoconductive and biocompatible material that facilitates bone healing and regeneration. The Young’s modulus and yield strength of PLLATCP scaffolds have been shown to be about 22 MPa and 1.5 MPa, respectively.[68]

FUTURE OF BIODEGRADABLE POLYMER IMPLANTS The market for biodegradable orthopedic implants in the US and Europe exceeded over 250 million in revenues in 2006 with growth rates expected to increase further.[69 2006] As

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the popularity of biodegradable implants continues to grow, researchers are looking to develop novel low cost smart polymers that can complement the polymers already in use. An excellent example of this is the synthesis of energy-storing aliphatic polyesters, such as polyhydroxyalkanoates (PHA), by microorganisms. Two PHA materials, polyhydroxybutyrate (PHB) and polyhydroxyvalerate (PHV) have been researched extensively in the field of biodegradable plastics.[70-72] More recently, PHB has been studied for use in drug delivery systems, artificial skin grafts and as bioabsorbable sutures.[73] The sutures produced by recombinant DNA technology are FDA approved.[74] The PHB homopolymer is highly crystalline (~80%) and brittle, and its major degradation byproduct, hydroxybutyric acid, is found naturally in human blood. The Tm of PHB is around 177 ºC, with a recorded tensile and flexural strength of 40 MPa and 3.5 GPa, respectively.[70] The brittleness of the polymer can be overcome by copolymerizing it with PHV (80:20 PHBPHV) to reduce the crystallinity to 35% and increase failure strain from 8 to 50%.[70, 75-77] The advent of video assisted minimally invasive surgery has resulted in less pain for patients, reduced scarring and tissue injury, shortened hospital time and increased accuracy of the procedure. The major disadvantage with the technique, however, is that the smaller work area for the surgeon may increase the challenge in implanting a larger implant or knotting a suture. Groundbreaking research in material science has allowed for the development of degradable elastic shape memory polymers that can take a certain shape upon thermal or light induction.[78, 79] These polymers may be introduced into the body in a compressed form prior to application of the appropriate stimulus to regain the desired shape. Several different biodegradable polymers are currently investigated for their potential as shape memory polymers including PLLA-co-PCL (PCLA), PDLLA/hydroxyapatite composites, 82:18 PLAPGA, tert-butyl acrylate (tBA)/poly ethylene glycol dimethacrylate (PEGDMA), and copolyester-urethane networks.[80-83] Advances in computational modeling techniques utilizing high throughput screening may now eliminate the need for detailed characterization of individual polymers. The traditional mode of materials design begins with the synthesis of a new material, followed by its characterization, and eventual identification of a suitable application. The goal of the computational approach is to accurately predict the behavior of polymers under different conditions and create virtual biomaterial libraries containing thousands of individual compositions.[84] Using these libraries, prediction of material characteristics such as elastic modulus, glass transition temperature, hydrophobicity and degradation time will be facilitated.[84, 85] Several design factors can be controlled to ensure that the implants will survive in the host and provide appropriate fixation. These include shape, type of implantabutment mating, presence of threads, thread design, surface topography and chemical composition of the material.[86, 87] In addition, customized implants can now be created directly from computer data by using solid freeform fabrication [88-91] and combinatorial polymer scaffold libraries.[92] Solid freeform fabrication is a technique that allows for the three dimensional printing of custom designed objects created using a CAD drawing. The uniqueness of this technique stems from the ability to add and bond materials in layers. This provides the user with precision control and the option to create an object with multiple materials. Combinatorial polymer scaffold libraries provide the user with a tool to determine the appropriate cell-scaffold variations that are most likely to increase tissue formation on the construct for tissue engineering applications. With a plethora of biomaterials currently being researched for tissue engineering of load bearing tissues, such as articular cartilage and the

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knee meniscus, a predictive tool such as this could be used to screen cell-polymer combinations to streamline the tissue engineering process.

CONCLUSION We can speculate that advancements in polymer science will continue to fuel the development of novel biocompatible polymeric implants, with or without the incorporation of bioactive agents, toward fixation as well as healing and repair of the injured tissue. Using techniques such as copolymerization and self-reinforcement, implant tensile, flexural and shear strengths, crystallinity, and in vivo degradation profiles can be controlled. In addition, computational modeling techniques can now be used to predict polymer properties prior to implantation. Finally, novel biodegradable elastic smart polymers, that can change shape by triggering changes in the immediate environment, may potentially be used in minimally invasive surgeries as fixation devices.

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CONTRIBUTORS Tuija Annala Institute of Biomaterials Tampere University of Technology P.O.Box 589, FI-33101 Tampere, Finland [email protected]

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Kyriacos A. Athanasiou Rice University Department of Bioengineering: MS-142 P. O. Box 1892 Houston, TX 77251-1892, USA [email protected]

Ole M. Böstman Dept. of Orthopaedics and Trauma Surgery Helsinki University Hospital, Finland Tiirasaarent. 24, FIN-00200 Helsinki, Finland [email protected]

Pieter Buma Orthopedic Research Laboratory, Radboud University Nijmegen Medical Centre, P.O. Box 9101, 6500 HB, Nijmegen, The Netherlands. [email protected]

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Contributors Sudhir S. Chakravarthi AstraZeneca, Wilmington, Delaware, USA University of Nebraska Medical Center 986025 Nebraska Medical Center Omaha, NE 68198-6025, USA [email protected]@unmc.edu

Ying Deng Biomedical Engineering Program and Department of Basic Biomedical Sciences University of South Dakota, Sioux Falls, SD 57107, USA [email protected]

Jon Olav Drogset Norwegian University of Science and Technology, Faculty of Medicine; Department of Orthopedic Surgery and Rheumatology, Trondheim University Hospital N-7006 Trondheim, Norway

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Marc Dumonski Rush University Medical Center Department of Orthopaedic Surgery, Chicago, IL [email protected]

Uwe Eckelt Klinik und Poliklinik für Mund-, Kiefer- und Gesichtschirurgie Universitätsklinikum Carl Gustav Carus der TU Dresden Fetscherstr. 74, 01307 Dresden, Germany [email protected]

Lars Ejerhed Department of Orthopaedics Northern Älvsborg County Hospital, 461 85 Trollhättan, Sweden [email protected]

M. van der Elst Department of Surgery and Traumatology, Reinier de Graaf Gasthuis PO Box 5011, 2600GA, Delft, The Netherlands [email protected]

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Contributors Tom A.P. Engels Eindhoven University of Technology Dept. Mechanical Engineering, Section Polymer Technology PO Box 513, WH 4.105 5600 MB Eindhoven, The Netherlands [email protected]

Arthur de Gast Vrije Universiteit medical center Dept. Orthopaedic Surgery PO Box 7057, De Boelelaan 1117 1007 MB Amsterdam, The Netherlands [email protected]

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Leon E. Govaert Eindhoven University of Technology Dept. Mechanical Engineering, Section Polymer Technology PO Box 513, WH 4.143 5600 MB Eindhoven, The Netherlands [email protected]

Najmuddin Gunja Department of Bioengineering Musculoskeletal Bioengineering Laboratory Rice University, Houston, TX 77054, USA [email protected]

Abigail R. Hamilton, MD Massachusetts General Hospital Harvard Medical School Yawkey Center, Suite 2100 55 Fruit Street, Boston, MA 02114

Gerjon Hannink Orthopedic Research Laboratory Radboud University Nijmegen Medical Centre, P.O. Box 9101, 6500 HB, Nijmegen, The Netherlands [email protected]

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366

Tim U. Jiya Vrije Universiteit medical center Dept. Orthopaedic Surgery PO Box 7057, De Boelelaan 1117 1007 MB Amsterdam, The Netherlands [email protected]

Jesse B. Jupiter, MD Massachusetts General Hospital Harvard Medical School Yawkey Center, Suite 2100 55 Fruit Street, Boston, MA 02114 [email protected]

Jüri Kartus Department of Orthopaedics Northern Älvsborg County Hospital, 461 85 Trollhättan, Sweden [email protected]

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Minna Kellomäki Institute of Biomaterials Tampere University of Technology P.O.Box 589, FI-33101 Tampere, Finland [email protected]

Theo G. van Kooten Department of BioMedical Engineering, University Medical Center Groningen (UMCG) A. Deusinglaan 1, 9713 AV Groningen, The Netherlands [email protected]

Roel Kuijer Department of Biomedical Engineering University Medical Center Groningen P.O. Box 196, 9700 AD Groningen, The Netherlands [email protected]

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Contributors Cato T. Laurencin Departments of Biomedical Engineering and Chemical Engineering University of Virginia P.O. Box 400741, Charlottesville VA 22904-4741, USA [email protected]

Lennart Magnusson Västerås Ortopedpraktik AB Kopparbergsvägen 14, 722 13, Västerås, Sweden Christiaan J. van Manen Department of Surgery and Traumatology, Reinier de Graaf Gasthuis PO Box 5011, 2600GA, Delft, The Netherlands [email protected]

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Chaitanya S. Mudgal, MD, MS(Orth.), M.Ch(Orth.) Orthopaedic Hand Service Massachusetts General Hospital Harvard Medical School Yawkey Center, Suite 2100 55 Fruit Street, Boston, MA 02114. [email protected]

Eric de Mulder Orthopedic Research Laboratory Radboud University Nijmegen Medical Centre, P.O. Box 9101, 6500 HB, Nijmegen, The Netherlands.

Dennis H. Robinson University of Nebraska Medical Center 986025 Nebraska Medical Center Omaha, NE 68198-6025, USA [email protected]

Kurt Ruffieux DS Degradable Solutions AG Wagistrasse 23 8952 Schlieren, Switzerland [email protected]

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368

Contributors Carmen Scholz Department of Chemistry University of Alabama in Huntsville 301 Sparkman Dr., Huntsville, AL 35899, USA [email protected]

Kern Singh Rush University Medical Center Department of Orthopaedic Surgery 1725 W Harrison St., Suite 1063, Chicago, IL 60612, USA [email protected]

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Serge H.M. Söntjens Eindhoven University of Technology Dept. Mechanical Engineering, Section Polymer Technology PO Box 513, Helix STW 4.30 5600 MB Eindhoven, The Netherlands [email protected]

Theo H. Smit Vrije Universiteit medical center Dept. Physics and Medical Technology PO Box 7057, Van der Boechorststraat 7, A-115 1007 MB Amsterdam, The Netherlands [email protected]

K. Elizabeth Tanner Departments of Civil and of Mechanical Engineering James Watt (South) Building University of Glasgow, Glasgow, G12 8QQ, UK [email protected]

Erica D. Taylor Department of Orthopaedic Surgery University of Virginia Health System, Charlottesville, VA, USA

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Contributors G. Lawrence Thatcher TESco Associates, Inc 500 Business Park Drive Tyngsborough, MA 01879, USA [email protected]

Tony G. van Tienen Orthopedic Research Laboratory Radboud University Nijmegen Medical Centre, P.O. Box 9101, 6500 HB, Nijmegen, The Netherlands

Hideto Tsuji Department of Ecological Engineering Faculty of Engineering Toyohashi University of Technology Tempaku-cho, Toyohashi, Aichi 441-8580, Japan [email protected]

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Alexander R. Vaccaro Thomas Jefferson University and the Rothman Institute Department of Orthopaedic Surgery, 925 Chestnut Street, Philadelphia, PA 19107, USA [email protected]

Hay A.H. Winters Vrije Universiteit medical center Dept. Plastic and Reconstructive Surgery PO Box 7057, De Boelelaan 1117 1007 MB Amsterdam, The Netherlands [email protected]

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INDEX

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A absorption, 78, 80, 82, 101, 103, 106, 158, 166, 167, 181, 183, 198, 224, 255, 282, 288, 323, 357 accidents, 4 accuracy, 102, 333, 356 acetate, 5, 6 achilles tendon, 269 acidic, 109, 163, 165, 167, 168, 169, 170, 255, 345 acidity, 63 ACL, viii, 94, 95, 267, 268, 269, 270, 271, 272, 273, 274, 275, 276, 277, 278, 279, 280, 281, 282, 285, 288, 289, 290, 292, 297, 305, 358 ACL injury, 268 ACL reconstruction, 268, 280, 281, 282, 290, 297, 358 acrylate, 151, 153, 356 acrylic acid, 154 ACS, 19 activation, 147, 262, 283, 284, 285, 291 activators, 279 acute, 156, 175, 290 Adams, 150 adaptation, 36, 109, 230, 362 additives, 6, 49, 63, 101, 112, 148, 200, 354 adhesion, 14, 41, 43, 52, 65, 84, 85, 91, 131, 139, 140, 142, 143, 144, 145, 146, 147, 148, 149, 150, 151, 153, 154, 155, 173, 262, 273, 274, 322, 324, 334, 335, 354, 361 adhesions, 324, 325 adhesive interaction, 156 adipate, 42 adipose tissue, 148 adjustment, 111 adsorption, 141, 142, 149, 150, 151, 171, 362 adult, 158, 178, 214, 222, 223, 298, 325, 333, 334

adulteration, 123 adults, 207, 222, 224, 225, 257 adverse event, 124, 326, 328 AFM, 139, 144 age, 4, 28, 131, 200, 211, 212, 213, 223, 314, 319, 322, 325, 330, 333, 335, 349 ageing, 144 agent, 85, 101 agents, 100, 106, 128, 133, 179, 352, 357 aggregation, 137 aging, 28, 38, 69, 105, 153 aid, 322, 325, 339, 355 aiding, 324 air, 110, 117, 126, 129, 131, 143, 146, 176, 287 Alabama, 3, 368 alanine, 171 albumin, 176 alcohol, 6, 7, 44, 70, 123, 157, 165, 300 alcoholics, 192 alcoholism, 187 Alginate, 160 aliphatic polymers, 168 alkali, 173 alkaline, 52, 53, 58, 60, 61, 65, 67, 68, 69, 70, 84, 85, 147, 165, 167, 168, 171, 174 alkaline hydrolysis, 67, 167 alkaline media, 165, 168 alkaline phosphatase, 84, 85, 147 allergic reaction, 260 allergy, 322 allogeneic, 304 allograft, 217, 218, 226, 295, 300, 306, 309, 310, 315, 316 allografts, 138, 268, 271, 278, 300, 309, 310 alpha, 90, 161, 175, 184, 186, 187, 285, 333, 334, 361 alpha-1-antitrypsin, 285

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Index

alternative, 123, 125, 139, 149, 203, 242, 258, 277, 283, 299, 300, 301, 304, 309, 322, 325 alters, 138, 173, 305 aluminum, 80, 127 ambient air, 143 amendments, 25 amide, 11, 15, 55, 143, 154, 354 amine, 16, 143, 144, 154 amino, 3, 5, 11, 148, 172 amino acid, 3, 5, 11, 148, 172 ammonia, 130 amorphous, 9, 10, 12, 13, 18, 23, 31, 32, 35, 37, 38, 47, 49, 53, 54, 57, 58, 59, 60, 62, 69, 80, 82, 103, 104, 128, 163, 164, 165, 167, 170, 172, 229, 230, 238, 255, 288, 344, 345, 353, 355, 362 amorphous polymers, 9, 13, 18, 37, 38, 167 amplitude, 38 Amsterdam, x, xi, 21, 38, 65, 68, 211, 321, 349, 365, 366, 368, 369 anastomosis, 43 anatomy, 22, 270 angiogenic, 308 angulation, 186, 200, 203 animal models, 140, 306, 321 animal studies, 183, 273, 299, 300, 306, 321, 331, 332 animals, 177, 179, 183, 186, 299 anisotropy, 37 ankle joint, 213, 214, 223 annealing, 16, 28, 30, 32, 39, 51, 66, 101, 103, 111, 120, 130, 137, 189 anterior cruciate, 38, 186, 192, 193, 250, 267, 268, 277, 278, 279, 280, 290, 291, 292, 293, 297, 308, 312, 358 antibiotic, 20, 79 antibiotics, 14, 20, 79, 198 antigen, 262 apatite, 75, 76, 83, 84, 85, 90, 128, 147, 158, 159, 212 apatite layer, 75, 76 APL, 259 application, ix, 3, 11, 13, 14, 15, 16, 17, 20, 21, 25, 30, 35, 39, 72, 91, 142, 147, 154, 160, 172, 175, 202, 224, 229, 230, 233, 234, 235, 238, 270, 292, 299, 312, 313, 314, 321, 322, 324, 325, 328, 331, 332, 333, 341, 342, 346, 349, 354, 355, 356 aqueous solution, 81, 165 aqueous solutions, 81, 165 arginine, 140, 159 argon, 125, 130, 143 artery, 212 arthritis, 259, 269, 278, 308 arthrodesis, 178, 211, 225, 226, 257, 258, 264

arthroplasty, 259, 273 arthroscopy, 250, 307, 310, 360 articular cartilage, 20, 182, 184, 185, 191, 295, 296, 297, 298, 299, 300, 306, 307, 308, 309, 310, 312, 352, 356 aseptic, 191, 245, 257, 285, 291 aspartate, 140 aspect ratio, 99 aspirate, 87 aspiration, 257 assessment, 38, 91, 99, 107, 117, 157, 254, 260, 273, 310, 313, 318, 322, 326, 328, 332, 334 assessment tools, 107 ASTM, 34, 35, 100, 120 asymptomatic, 315, 317 athletes, 268 atomic force, 144, 176 Atomic Force Microscopy, 139, 144, 153, 172, 176 atoms, 10, 11, 14, 71, 143 atrophy, 231, 341 attachment, 82, 85, 141, 145, 147, 150, 153, 154, 155, 156, 159, 267, 271, 272, 273, 274, 276, 277, 279 attacks, 6 Australia, 22 authority, 129 autocatalysis, 163, 165, 166, 167, 168, 170, 173, 181, 353 autografts, 227, 268, 271, 290, 291 autologous bone, 314, 324, 325, 329, 343, 347 availability, 102, 104, 141, 187, 235, 254, 300 averaging, 105 avoidance, 177, 237 awareness, 4, 306 axonal, 325 axons, 325

B Bacillus, 171 back, 22, 25, 101, 109, 297, 300, 335, 342, 353 back pain, 22, 328, 335 bacteria, 9, 172, 347 bacterial, 9, 10, 14, 170, 255, 262, 268, 269 bacterial contamination, 170 bacterial infection, 262, 268, 269 bananas, 5, 6 barium, 349 barium sulphate, 349 barrier, 322, 324, 332, 334, 335 barriers, 186, 324, 325 basic fibroblast growth factor, 159 basicity, 63

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Index baths, 123 beam radiation, 98, 128, 129, 137, 334 behavior, 3, 13, 24, 25, 27, 28, 29, 30, 31, 37, 39, 54, 60, 63, 91, 92, 154, 157, 160, 176, 178, 261, 270, 274, 275, 276, 296, 340, 354, 356, 359, 360, 362 bending, 32, 77, 80, 86, 98, 112, 115, 125, 128, 186, 256, 260, 261, 344, 355 benefits, 188, 197, 260, 261, 263, 281, 282, 322, 323, 331 biaxial, 70, 174 biaxial orientation, 70, 174 bicarbonate, 170 binding, 111, 128, 141, 142, 150, 151, 160, 171, 256 bioactive materials, 75 biochemical action, 4 biochemistry, 307 biocompatibility, 3, 10, 19, 20, 21, 37, 80, 84, 90, 91, 134, 150, 151, 153, 154, 155, 160, 164, 170, 173, 175, 183, 184, 187, 190, 242, 255, 273, 274, 276, 281, 283, 322, 323, 333, 354, 357, 359, 360 biocompatible, 4, 10, 11, 15, 85, 86, 163, 164, 165, 173, 249, 271, 276, 298, 299, 339, 344, 348, 349, 351, 352, 355, 357 biocompatible materials, 249 biodegradability, 10, 16, 17, 42, 44, 68, 135 biodegradable materials, 11, 54, 124, 133, 143, 145, 230, 233, 236, 237, 238, 271, 275, 281, 354, 361 biodegradation, 3, 17, 58, 67, 152, 157, 164, 165, 166, 176, 179, 189 Bioglass, 72, 77, 84, 85, 91, 92, 161 bioinert, 88 biological activity, 150, 353 biological behavior, 178 biological interactions, 142 biological macromolecules, 141, 144 biological responses, 360 biological systems, 141, 143 biomass, 164, 361 biomaterial, 10, 57, 72, 135, 136, 149, 150, 151, 152, 153, 155, 191, 270, 346, 354, 356 biomaterials, 19, 20, 42, 66, 75, 79, 88, 90, 91, 128, 139, 144, 149, 152, 154, 160, 173, 270, 279, 283, 339, 340, 343, 349, 352, 356, 361, 362 biomechanics, 276, 277, 300, 311 biomedical applications, 4, 13, 14, 15, 41, 45, 51, 65, 72, 153, 160, 353, 359, 362 biomimetic, 91, 147, 158, 159 biopolymer, 164 biopolymers, 10, 19, 65, 66, 153 160, 346, 352 biopsies, 273, 299 biopsy, 187, 309 biosafety, 134

373

biosynthesis, 145, 156 biotechnology, 277 bladder, 160 bleeding, 112 blend films, 64 blends, 39, 49, 50, 54, 63, 64, 67, 70, 89, 102, 104, 131, 138, 147, 164, 172, 174 blindness, 231 blocks, 109, 224, 288, 290 blood, 14, 16, 73, 75, 141, 145, 160, 183, 214, 262, 272, 283, 299, 325, 343, 345, 356 blood plasma, 75 blood supply, 272 blood vessels, 16, 325 blowing agent, 49 body fluid, 75, 76, 144, 145, 170 body temperature, 34, 76, 245 bonding, 85, 90, 92, 304 bonds, 17 bone graft, 65, 67, 160, 186, 223, 225, 226, 227, 256, 257, 291, 313, 314, 322, 323, 324, 325, 341, 357 bone grafts, 223, 227, 291, 341, 357 bone growth, 4, 96, 125, 148, 171 bone marrow, 87, 148, 154, 159, 160, 276, 280 bone powder, 185 bone remodeling, 222, 224, 225 bone resorption, 170, 261, 289 bone scan, 222 Boston, 253, 278, 365, 366, 367 bounds, 78 bovine, 84, 154, 186, 305, 311 braids, 271, 275 brain, 231 branched polymers, 10, 44 branching, 110 breakdown, 85, 86, 96, 98, 101, 103, 146, 164 breeding, 322 broad spectrum, 140 buffer, 88, 163, 167, 168, 173, 174 bulk materials, 72 burn, 116 butane, 311 butyric, 14

C Ca2+, 75 CAD, 356 cadaver, 19, 186, 282, 358 calcium, 72, 78, 80, 81, 82, 83, 84, 85, 89, 90, 91, 125, 128, 134, 158, 159, 160, 161, 170, 184, 201, 212, 256, 314, 341, 343, 346, 355, 358, 362 calcium carbonate, 184

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374

Index

calvaria, 84, 158 calvarium, 233, 235 candidates, 14, 21, 141 capillary, 146, 272 caprine, 312 caprolactone, 3, 12, 14, 15, 20, 42, 43, 45, 47, 55, 58, 59, 67, 69, 70, 80, 89, 92, 102, 156, 157, 172, 299, 311, 312, 351, 352, 358, 359, 362 capsule, 87, 231, 259 carbohydrates, 11 carbon, 5, 6, 9, 10, 11, 13, 14, 71, 86, 130, 143, 155, 160, 164, 181, 255, 269, 278, 298, 300, 308, 309, 321, 322, 335 carbon atoms, 10, 11, 14 carbon dioxide, 130, 160, 255 carbonyl groups, 153 carboxyl, 16, 52, 60, 64, 143, 165, 166 carboxyl groups, 52, 60, 64, 143, 165, 166 carboxylic, 7, 8, 9, 60, 127, 144, 360 carboxylic acids, 7 carcinoma, 213 cardiopulmonary, 226 cargo, 4 carrier, 298 cartilage, 20, 140, 141, 142, 144, 147, 151, 179, 182, 184, 185, 191, 202, 242, 295, 296, 297, 298, 299, 300, 301, 302, 304, 305, 306, 307, 308, 309, 310, 311, 312, 352, 356, 358, 359 cartilaginous, 305 cast, 47, 48, 49, 64, 66, 67, 70, 117, 131, 197, 200, 203 casting, 67, 76, 77, 80, 178, 354 catabolic, 306 catalysis, 16 catalyst, 13, 163 catalytic effect, 54, 60, 65 catheters, 16 cavities, 112, 341 cell adhesion, 14, 41, 52, 131, 140, 142, 143, 144, 145, 146, 147, 148, 150, 151, 153, 154, 173 cell culture, 75, 80, 82, 84, 85, 87, 123, 140, 154, 156, 160, 300 cell cycle, 142 cell differentiation, 151 cell division, 274 cell grafts, 325, 334 cell growth, 140, 141, 145, 147, 150, 151, 152, 153, 155, 273 cell line, 85, 139, 145, 153, 156 cell surface, 142 cellulose, 4, 128, 143, 153, 191 cement, 158, 159, 160, 343, 358 Centers for Disease Control (CDC), 268, 278

ceramic, 71, 72, 73, 84, 87, 88, 91, 188, 346, 355 cerebrospinal fluid, 325 chain mobility, 25, 26, 54, 59, 145 chain scission, 101, 103, 106, 126, 128, 131, 146, 171 chain transfer, 100 channels, 129, 295, 298, 334 charge density, 142 chemical approach, 11 chemical interaction, 147 chemical properties, 128, 140 chemical stability, 354 chemotaxis, 159 children, 178, 191, 192, 203, 211, 214, 222, 223, 224, 225, 227, 341 chiral, 11, 14, 164 chitin, 5 chitosan, 72, 73, 80, 90, 128, 131, 138, 147, 158, 173 chloroform, 46, 82, 164 CHO cells, 143 chondrocyte, 157, 179, 358 chondrocytes, 20, 146, 155, 304, 305 chromatography, 58 chromium, 71 chronic inflammatory cells, 259 ciprofloxacin, 19 classes, 4, 72, 171 classical, 310 classification, 202, 203, 244 claudication, 328 cleaning, 112, 117 cleavage, 6, 16, 117, 163, 167, 171, 172 clinical assessment, 241, 259 clinical symptoms, 214 clinical trial, 36, 75, 177, 179, 183, 191, 197, 203, 251, 260, 291, 299, 306, 316 clinical trials, 75, 177, 183, 197, 299, 306, 316 clinically significant, 262, 285, 317 clinics, 130, 340 closure, 109, 214 CO2, 41, 49, 67, 85, 117, 165 coatings, 4, 16, 85, 89, 92, 145, 146, 147, 186 cobalt, 126 Cochrane, 207 cohort, 262 coil, 10, 18, 25 cold compress, 169 collaboration, 300 collagen, 91, 147, 148, 156, 157, 158, 159, 160, 161, 172, 173, 270, 271, 275, 276, 295, 296, 297, 299, 301, 302, 305, 307, 308, 309, 310, 312, 323, 324, 328, 360, 362 collateral, 261, 265, 274, 279

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Index community, 4, 188 compatibility, 89, 102, 279 complement, 283, 284, 285, 291, 292, 356 complement components, 283 complement system, 283, 291 complexity, 3, 16, 169, 173, 181, 270 compliance, 187 complications, 22, 187, 188, 198, 199, 201, 203, 211, 222, 223, 253, 254, 256, 257, 258, 259, 260, 263, 282, 285, 289, 313, 315, 321, 322, 326, 347, 352 components, 31, 111, 130, 178, 190, 283 composites, 14, 54, 71, 72, 73, 75, 76, 77, 78, 79, 80, 82, 84, 85, 86, 87, 88, 89, 90, 91, 92, 104, 123, 125, 126, 130, 137, 148, 155, 159, 172, 355, 356, 358, 359, 360, 361, 362 composition, 13, 65, 90, 100, 105, 108, 139, 142, 143, 146, 149, 151, 155, 164, 171, 241, 255, 305, 314, 331, 332, 355, 356, 359 compounds, 138, 143, 147, 178, 254, 352 compressibility, 109 compressive strength, 79, 82, 85, 325 computational modeling, 351, 356, 357 computed tomography, 254, 292 concentration, 32, 51, 52, 54, 59, 102, 143, 163, 168, 173, 174, 255, 284, 340 conceptualization, 93 concurrent engineering, 97 condensation, 5, 6, 7, 8, 13, 19 conditioning, 30, 38, 109, 110, 140, 141, 142, 147 configuration, 10, 11, 14, 18, 148, 232, 305 conflict, 112, 246 conflict of interest, 246 Connecticut, 267 connective tissue, 87, 181, 267, 273 consensus, 200 conservation, 309 consolidation, 197, 198, 200, 201, 203 constant load, 24, 37 constant rate, 24, 27, 165 construction, 24, 106, 116, 276 contamination, 106, 117, 124, 126, 170, 179, 362 continuity, 214, 222, 224, 357 control, 71, 73, 87, 101, 109, 110, 135, 137, 145, 149, 150, 158, 160, 175, 185, 190, 203, 224, 260, 269, 281, 324, 356, 360 control group, 224, 260, 324 controlled trials, 199, 329 conversion, 73, 97 cooling, 18, 99, 101, 103, 112, 117, 214 cooling process, 18 copolymer, 8, 14, 31, 59, 87, 101, 102, 103, 106, 125, 126, 128, 137, 143, 154, 155, 175, 185, 188, 189, 191, 212, 214, 256, 305, 355, 361

375

copolymerization, 39, 43, 67, 172, 352, 357 copolymers, 13, 14, 42, 43, 44, 46, 47, 48, 51, 58, 66, 69, 80, 88, 100, 125, 126, 127, 131, 134, 136, 149, 154, 155, 157, 164, 174, 177, 180, 187, 229, 230, 238, 255, 271, 305, 352, 355, 359 coral, 174 corona, 143 coronary arteries, 19 correlation, 65, 152, 249, 310 correlations, 284, 362 corrosion, 4, 109, 177, 178, 229, 351 cortex, 170 cost benefit analysis, 119 cost-effective, 11 costs, 72, 93, 97, 100, 112, 129, 198, 203, 237, 253, 300 cotton, 3 coupling, 106, 142 covalent, 25, 160 covalent bond, 25 covering, 341 CPS, 143 crack, 16, 32, 105, 111 cracking, 16 craniofacial, 171, 179, 229, 230, 231, 239, 358, 362 craniomaxillofacial, 254, 260 cranioplasty, 238 cranium, 229, 231, 232, 233, 238 CRC, 252 C-reactive protein, 247 creep, 24, 26, 27, 37, 38, 72, 98, 269, 275, 276, 344 creep tests, 27 critical value, 168 cross-linking, 16, 47, 131, 276 cross-sectional, 34 crystal growth, 125 crystal structure, 18, 102 crystallites, 18, 47, 54, 56, 57, 60, 68, 70, 86, 87, 344 crystallization, 10, 31, 39, 47, 54, 60, 69, 70, 99, 101, 102, 103, 111, 127, 135, 230, 288 crystallization kinetics, 99 crystals, 75, 99, 103, 170, 172 CT scan, 317, 327, 328, 333 cues, 142 culture, 67, 84, 87, 140, 145, 150, 153, 156, 175, 274, 304, 305, 311 culture conditions, 145 cutters, 110 cutting tools, 117 cycles, 100, 129 cyst, 290, 293 cysts, 241, 242, 243, 244, 245 cytokines, 142, 184

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Index

376 cytotoxic, 80, 300 cytotoxicity, 170

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D dacron, 145 debridement, 199, 257 decay, 5, 127, 133, 136, 198 decisions, 97 decomposition, 109, 167, 168, 186 decompression, 334 defects, 36, 108, 152, 157, 158, 178, 180, 185, 189, 191, 211, 223, 224, 301, 341, 343, 346, 357 defense, 140 defense mechanisms, 140 defibrillation, 275 deficiency, 277 deficit, 227 definition, 71, 124 deformation, 22, 23, 24, 25, 26, 27, 28, 29, 30, 31, 32, 33, 35, 39, 73, 76, 145, 174, 213, 221, 224, 270, 275, 323, 331, 343, 344, 345 deformities, 211 degenerative disease, 318, 319, 334, 335 degenerative joint disease, 257 degradable polymers, ix, xi, 3, 4, 15, 22, 23, 34, 68, 71, 72, 87, 143, 146, 147, 179, 187, 321, 351, 352, 354, 355, 359 degradation mechanism, 41, 43, 54, 55, 129, 146, 173, 181 degradation process, 85, 181, 255, 349, 355 degradation profiles, 357 degradation rate, 11, 14, 16, 17, 41, 44, 54, 58, 59, 60, 63, 64, 65, 71, 78, 80, 84, 88, 130, 151, 168, 170, 175, 181, 186, 199, 255, 273, 300, 353 degrading, 15, 18, 86, 108, 112, 144, 148, 171, 172, 177, 179, 183, 184, 185, 187, 273, 301, 325 degree of crystallinity, 10, 12, 18, 103, 117, 127, 352 dehydration, 13 dehydrogenase, 11 delivery, 4, 10, 14, 20, 96, 97, 100, 104, 107, 159, 164 demineralized, 133, 138, 317 demographics, 213 dendrimers, 173 densitometry, 357 density, 60, 61, 73, 85, 99, 112, 117, 126, 129, 147, 155, 157, 160, 172, 275, 291 dentist, 347 dentistry, 341, 343 dentures, 4 depolymerization, 4, 18, 101, 103, 106, 353

deposition, 75, 76, 83, 84, 141, 148, 270, 305, 359, 361 derivatives, 7, 9, 128, 153, 154, 156, 178, 181 designers, 101, 110, 116 desorption, 149 destruction, 130, 306, 308 detection, 102, 283 deviation, 214 dexamethasone, 148, 159 dialysis, 9 diamines, 15, 17 diaphysis, 214, 217 differential scanning, 103, 127, 129 Differential Scanning Calorimetry (DSC), 58, 101, 103, 105, 127, 129 differentiation, 140, 142, 147, 148, 149, 151, 158, 159, 160, 274, 299, 300, 301, 304, 305, 311 diffraction, 87 diffusion, 38, 59, 60, 64, 165, 166, 168, 170, 298 diffusivity, 170 dimer, 9, 43, 255 dimerization, 11 dimethacrylate, 356 direct measure, 107 directionality, 305 disability, 328, 330 discomfort, 187, 211, 257, 259 diseases, xi, 21 disinfection, 123 dislocation, 202, 230, 250, 251 dislocations, 247 dispersion, 79 displacement, 34, 146, 186, 226, 235, 261 dissociation, 7 distal radius, 186, 192, 199, 200, 256, 257, 264 distilled water, 82 distribution, 9, 54, 58, 60, 61, 76, 97, 99, 100, 102, 105, 109, 120, 131, 155, 165, 255, 355 division, 274 DNA, 18, 130 dogs, 299, 301, 308, 311 domain structure, 142 donor, 211, 214, 223, 226, 227, 235, 267, 268, 269, 300 doped, 91 drainage, 242, 245, 258 dressings, 3 drinking, 231 drug carriers, 135 drug delivery, 4, 9, 10, 13, 14, 15, 16, 19, 20, 42, 43, 44, 79, 89, 136, 164, 173, 175, 278, 306, 352, 353, 356 drug delivery systems, 13, 16, 42, 44, 173, 353, 356

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Index drug release, 43, 175, 340 drugs, 11, 44, 346 dry ice, 132 drying, 98, 100, 103, 175 ductility, 76, 79, 80, 88 duplication, 347 DuPont, 120 durability, 4, 5 duration, 96, 100, 107, 177, 225, 230 dynamic mechanical analysis, 76, 77, 85 Dysplasia, 213

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E ecological, 65 edema, 290 effluent, 51 elaboration, 273, 274 elastic deformation, 24 elasticity, 127, 185, 186, 313, 318 elasticity modulus, 127, 186 elastin, 172, 296 elbow, 178, 201, 205 electric field, 51, 52 electrodes, 116 electrolytes, 131 electron, 76, 77, 79, 123, 127, 128, 129, 135, 137, 190, 334 electron beam, 123, 128, 135, 137, 334 electron microscopy, 58, 76, 180, 183 electron paramagnetic resonance, 127 electrons, 128 electrophoresis, 128 electrospinning, 15, 354, 355, 360 electrostatic force, 142 electrostatic interactions, 142 ELISA, 292 elongation, 15, 47, 98, 103, 104, 255, 270, 353, 355 emboli, 43 embolization, 43 enantiomers, 355 encapsulation, 88, 99, 170 endothelial cell, 139, 143, 145, 147, 149, 150, 153, 154, 155, 156, 158, 160, 161, 298 endothelial cells, 143, 145, 149, 150, 154, 155, 156, 158, 161, 298 energy, 9, 103, 106, 108, 110, 128, 131, 347, 351, 356 engagement, 142 enlargement, 290 entanglement, 39, 54 Enthalpy, 46, 48 enthusiasm, 271

377

entrapment, 168 entropy, 277 environment, 4, 5, 22, 54, 75, 123, 140, 141, 142, 146, 149, 165, 166, 167, 170, 225, 255, 306, 352, 354, 357 environmental awareness, 4 environmental conditions, 129, 131 environmental factors, 163, 165, 173, 186 environmental issues, 99 enzymatic, 10, 14, 60, 61, 67, 68, 69, 70, 135, 164, 165, 171, 172, 176, 180, 247, 255 enzyme-linked immunosorbent assay, 292 enzymes, 5, 9, 16, 54, 149, 171, 172, 199, 255, 265 epidermal growth factor, 274 epiphyses, 208, 214 epiphysis, 203, 219, 220, 221, 223 epithelial cell, 145, 156, 160 epithelial cells, 145, 156, 160 EPR, 127 equilibrium, 9, 25, 32, 103 equilibrium state, 25, 32 erosion, 18, 54, 55, 58, 68, 130, 168, 273, 360 ester, 5, 6, 7, 14, 15, 16, 17, 18, 44, 52, 55, 64, 136, 137, 143, 152, 154, 163, 165, 168, 171, 172, 354, 360 ester bonds, 7, 17, 18, 168 esterase, 157 esterification, 6, 273 esters, 3, 6, 7, 16, 354 estradiol, 135 etching, 154 ethylene, 7, 15, 17, 19, 42, 49, 66, 67, 70, 80, 90, 123, 130, 131, 135, 137, 138, 139, 149, 151, 154, 160, 174, 189, 335, 356 ethylene glycol, 19, 66, 131, 137, 139, 149, 151, 174, 356 ethylene oxide, 49, 67, 70, 80, 90, 123, 130, 135, 137, 138, 151, 154, 189, 335 etiology, 297 Europe, 328, 355 evolution, 26, 29 examinations, 233, 244, 247, 281, 283, 285 excision, 259 excitation, 24, 285 exclusion, 58 excretion, 158 exophthalmos, 231 experimental condition, 184 exposure, 75, 126, 130, 140, 142, 143, 233, 236, 237, 260, 303, 304 extensor, 200, 212, 256, 257 extensor digitorum, 212

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Index

378

extracellular matrix, 20, 141, 142, 145, 149, 270, 273, 276, 277, 301 extraction, 49, 50, 97, 98, 99, 100, 106, 109, 125, 153, 341, 347 extrusion, 13, 77, 78, 103, 106, 107, 109, 110, 111, 112, 178, 186, 256, 276, 296, 297, 324, 352, 354 eyelid, 235

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F fabric, 72 fabricate, 51, 67, 305, 321, 332, 354 fabrication, 91, 181, 270, 271, 314, 332, 352, 353, 356, 362 facial nerve, 237 family, 44, 273 fascia, 212 fatigue, 24, 38, 41, 269, 271, 276 fatty acids, 16 FDA approval, 325 feedback, 94 feeding, 98, 109 femoral bone, 282, 288 femur, 171, 180, 181, 182, 184, 190, 213, 216, 217, 218, 267, 268, 281, 282, 285, 286, 288, 289, 296, 297, 304, 353 fermentation, 11, 164 fetal, 85 fever, 124 fiber, 46, 47, 49, 51, 61, 66, 68, 70, 90, 91, 92, 101, 106, 270, 271, 272, 275, 276, 278, 280, 304, 305, 308, 309, 311, 322, 335 fiber bundles, 271, 275 fibers, 15, 17, 20, 38, 47, 48, 66, 70, 105, 106, 169, 175, 259, 269, 270, 272, 273, 274, 275, 276, 278, 280, 298, 300, 302, 304, 306, 354, 360 fibrillar, 141, 142 fibrillation, 191, 312 fibrils, 270, 305 fibrinogen, 362 fibroblast, 150, 155, 156, 157, 159, 273, 274, 276, 279, 280 fibroblast growth factor, 159, 274 fibroblasts, 145, 149, 150, 155, 156, 161, 273, 276, 279 fibroma, 225 fibronectin, 141, 142, 147, 149, 150, 151, 173 fibrous tissue, 298, 301 fibula, 211, 212, 213, 214, 215, 219, 220, 221, 222, 223, 224, 225, 226, 227 filled polymers, 73, 88 filler particles, 76, 77

fillers, 43, 45, 47, 65, 71, 72, 73, 79, 81, 85, 88, 105, 346 film, 46, 61, 68, 80, 110, 131, 141, 151, 176, 324, 325, 334, 335 films, 47, 53, 56, 62, 64, 66, 67, 69, 70, 90, 128, 131, 143, 147, 153, 154, 155, 156, 160, 168, 172, 174, 176, 178, 324 filters, 109 filtration, 109 fines, 109 Finland, 72, 123, 177, 230, 260, 363, 366 first generation, 178 flexibility, 16, 353, 355 flexural strength, 47, 76, 78, 130, 354, 356 flow, 21, 26, 27, 30, 31, 33, 35, 98, 102, 109, 110, 111, 112, 113, 117 flow rate, 26, 30, 33 fluid, 75, 76, 81, 144, 170, 172, 198, 199, 281, 284, 289, 290, 306, 310, 325, 341 fluid transport, 144 fluidization, 98 fluorescence, 180 fluorine, 153 foams, 11, 15, 84, 91, 92, 160, 168, 175 focusing, 244 foils, 235 folding, 10, 103 food, 4, 6, 101, 347 food additives, 6 Food and Drug Administration (PDF), 13, 36, 101, 134, 269, 271, 317, 325, 352, 353, 356 fractionation, 5, 129 fragmentation, 269 France, 154 free energy, 61, 156 free radical, 101, 127, 136, 169 free radical scavenger, 101 free radicals, 127, 169 free volume, 103 freeze-dried, 185 freezing, 10 friction, 271, 304, 306, 347, 348 frontoorbital, 232 fuel, 357 fumarate, 80, 90, 143, 152 fumaric, 357 fungus, 172 fusion, xi, 21, 22, 23, 35, 36, 37, 38, 39, 48, 61, 68, 187, 212, 222, 224, 225, 226, 227, 257, 258, 313, 314, 315, 316, 317, 318, 319, 321, 322, 323, 324, 325, 327, 328, 329, 330, 331, 332, 333, 334, 335, 336, 343, 357

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Index

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G gait, 211 Gamma, 100, 126, 127, 128, 133, 135, 136, 138, 361 gamma radiation, 127, 136 gamma rays, 126, 128 gas, 104, 117, 123, 126, 128, 129, 130, 137, 143, 144, 149, 158 gases, 143 gastric, 172 gel, 98, 127, 129, 160, 305, 354 gel permeation chromatography (GPC), 58, 101, 102, 105, 127, 129 gelatin, 159, 172, 173, 312 gelation, 49 gels, 103, 109 gene, 142 general surgery, 358 generation, 66, 146, 189, 199, 305, 336, 351 genes, 4, 11, 142 Geneva, 330, 333 Germany, xi, 22, 65, 66, 113, 229, 348, 364 Gibbs, 61 Gingiva, 341 glass transition, 9, 25, 28, 31, 39, 76, 105, 111, 115, 124, 126, 127, 128, 129, 164, 165, 256, 352, 353, 356 glass transition temperature, 9, 25, 28, 31, 76, 105, 111, 124, 126, 127, 128, 129, 164, 165, 256, 352, 353, 356 glasses, 27, 28, 30, 38, 72, 73, 76, 85, 92 glassy polymers, 23, 28, 30, 32, 33, 35, 37, 38, 39 glassy state, 37 glucose, 9, 152 glucose metabolism, 152 glutamine, 147 glycerol, 16, 44 glycine, 140, 159, 255 glycol, 164, 357 glycolysis, 9 glycosaminoglycans, 141, 296 gold, 156, 257, 267, 268, 281, 282 gold standard, 267, 268, 281, 282 Gore, 361 government, iv GPA, 13 GPC, 58, 101, 102, 105, 127, 129 grafting, 52, 65, 68, 153, 160, 164, 172, 226, 314 grafts, 211, 212, 223, 225, 226, 227, 267, 268, 275, 276, 281, 282, 291, 310, 325, 334, 341, 356, 357 granules, 82, 125, 212, 214, 215, 219, 220, 224, 344, 346

379

groups, 13, 15, 16, 44, 52, 54, 60, 64, 72, 97, 101, 105, 108, 110, 118, 127, 140, 143, 144, 146, 148, 153, 154, 155, 165, 166, 167, 172, 183, 199, 201, 249, 255, 258, 271, 276, 281, 282, 283, 284, 290, 305, 324, 354, 360 growth factor, 133, 149, 151, 159, 225, 274, 277, 279, 298, 322, 328, 343 growth factors, 149, 225, 274, 277, 279, 298, 322, 328, 343 growth inhibition, 232 growth rate, 355 guidance, 334 guidelines, 107, 331

H half-life, 180 hallux valgus, 343 hamstring, 268, 269, 288, 293 hand surgeon, 254, 263 handling, 202, 233, 237, 238 hardening, 27, 29, 30, 32, 37 Harvard, 253, 365, 366, 367 harvest, 212, 226, 227, 267, 268 harvesting, 211, 223, 224, 225, 227, 268, 273 hazards, 177, 179, 183, 187 healing, ix, 10, 11, 22, 42, 83, 93, 96, 97, 98, 102, 106, 110, 140, 152, 155, 159, 178, 185, 198, 200, 203, 235, 242, 245, 247, 254, 260, 267, 268, 274, 276, 279, 293, 295, 297, 299, 308, 310, 313, 322, 324, 325, 331, 341, 351, 354, 355, 357 health, 105, 135 health care, 135 heat, 28, 61, 82, 101, 105, 106, 108, 109, 110, 111, 123, 125, 133, 256, 342, 346, 348, 362 heaths, 332, 333 heating, 236, 255, 256 height, 256, 314, 317, 329 Helix, 368 hematoma, 317 hemostasis, 43 hepatocytes, 145, 156 heterogeneity, 145 heterogeneous, 105, 166, 168, 175, 181, 183, 270 high pressure, 186, 354 high resolution, 331 high temperature, 25, 124, 129, 171 hips, 188 histochemical, 309 histological, 35, 88, 137, 152, 181, 183, 239, 278, 280, 309, 323, 325, 333, 335, 360 histology, 180, 328 homeostasis, 274

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380

Index

homogeneity, 117 homogenous, 168, 331 homopolymers, 13, 103, 355 horizon, 263 hospitals, 124, 125, 126, 130, 356 host, 10, 22, 100, 224, 255, 262, 269, 270, 273, 276, 318, 356 host tissue, 10, 22, 262, 270 human, xi, 4, 5, 16, 19, 35, 36, 41, 42, 54, 73, 85, 91, 133, 138, 149, 150, 153, 154, 155, 156, 157, 158, 159, 160, 161, 172, 186, 187, 270, 292, 300, 307, 309, 310, 314, 316, 319, 321, 322, 328, 332, 334, 356, 358, 360 human lung fibroblasts, 157 humans, 35, 177, 179, 183, 187, 188, 282, 288, 300, 335 humerus, 187, 199, 213 humidity, 23, 34, 129, 131 hyaline, 185, 305 hybrid, 15, 91, 157, 158, 159, 361 hybridization, 90 hydration, 175, 167, 296 hydro, 18, 52, 58, 60, 63, 64, 68, 70, 82, 127, 130, 141, 143, 145, 147, 150, 151, 156, 157, 167, 183, 255 hydrogels, 70, 72, 90 hydrogen, 130, 143, 170 hydrogen peroxide, 130 hydrolysis, 6, 11, 13, 14, 17, 18, 54, 55, 57, 60, 65, 66, 67, 68, 69, 70, 102, 127, 128, 146, 163, 165, 167, 168, 169, 171, 172, 173, 174, 176, 181, 186, 247, 255, 259, 262, 273, 331, 345, 353, 355 hydrolytic stability, 11, 17 hydrolyzed, 52, 165, 172 hydrophilic, 18, 52, 58, 60, 63, 64, 68, 70, 82, 127, 130, 141, 143, 145, 147, 150, 151, 156, 157, 167, 183, 255, 353 hydrophilic materials, 141 hydrophilicity, 13, 17, 41, 52, 58, 59, 63, 64, 68, 130, 145, 147, 156, 165, 166, 172, 354 hydrophobic, 17, 18, 64, 70, 127, 130, 141, 143, 145, 148, 150, 151, 156, 166, 352, 353, 360 hydrophobicity, 145, 168, 183, 255, 356 hydroxyacids, 333 hydroxyapatite, 14, 20, 72, 73, 77, 78, 86, 88, 89, 90, 91, 92, 127, 130, 135, 136, 145, 147, 148, 155, 159, 164, 170, 355, 356, 360, 361, 362 hydroxyl, 6, 7, 8, 9, 16, 52, 60, 143, 153 hydroxyl groups, 16, 60, 153 hydroxylapatite, 137, 343 hydroxypropyl, 147 hypothesis, 181, 183 hysteresis, 145, 146

I id, 83, 145, 203, 224, 257, 258, 301, 304 identification, 100, 356 IFN, 191 IGF, 274 IL-2, 262 IL-8, 284 images, 51, 83, 232, 331 imaging, 22, 36, 38, 177, 183, 190, 198, 232, 254, 285, 291, 292, 293, 310, 319, 322, 331, 332, 334, 351 imbalances, 112 immersion, 67, 79 immobilization, 160, 173, 253 immune response, 262, 271 immunogenicity, 276 immunohistochemical, 137, 152, 360 immunological, 177, 204 impact strength, 14, 47, 49, 77 implantology, 142, 339 implementation, 22, 254, 263, 277 impurities, 13, 123, 125, 166, 181 in situ, 90, 322, 343 in situ hybridization, 90 in vitro, 20, 36, 54, 66, 68, 70, 71, 75, 84, 87, 89, 91, 92, 128, 131, 132, 135, 137, 140, 141, 146, 149, 152, 154, 159, 160, 161, 174, 175, 176, 184, 188, 189, 191, 239, 255, 273, 275, 276, 278, 279, 291, 312, 322, 323, 332, 335, 345, 357, 359, 360, 361, 362 in vivo, ix, 20, 21, 36, 38, 54, 57, 63, 68, 71, 75, 84, 86, 88, 91, 96, 129, 135, 149, 152, 155, 158, 174, 175, 176, 179, 184, 187, 188, 189, 191, 192, 239, 255, 256, 263, 273, 274, 275, 278, 288, 291, 292, 310, 311, 312, 321, 322, 323, 328, 331, 332, 333, 335, 349, 350, 353, 354, 357, 359, 360, 361 incidence, 197, 198, 200, 204, 262, 277 inclusion, 9, 170 inclusion bodies, 9 incubation, 83, 262 incubation time, 83 indication, 34, 146, 229, 230, 326, 328, 344 indicators, 124 induction, 102, 262, 356 industry, 4, 202, 254, 261 inert, 72, 184, 253, 262 infancy, 253, 261 infants, 231, 238 infection, xi, 197, 198, 211, 223, 257, 260, 261, 262, 269 infections, 199, 260, 268 infectious, 262, 341

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Index infectious disease, 341 infinite, 47 inflammation, ix, 14, 38, 129, 140, 203, 229, 281, 282, 301, 344 inflammatory, 10, 19, 22, 31, 87, 98, 104, 124, 146, 177, 183, 184, 187, 198, 245, 257, 259, 262, 281, 283, 284, 285, 314, 323, 324, 325, 328 inflammatory response, 10, 22, 31, 87, 98, 104, 146, 281, 314, 323, 324, 325, 328 inflammatory responses, 10, 31 inhibition, 43, 232 inhibitors, 279 inhibitory, 298 initiation, 24, 30, 32 injection, 76, 78, 84, 100, 104, 107, 109, 111, 112, 116, 117, 123, 125, 130, 134, 256, 344, 349, 352, 354 injection moulding, 76, 78, 123, 349 injuries, 187, 198, 199, 201, 207, 208, 261, 268, 277, 278, 282, 295, 351 injury, 140, 178, 197, 198, 203, 209, 260, 268, 280, 283, 308, 356 insertion, 12, 96, 98, 170, 182, 185, 190, 214, 215, 230, 231, 235, 282, 288, 290, 291, 326, 343, 346, 349 insight, 25, 102, 105 inspection, 99, 106, 107 inspectors, 108, 111 instability, 14, 21, 193, 211, 214, 223, 224, 226, 227, 241, 251, 322, 325 instruments, 124, 130 insulin, 274 insulin-like growth factor, 274 integration, 71, 272, 282, 285, 287, 288, 290, 314, 354 integrin, 142, 147, 150, 151 integrity, ix, 24, 165, 170, 171, 173, 273, 296, 298, 322, 323, 324, 331 interaction, 3, 17, 63, 99, 104, 141, 145, 148, 150, 157, 158, 167, 171, 270 interactions, 26, 131, 140, 142, 143, 146, 147, 148, 149, 153, 156, 157 interface, 18, 73, 76, 77, 80, 86, 87, 92, 105, 125, 134, 156, 176, 184, 308 interfacial layer, 139, 149 interfacial properties, 150 interference, 38, 94, 96, 179, 189, 281, 282, 285, 288, 289, 290, 291, 292, 293, 322, 342, 358, 359 intermolecular, 26 intermolecular interactions, 26 internal fixation, 35, 37, 92, 179, 185, 187, 188, 190, 191, 192, 198, 200, 201, 256, 263, 291, 323, 357, 360

381

interosseous membrane, 212 interrelations, 97, 119 intervention, 212, 233, 235, 245, 258, 259, 268, 297, 298, 322 intracranial, 231, 232 intracranial pressure, 231 intraocular, 175 intraoperative, 359 intrinsic, 27, 28, 29, 30, 32, 146, 259 invasive, 306, 356, 357 inversion, 49, 82, 202 ionic, 5, 155, 163, 167, 173, 174 ions, 54, 84, 85, 153, 168, 170, 355 iron, 71 irradiation, 123, 126, 127, 128, 129, 131, 133, 135, 136, 137, 138, 256, 310, 325 irritation, 184, 200, 201, 254, 256, 260, 261 ISO, 134, 135, 137 isolation, 171 isomerization, 255 isomers, 11, 43, 167, 199, 255, 331, 353 isotherms, 103 isotopes, 126, 128 Israel, 39 Italy, 304 ivory, 4

J Japan, 41, 68, 369 Japanese, 68, 305 Jefferson, 313, 369 joints, 178, 197, 214, 257, 267, 300, 308, 310, 312 Jung, 153, 239

K K+, 75 kappa, 244 keratin, 172 keratinocytes, 154 kinematics, 268, 270, 331 kinetics, 28, 30, 31, 32, 33, 34, 39, 93, 98, 99, 102, 103, 104, 105, 106, 109, 112, 169, 176 King, 89, 278, 310 knee, 87, 178, 187, 188, 199, 261, 267, 268, 270, 273, 277, 278, 281, 283, 285, 291, 295, 296, 300, 306, 307, 308, 310, 311, 312, 340, 357 knee arthroplasty, 273 knees, 282, 284, 304, 305, 307, 308, 358 kyphosis, 212, 225

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L L1, 131 lactate dehydrogenase, 11 lamella, 60 lamellae, 18, 127 lamellar, 10 lamina, 324 laminectomy, 322, 324, 334 land, 110 landfills, 4 Langmuir, 154, 176 lattice, 167 LCA, 37 leaching, 82, 83, 85, 157, 158, 305, 354, 360 leakage, 325 leg, 214, 219, 220, 221, 223, 226, 300 lesions, 184, 187, 188, 241, 250, 295, 297, 299, 308, 309 leucine, 147 leukemic, 151 leukemic cells, 151 leukocytes, 154 life cycle, 93, 94, 96, 97, 106 life span, 177, 179, 183 lifespan, 25 lifestyles, 269 lifetime, 24, 30, 32, 33, 34 ligament, 38, 135, 155, 186, 189, 192, 193, 250, 259, 261, 264, 265, 267, 268, 269, 270, 271, 273, 274, 275, 276, 277, 278, 279, 280, 290, 291, 292, 293, 297, 305, 308, 312, 324, 340, 342, 352, 358 ligand, 147, 173 Light-induced, 362 lignin, 3, 5 likelihood, 244 limitation, 202, 254, 269 limitations, 119, 178, 183, 202, 214, 223, 267, 268, 275, 295, 340, 359 linear, 11, 24, 46, 55, 128, 275, 276, 352, 355 linear polymers, 11 linkage, 16, 17 lipases, 171, 176 lipids, 138 liquid nitrogen, 84 liquid phase, 67 liquids, 131, 145 loading, ix, 22, 23, 24, 26, 27, 30, 31, 34, 35, 38, 41, 253, 261, 290, 296, 297, 301, 306, 321, 324, 331, 332, 351, 354, 355 localization, 24, 27, 30 location, 22, 104, 110, 170, 223, 351 locus, 262

logistics, 300 London, 70 long period, 54, 57, 144 losses, 85, 101 Louisiana, 278 low back pain, 328 low molecular weight, 102, 165, 166, 168, 171, 174 low-temperature, 123, 127, 130, 134, 137, 361 lubrication, 306 lumbar, 35, 36, 38, 314, 319, 321, 322, 323, 324, 325, 327, 328, 330, 331, 332, 333, 334, 335 lumbar spine, 314, 322, 324, 325, 328, 331, 332, 334, 335 lung, 157 lymph, 183, 323 lymph node, 183, 323 lymphatic, 183, 184 lymphatic system, 183, 184 lymphocyte, 262 lysine, 142 lysozyme, 285

M machines, 109, 111 macromolecular chains, 144 macromolecules, 4, 139, 141, 144, 149, 242 macrophage, 91, 157, 184, 361 macrophages, 84, 183, 255 Madison, 18, 35 magnetic, 22, 36, 38, 129, 177, 183, 232, 254, 292, 310, 319, 331, 334 magnetic resonance, 22, 36, 38, 177, 183, 232, 254, 292, 310, 319, 331, 334 magnetic resonance imaging, 22, 36, 38, 177, 183, 232, 292, 310, 319, 331, 334 maintenance, 36, 116, 296 Mammalian, 156 mammalian cells, 4 management, 260, 268 mandible, 226, 229, 236, 237, 238 mandibular, 160, 175, 181, 230, 234, 235, 236, 237, 238, 239, 357 manifold, 4, 111 manifolds, 112 manufacturer, 230 manufacturing, 3, 19, 93, 96, 97, 100, 101, 103, 104, 105, 106, 107, 108, 110, 111, 112, 117, 119, 134, 164, 185, 187, 189, 256, 360 market, 4, 97, 249, 339, 340, 343, 355 marketing, 94 marrow, 149, 158, 277 mass loss, 76, 345

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Index Massachusetts, 93, 253, 365, 366, 367 mastication, 230, 234 material degradation, 148 material surface, 140, 142, 145, 146, 148 materials science, 67, 254 matrix, 20, 72, 73, 74, 76, 78, 90, 105, 106, 133, 141, 142, 145, 147, 148, 149, 160, 165, 166, 167, 168, 172, 178, 186, 256, 270, 271, 272, 273, 274, 276, 280, 285, 296, 298, 301, 304, 305, 306, 317, 324, 354, 360 matrix metalloproteinase, 298 matrix protein, 141, 142, 145, 160 maturation, 273, 274 maxillary, 235 maxillary sinus, 235 MCL, 279 measurement, 107, 180 measures, 326, 328, 331 mechanical behavior, 275, 276 mechanical stress, 229, 238, 255 mechanical testing, 35 media, 54, 58, 60, 63, 69, 82, 165, 167, 168 medial collateral, 274, 279 medial meniscus, 307, 309, 310, 311 median, 257 mediation, 89 medication, 257 medicinal plants, 3 medicine, 19, 140, 342, 351 melt, 10, 11, 12, 13, 18, 31, 39, 47, 49, 57, 61, 66, 69, 97, 98, 99, 101, 102, 103, 105, 106, 107, 109, 110, 111, 112, 117, 157, 174, 178, 186, 256, 354, 360 melting, 9, 12, 13, 14, 43, 45, 46, 48, 98, 101, 103, 105, 108, 109, 110, 112, 116, 125, 127, 128, 164, 165, 172, 230, 347, 352, 353 melting temperature, 43, 45, 46, 125, 127, 165, 172 melts, 238, 348 membranes, 51, 52, 67, 68, 155, 158, 185, 340, 341 memory, 98, 151, 356, 362 mental disorder, 187 mercury, 131 mesenchymal stem cells, 148, 159, 225, 305 meta-analysis, 193, 199, 200 metabolic, 100, 101 metabolism, 152, 180 metabolite, 14 metals, ix, 22, 24, 37, 42, 72, 322, 339, 352 metatarsal, 199 methacrylic acid, 85 methanol, 67, 125 methyl group, 17, 172, 183, 255 methyl groups, 172, 183, 255

383

methylene, 15 Mg2+, 75 MHC, 262 micelles, 11 microbes, 65, 171 microbial, 156, 164, 165, 172 micrometer, 346 micro-molding, 109 microorganisms, 5, 134, 356 micro-organisms, 124, 130, 131 microparticles, 135, 148, 159, 168 microscope, 129 microscopy, 58, 76, 176, 180, 181, 183, 323 microspheres, 20, 127, 129, 136, 137, 160, 168, 173, 174, 175, 340, 358 microstructure, 49, 225 microvascular, 227 migration, 22, 183, 190, 225, 229, 240, 257, 258, 260, 262, 274, 314, 318, 321, 354 milk, 323 mimicking, 180 mineralization, 84, 91, 147, 161 mirror, 116 misleading, 325 mitogenic, 274 mixing, 49, 80, 82 MMP, 298 MMPs, 298 mobility, 21, 25, 26, 27, 30, 35, 54, 59, 60, 142, 145, 167, 172, 176 modalities, 267, 268 modality, 269 modeling, 35, 351, 352, 356, 357 models, 35, 74, 75, 140, 260, 276, 306, 321, 332 modulus, 13, 14, 15, 21, 26, 28, 46, 47, 49, 50, 72, 73, 76, 77, 79, 80, 81, 82, 83, 85, 86, 103, 125, 126, 128, 130, 186, 255, 276, 279, 301, 313, 318, 321, 323, 353, 355, 356 moieties, 102, 104, 105 moisture, 124, 125, 129 mold, 94, 103, 107, 110, 111, 112, 114, 116, 117, 305 mole, 4, 9, 61 molecular mass, 163, 170, 171 molecular mobility, 21, 26, 27, 30, 35, 60, 167, 172, 176 molecular orientation, 47, 49, 98, 99 molecular structure, 4, 44, 54, 58, 129 molecular weight distribution, 9, 60, 97, 102, 105, 120, 131, 165 molecules, 4, 9, 10, 17, 18, 25, 54, 60, 64, 65, 71, 125, 126, 131, 141, 142, 144, 165, 172, 270, 275 molybdenum, 71

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384

monoamine, 154 monoclonal, 150, 191, 283, 292 monoclonal antibodies, 150, 191, 283, 292 monocytes, 157, 262 monolayer, 176 monolayers, 144, 151, 154 monomer, 7, 8, 9, 41, 58, 69, 97, 99, 100, 101, 106, 108, 109, 112, 118, 155, 163, 164, 166, 255, 353, 355 monomers, 5, 7, 9, 11, 54, 58, 60, 65, 105, 128, 155, 156, 163, 172, 181, 199, 344, 354, 355 mononuclear cell, 262 mononuclear cells, 262 Moon, 137, 300, 311 morbidity, 211, 223, 226, 227, 258, 267, 268, 269 morphological, 67, 127, 128, 160, 273 morphology, 45, 47, 51, 54, 58, 63, 65, 66, 67, 68, 69, 70, 79, 90, 98, 100, 103, 105, 111, 117, 140, 152, 290, 310, 354, 361, 362 motion, 9, 22, 203, 214, 223, 259, 260, 267, 271, 314, 323 moulding, 76, 78, 84, 123, 186, 200, 349 mouse, 158 movement, 230, 233, 296, 306, 347 MRI, 23, 168, 175, 198, 216, 242, 281, 282, 285, 286, 287, 288, 290, 292, 307, 315, 331, 332, 357 muscle, 11, 153, 160, 170, 201, 268, 269, 288 muscle cells, 153, 160 muscle strength, 269 muscles, 212, 214 musculoskeletal, 199, 207, 267, 274, 277, 278, 279, 307, 351, 352 myelination, 325

N NA, 174, 189 Na+, 75 NaCl, 83 nanocomposites, 89, 90 nanocrystalline, 90, 91, 145, 155 nanofibers, 14, 51, 68 nanometer, 11, 144, 146, 155 nanoparticles, 172, 176 Nanostructures, 90 nanotechnology, 144, 155 natural, 4, 11, 67, 80, 143, 147, 157, 201, 270, 271, 275, 276, 293, 343, 361 natural polymers, 4 Nebraska, 163, 364, 367 neck, 24, 191, 199, 240 necrosis, 98, 301 needles, 81

negative influences, 273 neonatal, 240 nerve, 231, 237, 257, 333 nerves, 179 Netherlands, x, xi, 22, 66, 67, 68, 139, 197, 211, 295, 363, 364, 365, 366, 367, 368, 369 network, 20, 27, 142, 157 neural tissue, 321, 324, 325 neurogenic, 328 neurologic symptom, 231 neutrophil, 283 New York, 19, 20, 35, 37, 65, 66, 134, 278, 279, 307 New Zealand, 171, 324 nickel, 71, 357 Nielsen, 308, 309, 311 nitrogen, 84, 130, 150 NMR, 100, 129 Nobel Prize, 4 non toxic, 82 non-infectious, 262 non-random, 191, 242, 251 non-steroidal anti-inflammatory drugs, 198 non-uniform, 9 non-union, 22, 202, 208, 251, 257, 258, 315, 322, 330 non-vascular, 211, 224 normal, 86, 142, 179, 181, 185, 200, 201, 203, 214, 223, 270, 275, 277, 279, 300, 307, 311 North America, 328, 333 Norway, 281, 364 novel materials, 4 novel polymers, 351 nuclear, 126, 129 nuclear magnetic resonance, 129 nucleotide sequence, 277 numerical analysis, 311 nutrients, 298 nutrition, 341

O observations, 28, 31, 141, 145, 225, 265, 268, 285 OCD, 284 O-D, 37 oil, 5, 6, 127, 169 oils, 117 oleic acid, 79 oligomer, 163 oligomeric, 167 oligomers, 54, 58, 60, 65, 163, 165, 166, 167, 168, 169, 170, 181, 273, 325 oligopeptide, 148 oligosaccharides, 147

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Index omission, 203 optical, 11, 58, 102, 231, 255, 353 optical microscopy, 58 optimization, 273, 277, 279 oral, 42, 175, 189, 235, 239, 242, 361 organ, 10, 41, 45 organic, 5, 151, 164, 178 organic solvents, 164 organism, 124 orientation, 32, 39, 57, 60, 69, 112, 113, 200, 270, 271, 272, 297 orthopaedic, xi, 36, 37, 69, 88, 140, 142, 143, 144, 153, 159, 163, 179, 187, 189, 208, 263, 295, 361 orthopedic surgeon, 211 osmotic, 242, 243, 245 ossification, 231 osteoarthritis, 184, 257, 259, 263, 264, 297, 299, 300, 306, 307, 312 osteoblasts, 79, 84, 85, 146, 147, 158, 159, 160, 175, 360 osteocalcin, 147 osteochondritis dissecans, 188, 291 osteoclasts, 84, 147 osteoinductive, 133, 191, 225, 241 osteomyelitis, 14, 20, 212 osteopenia, 177, 178, 254 osteoporosis, 200 osteosarcoma, xi, 84 osteotomies, 88, 180, 186, 189, 190, 191, 212, 264, 357 osteotomy, 178, 180, 181, 186, 188, 232, 259, 260 outpatient, 214 overtime, 256 oxidants, 20 oxidative, 157 oxide, 49, 67, 70, 80, 90, 123, 129, 130, 135, 137, 138, 151, 154, 189, 335 oxygen, 5, 11, 126, 127, 130, 146, 150, 153, 154, 301

P packaging, 4, 5, 66, 94, 97, 99, 110, 123 pain, 22, 198, 200, 214, 222, 223, 227, 245, 257, 258, 259, 268, 269, 281, 283, 284, 297, 322, 328, 335, 343, 356 palpation, 229, 230 pancreatic, 172 paramagnetic, 127 parameter, 18, 46, 47, 102, 110, 145, 146, 274, 283 particles, 72, 76, 77, 78, 80, 82, 84, 85, 89, 90, 105, 112, 131, 183, 301, 323, 328, 349 patella, 199, 208, 268, 281, 282

385

patellar tendon, 186, 268, 276, 282, 291 pathogenesis, 187, 231 pathology, 359 pathways, 157 patterning, 147 PBT, 143, 154, 311 PDGF, 274 PDI, 9 pediatric, 96, 358 PEEK, 153, 330 pelvis, 213 penetrability, 128 peptide, 11, 147, 151, 152, 158 peptides, 5, 142, 147, 172, 173, 277, 283 perception, 257, 352 perforation, 242, 244, 318 periodontal, 155 periodontitis, 341 periodontium, 341 periosteum, 212 peripheral blood, 262 peripheral blood mononuclear cell, 262 permeability, 70 permeation, 58, 127, 129 permit, 298 peroxide, 98, 130 PET, 17, 18, 42, 139, 145, 154 petroleum, 5 pH, 16, 56, 57, 58, 59, 62, 64, 71, 82, 83, 85, 98, 144, 163, 165, 166, 168, 169, 170, 171, 172, 173, 174, 175, 184, 190, 249, 345 pH values, 184 phagocyte, 20 phagocytic, 183, 258 phagocytosis, 98, 99, 190 phalanges, 199 pharmaceutical, 41, 42, 43, 44 pharmacological, 10 phase boundaries, 67 PHB, 9, 10, 14, 42, 45, 46, 72, 73, 76, 77, 83, 84, 86, 87, 88, 91, 129, 333, 356 phenomenology, 23 phenotype, 158, 160 Philadelphia, 35, 277, 313, 369 phosphate, 54, 55, 56, 57, 58, 59, 60, 61, 62, 64, 67, 68, 70, 72, 73, 76, 78, 82, 83, 85, 88, 89, 90, 91, 92, 125, 128, 134, 148, 158, 159, 160, 161, 168, 170, 201, 211, 212, 227, 256, 314, 341, 343, 346, 355, 358, 359, 361 phosphate glasses, 92 phosphates, 78, 82, 84, 85, 343 phosphorous, 82 phosphorylation, 150, 151

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Index

photochemical, 131 photons, 130, 131, 357 physical activity, 223 physical aging, 28, 69 physical and mechanical properties, 14, 125, 163, 164, 352 physical force, 111 physical health, 331 physical properties, 3, 10, 15, 16, 43, 44, 45, 52, 70, 75, 103, 152, 163, 178, 352, 361 physical therapy, 269 physicochemical, 142, 181, 184 physico-chemical characteristics, 164 physicochemical methods, 142 physico-chemical properties, 128, 140 physics, 70 physiological, 75, 81, 144, 147, 325 piezoelectric, 14, 76 piezoelectric properties, 14 pigs, 240, 260 pilot studies, 259 pilot study, 319, 333, 334 pitch, 116, 202 planning, 105, 282 plantar, 211, 214, 223 plantar flexion, 211, 214, 223 plants, 3, 6 plaques, 76 plasma, 52, 75, 80, 82, 123, 130, 131, 132, 135, 137, 143, 144, 149, 150, 151, 153, 154, 156, 284, 292, 305, 312 plasminogen, 279 plastic, 4, 5, 21, 22, 23, 24, 26, 27, 28, 29, 30, 31, 33, 35, 37, 111, 112, 124, 127, 242, 252, 275, 331, 345 plastic deformation, 22, 23, 24, 27, 28, 29, 30, 35, 275, 331, 345 plastic strain, 26, 27, 30 plasticity, 38 plasticization, 32, 39 plasticizer, 54 plastics, 4, 65, 101, 103, 104, 105, 106, 107, 109, 110, 111, 112, 356 plastics processing, 109 platelet, 156, 274, 305, 312 platforms, 346 PLGA, 13, 19, 80, 82, 85, 126, 127, 128, 136, 137, 140, 143, 146, 147, 148, 155, 157, 158, 159, 160, 173, 175, 199, 229, 230, 238, 334, 357, 360 PM, 208, 308, 312, 319 PMMA, 37 polarity, 10 polarized light, 111, 323

polarized light microscopy, 323 poly(2-hydroxyethyl methacrylate), 145, 156 poly(3-hydroxybutyrate), 88, 89, 333, 354 poly(ethylene terephthalate), 7, 42, 160 poly(glycolide), 3, 45 poly(lactic-co-glycolic acid), 157, 159, 358 poly(L-lactide), 12, 31, 36, 39, 44, 66, 67, 68, 69, 70, 92, 134, 135, 137, 174, 176, 190, 250, 308, 332, 334, 349, 359, 361 poly(methyl methacrylate), 154, 155, 360 polyamides, 5 polyamine, 154 polycarbonate, 20, 23, 24, 25, 26, 28, 30, 37, 38, 39, 156, 157 polycarbonates, 5, 15, 16, 18, 143 polycondensation, 43, 164 polydispersity, 166 polyester, 7, 8, 15, 16, 17, 18, 22, 41, 67, 161, 165, 176, 184, 185, 186, 301, 309, 352, 353 polyesters, 3, 5, 6, 7, 8, 9, 10, 11, 14, 15, 16, 17, 18, 42, 44, 46, 52, 55, 64, 65, 67, 68, 69, 70, 90, 175, 186, 187, 199, 351, 353, 356 polyesterurethanes, 152 polyether, 15 polyethylene, 72, 120, 164, 183, 190, 269 polyethylene terephthalate, 269 polyglycolic acid, 73, 134, 135, 178, 191, 199, 200, 271, 285, 288, 292, 314, 331, 355, 358, 359 polyhydroxybutyrate, 20, 72, 88, 92, 356 polymer blends, 49, 66, 71, 73, 104, 105, 147, 173 polymer chains, 9, 18, 32, 72, 103, 111, 117, 165, 167, 172, 345 polymer composites, 71, 72 polymer density, 73 polymer film, 80, 131, 151 polymer materials, 300 polymer matrix, 78, 90, 133, 165, 178 polymer mixing, 49 polymer molecule, 126 polymer networks, 92, 151 polymer properties, 133, 137, 173, 357 polymer structure, 103, 110, 146, 256, 362 polymer synthesis, 97 polymer systems, 38, 142, 147, 148 polymerase, 5 polymeric blends, 164 polymeric materials, 3, 4, 45 polymeric matrices, 271 polymerization, 5, 7, 9, 12, 13, 14, 15, 19, 32, 43, 97, 99, 100, 101, 102, 103, 104, 105, 106, 156, 163, 164, 166, 173, 174, 255, 351 polymerization process, 15, 101 polymerization temperature, 102, 166

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Index polyolefins, 5 polyorthoesters, 128 polypeptide, 156 polypeptides, 283 polyphenols, 5 polypropylene, 269 polysaccharide, 3, 4, 147 polystyrene, 29, 38, 42, 140, 143, 150, 151, 153, 154 polytetrafluoroethylene, 149, 269, 278 polyurethane, 15, 16, 17, 152, 153, 157, 160, 276, 280, 295, 299, 300, 308 polyurethane foam, 152 polyurethanes, 5, 15, 16, 17, 127, 130, 136, 137, 143, 157, 300, 311 polyvinyl alcohol, 310, 311 poor, 72, 76, 110, 131, 170, 181, 187, 258, 261, 269, 287, 288, 300, 304, 332, 353 popliteus, 304 population, 134, 200, 277, 331 pore, 49, 51, 54, 65, 82, 83, 144, 145, 148, 161, 212, 225, 271, 272, 301, 305, 311 pores, 143, 144, 146, 148, 168, 272, 301 porosity, 57, 82, 83, 85, 129, 145, 170, 181, 212, 255, 271, 272, 273, 300, 301, 305, 347, 361 porous, 10, 49, 50, 65, 67, 69, 70, 79, 82, 85, 89, 90, 92, 129, 144, 148, 152, 158, 159, 160, 168, 169, 170, 175, 181, 189, 212, 271, 276, 279, 298, 300, 301, 308, 311, 312, 346, 347, 357, 360, 362 porous materials, 129, 301 postoperative, 222, 227, 313, 324, 326, 327, 329, 334 pouches, 127 powder, 77, 84, 185, 344 powders, 109 power, 130, 234, 236 PPD, 262 pragmatic, 93 precipitation, 67, 105 preclinical, 263, 295, 299, 321, 325, 331 pre-clinical, 22, 179 prediction, 30, 34, 37, 356 predictive model, 35 predictive models, 35 pre-existing, 184, 344 president, xi press, 189, 239, 292 pressure, 4, 49, 109, 110, 129, 130, 201, 231, 345, 347 prevention, 43, 224, 276, 299, 300 prices, 126 printing, 83, 356 probability, 60 producers, 76 product design, 94, 100, 110, 111

387

product life cycle, 93, 96, 107 production, 11, 20, 52, 72, 75, 79, 82, 87, 91, 97, 109, 110, 111, 117, 123, 160, 166, 167, 178, 184, 202, 230, 253, 254, 256, 273, 276, 277 progenitor cells, 298 prognosis, 307 program, 116, 134 proinflammatory, 91, 283, 361 prolapse, 324 proliferation, 43, 91, 139, 140, 142, 145, 147, 148, 149, 151, 154, 155, 157, 159, 268, 269, 271, 273, 274, 276, 279, 361 propagation, 105 property, 3, 9, 41, 47, 65, 68, 69, 72, 125, 143, 304, 324, 362 propylene, 139, 143, 152, 357 prostheses, 120, 270, 273, 301 prosthesis, 270, 276, 279, 280, 301, 311 proteases, 157, 171 protection, 43, 177, 178, 299, 300 protein, 82, 131, 140, 141, 142, 149, 151, 154, 158, 160, 179, 247, 314, 319, 328, 334 proteinase, 60, 61, 171, 172, 176 proteins, 3, 4, 5, 141, 142, 144, 145, 148, 149, 150, 151, 156, 173, 274, 296 proteoglycans, 141 protocols, 35, 94, 106, 107, 108, 110, 112, 141, 260 protons, 6 prototype, 101, 109, 117 prototyping, 116 PTFE, 361 purification, 4, 13, 171 PVA, 52, 53, 63, 70, 310, 311 PVC, 31, 111 PVC polymers, 111 PVP, 131 pyruvate, 9

Q quadriceps, 268, 269 quality assurance, 94, 97, 107 quality control, 102, 107 quarantine, 126 quartz, 176 questionnaires, 328, 331 quinine, 258 quinone, 262

R race, 11, 12, 18, 31, 43, 167, 171, 325

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Index

radiation, 97, 98, 126, 127, 128, 129, 131, 135, 136, 137, 156, 223, 334 radiation therapy, 223 radical polymerization, 9 radiculopathy, 330, 335 radio, 35, 153, 258 radiofrequency, 150 radiography, 22, 180, 223, 333 radiological, 22, 231, 328 radiopaque, 135, 316, 322 radius, 184, 186, 192, 199, 200, 256, 257, 258, 264 random, 8, 10, 14, 18, 25, 43, 101, 102, 147, 158, 163, 165, 168, 255, 355 range, 10, 13, 27, 49, 57, 60, 63, 82, 86, 88, 104, 109, 110, 127, 145, 148, 198, 212, 214, 223, 259, 260, 271, 301, 314, 328, 339, 353, 355 Rapid Prototyping, 305 rat, 84, 85, 92, 158, 181, 185, 308, 325, 333, 334 rats, 87, 88, 170, 179, 190, 288, 312, 332, 345, 360 raw materials, ix, 101, 103, 105, 107, 125 reaction rate, 38 reaction temperature, 12 reactive groups, 144 reactive oxygen, 146 reactive oxygen species, 146 reactivity, 262, 354 recall, 124 receptors, 141 recombinant DNA, 356 recombination, 110 reconstruction, 43, 89, 152, 157, 158, 189, 192, 193, 198, 211, 212, 214, 217, 218, 219, 222, 224, 226, 227, 235, 241, 242, 246, 249, 250, 251, 259, 264, 265, 277, 278, 279, 280, 285, 288, 289, 290, 291, 292, 293, 308, 311, 358 reconstructive surgery, 223, 292 recovery, 98, 268, 269 recreational, 200 recycling, 69, 174 refineries, 5 refractory, 328 regenerate, 67, 223, 225, 276 regeneration, 20, 22, 42, 44, 91, 140, 151, 152, 158, 159, 160, 179, 185, 211, 212, 223, 224, 227, 267, 270, 272, 273, 275, 276, 277, 279, 295, 300, 304, 305, 308, 310, 311, 312, 324, 325, 333, 334, 355, 358, 361 regenerative capacity, 295, 300 regenerative medicine, 140 regression, 184 regular, 10, 126, 145, 164, 214 regulation, 129, 134, 298 regulations, 123

regulatory bodies, 124 regulatory requirements, 99 rehabilitation, 98, 247, 259, 260, 297 reinforcement, 78, 84, 105, 106, 276, 298, 352, 354, 357 reinforcing fibers, 105, 106 rejection, 87 relationship, 61, 128, 142, 150, 151, 291, 355 relationships, 3, 69, 145, 270, 306 relaxation, 39, 98, 100, 103, 111, 114 relevance, 107 reliability, 4, 177, 187, 333 Reliability, 333 remodeling, 114, 171, 218, 219, 220, 222, 224, 225, 298, 301, 323 remodelling, xi, 323 renewable resource, 41, 65 repair, 36, 91, 171, 179, 188, 189, 192, 197, 199, 242, 244, 249, 250, 251, 261, 268, 269, 270, 278, 279, 280, 285, 292, 295, 297, 298, 299, 306, 307, 308, 351, 352, 357, 358 repeatability, 107 research and development, 177, 267, 275 resection, 212, 223, 226, 227, 297, 307 reservation, 314 residuals, 126 residues, 54, 55, 56, 57, 60, 68, 70, 151, 166 resin, 100 resistance, 72, 109, 131, 151, 167, 272, 344 resolution, 86, 331, 334 resources, 41, 65 respiration, 180, 255 restitution, 182 retention, 88, 102, 131, 135, 177, 178, 184, 185, 186, 191, 192, 230, 355, 359 revascularization, 298 rheological properties, 90 rheology, 98, 102, 103, 105, 110, 112 rheumatoid arthritis, 257 rice, 363, 365 rigidity, 126, 178, 200, 235, 260, 261, 314, 344, 348, 351 risk, 22, 79, 94, 106, 124, 125, 185, 187, 197, 198, 200, 201, 203, 237, 245, 246, 260, 262, 267, 271, 321, 322, 324, 331, 341, 347 risk assessment, 94 risks, 170, 173, 177, 188, 197, 198, 269, 328 RNA, 4 rodents, 183 rods, 76, 77, 80, 86, 88, 90, 92, 129, 135, 170, 178, 179, 181, 186, 188, 191, 192, 198, 208, 254, 260, 262, 263, 264, 282, 285, 288, 292, 324, 331, 333, 349, 352, 353, 360

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Index roentgen, 331 room temperature, 33, 34, 76, 178, 256, 353 ROP, 43, 44 rotator cuff, 241, 261, 352, 359 roughness, 82, 127, 130, 144, 145, 146, 155 RRM, 134, 189, 190 rubber, 4, 5, 32, 39, 67, 353 rubbery state, 9

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S SAD, 87 safety, 126, 129, 164, 289, 325, 332 saline, 77, 82, 83, 111 salt, 82, 83, 85, 158, 305 salts, 164, 170 sample, 24, 27, 28, 102, 126 sampling, 107 SBF, 75, 76, 77, 78, 82, 83, 84, 85 scaffolding, 20, 106 Scandinavia, 199 scanning calorimetry, 58, 127 Scanning electron, 76, 77, 79 scanning electron microscopy (SEM), 51, 58, 76, 81, 83, 129 scaphoid fracture, 208, 258 scar tissue, 324, 332 scatter, 254 scattering, 10 scepticism, 197, 339 Schwann cells, 333 scientific community, 4 scoliosis, xi, 212 search, 300 secretion, 3 security, 282, 293 seed, 304 seeding, 150, 158, 273, 279, 304 seizures, 231 selecting, 106 self-assembling, 9 self-assembly, 278, 305 semi-crystalline polymers, 9, 164 sensitivity, 38, 100, 112 separation, 49, 67, 84, 111, 186 septic arthritis, 268 septum, 212 sequelae, 19, 317 sequencing, 10 series, 23, 38, 107, 165, 183, 200, 203, 224, 225, 227, 254, 261, 271, 283, 316, 322, 325, 328, 333 serine, 157, 171 serum, 82, 141, 149, 150, 154, 156

389

serum albumin, 149 shape, 11, 18, 43, 54, 65, 76, 102, 107, 111, 142, 165, 178, 214, 223, 288, 347, 356, 357, 362 shape-memory, 362 shaping, 158, 178 shear, 29, 30, 80, 86, 98, 102, 106, 108, 109, 110, 112, 115, 125, 186, 296, 357 shear strength, 86, 125, 186, 357 sheep, 20, 170, 179, 190, 192, 227, 239, 282, 290, 299, 304, 309, 310, 312, 313, 314, 319, 324, 334 Sheep, 311 shipping, 94, 111 SHM, 39 short period, 245 short-term, 21, 24, 34, 301, 318 shoulder, 187, 191, 193, 241, 242, 243, 244, 245, 247, 249, 250, 251, 261, 284, 292, 340 shoulders, 251 sign, 281 signaling, 150, 151 signalling, 142 signals, 76, 142, 274 signs, 183, 214, 223, 231, 235, 247, 261, 297, 300, 301, 305, 314, 316, 323, 325, 328, 330, 331 silicon, 82, 84 silk, 3, 85, 143, 152, 171, 172, 271, 276, 280, 361 similarity, 171 simulated body fluid, 75, 76, 170 Singapore, 361 sinus, 198, 201, 235, 254, 257, 258 sinuses, 200 sinusitis, 199 SIS, 140, 148 sites, ix, 14, 142, 199, 200, 223, 224, 262, 268, 272, 351 skeleton, 22 skills, 339 skin, 14, 104, 117, 124, 155, 212, 214, 229, 230, 260, 276, 356 sleep disorders, 231 small intestine, 157 smooth muscle, 153, 160 smooth muscle cells, 153, 160 sodium, 80, 170, 312 soil, 5, 172 soleus, 212 sol-gel, 360 solid phase, 150 solubility, 76, 165, 167, 170 solvation, 151 solvent, 9, 49, 51, 67, 77, 80, 82, 98, 99, 153, 354 solvents, 11, 49 South Dakota, 351, 364

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390

Index

spacers, 38, 317, 319, 322, 333 spatial, 25, 181, 270 species, 130, 146, 177, 179, 187 specificity, 150, 151, 157 spectroscopy, 129 spectrum, 140, 142, 305, 344, 346, 351 speed, 34, 116, 126, 128, 277, 341 spheres, 14, 82 spin, 131 spinal cord, 325, 333, 334 spinal fusion, xi, 36, 68, 321, 322, 323, 324, 325, 328, 331, 333, 334, 335, 336 spine, xi, 22, 23, 31, 35, 225, 226, 313, 314, 316, 317, 318, 319, 321, 322, 324, 325, 328, 330, 332, 333, 334, 335, 349, 361 splint, 3 splinting, 257 spondylolisthesis, 225, 226, 322, 325, 330, 334, 335 sporadic, 183 sports, 19, 203, 214, 223 sprue, 111, 112 stability, 11, 17, 72, 85, 98, 99, 101, 102, 126, 127, 134, 136, 137, 167, 168, 176, 198, 200, 212, 229, 230, 233, 236, 237, 238, 239, 241, 244, 249, 258, 259, 260, 261, 267, 269, 271, 277, 282, 323, 324, 339, 354, 358 stabilization, 35, 191, 223, 226, 227, 250, 251, 275, 284, 292, 296, 318 stabilize, 257, 259, 324, 341 stabilizers, 101 stages, 35, 97, 102, 106, 111, 123, 255, 268 stainless steel, 71, 186, 208, 324, 357 standardized testing, 35 standards, 101, 108, 123, 126 starch, 5, 164 steady state, 27, 112 steel, 71, 117, 186, 208, 324, 357 stem cell therapy, 305 stem cells, 148, 152, 159, 225, 305 stenosis, 330, 334, 335 stereospecificity, 163 steric, 59 sterile, 97, 99, 101, 124, 130, 132, 134, 198, 199, 200, 203, 214, 261, 262, 285 sterilisation, ix, 137, 181, 189 sterilization, 94, 97, 98, 100, 111, 123, 124, 125, 126, 127, 128, 129, 130, 131, 133, 134, 135, 136, 137, 138, 175, 255, 256, 268, 269, 324, 325 stiffness, 22, 36, 71, 72, 77, 78, 79, 80, 81, 82, 87, 201, 270, 285, 296, 300, 301, 314, 322, 335, 352, 353 stimulus, 356 stock, 112, 116

storage, 9, 76, 77, 85, 94, 111, 198 strain, 9, 14, 23, 24, 25, 26, 27, 28, 29, 30, 31, 32, 33, 35, 37, 51, 52, 73, 74, 75, 76, 81, 108, 275, 276, 344, 356 strains, 11, 30, 74, 171, 172, 239, 275 strategies, 140, 211, 279, 295, 351, 352 stratification, 161 stress, 9, 16, 21, 22, 23, 24, 25, 26, 27, 28, 29, 30, 31, 32, 33, 34, 35, 37, 38, 39, 49, 51, 73, 74, 75, 77, 98, 99, 103, 111, 113, 114, 117, 177, 178, 200, 229, 238, 255, 256, 267, 269, 270, 273, 275, 276, 296, 318, 322, 344, 351, 357 stress level, 24, 25, 26, 35, 77 stress-strain curves, 25, 27, 30, 31 stromal, 148, 149, 154, 158, 159, 160, 276, 280 stromal cells, 148, 149, 158, 159, 160, 276 strong interaction, 49, 64 structural changes, 47 structural relaxation, 39 sub-cellular, 147 subcutaneous tissue, 92, 186, 192, 255, 260 subjective, 211, 281, 283 submucosa, 157 substances, 54, 283, 306 substitutes, 67, 90, 309, 322, 341 substitution, 155, 201, 224, 310 substrates, 149, 150, 154, 271 subtilisin, 171 success rate, 224, 300 sucrose, 164 Sudan, 303 suffering, 112 sugar, 82 sulphate, 306, 349 sulphur, 4 Sun, 151, 158, 160, 307, 359, 362 supercritical, 49, 85, 160 supercritical carbon dioxide, 160 superiority, 202, 326 suppliers, 100, 101, 126, 129 supply, 59, 60, 64, 117, 181, 183, 211, 267, 269, 271, 272, 301 surface area, 105, 110, 146, 148, 165, 168, 175, 181, 262, 272 surface chemistry, 18, 150, 151, 156 surface energy, 18, 82 surface modification, 18, 139, 141, 143, 146, 147, 148, 149, 158, 160, 172, 176 surface properties, 3, 41, 65, 131, 140, 141, 146, 165 surface roughness, 82, 127, 130, 145, 146, 155 surface structure, 43, 144 surface treatment, 52, 65, 80 surfactant, 89, 147, 158

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Index surgeons, 22, 198, 239, 254, 261, 263, 268, 278, 279, 349 surgeries, 22, 268, 357 surgical, ix, 22, 92, 93, 96, 97, 98, 124, 170, 177, 178, 179, 188, 189, 197, 198, 200, 201, 202, 203, 212, 223, 225, 232, 233, 235, 236, 237, 238, 251, 254, 255, 257, 259, 263, 267, 268, 269, 284, 297, 298, 305, 306, 309, 318, 321, 326, 328, 333, 341, 352, 358, 359 surgical intervention, 198, 212, 233, 235, 259, 268, 297, 298, 306 surprise, 268 surveillance, 94 survival, 354 susceptibility, 11 suspensions, 145 suture, 19, 85, 94, 96, 97, 178, 231, 233, 234, 235, 241, 242, 249, 250, 251, 254, 261, 299, 340, 342, 348, 352, 355, 356, 361 Sweden, 38, 241, 244, 259, 364, 366, 367 swelling, 9, 145, 166, 191, 198, 200, 222, 254, 257, 259, 261, 262, 284, 315, 345 Switzerland, 330, 339, 348, 367 symbols, 26 symmetry, 10 symptom, 231, 330 symptoms, 22, 211, 214, 223, 231, 284, 316 syndrome, 240 synergistic, 274 synergistic effect, 274 synovial fluid, 81, 284, 306 synovial tissue, 299 synovitis, 245, 247, 269, 270, 285, 291, 292, 301, 310 synthesis, 5, 7, 8, 10, 14, 15, 17, 43, 80, 91, 97, 100, 152, 157, 159, 161, 163, 164, 236, 311, 356 synthetic polymers, 4, 65, 190, 250, 276, 339 systems, 13, 16, 39, 42, 44, 49, 67, 142, 143, 146, 157, 159, 173, 181, 209, 230, 233, 237, 239, 253, 254, 279, 340, 341, 353, 356

T tacticity, 10, 69, 172, 173 TCC, 283, 284 TCP, 72, 78, 80, 81, 83, 84, 90, 140, 142, 143, 145, 224, 225, 344, 355, 361 technological progress, 4 technology, xi, 68, 160, 254, 261, 263, 318, 321, 322, 325, 346, 352, 356, 358 teeth, 341 Teflon, 142, 145, 300, 310, 311 TEM, 58, 86, 87

391

temperature, 9, 23, 25, 26, 28, 30, 31, 34, 37, 46, 47, 57, 58, 68, 76, 83, 98, 100, 103, 110, 112, 123, 124, 125, 126, 127, 129, 130, 131, 134, 137, 163, 167, 168, 171, 172, 173, 174, 214, 245, 255, 348, 353, 354, 360, 361 temperature dependence, 37 tendon, 186, 198, 200, 256, 259, 264, 265, 268, 276, 278, 279, 281, 282, 291, 299, 304, 309, 340, 342 tendons, 198, 260, 269, 288, 342 tenosynovitis, 200, 257 tensile, 4, 13, 15, 24, 27, 33, 37, 47, 49, 58, 59, 62, 64, 76, 77, 79, 80, 127, 130, 132, 171, 255, 269, 275, 296, 304, 345, 352, 353, 355, 356, 357, 359 tensile strength, 4, 13, 15, 47, 58, 59, 62, 64, 76, 79, 80, 127, 130, 132, 171, 255, 269, 275, 304, 352, 353, 355, 359 tensile stress, 275, 296 tension, 30, 32, 37, 96, 109, 271, 276, 324 teratogenic, 177, 183 terephthalic acid, 17 test data, 117 textile, 17, 38, 129 TGF, 133, 274 therapy, 197, 305, 312, 347 thermal degradation, 110 thermal energy, 231 thermal properties, 105, 126, 127, 128, 130, 133 thermal stability, 72, 85, 98 thermal treatment, 28, 30, 125, 135, 192 thermally induced phase separation, 84 thermodynamic, 28, 101 thermodynamic equilibrium, 28, 101 thermodynamics, 151 thermoplastic, 9, 11, 92 thermoplastics, 37, 38 thin film, 156, 168, 176 thin films, 156, 168, 176 thinking, 103 thioridazine, 63 Thomson, 61, 67 thoracic, 225, 325 three-dimensional, 10, 11, 69, 89, 91, 125, 143, 145, 148, 158, 159, 160, 270, 271, 272, 273, 276, 362 three-dimensional space, 148 threshold, 324 tibia, 213, 219, 220, 221, 223, 227, 267, 268, 286, 288, 290, 291, 296, 297, 304, 353 time consuming, 233 time frame, 297 time periods, 125 time use, 4 time-frame, 4 Timmer, 152

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392

Index

tin, 13, 166, 181 TiO2, 85 tissue engineering, xi, 10, 14, 15, 19, 20, 66, 69, 71, 90, 91, 92, 131, 137, 140, 143, 152, 154, 157, 158, 159, 160, 161, 172, 267, 271, 273, 275, 276, 277, 278, 304, 305, 306, 310, 312, 340, 356, 358, 361 titania, 85, 91, 155 titanium, 85, 189, 208, 229, 233, 234, 235, 236, 237, 239, 250, 259, 261, 282, 290, 291, 292, 313, 317, 323, 335, 339, 343, 347, 357 Titanium, 239 titanium dioxide, 85 TMC, 259, 264, 353 Tokyo, 68 tolerance, 124 torque, 96, 98, 109, 261, 282, 291 total costs, 97 toughness, 4, 14, 32, 72, 272, 353, 355 toxic, 10, 11, 15, 82, 126, 129, 138, 164, 169, 177, 183, 300, 352 toxic effect, 169 toxic products, 300 toxicity, 13, 41, 42, 134, 169, 175 trabeculae, 314 trabecular bone, xi, 36, 183, 185, 224 tracking, 107 traction, 203 trade, 13 tradition, 140 training, 297 trajectory, 214 transcription, 4, 142 transection, 308 transesterification, 125 transesterification reaction, 125 transfer, 65, 73, 100, 109, 211, 226, 227, 270, 273, 296, 297, 322, 323, 345 transference, 98 transformation, 81, 224 transformations, 348 transforming growth factor, 133, 274 transition, 9, 31, 37, 133, 185, 255, 270, 354 transition temperature, 9, 25, 28, 31, 76, 105, 111, 124, 126, 127, 128, 129, 133, 164, 165, 185, 255, 256, 352, 353, 354, 356 transitions, 9 transmission, 46, 58, 180, 183, 267, 268, 269, 271, 297, 306 transmission electron microscopy, 58, 180, 183 transparent, 108, 164, 349 transplant, 300 transplantation, 227, 235, 300, 304, 309, 310

transport, 144 trapezium, 259 trauma, ix, 21, 22, 197, 202, 203, 204, 205, 207, 211, 284, 297, 306, 341 trial, 36, 179, 188, 191, 200, 202, 203, 208, 250, 251, 257, 260, 291, 292, 299, 333 tricarboxylic acid, 165 tricarboxylic acid cycle, 165 triggers, 348 tritium, 181 trypsin, 171 tubular, 224, 260, 324, 325 tumor, 211, 212, 213 tumour, 335 two-dimensional, 159, 160 tyrosine, 15, 20, 150, 151, 276

U ulceration, 260 ultrasound, 230, 238, 239 ultrastructure, 92, 307, 309 ultraviolet, 130, 131 uniform, 9, 74, 79, 164, 311, 326, 328 unions, 22, 257, 258, 315 United Kingdom (UK), 71, 257, 264, 368 United States, 101, 268, 269, 275, 278 urea, 16, 17, 264, 276, 280 urethane, 15, 16, 17, 39, 152, 157, 264, 356, 362 urine, 170 UV irradiation, 131, 137, 138 UV radiation, 131

V vacuum, 126, 128 valgus, 211, 213, 214, 221, 223, 224, 225, 227 validation, 93, 94, 102, 107, 123, 124, 126, 129, 134, 135, 362 values, 27, 47, 49, 57, 58, 59, 60, 61, 62, 63, 64, 74, 76, 78, 142, 145, 184, 261, 276, 284, 324, 331 vapor, 46 variability, 268, 277 variables, 101, 169, 354 variation, 102, 112, 124, 127, 143, 184 vascular endothelial growth factor (VEGF), 298 vascularization, 22, 98, 155, 298 vehicles, 4, 11, 14 vein, 150, 156 velocity, 34, 109 vertebrae, 343 vessels, 179, 212, 225, 298

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Index viscosity, 47, 80, 99, 101, 104, 105, 106, 109, 112, 128, 132, 181, 324 visible, 25, 87, 108, 119, 219, 241, 242, 245, 246, 247, 249, 285, 287, 289, 290, 331, 349 vision, 231 visualization, 190 voids, 106, 165 vortex, 117 vulcanization, 4

393

wetting, 77, 105, 106 windows, 101, 102, 110, 111 wires, 42, 198, 201, 202, 240, 254, 261 women, 212 wool, 3 workers, 147 wound healing, 140 wound repair, 306

X W xenograft, 300 X-ray crystallography, 10

Y yarn, 275 yeast, 171 yield, 5, 8, 10, 15, 17, 21, 24, 25, 26, 27, 28, 29, 30, 32, 33, 34, 37, 39, 41, 43, 49, 57, 98, 164, 268, 275, 355

Z zeta potential, 168 zinc, 166, 174, 181 zygomatic, 36, 190, 234, 235, 250, 349

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walking, 214, 223, 277 Warsaw, 314 water, 6, 7, 8, 16, 17, 18, 34, 44, 49, 50, 54, 58, 59, 60, 64, 65, 67, 78, 80, 81, 82, 85, 103, 106, 117, 129, 131, 141, 145, 153, 163, 164, 165, 166, 167, 168, 169, 170, 172, 176, 214, 255, 331, 345 water absorption, 78, 80, 106, 166, 167 water diffusion, 59, 60 water vapour, 129, 153 water-soluble, 16, 58, 60, 65, 163, 165, 166, 168 wavelengths, 131 weakness, 111, 211, 214, 223, 242, 331 wear, 22, 72, 107, 109, 269, 270, 272, 300 weight loss, 58, 83, 106, 166, 167, 168, 273, 345 weight ratio, 85 welding, 239, 347 wells, 112 wettability, 139, 143, 145, 146, 147, 149, 150, 155, 156

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