Bioactive Glasses: Materials, Properties and Applications (Woodhead Publishing in Materials)

Bioactive glasses

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Related titles: Biomedical composites (ISBN 978-1-84569-436-4) Biomedical composites provides a thorough review of the current status, recent progress and future trends in composites for biomedical applications. Part I discusses the fundamentals with chapters on natural composites, design and fabrication, and general hard and soft tissue applications. Part II reviews specific applications. Chapters discuss bone repair, coatings for implants, spinal implants, injectable composites and composites for tissue engineered scaffolds. Part III covers the biocompatibility, mechanical behaviour and failure of biocomposites with such topics as cellular response, testing of biocomposites and tribology. Part IV analyses the future for biocomposites with chapters on nanostructured biocomposites, developing biocomposites as scaffolds and biocomposites in tissue engineering and regenerative medicine. Bioceramics and their clinical applications

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Bioactive glasses Materials, properties and applications Edited by Heimo O. Ylänen

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Published by Woodhead Publishing Limited, 80 High Street, Sawston, Cambridge CB22 3HJ, UK www.woodheadpublishing.com Woodhead Publishing, 1518 Walnut Street, Suite 1100, Philadelphia, PA 19102-3406, USA Woodhead Publishing India Private Limited, G-2, Vardaan House, 7/28 Ansari Road, Daryaganj, New Delhi – 110002, India www.woodheadpublishingindia.com First published 2011, Woodhead Publishing Limited © Woodhead Publishing Limited, 2011 The authors have asserted their moral rights. This book contains information obtained from authentic and highly regarded sources. Reprinted material is quoted with permission, and sources are indicated. Reasonable efforts have been made to publish reliable data and information, but the authors and the publishers cannot assume responsibility for the validity of all materials. Neither the authors nor the publishers, nor anyone else associated with this publication, shall be liable for any loss, damage or liability directly or indirectly caused or alleged to be caused by this book. Neither this book nor any part may be reproduced or transmitted in any form or by any means, electronic or mechanical, including photocopying, microfilming and recording, or by any information storage or retrieval system, without permission in writing from Woodhead Publishing Limited. The consent of Woodhead Publishing Limited does not extend to copying for general distribution, for promotion, for creating new works, or for resale. Specific permission must be obtained in writing from Woodhead Publishing Limited for such copying. Trademark notice: Product or corporate names may be trademarks or registered trademarks, and are used only for identification and explanation, without intent to infringe. British Library Cataloguing in Publication Data A catalogue record for this book is available from the British Library. Library of Congress Control Number: 2011932770 ISBN 978-1-84569-768-6 (print) ISBN 978-0-85709-331-8 (online) The publisher’s policy is to use permanent paper from mills that operate a sustainable forestry policy, and which has been manufactured from pulp that is processed using acid-free and elemental chlorine-free practices. Furthermore, the publisher ensures that the text paper and cover board used have met acceptable environmental accreditation standards. Typeset by RefineCatch Limited, Bungay, Suffolk Printed by TJI Digital, Padstow, Cornwall, UK

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Contents

Contributor contact details Introduction Part I

1

ix xiii

Materials and mechanical properties of bioactive glass

1

Melt-derived bioactive glasses

3

L. HUPA, Åbo Akademi University, Finland

1.1 1.2 1.3 1.4 1.5

Introduction Manufacture and physical properties Chemical properties and bioactivity Future trends References

3 6 13 23 23

2

Surface modification of bioactive glasses

29

J. CHANG and Y. L. ZHOU, Chinese Academy of Sciences, China and Y. ZHOU, Shanghai Jiao Tong University, China

2.1 2.2

2.5 2.6

Introduction Surface modification of bioactive glasses to improve bioactivity Surface modification of bioactive glasses using organic molecules to improve dispersivity Surface modification of bioinert materials using bioactive glasses Conclusions and future trends References

39 43 44

3

Cell interaction with bioactive glasses and ceramics

53

2.3 2.4

29 30 35

R. P. K. PENTTINEN, University of Turku, Finland

3.1 3.2

Introduction Biology of bioactive glasses

53 54 v

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Contents

3.3 3.4 3.5 3.6 3.7

Reaction of cells with glasses and related ceramics Effect of silica on bone formation Future trends References Appendix: list of abbreviations

57 66 71 72 84

4

Regulatory aspects of bioactive glass

85

S. LINDGREN, T. PÄNKÄLÄINEN, J. LUCCHESI and F. OLLILA, BonAlive Biomaterials Ltd, Finland

4.1 4.2 4.3 4.4 4.5

Introduction General requirements Indication areas Market approval process in some geographical areas References

Part II Applications of bioactive glass 5

Bioactive glass and glass-ceramic scaffolds for bone tissue engineering

85 86 94 98 103 105

107

X. CHATZISTAVROU, University of Erlangen-Nuremberg, Germany, P. NEWBY, Imperial College London, UK and A. R. BOCCACCINI, University of Erlangen-Nuremberg, Germany and Imperial College London, UK

5.1 5.2 5.3 5.4 5.5 5.6 5.7 5.8

Introduction Requirements for bone tissue scaffolds Bioactive glasses and glass-ceramics in bone tissue engineering Bioactive glass-based scaffolds: fabrication technologies Scaffolds from boron-containing bioactive glass Polymer-coated composite scaffolds Conclusions References

107 108 109 110 114 119 123 124

6

Nanoscaled bioactive glass particles and nanofibres

129

M. EROL, Istanbul Technical University, Turkey and A. R. BOCCACCINI, University of Erlangen-Nuremberg and Imperial College London, UK

6.1 6.2 6.3 6.4 6.5 6.6 6.7

Introduction Characteristics of nanoscale bioactive glasses Fabrication of bioactive glass nanoparticles and nanofibres Applications of nanoscale bioactive glasses Conclusions Acknowledgement References

© Woodhead Publishing Limited, 2011

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Contents

7

Bioactive glass containing composites for bone and musculoskeletal tissue engineering scaffolds

vii

162

S. VERRIER, AO Research Institute Davos, Switzerland, J. E. GOUGH, University of Manchester, UK and A. R. BOCCACCINI, University of Erlangen-Nuremberg, Germany and Imperial College London, UK

7.1 7.2 7.3 7.4 7.5 7.6 7.7

Introduction Composite materials approach to tissue engineering scaffolds In vitro and in vivo evaluation Discussion Conclusions and future trends References Appendix: list of abbreviations

162 165 169 179 180 181 188

8

Use of bioactive glasses as bone substitutes in orthopaedics and traumatology

189

J. HEIKKILÄ, Sports Clinic and Hospital Mehiläinen Turku, Finland

8.1 8.2 8.3 8.4 8.5 8.6 8.7 8.8 8.9 8.10 8.11 8.12 8.13 8.14 8.15 9

Introduction Glass surface reactions The bonding of bioactive glass and bone formation The biocompatibility of bioactive glasses The strength of bioactive glass Bone formation Clinical use for benign bone tumors Bioactive glass and infection Bioactive glass in cancellous bone and metaphyseal fractures Bioactive glass in diaphyseal bone fractures Bioactive glass in spinal surgery Arthroplasty Summary of applications in orthopaedics and traumatology Future trends References Bioactive glass S53P4 as a bone graft substitute in the treatment of osteomyelitis

189 192 194 195 196 197 200 201 201 202 202 203 203 204 204 209

N. C. LINDFORS, Helsinki University Central Hospital, Finland

9.1 9.2 9.3 9.4 9.5

Introduction Bone grafts in the treatment of osteomyelitis Antibacterial properties of bioactive glass S53P4 Vascularization-promoting properties of bioactive glasses Bioactive glass S53P4 in the treatment of osteomyelitis: a multicentre study

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Contents

9.6 9.7

Conclusions References

213 215

10

Bioactive glass for maxillofacial and dental repair

217

M. J. PELTOLA and K. M. J. AITASALO, Turku University Hospital, Finland

10.1 10.2 10.3 10.4 10.5 10.6 11

Introduction Current materials and requirements in maxillofacial reconstruction Properties of bioactive glass Clinical applications of bioactive glass in maxillofacial reconstruction Clinical applications of bioactive glass in dentistry References Bioactive glass and biodegradable polymer composites

217 217 219 220 223 224 227

T. NIEMELÄ and M. KELLOMÄKI, Tampere University of Technology, Finland

11.1 11.2 11.3 11.4 11.5 11.6 11.7 11.8

Introduction Biodegradable polymers Manufacturing of the composites Bioactive glass particle composites Bioactive glass fiber composites Coatings Future trends References

227 228 229 231 235 238 241 241

12

Bioactive glasses for wound healing

246

M. SHAH MOHAMMADI, C. STÄHLI and S. N. NAZHAT, McGill University, Canada

12.1 12.2 12.3 12.4 12.5 12.6

Introduction Silicate-based versus phosphate-based bioactive glasses Antibacterial properties of bioactive glasses Stimulation of angiogenesis Conclusions References

246 246 250 254 259 260

Index

267

© Woodhead Publishing Limited, 2011

Contributor contact details

(* = main contact)

Editor and Introduction H. O. Ylänen Tampere University of Technology Department of Biomedical Engineering Hermiankatu 12A PL 692 33101 Tampere Finland E-mail: [email protected]

Chapter 1 L. Hupa Åbo Akademi University Biskopsgatan 8 20500 Turku Finland E-mail: [email protected]

Chapter 2 J. Chang* and Y. L. Zhou State Key Laboratory of High Performance Ceramics and Superfine Microstructure Shanghai Institute of Ceramics Chinese Academy of Sciences 1295 Dingxi Road Shanghai 200050 China

Y. Zhou Med-X Research Institute Shanghai Jiao Tong University 1954 Hua Shan Road Shanghai 200030 China

Chapter 3 R. P. K. Penttinen Department of Medical Biochemistry and Genetics University of Turku Kiinamyllynkatu 10 20520 Turku Finland E-mail: [email protected]

Chapter 4 S. Lindgren*, T. Pänkäläinen, J. Lucchesi and F. Ollila BonAlive Biomaterials Ltd Biolinja 12 20750 Turku Finland E-mail: [email protected]

E-mail: [email protected]

ix © Woodhead Publishing Limited, 2011

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Contributor contact details

Chapter 5

Chapter 7

X. Chatzistavrou and A. R. Boccaccini* Institute of Biomaterials Department of Materials Science and Engineering University of Erlangen-Nuremberg 91058 Erlangen Germany

S. Verrier Musculoskeletal Regeneration AO Research Institute Davos 7270 Davos Switzerland

E-mail: aldo.boccaccini@ww. uni-erlangen.de

P. Newby Department of Materials Imperial College London Prince Consort Road London SW7 2BP UK

Chapter 6 M. Erol Department of Chemical Engineering Istanbul Technical University Maslak 34469 Istanbul Turkey E-mail: [email protected]

A. R. Boccaccini* Institute of Biomaterials Department of Materials Science and Engineering University of Erlangen-Nuremberg 91058 Erlangen Germany

J. E. Gough School of Materials University of Manchester Manchester M1 7HS UK A. R. Boccaccini* Institute of Biomaterials Department of Materials Science and Engineering University of Erlangen-Nuremberg 91058 Erlangen Germany E-mail: aldo.boccaccini@ww. uni-erlangen.de

Chapter 8 J. Heikkilä Assistant Professor in Orthopaedics and Traumatology Sports Clinic and Hospital Mehiläinen Turku Kauppiaskatu 8 20100 Turku Finland E-mail: [email protected] [email protected]

E-mail: aldo.boccaccini@ww. uni-erlangen.de

© Woodhead Publishing Limited, 2011

Contributor contact details

Chapter 9

Chapter 11

N. C. Lindfors Department of Orthopaedic and Hand Surgery Helsinki University Central Hospital Töölö Hospital Topeliuksenkatu 5 00260 Helsinki Finland

T. Niemelä and M. Kellomäki* Department of Biomedical Engineering Tampere University of Technology PO Box 692 33101 Tampere Finland

E-mail: [email protected]

Chapter 12

xi

E-mail: [email protected]

M. J. Peltola* and K. M. J. Aitasalo Department of Otorhinolaryngology – Head and Neck Surgery Turku University Hospital 20521 Turku Finland

M. Shah Mohammadi, C. Stähli and S. N. Nazhat* Department of Mining and Materials Engineering McGill University 3610 University Street Montreal Quebec H3A 2B2 Canada

E-mail: [email protected]

E-mail: [email protected]

Chapter 10

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Introduction

Today, millions of prostheses, implants and bone grafts are needed to maintain the quality of life of the aging population. The increased human lifespan alone has created enough problems in this regard, even if we discount the need for repairing or replacing body parts damaged by an individual’s own actions. As a result, material scientists have for decades faced the challenge of developing advanced biomaterials to repair the defects of the human body or to replace damaged parts. A significant advance in the search for better biomaterials was provided by the introduction of two new synthetic biomaterials during the early 1970s. The materials were developed independently and almost simultaneously by several groups of materials scientists. The new synthetic biomaterials were able to bond to host tissue through chemical processes occurring on the materials’ surface, and the materials were termed bioactive ceramics. A comprehensive review of the state of the art in bioceramics from basic science to clinical applications is presented in The Handbook of Bioceramics and Their Applications, edited by Professor Tadashi Kokubo (Woodhead Publishing Limited, 2008). The current book focuses on a special subgroup of bioactive ceramics, namely bioactive glasses. Systematic research into bioactive glasses was started by Professor Larry Hench in 1969, when he introduced the concept of a strong bonding between bone and synthetic material brought about by chemical reactions occurring on a glass surface. The innovation concerned the chemical reactivity of the surface of a silica-based material that had the amorphous structure of silicate glass. Hench introduced the material, a bioactive glass, in the early 1970s. Bioactive glasses bond firmly to bone through chemical reactions and can ultimately be replaced by bone: these properties make them extremely promising as a material for medical applications. Most importantly, the constituents in bioactive glass are physiological chemicals found in the body, typically silicon, sodium, potassium, magnesium, oxygen, calcium and phosphorus. According to several studies, during the bonding and formation of bone, the concentration of the chemicals never rises to levels that could disturb the adjacent tissues. The use of bioactive glass as an implant material or in manufacturing medical devices is limited, however, by the mechanical properties of glass. Glass is brittle xiii © Woodhead Publishing Limited, 2011

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Introduction

and cannot therefore be used in positions where load-bearing properties are required. Glass can be cast to plates, rods or simple devices; alternatively it can be formed by sawing or grinding cast rods to rigid medical devices. It can also be used as a filler material in the form of particulate. In order to widen the manufacturing of bioactive glass products to technically more demanding clinical applications, new bioactive glasses were developed. Today, new bioactive glasses can be tailor-made to a variety of clinical applications, to different shapes, fibers, microspheres and to show optimal bioactivity in different physical conditions found in the human body. Novel applications of bioactive glasses require not only tight control of the bioactivity but also a thorough knowledge of the influence of the composition of glass on its manufacture and its formation into different products. The tailoring of glasses to certain applications is thus based on understanding and mastering a wide range of properties important for both medical applications and for the manufacture of glass. The manufacturing of conventional melt-derived bioactive glasses demands extremely high temperatures and careful annealing procedures. To avoid this, new technologies were developed for manufacturing new types of bioactive glasses such as those derived from sol-gel, spun bioactive glass fibers and bioactive glass nanoparticles. Today, the use of bioactive glass as a component of biomaterial composites is one of several interesting options in the development of a variety of clinical applications of this material. In the current book, the development of different bioactive glasses is reviewed by globally distinguished experts and scientists. The book starts with an introduction to different types of bioactive glasses, the influence of surface modification on the properties of these special glasses, and the process of bonding to different types of host tissues. It also discusses cell interactions at the interface of bioactive glasses and tissues. When marketing any products containing bioactive glass, regulations imposed by authorities are of crucial importance. A chapter addressing this issue is written by a company specializing in products made of bioactive glass. Having been developed over several decades, bioactive glass products are today in clinical use across the world. This means that it is now possible to discuss experiences of the clinical applications of these products. The current book also covers this interesting area of bioactive glasses. Earlier works have been published that have covered the area of bioceramics and even their clinical applications. However, to my knowledge, the current book is the first one to discuss solely bioactive glasses and their application. The book should become a standard textbook in both the fields of materials sciences and medical sciences. I hope that by publishing this book we can encourage an interest in the development of bioactive glasses in students and researchers the world over. Heimo O. Ylänen

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1 Melt-derived bioactive glasses L. HUPA, Åbo Akademi University, Finland

Abstract: This chapter discusses the properties of melt-derived bioactive glasses from the material technology point of view. The non-crystalline structure of bioactive glasses offers the possibility of adjusting their physical and chemical properties by altering their oxide composition within certain limits. Thus, understanding the relationships between the oxide composition and the relevant properties is essential when the glass composition is tailored for novel clinical applications. The focus is to summarize some published data on the in vitro and in vivo bioactivity. The restrictions put on glass composition by the manufacturing process are also discussed. The aim is to provide some fundamental tools for further studies and the development of melt-derived glasses to desired product forms for various clinical applications. Key words: melt-derived bioactive glasses, viscosity, crystallization, in vitro bioactivity, in vivo bioactivity.

1.1

Introduction

A new era in the development of materials for use in medicine began in the 1970s, when Professor Larry Hench discovered glasses capable of forming interfacial bonding with bone (Hench and Paschall, 1973). The compositions showing this special property were called bioactive glasses. In developing the glasses Professor Hench’s leading idea was to find a material that, rather than forming an interfacial layer of scar tissue, would instead form a living bond with the host tissues (Hench 2006). The hypothesis behind the glass development was simple but ingenious: as bone contains hydroxyapatite, HA, the implant material should be able to form an HA layer on the surface in biological solutions. Such a material would not be rejected by the body but be bonded directly with the tissue. Hench and co-workers tested whether phosphate containing soda-lime silicate glasses could fulfil the criterion of tissue bonding. Glasses within this system would then contain two important components of the hydroxyapatite (Ca5(PO4)3OH), namely Ca2+ and PO43− ions. The two other cations, Na2+ and Si4+, are also common components of the human body. Implants of Glass 45S5, one of the first compositions tested, were found to bond to rat femur. This glass, known also as Bioglass®, is still one of the most bioactive glasses known. The selection of the composition was ideal; the low silica content makes the glass easy to melt but also gives it much lower chemical durability than commercial soda-lime glasses in aqueous solutions. A low chemical durability and the ability of the composition to form a dual layer of silica and amorphous calcium phosphates on the surface are key features of 3 © Woodhead Publishing Limited, 2011

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Bioactive glasses

bioactive glasses (Hench, 1991, 1992). The subsequent reactions in the surface layers lead to bonding between the glass and surrounding tissue. Ideally, the glass will dissolve and be replaced by new tissue in time. The kinds of materials that were capable of bonding directly to living tissue were called bioactive materials. Bioactive materials are defined as (1) materials that have been designed to induce special biological activity; or (2) biomaterials that are designed to elicit or modulate biological activity (Williams 1999). Bioactive glasses are capable of forming a bond with both hard and soft tissue in vivo or in vitro environments by developing a surface layer of hydroxycarbonate apatite by release of ionic species from the bulk material (Williams 1999). Bioactive glass research deals in large part with developing a fundamental understanding of the dissolution and surface reactions of the glass and the tissue response to the dissolving material. Over the years, the interactions of several glass compositions with biological solutions have been studied. However, the research has been concentrated mainly on compositions close to that of the bioactive glass 45S5. Li et al. (1991) reported that sol-gel-derived glasses within the system Na2O-CaO-SiO2 show bioactivity within a much larger composition range than melt-derived glasses. Since then, sol-gel derived glasses have been studied intensively. Jones (2008, 2009) has summarized the use of sol-gel glasses as materials for nanostructured bioactive scaffolds. This chapter deals with the properties of melt-derived bioactive glasses. The standpoint is more in glass science and technology than in biological sciences. The goal is to summarize various criteria to be considered when developing glass compositions for new products to various clinical applications. Controlled bioactivity is the basis for development of glass compositions. However, the tissue-engineering approach to manufacture porous thin-walled scaffold structures of glasses or to use glasses as constituents in composites calls for a better understanding of the overall properties of the glasses. Generally, glasses have good chemical durability. However, one of the most important properties of bioactive glasses is their controlled reactivity in body solutions. The reactivity of glasses in aqueous solutions is strongly dependent on the glass composition and thus one of the key factors for choice of composition. Chemical durability as well as several other glass properties can be adjusted smoothly by the composition within a certain range. Further, when using bioactive glasses in clinical applications, other properties than chemical durability should be considered. Mechanical strength and especially the ability to sustain a certain mechanical impact and loading are important properties during surgery. Basically, the strength of glasses is high, but due to their brittle nature they cannot be used in load-bearing applications. As mechanical properties are basically more dependent on the surface condition of the glass than on the glass composition, they are not essential for the composition choice. Today the medical uses of bioactive glasses are based mainly on crushed fractions. The intense research on bioactive and biodegradable glasses as

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Melt-derived bioactive glasses

5

components of composites is likely to give birth to novel applications. For example, when using glass fibres in composites together with organic polymers the inherent mechanical strength of glasses is utilized in order to reinforce the composite structure. The mutual reactions between bioactive glasses and various biodegradable polymers are not well established. When more information is available the reactivity of glasses with other materials used in medicine should also be taken into account when choosing the composition. The specific criteria of reactivity in body solutions limit the composition range for the glasses. Within this range, the glass composition should be chosen so that it can be melted and formed into specific shapes with available methods. One of the most important characteristics for glass manufacture is the viscosity– temperature relationship, as it defines the methods that can be utilized in glass forming. Also the liquidus temperature should be considered; all practical melt forming operations have to be carried out at higher temperatures than the liquidus. Both viscosity and liquidus depend on glass composition. For traditional glasses the crystallization characteristics are not usually critical. However, the specific composition range of glasses showing bioactivity brings a risk of rapid crystal growth during glass forming. Another example is the manufacture of specific porous structures via sintering of glass particulates. In the manufacture of glassy structures, crystallization during sintering is an undesired phenomenon. On the other hand, controlled crystallization can be utilized for achieving bioactive glassceramics with specific properties. In both cases, the processing parameters are mastered via a good knowledge of the crystallization characteristics. Although crystallization in thermal treating of glasses is partly a kinetic phenomenon, it can be controlled by the choice of the glass composition. Thermal expansion of glasses should be considered when the glass is applied as a bioactive coating, for example on a metal prostheses. The glass composition should be adjusted to give a compatible adhesion with the metal for achieving a good adherence without any chipping or crawling of the coating on the metal. The choice of the glass composition for a specific application should be based on a firm knowledge on the influence of all major components on the most relevant properties of the glass with regard to both the final use and the manufacture of the product. Despite extensive research during the past 40 years, only a few glass compositions have been accepted for clinical use. The two US Food and Drug Administration FDA approved melt-derived compositions 45S5 (Hench and Paschall, 1973) and S53P4 (Andersson et al., 1990) consist of four oxides, SiO2, Na2O, CaO and P2O5. In general, a great number of elements can be dissolved in glasses. The effect of Al2O3, B2O3, Fe2O3, MgO, SrO, BaO, ZnO, Li2O, K2O, CaF2 and TiO2 on the in vitro or in vivo properties of certain compositions of bioactive glasses has been reported (Andersson et al., 1990; Vrouwenvelder et al., 1994; Brink et al., 1997; Haimi et al., 2009; Lusvardi et al., 2009; Zhang et al., 2009; Gentleman et al., 2010; Watts et al., 2010). However, the effect of the composition on the properties of bioactive and biodegradable glasses is not fully understood.

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Bioactive glasses

1.2

Manufacture and physical properties

1.2.1 Melting and forming Melt-derived bioactive glasses are melted and formed with methods similar to traditional soda-lime glasses. However, the requirements at the processing must meet the standards for materials used in medical applications. The batches are mixed of high purity analytical- and reagent-grade raw chemicals and thus the content of trace impurities in the glasses is low. Bioactive glasses are produced by melting batch components at an elevated temperature, typically 1350 to 1450°C, in electrically heated furnaces. The glasses are melted in platinum crucibles to avoid any contamination from oxide crucibles. Usually, no fining agents are added to the batches. The low viscosity of typical bioactive glass compositions at the melting temperature aids in eliminating gaseous inclusions from the melt. Melting times of small batches for laboratory testing varying from 1 to 24 hours have been employed. The glasses are often melted twice in order to increase homogeneity. Volatilization of components with high vapor pressures at high temperatures should also be taken into account. In bioactive glasses alkalis, boron, phosphorus and fluorides may vaporize. The glasses can be melted in covered crucibles to minimize losses. The vaporization in a certain process can also been taken into account by adjusting the batch composition. Forming and shaping procedures vary depending on the product type; casting into monoliths and drawing into rods or fibres are the main forming processes for bioactive glasses. After forming, the glass is annealed at a temperature corresponding to the viscosity 1013 dPa·s (1013 Poise), to remove residual stresses caused by cooling after forming. Granulates and powdered glass are produced by crushing and sieving the annealed plates into desired particle fractions. Also quenching the melt between stainless steel plates or pouring the melt into deionized water are further steps in the process of granulate fabrication. However, bioactive glasses start to react easily in aqueous solutions, which might affect the composition of the particle surfaces. Crushing and sieving increase the risk of contamination from the equipment used in the particle manufacture. Thus, in all processing of bioactive glasses into specific shapes, care should be taken in order to minimize any contamination.

1.2.2 Viscosity Viscosity is important in determining the melting parameters for achieving a bubble-free and homogeneous melt. Glasses are usually melted at temperatures corresponding to the viscosity value 10 to 100 dPa·s. The low viscosity facilitates easy elimination of the gases by buoyancy from the melt. Viscosity and its change with temperature is the most crucial factor in determining the forming and shaping procedures that can be used for a particular composition. The approximate viscosity values of interest in forming bioactive glasses into various shapes are summarized in Table 1.1.

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Melt-derived bioactive glasses

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Table 1.1 Approximate viscosity values (dPa·s) for bioactive glass forming processes Processing

Viscosity (η) in dPa·s

Melting Pressing Drawing of continuous fibres Sinter glass powder to porous body Annealing

10–102 104–106 102.5–103.5 108–109 1012–1013

1.1 Viscosity–temperature points for 45S5, S53P4 and 13-93 (Vedel et al., 2008).

Figure 1.1 shows measured viscosity–temperature points for three bioactive glasses, 45S5, S53P4 and 13–93 at the low and high temperature ranges. The measured values are according to Vedel et al. (2008). The oxide composition of the glasses are given in Table 1.2. The dashed lines between the low and high temperature ranges give typical viscosity–temperature curves for glass forming melts. However, at the intermediate temperatures bioactive glasses crystallize and melt viscosity does not exist. The viscosity values at different temperatures are important criteria in glass forming. The high temperature values correlate with melt-forming processes, while the low-temperature values specify the suitability of the glass, for example for sintering into porous bodies or firing as a coating on metal implants. The high temperature values of glasses 45S5 and S53P4 in Fig. 1.1 could be measured by rotational viscometer only at values below 100 dPa·s (Vedel et al., 2008). Thus, it is likely that for these glasses the liquidus temperature, i.e. the temperature at

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Table 1.2 Nominal oxide compositions 45S5, S53P4 and 13–93 in wt% (mol%) Oxides in wt% (mol%) Glass

Na2O

45S5 S53P4 13–93

24.5 (24.4) 23 (22.7) 6 (6)

K2O

12 (7.9)

MgO

CaO

P2O5

SiO2

5 (7.7)

24.5 (26.9) 20 (21.8) 20 (22.1)

6 (2.6) 4 (1.7) 4 (1.7)

45 (46.1) 53 (53.8) 53 (54.6)

Sources: Hench et al., 1973 (45S5), Andersson et al., 1990 (S53P4) and Brink et al.,1997 (13–93).

which crystallization commences on cooling, is close to the lowest experimental high temperature values shown in Fig. 1.1. Further, the low viscosity values at their liquidus suggest that these glasses can only be formed by casting. At the low temperature range, 45S5 and S53P4 crystallize at around 109 dPa·s. This means that the compositions cannot be sintered through viscous flow into porous bodies without extensive crystallization. The strong tendency to crystallize within a large temperature range has been utilized in the manufacture of various glass-ceramics of the parent glass 45S5 (Chen and Boccaccini, 2006; Chen et al., 2006b; Boccaccini et al., 2007). The high-temperature viscosity values of glass 13–93 in Fig. 1.1 suggest that it can be formed without extensive crystallization to around 104 dPa·s. This composition has be pressed, blown and drawn into continuous fibres (Brink, 1997; Pirhonen et al., 2006). At the low temperature range crystallization starts below 108 dPa·s, which indicates that the glass can be sintered into porous glassy bodies (Ylänen et al., 2000; Fu et al., 2010). The bioactive glasses 45S5 and S53P4 that crystallize easily on thermal treatments contain only four oxides. In an attempt to decrease the crystallization tendency potassium oxide, magnesia and boron oxide have been added to the formulations. The viscosity of glasses within the system Na2O-K2O-MgO-CaOB2O5-P2O5 has been discussed in three extensive studies (Brink, 1997: Karlsson and Rönnlöf, 1998; Vedel et al., 2008). The studies had two goals: to get reliable data on the high-temperature properties of bioactive glasses and to meet the need to find compositions that can be formed by other methods than casting. Both Karlsson/Rönnlöf and Vedel et al. suggest for the viscosity–temperature relationship an Arrhenius-type expression ˚I nη = A + B/T, where the constants A and B are expressed by the oxide composition of the melt. Vedel et al. also give models for calculating the constants A, B and T0 in the Vogel–Fulcher–Tamman equation for the viscosity–temperature relationship, logη = –A + B/(T – T0). However, as no experimental points exist for the intermediate temperature range, a single relationship over the whole temperature range is questionable. Models for calculating temperatures for certain high- and low-temperature viscosity points

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additively from the oxide composition are likely to be of reasonable accuracy (Vedel et al., 2008). Temperatures T at certain viscosities η were expressed as additive functions of the oxide content of the glass in wt%. The temperature values were based on certain experimental points when measuring the viscosity by hot-stage microscopy and rotational viscometer. The composition dependency was calculated for 30 statistically chosen glass compositions within the system Na2O-K2O-MgO-CaO-B2O3-P2O5-SiO2. The influence of composition on the temperature at a typical viscosity for sintering of porous bodies is given in equation 1.1. Equation 1.2 can be used to estimate the temperature at a viscosity value typical for glass melting. The rather low viscosity value was explained by the strong tendency of several glasses to crystallize at a relatively low viscosity (Vedel et al., 2008). Tη =1010 dPas (°C) = −148.036 + 3.566·xNa O + 7.071·xMgO 2

+ 9.740·xCaO + 7.770·xB O + 8.347·xP O 2 3

2 5

+ 9.287·xSiO

[1.1]

2

Tη =1015 dPas (°C) = 117.816 + 7.730·xK O + 6.078·xMgO 2

+ 18.469·xP O + 19.150·xSiO 2 5

2

[1.2]

The equations are valid for the composition range (wt%): Na2O (5 to 25), K2O (0 to 15), MgO (0 to 6), CaO (15 to 25), B2O3 (0 to 4), P2O5 (0 to 4), SiO2 (50 to 65).

1.2.3 Thermal expansion Bioactive glasses have been studied as coatings on metal prostheses in order to provide biological fixation to bone (Lacefield and Hench, 1986; Hench and Andersson, 1993b; Andersson et al., 1995; Bloyer et al., 1999; Moritz et al., 2004a, 2004b; Krause et al., 2006; Borrajo et al., 2007; Lopez-Esteban et al., 2009). The rapid surface reactions of bioactive glasses can limit their use as thin coatings. In such applications, glass compositions with a lower reactivity are preferred. If the coating is applied via traditional enamelling firing, the thermal expansion of the glass should be compatible with the metal. The coefficient of thermal expansion can be calculated additively from the composition with the factors suggested by Appen (1974). A compositional model for the linear thermal expansion α between 20 and 300°C for bioactive glasses was suggested by Karlsson and Rönnlöf (1998): see equation 1.3.

α (10−6 K−1) = 3.625 + 0.345·xNa O + 0.266·xK O 2

2

+ 0.098·xCaO + 0.064·xP O

2 5

[1.3]

The equation is valid for compositions with (wt%) 5 to 25 Na2O, 0 to 15 K2O, 0 to 5 MgO, 10 to 20 CaO, 0 to 3 B2O3, 0 to 6 P2O5 and 39 to 70 SiO2. © Woodhead Publishing Limited, 2011

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Bioactive glasses

1.2.4 Crystallization characteristics The importance of the liquidus temperature for glass melts arises from the fact that all melt-forming should be performed at temperatures higher than the liquidus. Surprisingly little information is available on accurate liquidus temperatures of bioactive glasses. According to Hench, the choice of the first compositions of bioactive glasses was partially based on finding a composition that is close to a ternary eutecticum in the system Na2O-CaO-SiO2 (Hench, 2006). This ternary system is commonly used to interpret liquidus temperatures and crystallization paths for soda-lime-silica glasses, it is also used when they contain some additional components to the three oxides. Figure 1.2 shows the phase equilibria and some liquidus surfaces in the silica-rich corner of the Na2O-CaO-SiO2 system (Morey and Bowen, 1964). Compositions of bioactive glasses 45S5, S53P4 and 13–93 are superimposed in the ternary system by calculating all alkalis, alkaline earths and glass formers in wt% into Na2O, CaO and SiO2, respectively. Figure 1.2 shows that a simple superimposing of the four component glasses 45S5 and S53P4 as well as glass 13–93 containing also the oxides of potassium and magnesium (c.f. Table 1.2) would suggest liquidus temperatures around or higher than 1200°C for all compositions. According to thermal analysis the crystals formed in heating glass 45S5 above transition temperature start to melt between 1080 and 1160°C (Chatzistavrou et al., 2006; Hall, 2007; Arstila et al., 2008a; Bretcanu et al., 2009) and the offset of the melting is between 1250 and 1260°C (Chatzistavrou et al., 2006; Arstila et al., 2008a). The offset temperatures for melting of crystals in S53P4 and 13–93 are 1210 to 1230°C and 1020 to 1180°C, respectively (Arstila et al., 2008a). The offset temperatures for 45S5 and S53P4 are close to the values suggested by the liquidus surface in Fig. 1.2. Until more accurate data is available for systems containing some phosphorus oxide, the ternary Na2O-CaO-SiO2 system seems to give an acceptable first approximation of liquidus. The larger differences in the measured liquidus and the value approximated from the simple three-component phase diagram for glass 13–93 containing also potassium oxide and magnesia indicate that other approaches to estimate liquidus are needed for glasses consisting of several oxides. One way could be composing empirical models for calculating the liquidus from the oxide composition of the glass. In that case, separate models for different primary phase fields should be considered as suggested by Karlsson et al. (2002). Thermodynamic modelling is likely to provide the most appropriate method for estimating the liquidus in future. So far, reliable thermodynamic data of the liquid phase in the compositional area of bioactive glasses are not available. In the ternary Na2O-CaO-SiO2 (Fig. 1.2) the compositions of 45S5 and S53P4 fall into the primary phase field of Na2Ca2Si3O9, while 13–93 is within the CaSiO3 field. Interestingly, these phases seem to be the most commonly observed primary phases formed in the thermal treatment of bioactive glasses within the compositional range given for equations 1.1 to 1.2 (Arstila et al., 2008b). The

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1.2 Liquidus surfaces and phase equilibria in the system Na2O-CaOSiO2. Composition points of glasses 45S5, S53P4 and 13–93 are superimposed into the diagram by assuming alkalis, alkaline earths and P2O5 by wt% in Na2O, CaO and SiO2, respectively (redrawn from Morey and Bowen, 1964).

sodium oxide rich compositions form sodium calcium silicates, while glasses with less sodium oxide form mainly CaSiO3 (Arstila et al., 2008b). Na2Ca2Si3O9, Na2CaSi3O8 and Na2CaSi2O6 are the most often suggested compositions of the primary phase observed in the thermal treatment of 45S5 (Rizkalla et al., 1996; Chen et al., 2006; Lefebvre et al., 2007, 2008; Arstila et al., 2008b; Bretcanu et al. 2009). The discrepancy observed in the primary phase can partially be explained by the solid solutions of Na2Ca2Si3O9 and Na2CaSi3O8. During crystallization of a stoichiometric glass Na2Ca2Si3O9 the nucleation did not start with the stoichiometric composition but approached to that during crystallization (Fokin and Zanotto, 2007). The final composition of the fully crystallized glass was close to that of the parent glass. The primary phase, PCT, as well as the crystallization temperature, Tx, in the thermal treatment of bioactive glasses was predicted by additive functions from the oxide composition, equations 1.4 and 1.5 (Arstila et al., 2008b). Numerical values of PCT lower than 1.5 indicate sodium calcium silicate crystals, while values higher than 1.5 suggest the formation of CaSiO3.

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Bioactive glasses PCT = 0.018−0.107·xNa O + 0.002·x2Na O + 0.052·xMgO 2

+

0.032·x2

P2O5

2

+ 0.038·xSiO

[1.4]

2

Tx (°C) = 191.70 − 9.51·xNa O + 5.91·xMgO + 5.93·xCaO 2

+ 15.29·xB O + 12.92·xP O + 9.77·xSiO 2 3

2 5

2

[1.5]

The exact crystallization temperature depends on the heating rate. However, the crystallization temperature suggested by equation 1.5 can be used to compare the crystallization characteristics of different compositions and to give approximate values for the onset of crystallization in the thermal treatment of glasses. Generally, glasses forming CaSiO3 crystallize at temperatures 100 to 200°C higher than the glasses showing sodium calcium silicate crystals (Arstila et al., 2008b). On cooling the properties of glass forming melts turn into solid like properties at the glass transition temperature, Tg. As the exact transition temperature for a certain composition depends on the cooling rate, changes in several properties can be observed over a temperature range known as the glass transformation region. The transition temperature thus determines the lower limit for thermal processing of glass melts. The temperature span for sintering through viscous flow ranges from glass transformation up to crystallization temperature. The stability of glasses against crystallization upon thermal treatment is often described with relative values calculated from the temperatures of glass transition, liquidus and crystallization (Hruby, 1972; Zanotto, 1987). The reduced glass transition temperature Tgr, given by the ratio of the glass transition Tg to liquidus Tl, correlates with the general trends of nucleation in glass forming melts. A good glass former has Tgr = Tg/Tl higher than or equal to two-thirds, while lower values suggest crystallization. The crystallization kinetics depend also on the numerical value of Tgr; higher values than 0.58 indicate surface crystallization, while for lower values volume nucleation dominates (Zanotto, 1987). The compositional dependence of glass transformation temperature Tg of bioactive glasses has been described as an additive function of the oxide composition (Andersson, 1992; Karlsson and Rönnlöf, 1998; Arstila et al., 2008b). Equation 1.6 gives the glass transformation temperature based on dilatometric data (Karlsson and Rönnlöf, 1998), while the factors in equation 1.7 are based on thermal analysis (Arstila et al., 2008b). Tg (°C) = 635.9 − 5.26·xNa O + 3.18·xK O − 3.16·xMgO 2

2

− 0.47·xNa O·xK O − 0.12·xK2O·xCaO 2

[1.6]

2

Tg (°C) = −122.85 + 2.59·xK O + 5.36·xMgO + 9.37·xCaO 2

+ 4.08·xB O + 6.45·xP O + 8.04·xSiO 2 3

2 5

2

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Table 1.3 Calculated and measured values of some physical properties for glasses 45S5, S53P4 and 13–93 Property

45S5

S53P4

13–93

Reference

T (°C) at η = 1010 dPa·s T (°C) at η = 101.5 dPa·s Tliquidus(°C) PTC Tx (°C) Tg (°C)

559 (570) – (–) – (1210) 1.5 (NCS) 621 (647) 506 (–)

572 (587) 1233 (1235) – (1188) 1.1 (NCS) 661 (693) 514 (–)

651 (655) 1330 (>1330) – (–) 2.2 (CS) 852 (853) 564 (–)

Tg (°C) α (10−6K−1)

507 (530) 14.9 (–)

516 (541) 13.8 (–)

574 (600) 11.1 (–)

α (10−6K−1)

15.2 (–)

14.1 (–)

11.7 (–)

Vedel et al., 2008 Vedel et al., 2008 Arstila et al., 2008a Arstila et al., 2008b Arstila et al., 2008b Karlsson and Rönnlöf, 1998 Arstila et al., 2008b Karlsson and Rönnlöf, 1998 Appen, 1974

Notes: The measured values are given in parenthesis. – means not available using the model/not measured.

In both models the oxide composition is given in wt%. The validity range of equation 1.6 is the same as for equation 1.3. The validity of equation 1.7 is according to equation 1.1. Some measured and calculated physical property values of bioactive glasses 45S5, S53P4 and 13–93 are summarized in Table 1.3. The differences between the calculated and measured values indicate that the equations can be used to estimate the physical property values of the bioactive glasses.

1.3

Chemical properties and bioactivity

1.3.1 Surface reaction mechanisms Controlled surface reactivity leading to tissue bonding is the key characteristic of bioactive glasses. Compared to soda-lime-silica glasses they have poor chemical durability in aqueous solutions. The low silica and high sodium oxide content contribute to the low chemical durability of bioactive glasses. The reactions of glasses in aqueous solutions are generally described by two main mechanisms: the exchange of alkali ions in the surface with H+ and H3O+, and network dissolution through the attack of hydroxyl ions on the silica structure. Therefore, the reactions are controlled by the pH of the surrounding and interfacial solutions. The chemical durability is commonly measured with methods developed for characterizing soda-lime-silica glasses, e.g. durability in water (ISO 719) and alkaline solutions (ISO 695). Water resistance according to ISO 719 is measured as the volume of hydrochloric acid needed to neutralize the pH increase of pure water after 1 hour’s contact time with a certain fraction of glass particles at 98°C.

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Bioactive glasses Table 1.4 Consumption of HCl (ml, 0.01 M), and hydrolytic resistance class (HGB) according to ISO 719. pH values of solutions were measured before the acid titration at room temperature Sample

VHCl (ml)

HGB

pH

E-glass* Float glass 1–98** 13–93 S53P4 45S5

0.3 ± 0.01 1.0 ± 0.03 2.3 ± 0.1 2.8 ± 0.1 3.4 ± 0.1 5.3 ± 0.2

3 4 5 5 5 >5

9.18 ± 0.02 9.88 ± 0.10 10.24 ± 0.01 10.43 ± 0.05 10.71 ± 0.03 10.79 ± 0.01

Notes: * E-glass (wt%): 0.1 Na2O, 0.7 K2O, 0.7 MgO, 23.5 CaO, 6.4 B2O3, 14.1 Al2O3, 53.9 SiO2. **1–98 (wt%): 6 Na2O, 11 K2O, 5 MgO, 22 CaO, 1 B2O3, 2 P2O5, 53 SiO2 (Itälä et al., 2002). Source: Taipale et al., 2008; Fagerlund et al., 2010.

The acid consumed is related with the amount of Na+ ions leached from the glass expressed as mass of Na2O. The Na2O amount is compared with a relative scale from 1 for glasses with very good water durability to 5 for glasses with poor durability. Table 1.4 gives the hydrolytic resistance, HGB, of e-glass and bioactive glasses 45S5, S53P4 and 13–93 and 1–98. E-glass fibres have been used to develop porous fibre-reinforced composites for use as load-bearing orthopedic implants (Mattila et al., 2009). The bioactivity of glass 1–98 has been verified by in vivo experiments (Itälä et al., 2002). E-glass has medium resistance, while the bioactive glasses have very low resistance. The acid consumption of 45S5 is higher than the limit given for HGB = 5 in the ISO 719 standard. Table 1.4 also shows the pH of the solutions before the titration. The pH values of the bioactive glasses are very high, thus indicating network dissolution at particle surfaces. The large differences in the hydrochloric acid consumption between the bioactive glasses suggest that a procedure according to ISO 719 gives a rapid method of comparing their overall biodegradability. However, the relative 1 to 5 HGB values fail to give any relevant information on the bioactive glasses. Understanding the surface reactions of glasses is of utmost importance when selecting compositions for different clinical applications. Silicate glass surfaces have been divided into five different classes according to their reactivity in different environments (Fig. 1.3; Hench and Clark, 1978; Hench, 1992). The type I surface, typical for high silica surfaces in neutral solutions develops only a very thin hydrous layer. Type II surfaces are typical for commercial soda-lime glasses. The alkalis in the surface layer leach in solutions with pH <9 and a protective surface layer forms. In body fluids these glasses form a fibrous capsule upon implantation. Type III surfaces forming multiple layers are divided into two

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1.3 Different characteristic surface reaction types of silicate glasses in aqueous solutions (redrawn from Hench, 1992).

subclasses. Type IIIA surfaces form dual protective films on glasses, while type IIIB surfaces form when metal ions in the glass or solutions precipitate to multiple films on the surface. Only type IIIA surfaces forming a dual film of silica and calcium phosphate are bioactive and bond to tissue. Glasses with poor chemical resistance have type IV surfaces forming a thick but non-protective silica-rich layer, as alkalis are rapidly leached from the surface. The porous and depolymerised silica-rich layer does not protect the surface from further reaction

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Bioactive glasses

and the glass dissolves with time. In alkaline solutions, pH >9 to 10, the network structure of silicate glasses is attacked and the glass surface undergoes a congruent dissolution as described by the type V surface. Bioactive glass surfaces might show type V reactions if surrounding solution pH increases to high values due to insufficient solution circulation. Thus, the surface area to volume (SA/V) affects the reaction type.

1.3.2 Bioactive glasses in vitro The tissue bonding of bioactive silicate glasses can be related with their ability to form a dual layer of silica gel and hydroxyapatite on the surface in body fluids. The time-dependent reaction stages of interfacial reactions leading to tissue bonding of bioactive glasses have been discussed in detail by Hench and co-workers (Hench, 1991; Hench and Andersson, 1993a; Hench and Best, 2004). The reactions stages in biological solutions before the interaction of the surface with proteins and cells are described in five stages at the interface between the glass and solution in Table 1.5. The special characteristics of the hydroxyapatite layer formed on the glass in the reaction stages 1 to 5 allow biochemical adsorption of growth factors and other biological moieties resulting in rapid formation of new bone as discussed by Hench (1998). Depending on their ability to lead to both osteoconduction and osteoproduction or only osteoconduction, the bioactive materials are further divided into Class A and Class B (Hench, 1998, 2006). In this chapter, however, only factors affecting the formation of the calcium phosphate rich layer and crystallized hydroxyapatite on the glass surface are discussed. These inorganic Table 1.5 Reaction stages of bioactive glass surfaces in biological solutions Stage

Reaction

1

Exchange of alkalis from the glass surface with H+ or H3O+ in the solution. The rapid reaction is diffusion controlled and proportional to the square root of the time (Douglas and El-Shamy, 1967; Hench and Clark, 1978). Breaking of siloxane bonds in the glass interface leading to loss of soluble silica in the form of Si(OH)4 to the solution. The loss of silica is directly proportional to time. Condensation and repolymerization of a SiO2 rich layer on the surface. In this reaction stage the thickness of the silica rich layer increases. Migration of Ca2+ and PO42− ions to the surface through the SiO2 rich layer to an amorphous calcium phosphate rich layer on the top of the SiO2 layer. The layer grows by incorporation of Ca2+ and PO42− ions from the solution. Crystallization of the amorphous calcium phosphate layer by incorporation of OH−, CO32−, or F− anions from solution to form a mixed hydroxyl, carbonate, fluorapatite layer.

2

3 4

5

Source: Hench, 1991.

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reaction stages are also observed on bioactive glass surfaces in aqueous solutions such as Tris buffer (Hench, 2006). In solutions containing phosphate, the amorphous calcium phosphate-rich layer might form also on P2O5-free glasses (Ogino et al., 1980). The rate of calcium phosphate formation (stage 4) and the time of onset of crystallization (stage 5) vary greatly with composition. If these two stages take place slowly, the material is not bioactive (Hench, 1998). Surface reactions of the first bioactive glasses were studied in water or Tris buffer solution (Clark et al., 1976; Hench and Clark, 1978; Andersson and Kangasniemi, 1991). Since Kokubo and co-workers (Kokubo et al., 1990; Kokubo, 1991) developed the so-called simulated body fluid (SBF), it has been used widely to measure the bioactivity of glasses in vitro. Simulated body fluid is designed to contain the inorganic constituents of human blood and is thus assumed to provide similar conditions to those found in vivo. Over the years other similar types of solutions have also been suggested, and the composition of SBF has been adjusted (Oyane et al., 2003; Takadama et al., 2004; Kokubo and Takadama, 2006). When immersing bioactive glasses in SBF the five reaction stages described in Table 1.5 can be identified. The concentrations of different elements leached from the glasses into the dissolution medium at different immersion times have been analysed spectrophotometrically or with inductively coupled plasma analysis (ICP) (Jones et al., 2001; Clupper et al., 2003; Cerruti et al., 2005a; Zhang et al., 2009; Zhang et al., 2010). Surface layers formed on the glass at different immersion times have been analysed, e.g. by FTIR, SEM-EDXA, Raman spectroscopy, TF-XRD (Clark et al., 1976; Hench, 1991; Andersson and Kangasniemi, 1991; Ohtsuki et al., 1992; Kim et al., 1995; Rehman et al., 1998; Jones et al., 2001; Sepulveda et al., 2002; Notingher et al., 2003; Clupper et al., 2003; Cerruti et al., 2005a; Zhang et al., 2009). Formation of a bone-like apatite in SBF is often taken as an indication of in vivo bioactivity of the glass (Kokubo and Takadama, 2006). Glass 45S5 is a common reference for bioactive glass studies. Most studies on the influence of composition on the bioactivity deal with the effect of changing one or two components in the original 45S5 composition (Hench, 1991; Hench and Andersson, 1993a; O’Donnell et al., 2009; Gentleman et al., 2010). General approaches based on the mean number of non-bridging oxygen ions in the silica tetrahedron have been suggested to correlate with the bioactivity of glasses (Strnad, 1992; Strnad and Koga, 1999). Zhang et al. (2009) used changes in simulated body fluid and in glasses after immersion to characterize and give basis to models for the in vitro bioactivity as functions of the oxide composition. The coefficients in the additive models for pH of simulated body fluid as functions of the oxide composition at different immersion times are given in Table 1.6. At the shortest immersion times, the best prediction of the pH was given by models in which only the sums of alkalis and alkaline earths expressed as weight percentage were taken into account. At longer immersion times the contribution of the network-forming components B2O3, P2O5 and SiO2 affected the predicted pH of

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Bioactive glasses

Table 1.6 Factors for calculating pH additively from glass composition at different immersion times in SBF Factor for pH at

Constant Na2O + K2O MgO + CaO B2O3 B2O32 P2O5 SiO2 SiO22

4h

8h

24 h

72 h

168 h

6.876 1.778·10−2 1.611·10−2

6.852 2.114·10−2 1.729·10−2

7.017 6.465·10−4 1.952·10−2

4.416 1.977·10−3 6.739·10−2 6.873·10−2

6.482 1.258·10−3 3.405·10−2 8.562·10−3

5.908·10−2

3.609·10−2

6.641·10−3 1.089·10−2

2.867·10−4

Notes: Validity range according to equation 1.1. The glass composition is given in wt%. Source: Zhang et al., 2009.

Table 1.7 Calculated and experimental pH of SBF at 4 and 72 h immersion of glasses 45S5, S53P4 and 13–93 in SBF at 37°C, SA/V = 0.4 cm−1 pH of SBF

At 4 h At 72 h

45S5

S53P4

13–93

calc.

exp.

calc.

exp.

calc.

exp.

7.77 8.19

7.62 8.17

7.65 7.85

7.60 7.85

7.64 7.78

7.64 7.78

Source: Zhang et al., 2009.

SBF. Thus, the models based on regression analysis of measured pH nicely correlate with the reaction stages of bioactive glasses given in Table 1.5. Calculated and experimental values for the pH of SBF at 4 and 72 h immersion of glasses 45S5, S53P4 and 13–93 are given in Table 1.7. Zhang et al. (2009) also suggested a model for the silica gel thickness formed on glasses at 72 h (TLSi) immersion in SBF. The gel thickness was expressed by using relative values from 1 to 5, equation 1.8. The thickness ranges are 1 for no silica layer at all, 2: 0.1 to 3.3 µm, 3: 3.3 to 6.5 µm, 4: 6.5 to 9.8 µm, and 5 for a silica layer thicker than 9.8 µm. The oxides in the equation are given in wt% (Zhang et al., 2009). TLSi = 6.5387 + 7.847·10−2·xNa O + 8.725·10−3·xK2 O 2

2

− 5.747·10−1·xB O + 1.529·10−1·xB2 O − 7.753·10−1·xP O 2 3

2 3

+ 2.207·10−1·xP2 O − 1.573·10−3·x2SiO 2 5

2

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2 5

[1.8]

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A model has been suggested also for formation of the hydroxyapatite layer on the glass surface as a function of the oxide composition of the glass at 72 h immersion is SBF (Zhang et al., 2009). The layer formation is described by a relative number, bioactivity number (BN): see equation 1.9. Glasses with a uniform hydroxyapatite layer with a measured thickness > 0.9 µm have BN = 3. Glasses with incompletely developed hydroxyapatite layers with a layer thickness 0.1 to 0.9 µm are described as BN = 2. Glasses showing only spot like hydroxyapatite formation have BN = 1. When using equation 1.9, BN > 2.5 indicates high bioactivity and BN < indicates 1.5 low bioactivity. Intermediate bioactivity is suggested by values 1.5 < BN < 2.5. BN = 124.680 − 1.545·xNa O − 4.500·10−3·x2Na O − 1.681·xK O 2

2

2

− 2.067·xMgO + 4.199·10−2·x2MgO − 1.661·xCaO − 2.238xB O

2 3

+

1.314·10−1·x2B O3 2

− 1.680·xP O − 2 5

1.551·10−2·x2SiO 2

[1.9]

The models of in vitro bioactivity can be utilized as a first estimation of the reactivity of the glass when developing new compositions. Layer development on glass surfaces in SBF is commonly used as an indication of bioactivity. As SBF is a supersaturated solution towards calcium phosphate precipitation, the glass surface condition might affect the layer formation (Karlsson et al., 2002; Bohner and Lemaitre, 2009). The sample form, surface condition, surface area to volume ratio and fluid circulation also affect the reactions (Greenspan et al., 1994; Jones et al., 2001; Cerruti et al., 2005b; Zhang et al., 2008a, 2008b).

1.3.3 Bioactive glasses in vivo The tissue bonding properties of glasses and glass-ceramics within the system Na2O-CaO-P2O5-SiO2 system have been studied by Hench and co-workers. The compositional dependence of bonding type is illustrated in the Na2O-CaO-SiO2 phase diagram, Fig. 1.4 (Hench, 2006). The bioactivity inside the region A is given for compositions containing 6 wt% P2O5. These glasses bond to bone. The surface reactions of these glasses are described by IIIA in Fig. 1.3. The level of bioactivity increases to the middle of the region A. The compositions inside the region S bond also to soft tissue. The composition of 45S5, E, is in the middle of this region. Glasses in the silica-rich corner, region B, are almost inert and elicit a fibrous tissue around the implant, while glasses C resorb and disappear within 10 to 30 days of implantation (Hench, 1991). The silica-rich compositions B correspond to Type I surface in Fig. 1.3. The dashed lines in the figure give the overall glass-forming tendency of different compositions in the Na2O-CaO-SiO2 system. The glasses in the silica-rich corner are difficult to melt with conventional methods, while compositions below the lowest dashed line do not form glasses. The most bioactive compositions crystallize easily, thus preventing manufacture of other than quenched or cast products as discussed above. Most commercial

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1.4 Compositional dependence (in wt%) of bone bonding and softtissue bonding of bioactive glasses and glass-ceramics in the Na2OCaO-SiO2 system (redrawn from Hench, 2006). Notes: All compositions in region A contain 6 wt% P2O5. Tissue bonding of compositions within the different areas: A – bone bonding, B – nonbonding (reactivity too low), C – non-bonding (reactivity too high), S – soft tissue bonding, E – Bioglass® 45S5 composition, square – commercial soda-lime glasses. The dashed lines limit approximate composition ranges of glass-forming tendencies in the ternary system.

soda-lime glasses have compositions close to the region marked with the black square in the figure. Compositions within the soda-lime glass range can be manufactured into fibers, blown and pressed products, etc. It should be observed that the soda-lime glasses also contain other components but the three-component diagram gives a simplified overall illustration of the system. Two models have been developed for estimating in vivo bioactivity from the oxide composition of the glass (Andersson et al., 1990; Brink et al., 1997). Both models are based on observing the bone contact of glass cones implanted in rat tibia. The layer formation was studied with SEM-EDXA of the cross-sections at 8 weeks’ implantation. The model by Andersson et al. is valid within the compositional range (wt%) 15 to 30 Na2O, 10 to 25 CaO, 0 to 3 B2O3, 0 to 8 P2O5, 0 to 3 Al2O3 and 45 to 65.5 SiO2. An alumina addition was found to inhibit the bone bonding. Depending on the formation of SiO2, the calcium phosphate-rich layers and the bone bonding, the in vivo reactions were divided into five classes. These were given relative reaction numbers (RN) (Andersson et al., 1990): • •

nearly inert glasses showing only small changes in the surface and hardly any bone contact: RN = 1 fairly high solubility and bone contact but not bonding. The glasses show formation of a SiO2-rich layer but no calcium phosphate accumulation: RN = 2

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fairly high solubility and bone contact but not bonding. Formation of a SiO2rich layer and limited calcium phosphate accumulation: RN = 3 formation of a calcium phosphate-rich surface layer but no bone bonding: RN = 4 formation of a calcium phosphate-rich surface layer and bone bonding: RN = 6.

The relative reaction numbers, RN, correlate with the bone response so that glasses with BN >5 are bioactive. Equation 1.10 expresses the reaction number from the oxide composition given in wt% (Andersson et al., 1990): RN = 88.3875 − 0.980188·xNa O − 1.12306·xCaO 2

− 0.560527·x2B O − 1.20556·xP O − 2.808689·xAl O 2 3

2 5

2 3

− 0.0116272·x2SiO

[1.10]

2

The calculated RN numbers for 45S5 and S53P4 indicate bone bonding. The composition of 13–93 is outside the validity range of equation 1.10. Andersson and his co-workers suggested that calcium phosphate forms in vivo within the silica-rich layer, not on the top of it. Bone contact was good for glasses with a thick silica-rich outer layer. Such compositions were suggested to be biocompatible but do not bone with bone. Further, the formation of a calcium phosphate-rich layer on the surface was found insufficient as the only indication of bioactivity. Later, Andersson et al. verified the bone bonding of bioactive glasses with pushout tests in vivo (Andersson et al., 1992). Table 1.8 summarizes the results for glass surface reactions and the push-out test forces when using titanium cone as control. The inert glasses do not show any marked changed in the surface composition. These glasses are not chemically bonded to bone but encapsulated in connective tissue. Accordingly, the push-out strength of inert glasses is low.

Table 1.8 Summary of layer formation and push-out strengths at 8 weeks in rabbit tibia. The glass codes give the silica and phosphorus pentoxide content in the glasses by wt%. The other constituents are Na2O, CaO, B2O4 and Al2O3 Sample

Silica layer

HCA layer

Bone response

Strength (MPa)

Glass reaction

S65.5P1 S52P3 S52P8 S45P7 S46P0 S55.5P4 Titanium

Thin Yes Yes Yes Yes Yes

No No Yes Yes Yes Yes

None Contact Contact Bonding Bonding Bonding Contact

0.5 ± 0.4 3.6 ± 0.9 3.0 ± 0.5 23.0 ± 2.9 16.4 ± 3.9 19.9 ± 4.0 2.2 ± 0.6

Almost inert Soluble, single layer Soluble, dual layer Soluble, dual layer Soluble, dual layer Soluble, dual layer

Source: Andersson et al., 1992.

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Glasses with silica-rich layer formation but no hydroxyapatite precipitation show bone contact but no chemical bone bonding. The push-out force is comparable to titanium. Glasses with a silica-rich layer in the surface are biocompatible. Bioactive glasses forming a dual layer of silica and hydroxyapatite show bonding forces varying from one- to ten-fold values compared to titanium. In vivo bioactivity by layer formation on glasses within the system Na2O-K2OMgO-CaO-B2O3-P2O5-SiO2 has been expressed also by another relative number, the index of surface bioactivity (ISA) (Brink et al., 1997). The ISA number is related with the formation of surface layers on glasses in vivo: inert glass = 1, silica-rich layer = 2, layered structure = 3, bioactive = 4: equation 1.11. The compositional range of the equation is according to equation 1.1. The oxides are expressed in wt%:

[1.11] The calculated ISA values for 45S5 and S53P4 suggest that the glasses are bioactive, while 13–93 has an ISA value 3.2. This glass was, however, found to bond to bone and thus to be bioactive (Brink et al., 1997). The observations suggest that the models for in vivo bioactivity can be used to get a first estimation of the bioactivity of a specific glass composition. Also the different models describing the in vitro bioactivity correlate with the in vivo observations. The melt-derived bioactive glasses 45S5 and S53P4 are FDA-approved for certain clinical applications. When compared in vitro, the thickness of the dual layer was somewhat less in S53P4 than in 45S5 after 1 week in simulated body fluid (Hupa et al., 2010). Only small differences in the layer thickness on implants of the two glasses were observed after 8 weeks in the soft tissue of rats (Hupa et al., 2010). Both glasses bonded to bone but the layer thickness at 8 weeks was slightly less in S53P4. In clinical applications these glasses are used mainly in different bone-filling applications as glass particulates. The bioactive glasses have been found safe and promising as bone substitutes for the treatment of benign bone tumours (Lindfors, 2009; Lindfors et al., 2009; Lindfors et al., 2010). In a clinical follow-up study granules of S53P4 used to fill bone cavities were found to increase the cortical thickness (Lindfors et al., 2009). Even after 14 years some remnants of glass particles could be identified (Lindfors et al., 2010). The long-term clinical study demonstrated, however, the potential of bioactive glasses in benign bone tumour surgery both in children and adults. However, the observations suggest that the dissolution of bioactive glasses is slow. Although the in vitro, sin vivo and clinical observations give similar reaction trends, a physiochemical reaction mechanism and dissolution of bioactive glasses is not fully understood.

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Future trends

Bioactive glass research during the past 30 years indicates that the bioactivity and physical as well as chemical properties can be adjusted by changes in the oxide composition of the glasses. Several properties describing the relationship between the oxide composition of the glass and in vitro, in vivo and other properties are available. These models are valuable when tailoring glass compositions for various clinical applications. Much bioactive glass research has dealt with the development of compositions that bond to tissue. Although detailed long-term physicochemical reactions are still obscure, an essential understanding of the mechanisms behind the interaction of the biological environment and the glass surface reactions has been established. As glasses are brittle materials, their increasing use in medical applications is restricted by their unpredictable mechanical behaviour in load bearing applications. Therefore, the use of glasses in composites together with biodegradable organic polymers has been one of the main focuses during the recent years. This often means the use of glasses as particulates, thin fibres or thin-walled sintered structures. Due to the large surface area the glasses are likely to react rapidly, and thus they lose their mechanical reinforcing capacity. This gives rise to the need of finding compositions that would be osteoconductive but not resorb too rapidly. Although much of the reactivity of glasses in biological solutions is understood, the mutual interaction of glasses, for example with bioresorbable polymers is not fully understood. Thus, research and development into bioactive glass compositions for novel medical applications will still be topical for some time to come.

1.5

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Taipale, S., Ek, P., Hupa, M., Hupa, L., 2008. Continuous measurement of the dissolution rate of ions from glasses. Advanced Materials Research, Vols. 39–40, pp. 341–6. Takadama, H., Hashimoto, M., Mizuno, M., Kokubo, T., 2004. Round-robin test of SBF for in vitro measurement of apatite-forming ability of synthetic materials, Phosphorus Research Bulletin, 17, 119–25. Vedel, E., Arstila, H., Ylänen, H., Hupa, L., Hupa, M., 2008. Predicting physical and chemical properties of bioactive glasses from chemical composition. Part 1: viscosity characteristics. Glass Technology: European Journal of Glass Science and Technology, Part A, 49 (6), 251–9. Vrouwenvelder, W. C., Groot, C. G., de Groot, K., 1994. Better histology and biochemistry for osteoblasts cultured on titanium-doped bioactive glass: Bioglass 45S5 compared with iron-, titanium-, fluorine- and boron-containing bioactive glasses. Biomaterials, 15 (2), 97–106. Watts, S. J., Hill, R. G., O' Donnell, M. D., Law, R. V., 2010. Influence of magnesia on the structure and properties of bioactive glasses. Journal of Non-Crystalline Solids, 356, 517–24. Williams, D. F., 1999. The Williams Dictionary of Biomaterials. Liverpool, UK: Liverpool University Press. Ylänen, H., Karlsson, K. H., Itälä, A., Aro, H. T., 2000. Effect of immersion in SBF on porous bioactive bodies made by sintering bioactive glass microspheres. Journal of Non-Crystalline Solids, 275 (1, 2), 107–15. Zanotto, E. D., 1987. Isothermal and adiabatic nucleation in glass. Journal of NonCrystalline Solids, 89, 361–70. Zhang, D., Hupa, M., Aro, H. T., Hupa, L., 2008a. Influence of fluid circulation on in vitro reactivity of bioactive glass particles. Materials Chemistry and Physics, 111 (2–3), 497–502. Zhang, D., Hupa, M., Hupa, L., 2008b. In situ pH within particle beds of bioactive glasses. Acta Biomaterialia, 4 (5), 1498–505. Zhang, D., Vedel, E., Hupa, L., Aro, H. T., Hupa, M., 2009. Predicting physical and chemical properties of bioactive glasses from chemical composition. Part 3: in vitro reactivity. Glass Technology: European Journal of Glass Science and Technology, Part A, 50 (1), 1–8. Zhang, D., Leppäranta, O., Munukka, E., Ylänen, H., Viljanen, M. K., Eerola, E., Hupa, M., Hupa, L., 2010. Antibacterial effects and dissolution behavior of six bioactive glasses. Journal of Biomedical Materials Research, Part A, 93A (2), 475–83.

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2 Surface modification of bioactive glasses J. CHANG and Y. L. ZHOU , Chinese Academy of Sciences, China and Y. ZHOU , Shanghai Jiao Tong University, China

Abstract: The surface modification of biomaterials has exhibited great potential in biological applications by transforming the existing surface into more appropriate compositions and/or topographies. The surface of bioactive glasses plays a critical role in their performance, and many studies have been conducted on the surface modification of bioactive glasses or modification of other materials with bioactive glasses, which can combine the properties of the two materials together to obtain desirable functions. In this chapter, the surface modification of bioactive glasses and using bioactive glasses are broadly reviewed in the following three aspects: (1) surface modification of bioactive glasses to improve bioactivity; (2) surface modification of bioactive glasses surface with organic molecules to improve the dispersivity; (3) surface modification of bioinert materials with bioactive glasses. Key words: bioactive glasses, surface modification, biomineralization, coating techniques.

2.1

Introduction

Bioactive glasses (BGs), which were first synthesized 40 years ago, have been extensively studied as an artificial bone grafting material for bone repairs, and have gained great acceptance in clinical applications (Hench, 1998, 2006). As an implant for hard tissue repair, the surface of BGs plays an important role in their properties after implantation (Kenny and Buggy, 2003; Hench and Polak, 2002; Gabbi et al., 1995). In order to control the surface properties precisely, modification of the surface of a certain bioactive glass (BG) or the surface of a substrate with BG are usually employed which can combine the properties of both the coatings and the substrates to obtain desirable functions (Verne et al., 2009a; Wang, 2009). Surface modification of BGs has promising potential for converting the existing surface into more desirable compositions and/or topographies for biological application (Duan and Wang, 2006). Surface modification of BGs can be broadly classified into three categories: 1) treating the surface of BGs to improve the surface bioactivity through silanization, biomineralization and microroughening; 2) modifying BG particles to render their compatibility with another phase (Neouze and Schubert, 2008) through condensation or a grafting reaction (Slowing et al., 2008; Vallet-Regi, 2006; Vallet-Regi et al., 2007); 3) coating bio-inert materials (such as titanium alloy) with BGs by physical or chemical techniques such as plasma spray coating, physical vapor deposition, sol-gel process etc (Liu et al., 2008b). 29 © Woodhead Publishing Limited, 2011

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In this chapter, these three categories of the surface modification will all be introduced.

2.2

Surface modification of bioactive glasses to improve bioactivity

Surface characteristics are critical in developing a biomaterial implant, since the interaction between the biomaterial and the ambient physiological medium takes place at the surface of the implant (Iucci et al., 2004). After implantation of BGs into the body, numerous physiological reactions occur simultaneously in the intermediate layer between the targeted tissue and the implant. Taking the first synthesized bioactive glass, 45S5 Bioglass® (Table 2.1) as an example, the surface reactions that occurred at the Bioglass® surface are summarized in Table 2.2 (Stroganova et al., 2003; Hench, 1991; Lobel and Hench, 1996). These surface reactions occur within the first 12 to 24 hours after implantation. The first step is a rapid reaction of the release of sodium ions (Na+) from the surface of the glass via ion exchange with a hydrogen ion from the environment Table 2.1 The composition of 45S5 Bioglass® Compound

Percentage (wt%)

SiO2 CaO Na2O P2O5

45.0 24.5 24.5 6.0

Table 2.2 45S5 Bioglass® reaction stages with increasing time Increasing time Stage 11 10 9 8 7 6 5 4 2–3 1 0

Reaction event Crystallization of matrix Cellular attachment Differentiation of stem cells Attachment of stem cells Action of macrophages Adsorption of biological moieties (proteins, etc) Nucleation and crystallization of calcium phosphate to HCA Precipitation of amorphous calcium phosphate Dissolution and repolymerization of surface silica Sodium hydrogen ion exchange Initial glass surface

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(H+ or H3O+), which results in a negatively charged surface. Then a silica-rich layer is formed on the surface of Bioglass® as the loss of Na+ causes the breakdown of the silica network, with the resultant formation of Si(OH)4 groups. After that, an amorphous calcium phosphate (Ca-P) layer starts to form on the surface of the silica-rich layer and eventually incorporates the biological moieties, such as blood proteins, growth factors and collagens. The adsorption of the organic species from the body fluid occurs concurrently with the first three reaction stages, which is believed to contribute at least partially to the biological nature of the hydroxyapatite (HA) layer. Within about three to six hours in vitro, the Ca-P layer will crystallize into the HA layer, which has been described as the bonding layer. Since the composition of the HA layer on the surface of the Bioglass® is chemically and structurally similar to the mineral component of natural bone, it allows the regenerated tissue to attach directly onto the surface. As the reaction continues, this surface HA layer grows in thickness to form a bonding zone of up to 100 microns, which is essential and mechanically compliant to maintain the binding of the implant with the natural tissue. The HA layer is also able to adsorb biological moieties and support cell adhesion and growth. Furthermore, the releasing of ionic components from the glass surface has been shown to continue for a long period of time, and enhances the development of the surface reactive layers. In this way, BG participates in the whole repair process, leading to the creation of a direct bond of the material to the tissue. Therefore, the term ‘bioactive’ means that, through a series of interfacial ion exchange reactions, a silica-rich gel layer forms followed by the formation of the Ca-P layer on the BG surface, which promotes the interfacial bonding with tissues after implantation and enhance new tissue regeneration (Montanaro et al., 2002). Based on the description above, the surface characteristics, the silica-rich layer, biomineralization of the Ca-P layer, and immobilizing proteins and growth factors are essential for the bioactivity of BGs. Therefore, many works have been undertaken to modify the surface of BGs through silanization, deposition of Ca-P layer and adsorption with proteins, to improve the bioactivity of BGs. It is believed that a biomaterial’s surface characteristics can alter cell behavior at many levels (Lipski et al., 2008). Surface morphology can directly influence the cellular response, for example, by inducing the release of growth factors and cytokines from the adhered osteoblasts (Montanaro et al., 2002). A smaller-scale surface roughness (in the range from 10 nm to 10 mm) can increase the surface areas of BGs and may influence the biological interaction with cells and large biomolecules. Gough et al. compared smooth and roughened 45S5 Bioglass® in vitro and found that the roughened samples enhanced mineralized nodule formation in human primary osteoblast culture (Gough et al., 2004). Therefore, creating microroughness on the surface of BGs is an important approach to further enhance the bioactivity of the glasses. In this section, surface modification of BGs were discussed in three parts: silanization, biomineralization of Ca-P and the surface structure.

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2.2.1 Silanization The silica gel layer forming on the surface of BGs is regarded as the first step for bone–BGs bonding. On the silica layer, an amorphous Ca-P phase is precipitated, which evolves with time into HA (Gao et al., 2001; Lusvardi et al., 2009). Thereby the concept of silanization is important for the surface modification of BGs. One of the common non-toxic protein-coupling agents is 3-aminopropyltriethoxysilane (APTS), and it is used to silanize the surface of BGs to promote protein adhesion and cell growth on biological implants (Howarter and Youngblood, 2006). Since there are plenty of silanol on the surface of BGs (Andrade et al., 2004), APTS can be surface-modified on BG particles through a wet-chemical method in a dynamic inert nitrogen atmosphere (Chen et al., 2008c). The reaction process of Si-OH on the surface of BGs with APTS is shown in Fig. 2.1. Highly porous 45S5 Bioglass®derived glass-ceramic scaffolds were surface silane functionalized with APTS, which can favor the formation of HA and improve cell attachment and growth. The osteoblasts proliferated better on the functionalized glass-ceramic surface (Chen et al., 2006, 2008b). After surface modification with APTS, the introduced amino-groups can be used for protein grafting. One example is the immobilization of a model protein, carnosine, on the surface of BGs to bind human bone morphogenetic protein (BMP) (Verne et al., 2009b). BMPs are growth factors that are known to induce bone and bone marrow regeneration, among which BMP-2 and BMP-7 have already been applied clinically for bone regeneration due to high osteoinductive activity (Takahashi et al., 2005). Alkaline phosphatase (ALP) participates in the bone formation and mineralization process, and is widely used as a marker of osteoblast differentiation (Groeneveld et al., 1995). Verne et al. have demonstrated that ALP covalently grafted to APTS-modified BGs can enhance the ability of the material to induce HA precipitation in simulated body fluid (SBF) while maintaining the enzyme activity (Verne et al., 2010). Furthermore, protein-loaded BG scaffolds can also be used as a delivery system. For example, porous 45S5 Bioglass®-derived scaffolds were surface-functionalized with APTS and loaded with collagen. After functionalization, the stability of the collagen attachment and the release stability against the pH change in the biological environment were increased, to create an environment suitable for enhancing cell attachment (Chen et al., 2008a). Laminin can also be adsorbed on the APTS surface-modified

2.1 Modification process of bioactive glasses (BGs) with 3-aminopropyltriethoxysilane (APTS).

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binary 70S30 BGs (70mol%SiO2-30mol%CaO) and 58S BG (60mol%SiO236mol%CaO-4mol%P2O5) foams. Sustained and controlled release of laminin from the modified scaffolds, which has a beneficial effect on tissue formation (Lenza et al., 2003), was achieved over a 30-day period. It is known that nitric oxide (NO) plays an important role in regulating osteoblasts and osteoclasts in bone metabolism. APTS has also been used for preparation of NO-releasing BG materials (Pryce and Hench, 2004).

2.2.2 Biomineralization of calcium phosphate on the surface of bioactive glasses One critical event that facilitates bone formation is the interaction of BGs with physiological solutions and the subsequent formation of a Ca-P rich layer on the glass surface, which is important in the development of osteoconductive biomaterials for orthopaedic applications (Leonor et al., 2009). Apatite is the major mineral phase of which the hard tissues such as bone and dentin are composed (Kim et al., 2004; de Arenas et al., 2006). Therefore, the induction of apatite formation is another important approach to modifying the surface of BGs. For this purpose, different solutions have been applied for soaking the BGs in order to facilitate the formation the Ca-P layer on their surface. Simulated body fluid (SBF) (see Table 2.3) is a common solution for soaking BGs and apatite formation, since its ionic concentrations are nearly equal to those of human blood plasma (Kokubo and Takadama, 2006). In addition, Tris(hydroxymethyl) aminomethane (Tris) solution at different pHs (Cerruti et al., 2005) or with electrolytes typical for plasma (TE, a solution with 142.0 mM Na+, 5.0 mM K+, 1.5 mM Mg2+, 2.5 mM Ca2+, 148.8 mM Cl−, 4.2 mM HCO3−, and 1.0 mM HPO42−) can also be used to treat BGs for deposition Table 2.3 Ionic concentrations and pH of simulated body fluid and human blood plasma Concentration (mM)

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

Simulated fluid

Blood plasma

142.0 5.0 1.5 2.5 147.8 4.2 1.0 0.5 7.25

142.0 5.0 1.5 2.5 103.0 4.2 1.0 0.5 7.20–7.40

Source: Fujibayashi et al., 2003.

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of HA on the surface. Adsorption of serum proteins onto HA layer serves to enhance bone cell adhesion, proliferation, and function (El-Ghannam et al., 1997; Radin et al., 2005). Moreover, some investigations have demonstrated that certain proteins incorporated into the soaking solution can regulate the capability of forming Ca-P layer on the surface of BGs thus affecting the bioactivity of the BGs. A study has shown that the addition of fibronectin can reduce the electronegativity of the BG surface and consequently delay the formation of both the amorphous and the crystalline Ca-P layers (Lu et al., 2001). Radin and coworkers studied the influence of serum proteins on the formation of Ca-P layer on the surface of BG granules. Serum can promote porous surface structure of Ca-P layer on the Surface of BGs (Radin et al., 1997, 2000), while recombinant porcine amelogenin rP172 showed a modulation effect on the oriented growth of apatite crystals on 45S5 Bioglass® surface in supersaturated calcifying solution (SCS). Grafting functional groups on the surface of BGs can also alter the bioactivity of BGs. With the addition of the amino and carboxylic groups to mesoporous bioactive glasses (MBGs) through post-grafting process, the nucleation, growth rate and morphology of carbonated HA formed on the surface of the glasses were affected remarkably. MBGs functionalized by amino groups promote the formation of spherical HA particles, while the nucleation and growth rate of HA on MBGs functionalized by carboxylic groups decrease to a large extent with the increasing carboxylic group content (Sun et al., 2008). In brief, a Ca-P layer is osteoconductive, and vital for the bioactivity of BGs (Leonor et al., 2009). Surface modification of BGs by biomineralization of the Ca-P layer and subsequent immobilization of specific proteins can enhance cell attachment, proliferation and differentiation.

2.2.3 The role of surface structure Chemical treatment is an easy way to modify the surface morphologies and increase microroughening of the surface without changing the body properties (Li et al., 2009). BG microspheres (215 to 350 µm) were etched in different etching solutions (from pH = 1.0 to 13) and etching time (soaking time from a few seconds to 30 min) to achieve the desired surface appearance with an average roughness value of 0.35 to 0.52 µm (Itala et al., 2001). The microrough surface can accelerate the formation of Si-gel layer during the first hours of SBF and Tris immersion, and can enhance the attachment of human osteoblast-like MG-63 cells during the first 24 h of incubation in vitro (Itala et al., 2002). An in vivo study in a rabbit model also demonstrated that microroughened surface can enhance osteopromotive properties and bone-bonding response (Itala et al., 2003a). Furthermore, surface microroughness can cause a temporary change in the expression of specific genes (Itala et al., 2003b; Valimaki et al., 2005). Bioleaching is a soft biochemical approach that can obtain nanoscale surface under gentle condition. Borosilicate glasses were treated through fungus-based bioleaching and the surfaces were

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morpho-chemical modified with monodispersed ultrafine (~5 ± 0.5 nm) silicate nanoparticles through acidolysis (Kulkarni et al., 2008).

2.3

Surface modification of bioactive glasses using organic molecules to improve dispersivity

Although BGs show favorable bioactivity, lack of in situ moldability and relative brittleness limit their applications (Rich et al., 2002). Therefore, polymers/BG composites have been fabricated in order to obtain biomaterials with improved properties (Misra et al., 2008; Maquet et al., 2004; Jiang et al., 2005; Silva et al., 2004). Among the factors affecting the properties of the polymers/inorganic composites, the interface adhesion of inorganic particles to polymer matrix plays an important role (Hong et al., 2005; Supova, 2009). It has been found that some polymers/inorganic composites lost their strength rapidly in a physiological environment, and the failure occurred mainly at the interface between inorganic particles and the polymer matrix (Zhang et al., 2005). The main reason was the tendency for inorganic particles to agglomerate in the polymer matrix owing to their small dimensions and incompatible polarity with polymers (Cheng and Chang, 2006; Borum-Nicholas and Wilson, 2003; Liu et al., 2008a). Therefore, improving poor dispersion of inorganic particles in polymeric matrix was critical for preparing composite materials with improved properties. Several chemical reactions have been employed to modify BGs with organic molecules based on the reaction between Si-OH on the surface of BG particles and the functional group of organic molecules. The isocyanate-ended low-molecular-weight PLLA was reacted with BG particles (Si:P:Ca = 29:13:58 weight ratio) (particle size 40 nm) at 80°C for about 12 h to obtain surface-modified BG particles. The grafting resulted in an increase in phase compatibility, and the consequent improvement of tensile strength, tensile modulus and impact energy. An in vitro bioactivity test showed that, compared to pure PLLA scaffold, the BG/PLLA nanocomposite possessed a greater capability of inducing the formation of an apatite layer on the scaffold surface (Liu et al., 2008a). In another study, silane coupling agent 3-glycidoxypropyltrimethoxysilane was introduced to modify the surface of BG particles through condensation. The filler dispersion and phase compatibility between poly (D,L-lactide) (PDLLA) and BG particles were also improved (Zhang et al., 2009a). Stearic acid is widely used in the surface modification of inorganic materials, because its carboxyl groups can chemically bind with -Si-OH on the surface of BGs, and the long alcohol chain has high compatibility with polymers (Grassi et al., 2003; Saleema and Farzaneh, 2008). Dodecyl alcohol can also be used to modify BGs through an esterification reaction. The esterification of silanols with alcohols is well known as described in the following equation: -Si-OH + R-OH ← → Si-OH

HO-R← → Si-OR + H2O

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2.2 The SEM images (×500) of the composite films composed with unmodified and modified bioactive glass particles: (a) 45S5, (b) m-45S5, (c) mesoporous 58S, (d) m-mesoporous 58S, (e) 58S, (f) m-58S (Gao and Chang, 2009).

In this study, three kinds of BGs (45S5, 58S and mesoporous 58S BGs) were surface-modified with dodecyl alcohol through esterification at 260°C to improve the homogeneous dispersion of BG particles in polymeric matrix. A SEM observation (Fig. 2.2) illustrated that the modified BG particles were homogeneously dispersed in the PDLLA matrix. The modified composite films can still induce the formation of HA on its surface after immersion in SBF, and the

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2.3 SEM images (with ×500 and ×1000 images on the left column and ×60 000 images on the right column) of the surface of films after immersion in SBF for 7 days (a, b) PDLLA/58S composite film, (c, d) PDLLA/m-58S composite film (Gao and Chang, 2009).

distribution of HA was more homogeneous on the film (Fig. 2.3) (Gao and Chang, 2009). However, the disadvantage of the modification with dodecyl alcohol is the decrease of hydrophilicity, which may affect the biocompatibility of the composite materials, since increased surface hydrophilicity is known to be associated with enhanced protein adsorption and consequent cell adhesion and proliferation on biomaterials (Zhang et al., 2009b). Fortunately, this modification is reversible and the dodecyl alcohol can be removed after the achievement of homogenous dispersion of BG particles in composite materials by hydrolytic treatment in hot water. The properties (such as tensile strength) of the composite films after treatment will not be affected. Most importantly, cells on the composite films after hydrolysis show the highest proliferation rate and differentiation level (Fig. 2.4 and Fig. 2.5) (Zhou et al., 2010). In summary, non-toxic organic molecules, especially biocompatible molecules are useful to modify the surface of BGs for the improvement of the dispersivity in polymer matrix.

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2.4 The dMSC proliferation of the composite films. OD value on y-axis represents the number of living cells (p < 0.05) (Zhou et al., 2010).

2.5 ALP activity of dMSCs after culturing on the composite films for different periods (p < 0.05) (Zhou et al., 2010).

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Surface modification of bioinert materials using bioactive glasses

Owing to excellent bulk properties, such as relatively low modulus, good fatigue strength, corrosion resistance and biocompatibility, metallic materials including stainless steel, commercially available pure titanium (Ti), Ti alloys (such as Ti6Al4V) and nickel-Ti etc, are widely used for orthopedic applications (Ramaswamy et al., 2009). However, they are bioinert, owing to their absent or weak interaction with living tissues upon implant (Navarro et al., 2008). Interfacial movement under external stress leads to loosening and deterioration of the mechanical fit, which causes pain and eventually leads to clinical failure of the bioinert implants (Tilocca, 2009; Shi et al., 2002). One of the most general solutions is coating the bioinert substrates with bioactive materials, which can modify the surface bioactivity and maintain the mechanical properties of the metal implants at the same time (Vallet-Regi et al., 2003; Liu et al., 2008b). Several bioactive materials have been evaluated for implant coatings. BGs are one of the most promising candidates for implant coating because of high bonebonding ability through the formation of a HA layer on the surface and excellent resorbability in the body fluids (Lewandowska et al., 2007; Ballo et al., 2008). The bone-bonding ability of BGs is also helpful for providing right fixation of the implants. Three factors are essential for preparing BG coatings onto implants successfully: 1) the thermal expansion coefficients of BGs and the metal substrates should be similar in order to avoid the generation of large thermal stresses that can result in coating cracking or delaminating during fabrication; 2) the coating should have good adhesion to the substrates (Bolelli et al., 2007b); 3) the deposited coatings should be bioactive in SBF (Ma et al., 2006). Various coating methods have been employed to create BGs or BG composite layers onto metallic substrates, including the plasma-sprayed deposition technique, electrophoretic deposition, the sol-gel process and magnetron sputtering, and so on (Verne et al., 2005). Each of these coating techniques has its own strengths and weaknesses. In this section, different coating techniques using BGs are introduced.

2.4.1 Enameling technique The conventional enameling technique is a simple way to coat metal substrates by BGs. The advantages of this technique are low cost, ease of operation and optimization by changing the parameters. Experimentally, it is possible to prepare BG coatings with good adherence to Ti substrate by controlling the composition and process conditions (firing time, temperature and atmosphere) (Bloyer et al., 1999). For example, by carefully controlling the composition of BGs and the treatment of the enameling and glazing process, BG coatings with a thermal expansion coefficient congruent with the alumina or Ti alloy can be obtained (Gomez-Vega et al., 1999; Pazo et al., 1998). These kinds of BG coatings can

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induce the precipitation of a thick HA layer when soaked in SBF (Vitale-Brovarone and Verne, 2005). Compared to uncoated implants, samples coated with BGs showed osteoconductivity (Martorana et al., 2009) and enhanced osseointegration (Ignatius et al., 2005). A functionally graded structure has been applied to minimize crack propagation at the interface between the coating and the substrate. The multilayer approach was developed to achieve a compliant material that could withstand the stresses due to the expansion coefficient mismatch between the substrate and the coatings, and maintain the bioactivity of the outer layers (Brovarone et al., 2001). Bioactive glasses with higher silica were used to form the first layer of a composite coating, providing a strong bond with the metal substrates, whereas BGs with lower silica were used to form the outer layer to enhance bioactivity (Rahaman et al., 2008; Lopez-Esteban et al., 2009; Verne et al., 2004). Compared to uncoated Ti alloy, the functionally graded coatings showed good cytocompatibility (Foppiano et al., 2006) and can indirectly induce an increase of expression of Runx-2, a key marker of osteoblast differentiation (Foppiano et al., 2007). The employed intermediate layer between the metal and BGs can also avoid the contamination of the diffused metallic ions. Verne et al. found that an intermediate layer based on SiO2-CaO, which formed undesired additional phases with the metal substrate, was necessary to avoid the diffusion of metal ions (Verne et al., 2005). Besides changing the composition of BGs, other methods are also used to improve the bioactivity of the outer layers. MBGs have shown higher bioactivity than the conventional sol-gel BGs due to the higher specific surface area (Xia and Chang, 2006). A thin film of mesoporous silica was deposited on Ti6Al4V substrate coated with glasses, which induced apatite formation in SBF after 7 days (Gomez-Vega et al., 2001). In brief, it is possible to fabricate a layer of BGs onto bioinert implants with an adequate thermal expansion coefficient, good adherence ability and bioactivity using the enameling method by tailoring the composition of BGs and enameling processing conditions.

2.4.2 Plasma-sprayed deposition Plasma-spraying is a feasible way to coat metallic substrates with glass powders (Bolelli et al., 2007a). The high-temperature plasma (up to 10 000 to 30 000K) melts glass particles into droplets. Then the droplets are ejected at high velocity and sprayed on the substrate with rapid solidification (Kang et al., 2007). The advantages of the plasma-sprayed technique are in creating reasonably high coating bond strength and mechanical properties (Ding et al., 2001). Compared to the enameling technique, BG coating prepared by plasmaspraying has higher superficial mechanical strength. Schrooten et al. used reactive plasma spraying method to deposit the BG layer on a Ti6Al4V rod. This technique proved to produce a high-quality coating with the adhesion strength of 40.1 ± 4.8 Mpa in shear and 69.4 ± 8.4 MPa in tension (Schrooten et al., 1999). The BG

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coated metal implants showed good biocompatibility in vitro and in vivo. Gabbi et al. prepared amorphous BG-coated titanium and its alloys by plasma-spraying. Both in vitro and in vivo studies confirmed that BG coating was osteoconductive (Gabbi et al., 1995). In order to increase the mechanical properties of the coating, Goller et al. investigated the effect of bond coating layer, which can provide a good thermal expansion match between the substrate and the BG coating. In that study, they used a alumina-titania (60%Al2O3, 40%TiO2) bond coating and found this kind of coating showed better mechanical strength than the pure BG coating (Goller et al., 2003). In general, implants coated by BGs using plasma spraying showed superior properties such as bioactivity and long-term stability, and is generally accepted for orthopedic applications. However, the plasma spraying is a beeline process, so it is challenging to use it for coating complex substrates and to deposit graded coating (Ma et al., 2006).

2.4.3 Electrophoretic deposition (EPD) technique Electrophoretic deposition (EPD) is an electrochemical method which is usually carried out in a two electrode cells. When direct current (DC) electric field is applied, charged particles suspended in a suitable liquid move toward the oppositely charged electrode, and then the particles accumulate at the deposition electrode and create a relatively compact and homogeneous film. Therefore, EPD can be applied to any solid that is available as a fine powder (e.g. <~30 µm particle size) or as a colloidal suspension, including BG particles (Corni et al., 2008; Besra and Liu, 2007). Compared to plasma-sprayed deposition, this process seems to be very promising for developing glass and ceramic functionally graded coatings with different thickness (from less than 1 µm to more than 500 µm) on complex-shaped substrates. The BG-HA coating, which is composed of 55 mol%SiO2, 26 mol%CaO, 13 mol%MgO, 6 mol%P2O5 and HA was deposited on Ti6Al4V substrate by electrophoretic deposition. The coating thickness and morphology were controllable and the corrosion resistance of the coatings could be optimized by suitably adjusting the deposition potential and coating time (Balamurugan et al., 2009) to obtain desirable adhesive strength between the coating and the substrate (Li et al., 2001). In short, EPD offers the advantages of high purity of deposited material, high deposition rate, the possibility of deposition of uniform coatings on substrates of complex shape and large surface area with tailored composition, thickness and the structure of the coating. However, an appropriate solvent media is needed to ensure stability of the suspension and high electrophoretic mobility.

2.4.4 Sol-gel process The sol-gel process is an emerging technology to produce coating films at much lower temperature than traditional ceramic coatings (Conde et al., 2003). The

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attainments of films by the sol-gel process have been successfully used to prepare coatings on stainless steel, Ti alloy and aluminum in order to improve the oxidation and corrosion resistance of those metals (Ballarre et al., 2008). Many techniques are used to fabricate thin films by the sol-gel process, such as dip coating, spin coating, spray coating and so on (Attia et al., 2002). The advantages of the sol-gel process are the low processing temperature and easy to control the composition of the coatings (De Barros Coelho and Magalhaes Pereira, 2005). Vitreous SiO2CaO system was coated on Ti6Al4V substrates by the sol-gel method. The textural parameters (porosity and roughness) and thickness of the obtained films increased when the concentration of the precursor solutions was raised (Izquierdo-Barba et al., 2006). In another study, 316L stainless steel (SS) was coated with BG (57.44wt%CaO-35.42wt% SiO2-7.15wt%P2O5) through dipping in the sol. Crackfree and homogeneous BG coating was achieved and the BG-coated 316L SS showed more pitting corrosion resistance as compared with pristine samples (Fathi and Doostmohammadi, 2009). Through coating with BG layers, biocompatibility and bone osteointegration of the implants can be obtained simultaneously to reduce healing time (Fathi and Mohammadi, 2008). The sol-gel solution was prepared with BG powder (with composition of 65wt%SiO231wt%CaO-4wt%P2O5) mixing with titanic solution and then deposited onto commercial pure Ti dental implant. Compared to uncoated titanium, the BG coating resulted in surface roughness and was found to stimulate osteoblastic cells in producing a higher level of ALP and collagen (Ramires et al., 2003). Sol-gel BG composition (58S and 77S) coatings were prepared by dipping the alumina substrates in the sol and the coated samples were implanted in a rabbit model. The percentage of the bone in direct contact with the implant was greater for coated samples compared to the bulk alumina samples. In the case of 58S-coated implants, bone percentage significantly increased from 45.1% after 3 weeks up to 87.8% after 24 weeks of implantation (Hamadouche et al., 2000). In summary, the sol-gel method has promising potential for preparing BG coatings. However, the inorganic coatings are brittle and usually have low adhesion strength (Wang and Bierwagen, 2009).

2.4.5 Pulsed laser deposition Pulsed laser deposition (PLD) is a thin film deposition technique. The target materials are vaporized by a high-power pulsed laser beam in the vacuum chamber and deposited as a thin film onto the substrate. PLD has broad prospects for depositing mono- and multi-layer bioactive coatings on metallic substrates, and the advantages of this technique is the high adhesion, absence of contamination and pores, and stoichiometrical transfer of the target composition. D’alessio et al. demonstrated that it was possible to deposit films with the right stoichiometry by laser ablation of a Bioglass® target (D’alessio et al., 1999; D’alessio et al., 2001). Functional graded BG coatings can also be coated on the metallic substrate with

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good adhesion through PLD techniques (Tanaskovic et al., 2008; Tanaskovic et al., 2007), and the BG coated metal implants showed highly improved bioactivity (Berbecaru et al., 2009). In another study, Ti6Al4V metal implants were coated with BG42 (42%SiO2-20%Na2O-10%K2O-20%CaO-5%MgO3%P2O5) and implanted into paravertebral muscle of rabbit. Ca-P layers were developed after implantation, and the implants do not elicit any inflammatory response of the surrounding tissues (Borrajo et al., 2007). BG coatings on metal substrates can also be created by heating the BG particles placed on the surface of metal sheets using a focused laser beam. In this process, a Ti sheet was coated with BG particles by dip-coating in a glass suspension, and a focused CO2 laser beam was applied to heat-treat the glass to form a BG coating on the Ti sheet. The BG coatings showed good attachment to the substrate, maintained in vitro bioactivity (Moritz et al., 2004b), and demonstrated significantly higher osteoconductivity than the control Ti implants in vivo (Moritz et al., 2004a). Besides the techniques mentioned above, some new techniques for preparing BG coatings with desirable properties have been reported, such as the highvelocity suspension flame sprayed (HVSFS) technique (Bolelli et al., 2009) and the radio frequency (RF) magnetron sputtering technique (Stan et al., 2009). However, the feasibility of these techniques needs to be confirmed by further investigation.

2.5

Conclusions and future trends

Surface modification is providing increased opportunities for design, development and control of the properties of the biomaterials. In this chapter, an overview of surface modification methods used to improve the properties of inert materials by surface modification using BGs has been introduced. In addition, the methods used to improve the function of the BGs, or the properties of BGs on the purpose of preparation of suitable biocomposites for extensive biomedical applications were also reviewed. When a biomaterial is implanted into the body, the interaction at the interface is critical, and this is determined by the material’s chemistry and topography. At an early stage of the implantation, the adsorption of proteins on the surface is the key step, followed by the interaction of the adsorbed proteins with cells, which consequently affects the cell attachment, proliferation and differentiation, and determines the integration of the implants with surrounding tissues. Therefore, control of the surface chemistry, which can control the protein adsorption, and the micro-nano surface topography of BGs are two hot topics and will attract continued research interests. In particular, the combined effects of surface chemistry and topography will be investigated to optimize the surface properties of BGs. In addition, with the development of tissue engineering techniques, control of degradation rate of tissue engineering scaffolds is one of the problems that needs to be solved. Therefore, study on controlling the degradation of BGs by surface modification is also important.

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In developing clinical applicable orthopedic implant materials, although the techniques developed so far for BGs’ related surface modification have shown encouraging results, problems such as poor adhesion, cracking and lack of bioactivity still exist. From the view of composition, BG coatings with a silica content higher than 60 wt% have similar thermal expansion coefficients to Ti and Ti6Al4V alloys, which showed good mechanical stability but low bioactivity (Tanaskovic et al., 2007), while BG coatings with silica content below 60 wt% revealed high bioactivity (Lopez-Esteban et al., 2003) but poor mechanical properties. Functional graded BG coatings seem to be the right approach to solving this problem in the future, if the gradient composition and micro structure can be further precisely controlled. Therefore, more attention needs to be paid on the development of new technique to fabricate precisely controlled gradient coatings. Coating stability is a key issue, which determines the applicability of the coated materials. While increasingly advanced surface modification technologies are being introduced to improve the properties of titanium and Ti alloy, the long-term use of these coatings is questionable. Thus, the long-term stability evaluation of BG coatings on metal implants is yet to be established for clinical applications.

2.6

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properties of poly(L-lactide) composites’. Acta Biomater, 4, 1005–1015. DOI: 10.1016/j.actbio.2008.02.013 Liu X. Y., Morra M., Carpi A. and Li B. (2008b), ‘Bioactive calcium silicate ceramics and coatings’. Biomed Pharmacother, 62, 526–529. DOI: 10.1016/j.biopha.2008.07.051 Lobel K. D. and Hench L. L. (1996), ‘In vitro protein interactions with a bioactive gelglass’. J Sol-gel Sci Techn, 7, 69–76. DOI: 10.1007/BF00401885 Lopez-Esteban S., Gutierrez-Gonzalez C. F., Gremillard L., Saiz E. and Tomsia A. P. (2009), ‘Interfaces in graded coatings on titanium-based implants’. J Biomed Mater Res A, 88A, 1010–1021. DOI: 10.1002/jbm.a.31935 Lopez-Esteban S., Saiz E., Fujino S., Oku T., Suganuma K. and Tomsia A. P. (2003), ‘Bioactive glass coatings for orthopedic metallic implants’. J Eur Ceram Soc, 23, 2921–2930. DOI: 10.1016/S0955-2219(03)00303-0 Lu H. H., Pollack S. R. and Ducheyne P. (2001), ‘45S5 bioactive glass surface charge variations and the formation of a surface calcium phosphate layer in a solution containing fibronectin’. J Biomed Mater Res, 54, 454–461. DOI: 10.1002/10974636(20010305)54:3<454//AID-JBM200>3.0.CO;2-H Lusvardi G., Malavasi G., Menabue L., Aina V. and Morterra C. (2009), ‘Fluoridecontaining bioactive glasses: Surface reactivity in simulated body fluids solutions’. Acta Biomater, 5, 3548–3562. DOI: 10.1016/j.actbio.2009.06.009 Ma J., Chen C. Z., Yao L. and Bao Q. H. (2006), ‘Characterization of some methods of preparation for bioactive glass coating on implants’. Surf Rev Lett, 13, 93–102. DOI: 10.1142/S0218625X06007858 Maquet V., Boccaccini A. R., Pravata L., Notingher I. and Jerome R. (2004), ‘Porous poly(alpha-hydroxyacid)/Bioglass® composite scaffolds for bone tissue engineering. I: preparation and in vitro characterisation’. Biomaterials, 25, 4185–4194. DOI: 10.1016/j. biomaterials.2003.10.082 Martorana S., Fedele A., Mazzocchi M. and Bellosi A. (2009), ‘Surface coatings of bioactive glasses on high strength ceramic composites’. Appl Surf Sci, 255, 6679–6685. DOI: 10.1016/j.apsusc.2009.02.069 Misra S. K., Mohn D., Brunner T. J., Stark W. J., Philip S. E., Roy I., Salih V., Knowles J. C. and Boccaccini A. R. (2008), ‘Comparison of nanoscale and microscale bioactive glass on the properties of P(3HB)/Bioglass® composites’. Biomaterials, 29, 1750– 1761. DOI: 10.1016/j.biomaterials.2007.12.040 Montanaro L., Arciola C. R., Campoccia D. and Cervellati M. (2002), ‘In vitro effects on MG63 osteoblast-like cells following contact with two roughness-differing fluorohydroxyapatite-coated titanium alloys’. Biomaterials, 23, 3651–3659. DOI: 10.1016/S0142-9612(02)00098-4 Moritz N., Rossi S., Vedel E., Tirri T., Ylänen H. O., Aro H. T. and Närhi T. (2004a), ‘Implants coated with bioactive glass by CO2-laser, an in vivo study’, J Mater Sci Mater Med, 15, 795–802. DOI: 10.1023/B:JMSM.0000032820.50983.c1 Moritz N., Vedel E., Ylänen H. O., Jokinen M., Hupa M. and Yli-Urpo A. (2004b), ‘Characterization of bioactive glass coatings on titanium substrates produced using a CO2 laser’, J Mater Sci Mater Med, 15, 787–794. DOI: 10.1023/ B:JMSM.0000032819.64994.42 Navarro M., Michiardi A., Castano O. and Planell J. A. (2008), ‘Biomaterials in orthopaedics’. J R Soc Interface, 5, 1137–1158. DOI: 10.1098/rsif.2008.0151 Neouze M. A. and Schubert U. (2008), ‘Surface modification and functionalization of metal and metal oxide nanoparticles by organic ligands’. Monatshefte Fur Chemie, 139, 183–195. DOI: 10.1007/s00706-007-0775-2

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Pazo A., Saiz E. and Tomsia A. P. (1998), ‘Silicate glass coatings on Ti-based implants’. Acta Mater, 46, 2551–2558. DOI: 10.1016/S1359-6454(98)80039-6 Pryce R. S. and Hench L. L. (2004), ‘Tailoring of bioactive glasses for the release of nitric oxide as an osteogenic stimulus’. J Mater Chem, 14, 2303–2310. DOI: 10.1039/ b400922c Radin S., Ducheyne P., Falaize S. and Hammond A. (1997), ‘Transformation of bioactive glass granules into Ca-P shells in vitro’. Bioceramics, 10, 45–48. Radin S., Ducheyne P., Falaize S. and Hammond A. (2000), ‘In vitro transformation of bioactive glass granules into Ca-P shells’. J Biomed Mater Res, 49, 264–272. DOI: 10.1002/(SICI)1097-4636(200002)49:2<264//AID-JBM16>3.0.CO;2-2 Radin S., Reilly G., Bhargave G., Leboy P. S. and Ducheyne P. (2005), ‘Osteogenic effects of bioactive glass on bone marrow stromal cells’. J Biomed Mater Res A, 73A, 21–29. DOI: 10.1002/jbm.a.30241 Rahaman M. N., Yadong L., Bal B. S. and Wenhai H. (2008), ‘Functionally graded bioactive glass coating on magnesia partially stabilized zirconia (Mg-PSZ) for enhanced biocompatibility’. J Mater Sci Mater Med, 19, 2325–2333. DOI: 10.1007/s10856-007-3328-7 Ramaswamy Y., Wu C. T., Zreiqat H. (2009). ‘Orthopedic coating materials: considerations and applications’. Expert Rev Med Devic, 6, 423–430. Ramires P. A., Wennerberg A., Johansson C. B., Cosentino F., Tundo S. and Milella E. (2003), ‘Biological behavior of sol-gel coated dental implants’. J Mater Sci Mater Med, 14, 539–545. DOI: 10.1023/A:1023412131314 Rich J., Jaakkola T., Tirri T., Narhi T., Yli-Urpo A. and Seppala J. (2002), ‘In vitro evaluation of poly(epsilon-caprolactone-co-DL-lactide)/bioactive glass composites’. Biomaterials, 23, 2143–2150. DOI: 10.1016/S0142-9612(01)00345-3 Saleema N. and Farzaneh M. (2008), ‘Thermal effect on superhydrophobic performance of stearic acid modified ZnO nanotowers’. Appl Surf Sci, 254, 2690–2695. DOI: 10.1016/j. apsusc.2007.10.004 Schrooten J., Van Oosterwyck H., Vander Sloten J. and Helsen J. A. (1999), ‘Adhesion of new bioactive glass coating’. J Biomed Mater Res, 44, 243–252. DOI: 10.1002/ (SICI)1097-4636(19990305)44:3<243//AID-JBM2>3.0.CO;2-O Shi W., Kamiya A., Zhu J. and Watazu A. (2002), ‘Properties of titanium biomaterial fabricated by sinter-bonding of titanium/hydroxyapatite composite surface-coated layer to pure bulk titanium’. Mat Sci Eng A-struct, 337, 104–109. DOI: 10.1016/S09215093(02)00003-5 Silva G. A., Costa F. J., Coutinho O. P., Radin S., Ducheyne P. and Reis R. L. (2004), ‘Synthesis and evaluation of novel bioactive composite starch/bioactive glass microparticles’. J Biomed Mater Res A, 70A, 442–449. DOI: 10.1002/jbm.a.30099 Slowing, Ii, Vivero-Escoto J. L., Wu C. W. and Lin V. S. Y. (2008), ‘Mesoporous silica nanoparticles as controlled release drug delivery and gene transfection carriers’. Adv Drug Deliver Rev, 60, 1278–1288. DOI: 10.1016/j.addr.2008.03.012 Stan G. E., Morosanu C. O., Marcov D. A., Pasuk I., Miculescu F. and Reumont G. (2009), ‘Effect of annealing upon the structure and adhesion properties of sputtered bio-glass/ titanium coatings’. Appl Surf Sci, 255, 9132–9138. DOI: 10.1016/j.apsusc.2009.06.117 Stroganova E. E., Mikhailenko N. Y. and Moroz O. A. (2003), ‘Glass-based biomaterials: Present and future (a review)’. Glass Ceram, 60, 315–319. DOI: 10.1023/ B:GLAC.0000008235.49161.32 Sun J., Li Y. S., Li L., Zhao W. R., Gao J. H., Ruan M. L. and Shi J. L. (2008), ‘Functionalization and bioactivity in vitro of mesoporous bioactive glasses’. J Noncryst Solids, 354, 3799–3805. DOI: 10.1016/j.jnoncrysol.2008.05.001

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Supova M. (2009), ‘Problem of hydroxyapatite dispersion in polymer matrices: a review’. J Mater Sci Mater Med, 20, 1201–1213. DOI: 10.1007/s10856–009-3696–2 Takahashi Y., Yamamoto M. and Tabata Y. (2005), ‘Enhanced osteoinduction by controlled release of bone morphogenetic protein-2 from biodegradable sponge composed of gelatin and beta-tricalcium phosphate’. Biomaterials, 26, 4856–4865. DOI: 10.1016/j. biomaterials.2005.01.012 Tanaskovic D., Jokic B., Socol G., Popescu A., Mihailescu I. N., Petrovic R. and Janackovic D. (2007), ‘Synthesis of functionally graded bioactive glass-apatite multistructures on Ti substrates by pulsed laser deposition’. Appl Surf Sci, 254, 1279–1282. DOI: 10.1016/j.apsusc.2007.08.009 Tanaskovic D., Veljovic D., Petrovic R., Janackovic D., Mitric M., Cojanu C., Ristoscu C. and Mihailescu I. N. (2008), ‘Double-layer bioactive glass coatings obtained by pulsed laser deposition’. Bioceramics, 20, 361–363. Tilocca A. (2009), ‘Structural models of bioactive glasses from molecular dynamics simulations’. Pro R Soc A, 465, 1003–1027. DOI: 10.1098/rspa.2008.0462 Valimaki V. V., Yrjans J. J., Vuorio E. I. and Aro H. T. (2005), ‘Molecular biological evaluation of bioactive glass microspheres and adjunct bone morphogenetic protein 2 gene transfer in the enhancement of new bone formation’. Tissue Eng, 11, 387–394. DOI: 10.1089/ten.2005.11.387 Vallet-Regi M. (2006), ‘Ordered mesoporous materials in the context of drug delivery systems and bone tissue engineering’. Chem-Eur J, 12, 5934–5943. DOI: 10.1002/ chem.200600226 Vallet-Regi M., Balas F. and Arcos D. (2007), ‘Mesoporous materials for drug delivery’. Angew Chem Int Edit, 46, 7548–7558. DOI: 10.1002/anie.200604488 Vallet-Regi M., Ragel C. V. and Salinas A. J. (2003), ‘Glasses with medical applications’. Eur J Inorg Chem, 1029–1042. DOI: 10.1002/ejic.200390134 Verne E., Bretcanu O., Balagna C., Bianchi C. L., Cannas M., Gatti S. and Vitale-Brovarone C. (2009a), ‘Early stage reactivity and in vitro behavior of silica-based bioactive glasses and glass-ceramics’. J Mater Sci Mater Med, 20, 75–87. DOI: 10.1007/s10856-0083537-8 Verne E., Brovarone C. and Moisescu C. (2005), ‘Glazing of alumina by a fluoroapatitecontaining glass-ceramic’. J Mater Sci, 40, 1209–1215. DOI: 10.1007/s10853-005-69398 Verne E., Ferraris S., Vitale-Brovarone C., Spriano S., Bianchi C. L., Naldoni A., Morra M. and Cassinelli C. (2010), ‘Alkaline phosphatase grafting on bioactive glasses and glass ceramics’. Acta Biomater, 6, 229–240. DOI: 10.1016/j.actbio.2009.06.025 Verne E., Valles C. F., Brovarone C. V., Spriano S. and Moisescu C. (2004), ‘Double-layer glass-ceramic coatings on Ti6Al4V for dental implants’. J Eur Ceram Soc, 24, 2699– 2705. DOI: 10.1016/j.jeurceramsoc.2003.09.004 Verne E., Vitale-Brovarone C., Bui E., Bianchi C. L. and Boccaccini A. R. (2009b), ‘Surface functionalization of bioactive glasses’. J Biomed Mater Res A, 90A, 981–992. DOI: 10.1002/jbm.a.32153 Vitale-Brovarone C. and Verne E. (2005), ‘SiO2-CaO-K2O coatings on alumina and Ti6Al4V substrates for biomedical applications’. J Mater Sci Mater Med, 16, 863–871. DOI: 10.1007/s10856-005-3583-4 Wang D. and Bierwagen G. R. (2009), ‘Sol-gel coatings on metals for corrosion protection’. Prog Org Coat, 64, 327–338. DOI: 10.1016/j.porgcoat.2008.08.010 Wang S. B. (2009), ‘Ordered mesoporous materials for drug delivery’. Micropor Mesopor Mat, 117, 1–9. DOI: 10.1016/j.micromeso.2008.07.002

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3 Cell interaction with bioactive glasses and ceramics R. P. K. PENTTINEN, University of Turku, Finland

Abstract: This chapter reviews the biochemical and other interactions of cells with bioactive glass (BAG) and related ceramics. In recent years material development has been directed from melted glasses and their composites towards calcium-rich sol-gel glasses. It has also become apparent that for the clinical success of the material, its ‘bioactivity’, the generation of Si-rich, carbonated apatite layers in simulated body fluid, should correlate with favourable reactions in tissues and not only in cultured cells. At the end of the chapter some future aspects are presented, e.g. on the use of bioactive glasses in tissue engineering, gene therapy and drug administration. Key words: silica-rich apatite, sol-gel bioactive glasses, osteogenesis, silica toxicity, drug administration.

3.1

Introduction

Bioactive glasses (BAGs) (first presented by Hench in 1971; Rehman et al., 1994; Hench et al., 2004) are tissue-friendly, simple and non-toxic; they react with simulated body fluid (SBF) to create a hydroxyapatite layer, which adheres to bone and supports the growth of bone cells (Hata and Kokubo, 1995; Ducheyne and Qiu, 1999; Gough et al., 2004a, 2004b; Karlsson, 2004). Many functional bioactive glasses have been developed with different reaction speeds and porosity, and their dissolution products have a proven capacity to maintain the growth of osteoblasts, fibroblasts (Hattar et al., 2002; Verné et al., 2009; Alcaide et al., 2010) and chondroblasts (Aho et al., 2004; Bal et al., 2010). When implanted into bone defects, all bioactive glasses increase bone formation (proliferation of osteoblasts, alkaline phosphatase (ALP) activity, mineral deposition, and bone-specific protein expression (Schepers et al., 1991; Bosetti et al., 2003)). All those activities are clearly carried out as responses of the cells to the reaction components of the glasses. The osteoactivity correlates with the solubility of the glasses (e.g. Xynos et al., 2000), which suggests that at least some of the effectors are soluble products from the glasses. All components of bioactive glasses, except silica, are present in high concentrations in normal plasma. Thus silica can be considered to have an active role in the enhancement of osteoblastic differentiation. On the basis of numerous well-controlled studies it seems likely that the dissolution of bioactive glasses creates favourable local calcium and silicate concentrations for precipitating new apatite layers, either in solution or on almost any surface, including the old bone, 53 © Woodhead Publishing Limited, 2011

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in which the osteogenic cells then become buried. The earliest bioactive glasses were manufactured by melting the components. Recently the focus in silica-based glasses has been moved to sol-gel techniques that allow glass preparation at close to room temperature, and in gentle conditions (Arcos and Vallet-Regi, 2010). Since sol-gel glasses are also bioactive, silicate ions may have a function in the apatite precipitation (e.g. Pereira et al., 1994; Costantini et al., 2008) and calcium ions might cause the intracellular signalling events that together lead to the activation of selected genes.

3.2

Biology of bioactive glasses

3.2.1 Testing bioactive glasses with cells Various cells have been used to test the biocompatibility of bioactive glasses. Stable cell lines such as murine MC3T3-E1, 3T3-L1, human Saos-2, HeLa, Panc 10.05, etc. have been used because they are easy to handle, grow well and have been thoroughly characterized. The disadvantage of cancerous or transformed cells is that their biochemical reactions do not necessarily correspond to those occurring in normal tissues and many of them have lost hormonal or cytokine control. The results obtained with those cells cannot be interpreted as clinically valid. Standardized primary human cell cultures followed by in vivo animal experiments should be utilized in serious efforts to develop materials for medicine. The mechanism by which bioactive glasses promote osteoblastic activity in bone formation is not exactly known in spite of intensive research for over 20 years. In contact with water, SBF, or with the extracellular fluid in vivo, bioactive glasses react with a complex mechanism described in detail elsewhere in this book. This reaction, called ‘bioactivity’, is somewhat misleading because it occurs passively without any biological or live effectors, cells or macromolecules, takes place in all tissues (Andersson and Karlsson, 1991; Gatti et al., 1994) and depends on the pH and ionic strength of the environment (Cerruti et al., 2005; Chou et al., 2004). The melt-derived glasses release Na, K, Ca, phosphate and other ions depending on the components used in their manufacturing and generate silanol (SiOH) groups (e.g. Anderson et al., 1998). If calcium ions are abundant, the glass granules are rapidly coated with a thickening calcium phosphate layer resembling biological hydroxyapatite. The granules are finally separated into a silica-rich outer calcium phosphate layer and an inner silica gel compartment (Fig. 3.1). Large particles usually crack and allow cations of the inner part to leach, leaving a silica gel containing shell (Radin et al., 2000; Fig. 3.1). All this occurs without any cellular activity. Cells attached to apatite activate genes for bone ECM synthesis. The result is a mineralized collagen matrix that integrates with the pre-existing apatite layer as new osteoid. Cells invade the inner ‘empty’ silica gel shells and produce both bone and fibrous tissue. When glass particles are implanted in soft tissues with a rich vasculature, they turn completely

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3.1 Reaction of BAG granules (particle size 140 to 315 µm) implanted into rat femur. The cracks (arrows) and shell-form apatite surface (arrowheads) are clearly seen (courtesy of Dr. Matis Märtson, Tallinna Lastehaigla, Tallinn, Estonia).

to silica gel, which is slowly dissolved. In animal bone implants and after human sinus operations, remnants of BAG can be detected for several years (Aitasalo and Peltola, 2007; Peltola et al., 2008a). This means that melted glass is not removed by osteoclastic or other lytic cellular activity. Variations in the chemical structure of individual glasses alter the rate of the solubilization process. Hench (1994, 1998) divided the bioactive materials into osteoproductive (class A) and osteoconductive (class B) glasses. Those which contain less than 52 mol% SiO2 react rapidly. Based on Si release speed, apatite formation, ALP activity, transforming growth factor (TGF) beta release from cells in contact with the materials and bone bonding, Hench suggested that class A materials elicit both intracellular and extracellular responses at their interfaces whereas glasses between 52 and 58 mol% SiO2 react less vigorously (class B substances) and elicit only extracellular reactions. Melted glasses over 60 mol% SiO2 do not have bioactivity. Thus one should expect that cell membrane receptor(s) or ion channels are involved in the generation of the glass A responses. It has been suggested that dissolved calcium from BAG 60S increases the intracellular Ca++ concentration of osteoblasts and modulates their calcium-dependent glutamate release through the activation of an inositol 1,4,5-triphosphate and a ryanodine receptor (Valerio et al., 2009). If this concept is true the function of BAGs on bone cells resembles neural transmission. Bioactive glasses made by sol-gel techniques

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(Hench, 1997) give more flexibility and allow preparation of a variety of glass composites that can be tested in cell cultures for their biocompatibility. The techniques allow incorporation of biologically active molecules in the glass matrix. Most sol-gel glasses react more rapidly than melt-derived glasses (Li et al., 1992, 1995). Surface coating with biomimetic or other apatites (Kokubo et al., 1990) has become a widely accepted means to improve the association of metal prostheses with bone. In cell culture experiments osteoblast adhesion and proliferation on apatite has been shown to be rather low, however, if compared with tissue culture polystyrene (Puleo et al., 1991; Chou et al., 2005). These observations have often been neglected when apatite-based materials are studied or recommended as osteoblast growth scaffold. Another reason for good osteointegration might be the similarity of the bone mineral and crystalline apatites that grow in a natural way. Apatites adsorb proteins easily, including structural ECM proteins like collagens and fibronectin, growth factors and cytokines on their surfaces (Ekholm et al., 2005). Integrins (García et al., 1998) and other cell surface receptors, e.g. calcium dependent cadherins, bind these proteins (Siebers et al., 2005), transmit the signalling into the cytoskeleton using at least the MAP kinase pathway, activate cell divisions and induce the expression of specific genes (Anselme, 2000; Xynos et al., 2000, 2001; Hattar et al., 2002; Asselin et al., 2004; Christodoulou et al., 2005, 2006; Leonardi et al., 2010; Au et al., 2010). These PCR or microarray studies demonstrated a large number of genes that were either up- or downregulated by the exposure of BAG or BAG extracts. Klapperich et al. (2004) demonstrated activation of granulation tissue growth related genes (cell signalling, extracellular matrix (ECM) remodelling, inflammation, angiogenesis and hypoxia) in cells on the collagen–GAG mesh. Anderson et al. (2004) and Hook et al. (2010) reviewed array techniques to screen various biomaterials for their suitability and presented examples.

3.2.2 Bioactivity and biocompatibility The attachment, proliferation, differentiation and biosynthetic reactions of cells on BAG-derived apatite surfaces have been topics of numerous in vitro studies since BAGs were first presented. The main conclusion of those studies is that the cells attach, grow and differentiate on such surfaces to osteoblastic direction. Apatites derived biomimetically from BAGs contain silica, which is not present in high concentrations in the natural apatites. The exact role of silica in bone formation is not known. In BAG composites the onset of cell proliferation depends on the nature of the other compounds, e.g. the acid degradation products of lactic acid polymers buffer the alkalinity of the dissolution products and improve cell survival during the first days (Maquet et al., 2003). Only a few studies on the variation of cell responses towards different glasses have been published. Karpov et al. (2008) showed that 54CaO/40SiO2 sol-gel

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glass favoured osteoblastic and osteoclastic differentiation, whereas 16CaO/80SiO2 favoured osteoblastic differentiation. It could be expected that cellular reactions to various glasses in vivo are similar in principle; they follow the events of normal well-fixed bone fracture healing without the cartilaginous phase. In most published cases the materials that have shown good results in cell cultures have not been tested in vivo. It is impossible in this review to consider all articles from the vast literature on cells and BAGs and related ceramics. An attempt is made to collect selected examples.

3.3

Reaction of cells with glasses and related ceramics

3.3.1 Melted glasses with cultured cells Once the in vitro apatite formation has been detected and characterized, the materials are usually tested in cell cultures. In most BAG studies the focus has been in the formation and careful characterization of the apatite reaction layer. If cell cultures were included, they often were used only to test the toxicity of the material. Detailed biochemical studies on the interaction of the cells with the BAG surface are rare, but Perry et al. (2009) discussed this on silica-binding peptides. Several reports have considered biomaterial–cell interaction and cell adhesion by integrins (Siebers et al., 2005). Variations in the composition and reactivity of BAGs, length of precorrosion, composition of the growth medium and differentiation degree of the test cells create innumerable combinations that would need enormous experimental work to obtain optimal in vitro conditions for tissue engineering of bone. Direct contact with unreacted, melt-derived BAGs is harmful to most cells. When BAG is placed in a cell culture plate the reaction with medium water begins almost immediately. During dissolution in tissues the BAGs release OH− ions, which can be buffered by interstitial fluid anions and proteins. The cells may actively try to buffer the alkaline environment by secreting protons. Optimally, the culture medium should be replaced until the pH does not change before cells are applied to the BAG-containing cultures. Cells attach to various material surfaces by producing focal contacts or adhesion plaques containing integrins. The attachment of cells to BAG depends on the rapidity of the glass reaction because the adhesion base may be dissolved, and stable adhesions are made to apatite. Thus precorroded glass has been used in vivo (El-Ghannam et al., 2004) and also in cell cultures (El-Ghannam et al., 1997; Foppiano et al., 2007). It is interesting that when bone marrow cells containing osteoclasts were cultured on polished BAG S53P4 surfaces the areas covered by the osteoclasts remained smooth and protected from corrosion whereas the rest of the surface areas with other rapidly proliferating bone marrow cells showed increasing signs of erosion (Wilson et al., 2006). In cell culture conditions the reaction of BAG is rapid,

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whereas in bone, where tissue fluid flow is slow, the glass reaction and development of hydroxyapatite layer may last longer. Fibroblasts and related cells adhere well to borosilicate glass bottles that were used for cell cultures before polystyrene flasks became available. Most cultured cells can be detached from the glass surface by trypsin and less efficiently by EDTA or other calcium celating agents. Osteoclasts, however, adhere to glass surfaces with trypsin resistant bonds, which property was used for their enrichment from mixed cell cultures. These facts suggest at least three mechanisms by which cells can adhere to glass, including slow reacting BAGs. In animal experiments it was detected, however, that BAG cylinders attached to soft subcutaneous or muscular tissue of rats failed to show mechanically relevant soft-tissue bonding (Zhao et al., 2008). This result emphasizes the specificity of the BAG effect to osteoblasts. The adhesion depends on the duration of precorrosion, which increases the surface roughness and reduces alkalinity. The adhesive proteins contain a recognition sequence Arg-Gly-Asp (RGD) or its many variations, which are specific for different integrins. Osteoblasts express at least alpha1, 2, 3, 4, 5, 6, alphav, beta1, 3, and 5 integrin subunits (Siebers et al., 2005) but differences might exist between different cells and different glasses. Bone marrow stromal cells used alphav integrins but Saos-2 cells used both alphav and alpha5 integrins for attachment to hydroxyapatite (Kilpadi et al., 2004). The major binding sites of integrin alpha2beta1 to type I collagen have been reported in at least two protein domains: DGEA (Asp-Gly-Glu-Ala) and GFOGER (Gly-Phe-Hyp-Gly-Glu-Arg) of type I collagen alpha1(I) chain. Even though several studies have characterized the effect of DGEA on cultured cells, the role of DGEA in osteoblast signalling and differentiation is presently unclear (Marquis et al., 2008). Fibronectin-cell interaction on osteoblasts is mediated via alpha5beta1 integrin receptors (Moursi et al., 1996, 1997). This binding and signalling has also been claimed to be necessary for the differentiation of osteoblasts. On the other hand, it has been claimed that the Arg-Gly-Asp (RGD) cell-binding domain of fibronectin may inhibit bone formation in organ cultures (Gronowicz and Derome, 1994). In cell culture systems containing serum, and always in vivo, fibronectin is present and might function in the expression of osteoblastic phenotype by binding to reactive surfaces. The exact role of fibronectin in osteoid differentiation is still unclear. Clarke and Revell (2001) characterized the integrin pattern expressed in aseptically loosened prostheses. The main integrin was alpha2beta1. A comprehensive review on osteoblast adhesion on various biomaterials and on the molecules involved has been published (Anselme, 2000). In the following section some examples of glasses, cells and their reactions are reviewed. Four different bone substitutes were compared for their cell growth-maintaining characteristics and synthesis of marker proteins (Mayr-Wohlfart et al., 2001). The cells were grown on porous 7 mm thick material cubes (55% bioactive phosphate glass, tricalcium phosphate, neutralized glass ceramic and solvent-dehydrated

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bone) in 48 well plates for five days. The best results were obtained on plain plastic and the next best on dehydrated bone. It is difficult to estimate the results because exact treatments of the material surfaces and culture conditions were not given. Two granular melt-derived glasses, a reactive glass 45S5 and a non-reactive glass 60S, were treated with a trishydroxymethylaminomethan (THAM) buffer to create a reaction layer on the granules, and their biological effects (osteoid specific protein expression) were compared in MG63 cell cultures. The treated glass granules did not disturb the cell growth in any way, nor were any major differences observed (Hattar et al., 2002). In another study the same glasses, 45S5 and 60S (this time the 45S5 glass ± coated with commercial enamel matrix proteins) were compared in MC3T3-E1 fibroblast cultures. The cells grew well and produced proteins and ALP in all three cultures, the highest values in cultures coated with tooth enamel proteins (Hattar et al., 2005). Josset et al. (2003) used a BAG powder (50% SiO2, 20% Na2O, 16% CaO, 6% P2O5, 5% K2O and traces of Mg and Al oxides, mean particle diameter about 8 µm) and detected only a minor effect on cell proliferation and viability at 48 hours after exposure to culture medium treated with BAG in standard conditions. A very similar study has been carried out with glass 45S5 (Kaufmann et al., 2000). Glasses with various porosities showed no significant differences in cell proliferation and ALP activities. In addition, there were no differences in the growth of bone cells between control and BAG-substituted cultures, but ALP activities were initially lower than in cultures grown on polystyrene and exceeded those at 14 days indicating development of osteoblastic phenotype. Bosetti and Cannas (2005) differentiated bone marrow stem cells with three different granular BAGs (100 to 700 µm; 45S obtained by melting, 58S and 77S obtained by sol-gel technique). Clear osteopromoting effects were reported for the 45S and 77S glasses. 77S glass also inhibited the appearance of multinucleated, tartrate-resistant and acid phosphatase positive osteoclastic cells. Mouse peritoneal macrophages ingested the glass particles, released oxygen radicals (ROS) and cytokine tumour necrosis factor (TNF) alpha (Bosetti et al., 2002). Gough et al. (2004a) produced 45S5 BAG with a tape-casting method using polyvinyl butyral slurry in organic solvent mixture, and compared the osteoblast responses with cells grown on sintered glass. Rapid apatite formation caused increased apoptosis of human femoral head osteoblasts. The authors recommended slower reacting glasses for the optimal growth of cells. Foppiano et al. (2004) cultured MC3T3-E1.4 cells on two polished glasses (55S6P and 50S6P) and detected equally good attachment, proliferation and ALP secretion in cultures grown on both glasses and on titanium alloy but phosphate concentrations were five times higher in cell layers grown on tissue culture plastic indicating the highest mineralization. Cells proliferated best on 55S6P glass. Foppiano et al. (2006) also coated Ti6Al4V alloys with a double layer consisting of BAGs 55Si6P and 61Si6P. The materials were preconditioned in SBF for two weeks, which increased the stability of the coating. In later studies (Foppiano

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et al., 2007) the cells were harvested at seven days and RT-PCR was carried out on cell layer extracts. A doubling of Runx2-mRNA, a bone-specific transcription factor, was observed but type I collagen mRNA expression was reduced by 20%, which could not be explained. Polymer foam technique was used to prepare human bone resembling porous scaffolds from rapidly reacting 13–93 glass (Brink et al., 1997) with a pore size of 100 to 500 µm. MC3T3-E1 cells occupied the pores within six days (Fu et al., 2008). Varanasi et al. (2009) prepared conditioned media (ion extracts) using two BAGs (6P53-b and commercial Bioglass®) and added them with ascorbate to MC3T3 cell cultures. Collagen I synthesis, ALP, osteocalcin and Runx2 expressions were significantly increased at 5 to 12 days, even though the results varied and the expression of type I collagen alpha 1 and 2 chains (coding for the same molecule) differed temporarily. In conclusion, several different melted glasses have been developed and tested in cell cultures for their cell proliferation and osteoid differentiation-enhancing activities. There is a clear improvement in the cell growth and osteoblast differentiation but if compared with normal tissue culture polystyrene the cells usually proliferate more slowly. In recent years most cell culture studies have been carried out with sol-gel glasses that do not generate alkalinity, or glass extracts have been used.

3.3.2 Melted glasses in vivo Melted glasses were tested in animal experiments soon after their discovery. Thus far most clinical applications of silica-based BAGs have been presented for the melted glasses. BAG granules and plates have been used as bone and sinus cavity filler, orbital wall reconstruction and dental extraction socket material (Froum et al., 2002; El-Ghannam et al., 2004; Peltola et al., 2006, 2008a, 2008b; Arcos and ValletRegí, 2010). The success in those operations depends on the bacteriostatic (see later) and bone induction properties of BAG, and careful operation techniques. An important comparative study was carried out by Hamadouche et al. (2001), who prepared 4 × 6 mm cylinders from two sol-gel BAGs (58S and 77S), and implanted them with similar 45S5 Bioglass® cylinders into distal femurs of rabbits. Samples were collected after 3, 12, 24 and 52 weeks. All glasses showed in vivo bioactivity and the sol-gel glasses started to dissolve after 12 weeks. After 52 weeks about 40% of sol-gel glasses had disappeared whereas Bioglass® showed practically no degradation. Osteoclasts and apparent Howship lacunae were noted at the sol-gel-glass surface but not on the 45S5 glass. The same glasses were studied in mouse calvarial osteoblast and rat embryo fibroblast cultures by Silver et al. (2001), who concluded that the effect of glasses is due to medium or tissue alkalinization, which in turn increases collagen synthesis and cross-linking, and hydroxyapatite formation. Gil-Albarova et al. (2005) converted sol-gel glass

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(70% CaO, 30% SiO2) to glass-ceramic at 1100°C and compared it with sol-gel glass heated at 700°C in rabbit femur implants. The ceramics did not show any decay after six months, whereas sol-gel cylinders showed marked erosion. The adhesion of cells to various surfaces depends on the chemical composition: surface charge (Keselowsky et al., 2004), hydrophobicity (e.g. Anderson et al., 2004), topography and roughness (El-Ghannam, 2004). Evidence is accumulating that Ca+2 ions can regulate bone formation and turnover by interaction with plasma membrane Ca-sensitive receptors located on osteoblasts and osteoclasts (Matsuoka et al., 1999; Chang et al., 1999; Brown and Lian, 2008). Alkaline phosphatase may maintain high calcium levels by releasing calcium ions from apatite and organic phosphates. It can be expected that both hematopoietic and mesenchymal stem cells of bone marrow attach to the apatite layer by specific receptors. The hypothetical hydroxyapatite receptor has not been characterized, but one possibility is that the calcium receptor recognizes apatite. It has been shown to act in hematopoietic stem cell niche formation (Drüeke, 2006), and might function in bone marrow cell homing to apatite-coated scaffolds (Tommila et al., 2009). Once the osteogenic cell has adhered to the apatite of the BAG surface, it begins to secrete the organic components of bone ECM, type I collagen, small amounts of other fibrillar collagens, osteonectin, osteopontin, bone sialoprotein, etc., and ALP, which is involved in the generation of apatite. As a result, osteoblasts become buried in the new osteoid containing the apatite and structural proteins secreted and recognized by the cells with their plasma mebrane receptors, such as integrins. Thus the best results from BAG scaffolds are expected when the osteogenic cells are allowed to produce their own environment, in which natural cell-ECM contacts are established by cell receptors. The advantage of the BAG use is that the BAG granules adhere to osteoid without forming any soft connective tissue layer (Hench, 1998). Similar behaviour is seen even on the BAG granules in cellulose composite scaffolds in mineralizing areas (Holmbom et al. unpublished observations). When biomimetically apatite-coated cellulose sponges or porous polylactic acid (PLA)–PCLA scaffolds (coated using the method of Kokubo et al., 1990) were implanted into rat dorsal subcutaneous space, the material induced a vigorous inflammatory reaction (Ekholm et al., 2005; Tommila et al., 2008; Holmbom et al., 2005). Apatite-coated cellulose scaffolds implanted into femur defects induced a similar fibrogenic granulation tissue, and ossified poorly. The reason for the different behaviours of the apatite layer on BAG and apparently similar apatite layer precipitated biomimetically on cellulose in SBF solution from BAG is not known but might be related to the carbohydrate structure of cellulose. It is possible that subtle differences in the structure or surface pattern of apatite (Chou et al., 2004) or the foreign body reaction induced by cellulose are responsible for the variation in the responses. Bioactive glass microspheres were used as bone filler plugs in rabbit tibial condyles (Ylänen et al., 1999). Bone defect repair took place within expected

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time and surprisingly, the cartilage defect was also completely healed. These studies suggest that certain BAGs might have use in repair of accidentally injured cartilage, which is a severe clinical problem in trauma surgery. The ability of glass 45S5 to maintain cartilage phenotype has been shown by Asselin et al. (2004). 45S5 glass was used as substrate for nucleus pulposus cells that express type II collagen of cartilage and intervertebral disc. The phenotype of nucleus pulposus cells was maintained (Gan et al., 2000). Growth augmentation has been reported for cultured intestinal epithelial-like carcinoma cell line Caco-2 and human myofibroblast CCD-18Co cell line cultured in medium conditioned with 45S5 glass particles (diameter 4 µm) and glass-coated surfaces (Moosvi and Day, 2009). Keshaw et al. (2009) made porous 45S5–PLGA microspheres and tested the growth of CCD-18Co cell line in vitro and in vivo. A significant increase was observed in vascular endothelial growth factor (VEGF) secretion over a 10-day period compared with the PLGA spheres without glass. They recommended the microporous spheres for applications in regenerative medicine where tissue augmentation is required. The microporous spheres decreased the viability of cells and adding increasing concentrations of BAG improved the survival rate by up to 10%. Similar results on VEGF secretion were reported by Day et al. (2005) with porous PLGA-BAG composites in L929 cell cultures and subcutaneous implantations in mice.

3.3.3 Reaction of cells with sol-gel glasses and other SiO2 ceramics Sol-gel glasses and ceramics with cultured cells Tetraethyl silicate-based sol-gel techniques nearly revolutionized BAG research. The sol-gel glasses are believed to contain silanol (OH) groups (Perry and Keeling-Tucker, 2000) that nucleate apatites (Miyaji et al., 1998) in simulated tissue fluid (Li et al., 1992), or when implanted in tissues (Li et al., 1995). An excellent review on sol-gel glasses and their recent development has been published by Arcos and Vallet-Regí (2010). Anderson et al. (1998) coated Thermanox® coverslips with TEOS-derived silica gel and cultured human primary osteoblasts on them. The cells took up silica and formed osteogenic nodules at a much higher rate than if they were grown on plain polystyrene. Dieudonné et al. (2002) cultured bone marrow stromal cells on titanium discs with or without coating with sol-gel glass S53P4. The results were variable. Although the alkalinity was increased in culture medium, no cytotoxicity was observed with 45S5 glass that increased collagen synthesis and ALP activity compared with 58S and 77S sol-gel glasses (Bosetti et al., 2002). Gough et al. (2004b) cultured human primary osteoblasts on sol-gel-produced foam-like porous 58S BAG cubes. Both the glass and its DMEM (Dulbecco’s minimal

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essential medium) dissolution products produced mineralized nodules even without dexamethasone and glycerophosphate. Undiluted dissolution fluid was toxic, however, causing cell death, either by apoptosis or necrosis. Osteoblasts and osteogenic mouse embryonic stem cells profited from the dissolution products of bioactive sol-gel glass 58 (Xynos et al., 2000, 2001; Bielby et al., 2004a, 2004b, 2005). Gough et al. (2004c) continued studies on osteoblast attachment to smooth and rough surfaces of 45S5 BAG sol-gel monoliths. The rough surface of monoliths enhanced mineralized nodule formation. Culture time was only 48 hours, however. Han et al. (2007) compared the growth of MC3T3 cells on plain glass, glass functionalized with NaOH, and glass coated with 58S (58 mol% SiO2–38 mol% CaO–4 mol% P2O5) biomimetically coated with hydroxyapatite (HA) and fluoro-HA. No differences in cell viability were observed but the culture time was only 48 hours, which does not allow any detailed conclusions. Various calcium oxide and hydroxyapatite combinations of sol-gel glasses have been developed recently. The combinations avoid many difficulties detected in pure BAGs. They are easier to prepare in standard form, easier to handle, and avoid the alkalinity of BAGs. Alcaide et al. (2010) studied the interaction of mesoporous BAG (85SiO210CaO-5P2O5) with osteoblasts, fibroblasts and lymphocytes. No signs of cytotoxicity were detected, but a mild decrease in cell proliferation rate was obvious due to initial high calcium elution, whereas Si leaching had no effects. Extracts of the glass did not cause any inhibition or activation in the T-cell response. As an example of functionalization two glasses with 45 or 57 mol% SiO2 were grafted with ALP, which retained its biological activity (Verné et al., 2010). Other materials Vrouwenvelder et al. (1993, 1994) cultured rat fetal osteoblasts on BAGs, hydroxyapatite, titanium alloy and stainless steel. The best osteoblastic growth (amount of DNA and ALP activity) was detected in 8-day cultures grown on BAG and titanium-doped BAG. Courteney-Harris et al. (1995) used hard apatite and an apatite coating of Ti-alloy and, based on EM images, observed better growth of rabbit calvarial osteoblasts on coatings than on apatite. One of the first papers on calcium phosphate ceramics compared five different apatites, of which two contained 9% SiO2, and reported that osteoblasts proliferated on all of them but at a slower rate than on tissue culture polystyrene (Knabe et al., 1997). Unexpectedly the worst inhibitors were apatites containing SiO2. These inhibitory results (on CaNaPO4 or CaNaPO4 and MgSiO4) were interpreted as the toxic effects of high phosphate concentrations in the medium. The medium pH was slightly alkaline. However, cellular growth and surface coverage was best with plain glass Thermanox® coverslips. Since then hundreds of papers have been published containing information on the growth of cells on apatite surfaces.

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Ohgushi et al. (1996) cultured bone marrow cells on ceramics containing apatite–wollastonite and used bone Gla protein and ALP as osteogenic markers. The highest expressions were detected on material containing prefabricated apatite coating by the Kokubo method (1990). Fetal chondrocytes have been cultured on apatite–wollastonite glass ceramic and plastic coverslips (Loty et al., 1997). Chondrocytes attached themselves to the material surface by focal contacts that contained vinculin and beta1 integrin. In the control, glass coverslip cultures cells were attached more evenly and the cell spreading was faster. No significant differences between experimental and control cultures were noted in protein synthesis and ALP activity, and there were no specific advantages in using glass– ceramic compared with the coverslip cultures. Matsuoka et al. (1999) made a similar study on different cells (ROS17/2.8 osteogenic cells) in high Ca and Si concentrations and also implanted apatitized and non-treated hydroxyapatite and apatite–wollastonite cylinders in rabbit femoral distal metaphysis. Apatite was detected to be inferior to the composite in both experiments. Distinct stimulatory effects of both Ca and Si concentrations were detected on ALP and osteocalcin expressions. Loty et al. (2000, 2002) cultured fetal rat osteoblasts on apatite–wollastonite glass–ceramic with or without biomimetic treatment for 23 days. Apatite-coated wollastonite induced more ALP and mineralized nodules. In another study discs and granules of 55S BAG and 60S bioinert glasses were used to culture rat calvaria osteoblasts (Loty et al., 2001). Alkaline phosphatase activity and mineralized nodule formation were higher in the 55S cultures, as expected. Gupta et al. (2007a, 2007b) studied the growth of neonatal rat calvarial osteoblasts on calcium phosphate nanocomposites. Two days after the seeding of cells the composite containing c. 20% SiO2, 20% P2O5, 40% CaO and 20% Na2O showed the highest expressions of type I collagen, osteopontin and osteocalcin. On the other hand, Phan et al. (2003) showed negligible differences between gene expression patterns of silica containing phosphates at various Si concentrations and BAG 45S6P. Vitale-Brovarone et al. (2007) prepared sintered porous 45S BAG (CEL2) ceramic, and studied the growth and behaviour of MG-63 cells for 10 and 20 days in preconditioned scaffolds. Mineralizing nodule formation was observed and BMP-2 (bone morphogenetic protein) addition increased their numbers. Silica substituted apatites were created, and found to stimulate cell proliferation in human osteoblast cultures (Zou et al., 2009) even though the differences were small. Silicon-substituted calcium phosphates were reviewed by Bohner (2009). He concluded that there is no experimental evidence that Si ions are released from Si-substituted calcium phosphates at therapeutic concentrations. In addition, there is no study linking the improved biological performance of Si-substituted calcium phosphates to Si release. Phosphate glass ceramic (45P3S26Ca7Mg15Na4K) was studied as a growth substrate for mesenchymal stem cells (Leonardi et al., 2010). Proliferation, especially cell spreading, viability, osteogenic differentiation (ALP

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and collagen I synthesis) was lower than in cultures grown on plastic. Osteocalcin and collagen alpha1(I) chain gene expressions were, however, relatively high indicating some specificity in the osteogenic properties of this glass ceramic. Using related ceramics Sun et al. (2006) studied earlier the proliferation of bone marrow cells on akermanite (Ca-Mg-silicate) and found it better in osteogenic induction than tricalcium phosphate. Composites Composites have been made from melt-derived glass granules and polymer such as polylactide (Boccaccini et al., 2003; Maquet et al., 2003) or polylactide co-glycolide (Day et al., 2005; Keshaw et al., 2009). The latter composite was claimed to increase VEGF secretion in cell cultures but the finding is in doubt because the number of cells also increased. When implanted subcutaneously in mice slightly more granulation tissue grew around the composite implants at six weeks. It was expected that the polymer matrix would degrade first, revealing the bioactive ceramic scaffold for cell attachment and new bone formation. Porosity in the polymer can be achieved using salt leaching, solvent casting, gas expansion or freeze-drying processes (Boccaccini et al., 2003; Maquet et al., 2003). The procedures affect the structures of the pores, the amount of naked or covered BAG surface, and might strongly affect the attachment of cells. The granule sizes and glass content seem to affect the rapidity of the glass reaction and thus the cell attachment (Holmbom et al., unpublished observations). Maquet et al. (2003) prepared porous BAG (45S5, particle size <5 µm) composites with dimethyl carbonate. The pH of the solution in dissolution studies was dependent on the amount of BAG. The functionality of the polymer composites remained open because no cell culture or test animal studies were carried out. The addition of BAG to starch microparticles improved the adhesion of MC3T3-E1 cells and their expression of osteopontin, osteocalcin, Runx2 and type I collagen. Cell proliferation was much slower compared with cells grown on tissue culture polystyrene, however (Silva et al., 2007). Lu et al. (2003) prepared a similar three-dimensional bioactive polymer/BAG composite of PLGA (PLAGA)-45S5 glass. SaOS-2 cells grown on the composite expressed more type I collagen than if grown on plain PLGA or tissue culture polystyrene after two weeks. Thermoplastic poly(caprolactone-co lactide) glass S53P4 was shown to precipitate calcium phosphates on their surfaces (Jaakkola et al., 2004). Haimi et al. (2009) compared proliferation and osteogenic differentiation of adipose mesenchymal stem cells between PLA/BAG and PLA/beta-TCP composites. PLA/BAG composites were weaker than TCP-composites or plain PLA in their proliferative capacity but a little better than plain PLA and clearly better than TCP in directing the differentiation based on ALP activity of the cells. Plain polystyrene surfaces were not tested.

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3.3.4 Dissolution products The dissolution extracts are usually prepared by immersing BAGs in simulated tissue fluid that closely resembles the extracellular fluid. This extract is then added to cell cultures. The time at which the glass is extracted is critical, because early dissolution products contain large amounts of Ca and PO4 that are later precipitated on the glass material as the apatite layer. An improvement has been the use of a cell culture medium, into which 10% fetal calf serum is added before it is used in cultures. It can be argued that the dissolution should be carried out in a 10% serum-containing medium, or in whole plasma to best imitate the in vivo conditions. Osteoblasts and osteogenic mouse embryonic stem cells profit if grown with dissolution products of bioactive sol-gel glass 58S (Xynos et al., 2000, 2001; Bielby et al., 2004b, 2005). Compared with the medium supplemented with dexamethasone, which had previously been used to enhance osteoblast lineage derivation, the glass extracts were as effective in inducing the formation of mineralized nodules by murine ES cells. When glass extracts were used in combination with dexamethasone, a further increase in the number of nodules was observed. Thus an entirely inorganic material is able to stimulate differentiation of ES cells toward a lineage with therapeutic potential in tissue-engineering applications. Considerable amounts of apoptotic cells were detected in undiluted 58S glass foam 24-hour DMEM extract (Gough et al., 2004b). Using microarray techniques Christodoulou et al. (2006) showed that the mitogen activated protein kinase (MAPK) signalling pathway and IGF stimulation are involved in the response of BAG ionic dissolution products.

3.4

Effect of silica on bone formation

3.4.1 Biology of silica The theory that silica may be an important component in bone cell metabolism is based on experiments with chickens (Carlisle, 1972) and rats (Schwarz and Milne, 1972; reviewed by Carlisle, 1988). Many articles on BAGs begin with the words, ‘Evidence is accumulating that silicon plays a major role in bone formation’, but only a few serious papers have been published on the absorption, secretion and biological effects of silica. One theory to explain the effects of silica is that SiO2 acts as the primary nucleation centre in apatite formation (Cho et al., 1996), an evolutionary remnant of the silicon algae exoskeleton, which developed into apatites of the vertebrates. A possibility remains that the effects are due to a codeficiency of an unknown effector. Silica absorption in rats was studied by Adler et al. (1986) who reported that most of intracardially administered radioactive 31Si (sodium silicate) is secreted into urine within four hours. Liver and lung tissues accumulate severalfold concentrations in comparison with blood, and accumulation was detected

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in bone and skin. Silicon excretion from BAG implanted in muscle and bone has been reported (Lai et al., 2002, 2005). The interest in Si research has focused on its beneficial effects on collagen and glycosaminoglycan synthesis, which are critical for bone formation and maintenance, cardiovascular health, and wound healing. Jugdaohsingh et al. (2002) made a detailed study on silicon intake and absorption in humans, and the topic has recently been reviewed (Sripanyakorn et al., 2009; Nielsen, 2009). A clear correlation between silicon intake and urinary excretion was found. This indicates that cells are able to take up silicon and that silica, which is liberated from BAG, is absorbed and may have biological effects. The critical experiment showing that Si is essential for mammals, or defining its biochemical role, is still lacking (Nielsen, 2009). Silica might be important to human health by alleviating aluminum toxicity, and high intakes of Si may also facilitate the absorption or use of copper and magnesium, which are essential for bone growth and maintenance, cardiovascular health, and wound healing (Nielsen, 2009). However, aquaporins, like silicon transporters, have been described in plants (Ma, 2003; Ma et al., 2006; Yamaji et al., 2008; Mitani et al., 2009), and if really present in mammals might be involved in the biology of silicon.

3.4.2 Silica in skeletal health Jugdaohsingh et al. (2008) reported that it was hard to reproduce the grave growth disturbances in skeletal growth, skull deformities, shorter and more flexible limb bones and various skin and tooth abnormalities described by Carlisle (1972) and Schwarz and Milne (1972). Instead, silica deficiency caused only mild skeletal alterations and increased the growth of long bones. Biochemically most impressive were claims on impaired prolyl hydroxylase function (Si as a cofactor for prolyl hydroxylase) and lesser collagen fibril stability. Those findings have been reported thus far only by Jugdaohsingh’s group of researchers, and need confirmation from other investigators. It is also possible that proper digestion of food by chickens needs the grinding function of sand in the gizzard, which might be impaired in silica-deprived diets. Silicon-substituted calcium phosphates have better biological properties than pure apatites in bone formation (Bohner, 2009). This finding has brought BAG and apatite research closer together. Only a few studies have really shown any release of Si from Si-doped scaffolds (Bohner, 2009), and, as in the drug delivery field, the release rate, availability and therapeutic level of silica should be measured. This, however, might not be the case with sol-gel made Ca-doped apatites, phosphates and other ceramics, which are soluble in contrast to Si-doped and other apatites, which are nearly insoluble in the body. As is obvious from data presented above, the role of silica as a maintaining or stimulating agent of connective tissue is not yet fully established. An

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3.2 A cell-mediated mechanism of silica-activated ECM synthesis and fibrosis. Macrophages and monocytes take up and internalize silica in various forms. Intracellular silica activates macrophages, and damages lytic endosomes, mitochondria and plasma membrane. The macrophages secrete growth factors and cytokines, cause inflammation and activate ECM protein genes; new osteoid is laid down and mineralized by ions of the BAG. The final stage of the reaction depends on the physicochemical form of silica.

open question remains whether the effect of silica in BAGs is a cell-mediated process (monocyte-macrophage-linked) or has a direct effect on osteogenic cells (Fig. 3.2). Silica might also act as a passive nucleation agent for apatite as mentioned earlier. Very few papers have been published on the role of silica in the response of cells even though speculations on the osteogenic role of silica are frequently described. Silica has a negative surface charge and glycosylated cell membrane proteins are also negatively charged because of their terminal sialic acid residues (Gilberti et al., 2008). This may be important, for example in the binding of proteins.

3.4.3 Toxicity of silica Mesoporous nanosized silica particles produced from tetraethyl orthosilicate have been topics of recent widespread research interest. Cells, for example mesenchymal

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stem cells, may take up nanoparticles by receptor-independent, non-specific mechanisms and express aspect ratio and dose-dependent cytotoxicity estimated from apoptosis, actin fibre organization and the adhesive properties of cells (Huang D. M. et al., 2008; Huang X. et al., 2010; Vallet-Regí, 2010; Vivero-Escoto, 2010). The long-term effects, disappearance and degradation of intracellular nanoparticles have not been discussed, however. Some forms of silica are definitely toxic to cells. The toxicity of a given silica preparation depends on the presence of three main properties: physicochemical form (irregular and sharply cut but not necessarily crystalline particles); dangling bonds and the potential to release free radicals and reactive oxygen species (ROS); and the presence of hydrophilic patches (Gilberti et al., 2008; Ghiazza et al., 2010). Chrystalline SiO2, especially with a needle-like or elongated structure, is probably taken up by scavenger receptors of macrophages, which cannot digest it. The three major members of the scavenger receptor family are SR-AI and SR-AII, derived from alternative splicing, and MARCO (macrophage receptor with collagenous structure). The specificity of the scavenger receptor has been questioned, however, because cell death can occur in the absence of these receptors when silicosis is prevented (Gilberti et al., 2008). Ghiazza et al. compared vitreous silica, quartz and monodispersed silica spheres for their toxicity. Vitreous silica, which is totally amorphous, was as cytotoxic as quartz. Both materials show irregular shapes and sharp edges in contrast to silica spheres (Ghiazza et al., 2010). He et al. (2009) detected higher toxicity of mesoporous silica nano- and microparticles in kidney COS cells than in MDA-MB 468 cancer cells indicating cell specificity. Wang et al. (2009) also showed nanoparticle toxicity towards human embryonic kidney cells. Silica dust and other high-aspect ratio particles induce apoptosis or necrosis in macrophages and monocytes, and the dying cells release growth factors and induce fibroblasts to make a capsule surrounding the silica particles (Iyer et al., 1996; Hamilton et al., 2008; Chao et al., 2001). Circulating monocytes attach, mature and replace the dead macrophages and re-ingest the silica, and a vicious cycle is initiated (Fig. 3.2). Nanosized glass components have been made into poly(L-lactide) composite that was claimed to have good osteogenic properties (Liu et al., 2008). Toxicity was not observed and it is possible that the polymer prevents the uptake of nanoparticles before they are degraded. Di Pasqua et al. (2008) showed the remarkable cytotoxicity of oval silica nanoparticles MCM-41 towards human neuroblastoma cells, whereas two of its grafted (aminopropyl and mercaptopropyl) analogues were less toxic. Julien et al. (2010) studied nanowire toxicity and detected cell necrosis to be the main mechanism of cell death instead of apoptosis. Recently, silica nanowires with high aspect ratios were reported to be highly toxic in early zebrafish embryos (Nelson et al., 2010) and in cultured HeLa, 3T3-L1 and Panc 10.05 cells. Interestingly, similar aspect ratio dependent toxicity has been reported for nanosized apaptite preparations (Okada et al., 2010).

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Silica is dissolved in a soluble, non-toxic form from BAGs. This is supported by the results of Ghiazza et al. (2010), who detected that two forms of particulate silica, crystalline quartz and fully amorphous, ‘vitreous’ silica, did elicit a toxic response whereas monodispersed synthetic silica spheres had no harmful effects. The effect of the nanoparticle dimension is not limited to silica particles. In a detailed study Hamilton et al. (2009) reported that the aspect ratio determines the toxicity of titanium oxide nanoparticles. Their results suggest that any modification of a nanomaterial, resulting in a wire, fibre, belt or tube, should be tested for pathogenic potential. The toxicity changes dramatically as the shape of the material is altered into one that a phagocytic cell has difficulty in processing, resulting in lysosomal dysruption. Even though cell culture and in vivo studies seem to show effects of silicic acid on ECM production, the uptake of molecular silica, its intracellular localization and possible biological functions are poorly known. Treatment with Zeolite A, a silicon-containing compound, for 48 hours also enhanced DNA synthesis, ALP activity and osteocalcin release but the experiment was probably too short to observe any increased collagen synthesis (Keeting et al., 1992). In another cell culture study orthosilicic acid exposure for 36 hours did not alter the relative type I collagen gene expression, which also might be too short to induce mRNA synthesis. In contrast, however, the relative abundances of ALP and osteocalcin mRNAs increased significantly as a sign of osteoblastic differentiation (Reffitt et al., 2003). Based on the detection of cleaved type I procollagen C-terminal propeptide, the effect was dose-dependent at 10 and 20 µM concentrations. It is possible that the propeptide processing is inhibited in higher orthosilicic acid concentrations, because collagen protein synthesis increased slightly at 50 µM orthosilicate (Reffitt et al., 2003). This study is important because an established osteogenic cell line (MG-63 cells) without macrophage–monocyte contaminants was used and thus suggests a direct effect of silica on osteoblastic differentiation. In human Saos-2 cell cultures orthosilicate was only slightly cytotoxic in 10-day treatment at the highest soluble concentration (1700 µM). Thereafter colloidal particles were increasingly toxic. On the other hand, dimethylsilanediol, a common environmental contaminant, increased the proliferation of Saos-2 cells in similar concentrations, but the mechanism remains open (Duivenvoorden et al., 2008). Maehira et al. (2009) compared five nutritional calcium and silica sources (calcium carbonate without (control) or with (experiment) 50 ppm added silica, coral sand, eggshell and fossil stony coral) in the biochemical and mechanical properties of mouse bones. Silica addition and coral sand diets had the highest Si concentrations and also increased the strength and stiffness of femurs and the expression of bone and bone differentiation signalling proteins, and decreased the markers of adipogenesis.

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Taken together, it is obvious that silica in the form released from BAGs and in the concentrations created by their gradual dissolution is non-toxic, taken up by cells, secreted into circulation and urine, and may induce an intracellular osteogenic or at least proliferative response. It is not known, however, why silica in insoluble needle forms, sharp-edged quartz or nanosized particles with high aspect ratio can induce toxic and fibrotic reactions.

3.5

Future trends

3.5.1 Antibacterial effect of bioactive glasses Certain bacteria seem to be sensitive to BAGs. This topic has been studied in several papers (Stoor et al., 1998; Leppäranta et al., 2008; Munukka et al., 2007; Zhang et al., 2009; Mortazavi et al., 2010; El-Ghannam et al., 2010). The reason for this clinically important effect might be related to the ROS production of mitochondria. The antibacterial effect can become one of the main advantages of using BAG-based biomaterials in bone surgery.

3.5.2 Large-scale testing The testing of new biomaterials with cultured cells has been tedious and has needed a large number of cultures. A nanoliter scale array system was suggested as a useful alternative, which could test the suitability of hundreds of materials and modifications for human stem cell cultures in a single assay (Anderson et al., 2004; Hook et al., 2010). It remains to be seen whether similar systems could be used for BAG preparations and implanted in test animal tissues.

3.5.3 Bioactive glass materials in drug and gene delivery Sol-gel techniques allow the encapsulation of drugs and even living cells into three-dimensional matrices. About half of the cells stay alive through the process after 72 hours (Nieto et al., 2009). This technique may prove valuable for engineering of specific tissues. An example of drug delivery from sol-gel-prepared glass monoliths is presented in Fig. 3.3 that shows tetracycline fluorescence in tissues. Nanosized silica, if proven non-toxic to cells, could be useful in many clinical applications, for example in targeted cell-specific drug and gene delivery (Lu et al., 2007; Sun et al., 2008; Xiao et al., 2009; He et al., 2010). If silica nanoparticle treatments turn out to be practical in clinical medicine, the toxicity, fate, degradation and secretion of ingested silica particles from the cells and from the body should be resolved or carefully investigated. Obviously more studies with different cells and experiments in vivo are needed.

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3.3 Sol-gel glass monolith as a drug delivery vehicle in rat femoral defect. Tetracycline was incorporated in the monolith during the TEOS precipitation. The fluorescence after 14 days is seen in the surroundings of the glass monolith that does not have any fluorescence. NB = new bone.

3.6

References

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Valerio P., Pereira M. M., Goes A. M. and Leite M. (2009), ‘Effects of extracellular calcium concentration on the glutamate release by bioactive glass (BG60S) preincubated osteoblasts’, Biomed Mater, 4, 45011. July 28. DOI:10.1088/1748–6041/4/4/045011. Vallet-Regí M. (2010), ‘Nanostructured mesoporous silica matrices in nanomedicine’, J Intern Med, 267, 22–43. Review. DOI:10.1111/j.1365–2796.2009.02190.x. Varanasi V. G., Saiz E., Loomer P. M., Ancheta B., Uritani N., Ho S. P., Tomsia A. P., Marshall S. J. and Marshall G. W. (2009), ‘Enhanced osteocalcin expression by osteoblast-like cells (MC3T3-E1) exposed to bioactive coating glass (SiO2-CaO-P2O5MgO-K2O-Na2O system) ions’, Acta Biomater, 5, 3536–3547. DOI:10.1016/j. actbio.2009.05.035. Verné E., Bretcanu O., Balagna C., Bianchi C. L., Cannas M., Gatti S. and Vitale-Brovarone C. (2009), ‘Early stage reactivity and in vitro behavior of silica-based bioactive glasses and glass-ceramics’, J Mater Sci Mater Med, 20, 75–87. DOI:10.1007/s10856–0083537–8. Verné E., Ferraris S., Vitale-Brovarone C., Spriano S., Bianchi C. L., Naldoni A., Morra M. and Cassinelli C. (2010), ‘Alkaline phosphatase grafting on bioactive glasses and glass ceramics’, Acta Biomater, 6, 229–240. DOI:10.1016/j.actbio.2009.06.025. Vitale-Brovarone C., Verné E., Robiglio L., Appendino P., Bassi F., Martinasso G., Muzio G. and Canuto R. (2007), ‘Development of glass-ceramic scaffolds for bone tissue engineering: characterisation, proliferation of human osteoblasts and nodule formation’, Acta Biomater, 3, 199–208. DOI:10.1016/j.actbio.2006.07.012. Vivero-Escoto J. L., Slowing I. I. and Lin V. S. (2010), ‘Tuning the cellular uptake and cytotoxicity properties of oligonucleotide intercalator-functionalized mesoporous silica nanoparticles with human cervical cancer cells HeLa’, Biomaterials, 31, 1325–1333. DOI:10.1016/j.biomaterials.2009.11.009. Vrouwenvelder W. C., Groot C. G. and de Groot K. (1993), ‘Histological and biochemical evaluation of osteoblasts cultured on bioactive glass, hydroxylapatite, titanium alloy, and stainless steel’, J Biomed Mater Res, 27, 465–475. DOI:10.1002/jbm.820270407. Vrouwenvelder W. C., Groot C. G. and de Groot K. (1994), ‘Better histology and biochemistry for osteoblasts cultured on titanium-doped bioactive glass: Bioglass 45S5 compared with iron-, titanium-, fluorine- and boron-containing bioactive glasses’, Biomaterials, 15, 97–106. DOI:10.1016/0142–9612(94)90257–7. Wang F., Gao F., Lan M., Yuan H., Huang Y. and Liu J. (2009), ‘Oxidative stress contributes to silica nanoparticle-induced cytotoxicity in human embryonic kidney cells’, Toxicology in Vitro, 23, 808–815. DOI:10.1016/j.tiv.2009.04.009. Wilson T., Parikka W., Holmbom J., Ylänen H. and Penttinen R. (2006), ‘Intact surface of bioactive glass S53P4 is resistant to osteoclastic activity’, J Biomed Mater Res A, 77, 67–74. DOI:10.1002/jbm.a.30600. Xiao X., He Q. and Huang K. (2009), ‘Novel amino-modified silica nanoparticles as efficient vector for hepatocellular carcinoma gene therapy’, Med Oncol, E-pub, December 1. DOI:10.1007/s12032–009-9359–9. Xynos I. D., Hukkanen M. V., Batten J. J., Buttery L. D., Hench L. L. and Polak J. M. (2000), ‘Bioglass 45S5 stimulates osteoblast turnover and enhances bone formation in vitro: implications and applications for bone tissue engineering’, Calcif Tissue Int, 67, 321–329. DOI:10.1007/s002230001134. Xynos I., Edgar A. J., Buttery L. D. K., Hench L. L. and Polak J. M. (2001), ‘Geneexpression profiling of human osteoblasts following treatment with the ionic products of Bioglas® 45S5 dissolution’, J Biomed Mater Res, 55, 151–157. DOI:10.1002/1097– 4636(200105)55:2<151::AID-JBM1001>3.0.CO;2-D.

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Yamaji N., Mitatni N. and Ma J. F. (2008), ‘A transporter regulating silicon distribution in rice shoots’, Plant Cell, 20, 1381–1389. DOI:10.1105/tpc.108.059311. Ylänen H. O., Helminen T., Helminen A., Rantakokko J., Karlsson K. H. and Aro H. T. (1999), ‘Porous bioactive glass matrix in reconstruction of articular osteochondral defects’, Ann Chir Gynaecol, 88, 237–245. Zhang D., Leppäranta O., Munukka E., Ylänen H., Viljanen M. K., Eerola E., Hupa M. and Hupa L. (2009), ‘Antibacterial effects and dissolution behavior of six bioactive glasses’, J Biomed Mater Res A, 93, 475–483. DOI:10.1002/jbm.a.32564. Zhao D., Moritz N., Vedel E., Hupa L. and Aro H. T. (2008), ‘Mechanical verification of soft-tissue attachment on bioactive glasses and titanium implants’, Acta Biomater, 4, 1118–1122. DOI:10.1016/j.actbio.2008.02.012. Zou S., Ireland D., Brooks R. A., Rushton N. and Best S. (2009), ‘The effects of silicate ions on human osteoblast adhesion, proliferation, and differentiation’, J Biomed Mater Res Part B: Appl Biomater, 90B, 123–130. DOI:10.1002/jbm.b.31262

3.7

Appendix: list of abbreviations

ALP BAG BMP DMEM ECM MAPK PCLA PLA PLGA, PLAGA ROS SBF TGF THAM TNF VEGF

alkaline phosphatase bioactive glass bone morphogenetic protein Dulbecco’s minimal essential medium extracellular matrix mitogen activated protein kinase polycaprolacton polylactic acid polylactic acid–polyglycolide co-polymer reactive oxygen species simulated body fluid transforming growth factor trishydroxymethylaminomethan tumour necrosis factor vascular endothelial growth factor

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4 Regulatory aspects of bioactive glass S. LINDGREN, T. PÄNKÄLÄINEN, J. LUCCHESI and F. OLLILA , BonAlive Biomaterials Ltd, Finland

Abstract: This chapter will survey the regulatory requirements that medical device (including bioactive glass) manufacturers have to follow in order to set their products on the market. The first part provides an introduction to the critical requirements that need to be met to ensure the safety and effectiveness of the product, while the second part concentrates on the specific requirements that relate to different indication areas and the particular risks involved to the patient. The final part will give an overview of the market approval processes in the EU and the USA, together with a summary of authorities and guidelines for several other countries. Key words: bioactive glass, regulatory requirements, medical device classification, market approval process, risk management.

4.1

Introduction

Regulation in the medical field has one objective above all other: to optimize patient safety. As materials in contact with human tissue, bioactive glass products need to comply with regulatory requirements. Bioactive glass products fall under one of the two following regulation categories – medical devices or pharmaceuticals – with the largest part classified as medical devices, i.e. as materials intended to be used for treatment or alleviation of a disease or injury, and which do not achieve their principal intended action in or on the human body by pharmacological, immunological or metabolic means. The following text will survey the regulatory requirements that medical device manufacturers have to follow in order to set their products on the market. These requirements depend on the classification of the product in the geographical area in question. In general, the regulatory requirements of medical devices concern the safety and effectiveness of the product. The product classification criteria may be based, for example, on the vulnerability of the human body that necessitates taking into account the potential risks associated with the technical design and manufacture of the device (EU), or the premarket submission required for clearance to market a product while considering the intended use and risks posed to the patient (USA). The regulatory issues of bioactive glasses are divided here into three main sections. The first one, ‘General requirements’, provides an introduction to the critical requirements that need to be met to ensure the safety and effectiveness of the product. The second section, ‘Indication areas’, concentrates on the specific 85 © Woodhead Publishing Limited, 2011

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requirements that relate to different indication areas and the particular risks involved to the patient. The final section, ‘Market approval process in some geographical areas’, will give an overview of the market approval processes in the EU and the USA. In addition, the authorities and guidelines of several other countries are summarized. The regulatory field of medical devices is at a constant state of change, which consequently requires manufacturers to be on the alert for any forthcoming changes in requirements so as to keep in compliance with them. The resources needed for handling regulatory issues go hand-in-hand with the number of geographical areas covered. Usually the changes aim to improve the safety of the patient and are based on the experiences reported during the implementation of former and present requirements. From a manufacturer’s perspective, regulatory changes usually mean additional and/or more stringent requirements – and ultimately a higher obstacle for product market entry. The general and specific requirements presented in the following chapters mostly involve bioactive glasses. The classification of requirements for medical devices goes according to the risks presented in relation to the use of the device, not according to the material or device as such; therefore no specific bioactive glass-related requirements exist.

4.2

General requirements

The regulatory requirements regarding medical devices are intended to safeguard the health and safety of all parties involved in the utilization of the product – whether the patient, member of the medical staff, or any other user – by ensuring that manufacturers of medical devices follow specified procedures during design, manufacture and marketing. The base requirements are more or less the same for medical devices all over the world, but as the different regulatory authorities have not yet harmonized the assessment and control assigned for medical devices, the procedures for placing medical devices on the market in different countries do vary. However, harmonization of the requirements is ongoing in the EU, the USA, Canada, Japan and Australia (Global Harmonization Task Force, or GHTF). General regulatory requirements for medical devices concern all medical device products regardless of the classification of the device in question. The level of appliance of the requirements differs according to the risk-based classification. For example, when a medical device is meant for only temporary skin contact the requirements are less demanding than in cases where products are implanted permanently.

4.2.1 Quality system An established and implemented quality system is required for all medical device manufacturers. An effective quality system is a way to assure product safety and

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efficacy. Any given quality system comes with a set of general requirements that all devices need to comply with. In addition to these common requirements, there are specific requirements that depend on the risks related to the medical device in question. The requirements for a quality system are defined either with a standard, guidance or by a law. Two main systems for the requirements exist: • •

Standard ISO 13485 defines the requirements, e.g. in Europe; 21 CFR Part 820 Quality System Regulation (QSR) sets the requirements in the USA.

Both systems contain roughly the same general requirements, but there are differences in the nomenclatures and implementation. ISO 13485 is seen as a way to reach Good Manufacturing Practice. Current ISO 13485 specifies the requirements for organizations designing and manufacturing medical devices. Standard ISO 13485 ‘Medical devices – Quality management systems – Requirements for regulatory purposes’ is based on the ISO 9001 ‘Quality management systems – Requirements process model’ approach. The main difference between these two standards is that ISO 13485 requires the quality system to be implemented and maintained, whereas ISO 9001 demands continuous improvement. Other differences concern particular requirements, for example for controlling the work environment or for the validation of medical device processes. The objective of ISO 13485 is to ensure that the medical device organization will meet the regulatory quality management system requirements for medical devices worldwide. Medical device manufacturers who want to sell their products in the USA must comply with section 21 CFR Part 820 of the Code of Federal Regulation. This section of the law, Quality System Regulation (QSR)/Medical Device Good Manufacturing Practice (FDA GMP), requires medical device manufactures to establish and follow quality systems to ensure products consistently meet legal requirements and specifications. In cases where medical devices are marketed both in Europe and in the USA, the requirements of both systems have to be followed. The company must construct a process model quality system fulfilling the requirements of the standard ISO 13485 and a Standard Operation Procedure (SOP)-based quality system fulfilling the requirements of QSR/GMP. Recommendations for the structure of a quality system fulfilling both requirements do exist. To further complicate the establishment of the quality system, it includes such additional requirements as the Vigilance System in the EU (MEDDEV 2.12–1) and Medical Device Reporting (MDR) in the USA. The standard and the law both state the requirements for the quality systems only on a general level, since the requirements concern a huge number of different medical devices and related processes. Each organization is responsible for determining the necessity and extent of the elements. Some elements such as

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service and installation procedures in ISO 13485 can be excluded by the company manufacturing medical devices if no services or installation are provided, as is the case with bioactive glass products. The essential elements of a quality system consist of design, manufacturing, packaging, labeling, storage and delivery of medical devices. In practical terms, these elements include requirements, for example for documentation, resources, training, responsibilities, the organization, purchases, identification, traceability, equipment, validation, acceptance, statistics and customer-related issues. Compliance with the quality system by the company is regularly checked. In Europe the quality system is certified by a Notified Body. In the USA the operations are inspected by the Food and Drug Administration (FDA).

4.2.2 Safety and effectiveness Every medical device has a designed purpose. A device is clinically effective when it produces the effect (performance) intended by the manufacturer relative to the underlying medical condition. Clinical effectiveness is a good indicator of device performance and it is linked to product safety. A medical device that does not perform well could cause safety problems for the patient. For example if a bioactive glass product is intended for load-bearing use, but it does not fulfill the requirements for a load-bearing product (e.g. strength), the safety of the patient is endangered as the defect may collapse. It is important to consider the safety and effectiveness of a medical device simultaneously and in relation to each other. Risk management is an effective tool for total assessment of safety during the life cycle of any medical device. In the EU, the safety and effectiveness data can be collected according to Technical Documentation (NB-MED/2.5.1). In the USA, this technical documentation is called Design Master Record (DMR). The standards presented in the following sections are not compulsory for a medical device organization, but by following the guidelines set by them it is easier to make sure that the corresponding requirements are fulfilled. Risk management Risk management concerns the hazardous situations and harms caused by a medical device or a medical device family during the full life cycle of the device, including all the related activities. Risk management begins during the early development of the device and ends after the device is removed from the market. Risk management is part of the quality system, and defines how the company is to implement the requirements. Standard ISO 14971 ‘Medical devices – Application of risk management to medical devices’ specifies the manufacturer’s requirements for identifying the hazards associated with its medical device

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products, showing how to estimate and evaluate the associated risks, how to control these risks, and how to monitor the effectiveness of the controls. Standard ISO 31000 ‘Risk management – Principles and guidelines’ provides principles and generic guidelines on risk management, and can be applied to a wide range of activities including strategies, processes and products. This standard is intended to harmonize the risk management processes, for example by supporting the standard ISO 14971, which is specific to medical devices. The risk management plan specifies the approach, the management scheme, the components and the resources to be applied during risk assessment. The risk management plan can be written for each product or product family, for a process or project, or for the whole organization; but ultimately it has to cover at least the product(s) of the company. The plan defines the responsibilities, the documentation, the methods used, the criteria for risk estimation, and the risk control and risk/ benefit analysis. In practice the plan defines the courses of action recommended for the whole life cycle of the medical device – not only for its development. Identifying and estimating risks are performed by a team nominated to the task. The members of the team may vary during the life cycle of the device, but it is of most importance that specialists from each phase of the life cycle belong to the team. Consequently, there should be members at least from development, production, quality and marketing, as well as a medical advisor. Failure Mode and Effect Analysis (FMEA) is one method for analyzing potential problems. FMEA can be used to identify potential failure modes, determine their effect and identify actions needed to mitigate these failures. Potential failure modes are any errors or defects in the process, design, or the medical device itself that can affect the patient (or user). Effects are the consequences of failures. FMEAs can be divided into the following types: • • • • • •

System FMEA; global system functions; Design FMEA; device, components and subsystems during development; Process FMEA; manufacturing and assembly processes; User FMEA; issues arising during the use of a device (doctor/patient); Service FMEA; service functions; Software FMEA; software functions.

Only the Design, Process and User FMEAs are applicable to bioactive glass products. As the identified risks need to be mitigated to an approvable level in each development phase (design review), the early and consistent use of FMEAs in the design process ensures that potentially unacceptable risks are designed out. The controls needed to ensure the acceptability of the risks, for example, can be preclinical tests, clinical studies and other performance tests. FMEAs also capture historical information that can be utilized in future product improvements. The continuous use of FMEAs during the post-market phase ensures that necessary preventive actions are identified and implemented.

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Preclinical studies The goal of preclinical studies is to ensure the ultimate safety of a medical device before use in human patients. These studies are a combination of in vitro and in vivo tests for toxicity, biocompatibility and performance of the device. Preclinical studies must be implemented in a facility able to conduct studies in compliance with Good Laboratory Practice (GLP). When performing any in vivo tests, ethical principles for welfare of the animals have to be observed. In vivo studies need approval from the authorities of the country where the test is performed. The development of a medical device usually involves a number of preclinical tests. First tests may be carried out in the research phase to ensure the safety of every new material/component of the device. Since unnecessary in vivo studies are not allowed, a literature review is mandatory to elucidate any existing published data. Material suppliers doing business in the pharmaceutical/medical device industry have often conducted preclinical tests for their products, and this data can potentially be utilized during the development of a new device, which subsequently would decrease the amount of testing needed. Nevertheless, the final product still usually needs preclinical testing before entering clinical studies. Guidelines for defining the need of preclinical studies for a medical device are presented in standard ISO 10993–1: ‘Biological evaluation of medical devices – Part 1: Evaluation and testing within a risk management process’. The requirements for testing depend on the risk level. The standard categorizes products according to the duration of contact with the human body (less than 24 hours, 24 hours to 30 days and more than 30 days) and according to the type of the contact: (1) Surface device (skin/mucosal membrane/breached or compromised surface); (2) External communication device (blood path, indirect/tissue, bone, dentin/circulating blood); and (3) Implant device (tissue, bone/blood). The most comprehensive preclinical studies are required for the devices implanted for the longest duration and with the most potentially harmful contacts. Even though these requirements have been harmonized, interpreting the adequacy of the performed studies may vary in different countries. Different standards and guidelines for the preclinical tests can be applied. The series of ISO 10993 standards 1 to 18 are part of the harmonization process, and are thus generally used for evaluating the toxicology and biocompatibility of medical devices. The OECD test guidelines are also applied at least in parallel to the ISO 10993 standards. In addition, e.g. ASTM (American Society for Testing and Materials) has published a set of preclinical standards. A good example of the large number of different guidance options is the amount of test alternatives for measuring skin sensitization. The most commonly used tests for hypersensitivity are the guinea pig maximization test (GPMT), Closed Patch test (Buehler test) and the Murine Local Lymph Node Assay (LLNA). Standard and guideline alternatives for the skin sensitization are the following: ISO 10993–10 ‘Biological evaluation of medical devices – Part 10: Tests for irritation and skin sensitization’;

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OECD Test Guideline 429 (TG 429) ‘Skin Sensitization: Local Lymph Node Assay’; OECD Test Guideline 406 ‘Skin Sensitization: Description of the Buehler test method’ and ASTM F2148 ‘Standard Practice for Evaluation of Delayed Contact Hypersensitivity Using the Murine Local Lymph Node Assay (LLNA)’. Since bioactive glasses can be utilized in various products, the preclinical tests needed to ensure the safety and effectiveness may also vary considerably. According to ISO 10993–1, bioactive glass products in contact with bone for longer than 30 days require the following preclinical studies: cytotoxicity, sensitization, irritation or intracutaneous reactivity, systemic toxicity, subchronic toxicity, genotoxicity and implantation. In addition, supplementary studies may need to be considered. The required preclinical studies for bioactive glass products in contact with breached skin for less than 30 days are the following: cytotoxicity, sensitization and irritation or intracutaneous reactivity. Clinical evaluation The manufacturer of a medical device must demonstrate that the intended use/ purpose defined for the device, considering safety and performance, is met. To evaluate the compliance to this requirement, clinical data, i.e. safety and/or performance data generated from the use of the device, are needed. The Medical Device Directive (MDD) requires a clinical evaluation for all medical devices regardless of the device class of the product. Detailed guidance for clinical evaluation can be found in MEDDEV 2.7.1 and GHTF SG5/N2R8. A clinical evaluation is a critical examination of clinical data either from the relevant scientific literature or the results of all clinical investigations made, or a combination of both. The clinical evaluation does not necessarily comprise clinical investigation (clinical study), as for low-risk devices the clinical evaluation can be based on a literature review. When utilizing clinical data on comparable devices, the equivalency of the comparable device to the device in question must be justified. For devices in higher-risk classes a clinical study is required. Standard ISO 14155–1 ‘Clinical investigation of medical devices for human subjects – Part 1: General requirements’ defines procedures for the conduct and performance of clinical investigations of medical devices. Before starting a clinical study adequate evidence of the product’s non-clinical safety has to be available. Careful planning of the clinical study is important to ensure the safety of patients and that the objectives of the investigation are fulfilled. The requirements for a clinical study plan are stated in standard ISO 14155–2 ‘Clinical investigation of medical devices for human subjects – Part 2: Clinical investigation plans’. The clinical investigation plan must include a description of the device to be investigated, indications, justification and objectives of the study, the design and methods of the study, study end-points and variables to be measured, patient population and entry criteria, informed consent procedure, the investigational site(s) and investigator(s), responsibilities, labeling, risks,

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adverse event reporting, data collection procedures and data management, statistical methods and publication policy. Prior to commencing a clinical study, approval usually from the health authority and always from the ethics committee is needed in the country where the study is performed. After study completion (or premature termination) a clinical study report is completed. Advice for the content of the clinical study report can be found in ISO 14155–1. It is recommended to register all clinical studies in ClinicalTrials.gov, a registry of clinical studies conducted around the world. The test device used in the clinical study must be manufactured by a validated process. The clinical study itself is the product validation. In case the stated requirements for the clinical study/development of the product are achieved, the clinical study report will be the proof that the product acts according to the specifications. Labeling The labeling of the product communicates information about the medical device to users and/or patients. The information provided on the label can be written, printed or graphical. It includes all the information provided by the manufacturer for the users, i.e. instructions for use, package labels, brochures, and training material. Videos may be used as a way to strengthen the information contained in the message. The information must cover at least the identification, safety and performance of the device. The information needed for identification (e.g. product name, product description, unit size, batch/lot number, expiration date, reference number) and safe use (e.g. intended use, single use, warnings and precautions, residual risk, storage) of the device should be provided with the device. Standards and guidance for the content of package labels and instructions for use are available, e.g. in standard EN 1041 – ‘Information supplied by the manufacturer with medical devices’, in MDD (essential requirements) and in the FDA Device Labeling Guidance. The intended user has to be considered while planning the format, content, readability and location of the labeling on the particular device. Instructions for the professional user that has the required technical knowledge, experience and education should be on a different level than instructions for a lay user without any experience or technical knowledge. Supplemented drawings and diagrams may be used to make the instructions easier to understand. The readability of the information has to consider the language in use, the size of the text (e.g. MDD defines a minimum size for the CE-mark) and the preservation of the printing. The location of the information is important as it may affect how the user handles the products, e.g. the information needed before opening the package (for example about storage conditions) has to be on package labels, not in the ‘instructions for use’ inside the package. The use of symbols on labels is permitted. The use of standardized symbols is encouraged as they improve understanding of the instructions. The requirements

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for explaining a symbol vary from country to another. On product labels marketed in the EU, symbols presented in standard ISO 980 ‘Graphical symbols for use in the labeling of the medical devices’ do not need separate explanation. In the USA, descriptive text may be needed in connection with the symbol. In case nonstandardized symbols or symbols without obvious meaning to the device user are used, an explanation should always be provided. Essential requirements Essential requirements are applicable to products placed on the market in the EU. These requirements are the criteria for freely traded medical devices in the European market. The essential requirements are defined in ‘Annex I of Medical Device Directive’. Australia has similar requirements called Essential Principles. The essential requirements consist of safety and performance standards. They are divided into two parts: general requirements, and requirements regarding design and construction. Items 1 to 6, which discuss the general requirements, concern the safety and performance of the device. The design and construction requirements are dealt with by items 7 (Chemical, physical and biological properties), 8 (Infection and microbial contamination), 9 (Construction and environmental properties), 10 (Devices with measuring function), 11 (Protection against radiation), 12 (Medical devices connected to or equipped with an energy source) and 13 (Information supplied by the manufacturer). Evidence of meeting the essential requirements is documented in a checklist compiled of the requirements of MDD. A reference to documented evidence (test report, record, etc.) is recommended. A medical device has to fulfill the requirements that apply to it and are directly connected to its purpose (e.g. requirements under ‘protection against radiation’ are not applicable to bioactive glass products). The manufacturer may freely decide the method of how to satisfy the requirements. The use of standards may ease the process; for example, the risk-related requirements are fulfilled by applying the standard ISO 14971 ‘Medical Device – Application of risk management to medical devices’.

4.2.3 Post-market surveillance Post-market surveillance (PMS) is a systematic procedure for reviewing the experience gained from the devices at their post-production phase. The level of the requirements for the PMS should be in direct proportion to the risk associated with the device based on its intended use. The PMS system is based on information received from the field. It may include the following: customer feedback (from users, patients, distributors and sales representatives), complaint handling, vigilance, trending, customer surveys, patient cases, reports from authorities, literature reviews, and post-market clinical follow-ups. Most of these

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sources already exist, as they are being routinely generated while the manufacturer follows the general regulatory requirements. Guidance for post-market surveillance can be found in MEDDEV 2.12–2. The PMS system collects the information and analyzes it regularly. The aim of the system is to ensure the continuous safety and performance of the device. The system can provide, for example: • • • •

information about the performance/long-term performance of the device; confirmation of the risk analysis, i.e. the possible harms and their likelihood; feedback on the indications/instructions for use/customer satisfaction/needed product improvements; ideas for further development.

The results of the PMS may be utilized, for example in marketing (long-term safety) or in research and development (new/improved products). Post-Market Clinical Follow-up (PMCF) is performed through clinical studies and/or registries. The target for these studies can be long-term information on the safety and effectiveness of the device or identifying possible emerging risks. All PMCFs have to be planned, and they fall into one of the following categories: a follow-up of patients taking part in a pre-market study, or a prospective study after the device has already been placed on the market. National regulations on postmarket clinical studies need to be followed during PMCF. Guidance on PMCFs can be found in GHTF/SG5/N4.

4.3

Indication areas

Bioactive glasses are widely used in different indication areas. This means that the classification, the nomenclature code and the regulatory requirements may all vary for bioactive glass products. Some products may also be combination products, where the bioactive glass brings an additional element to the original purpose of the product (e.g. in wound treatment). As a rule of thumb, the regulatory requirements are proportional to the level of risk associated with the medical device. The level of control increases when the degree of risk is increased. The risk presented by a particular device depends substantially on its intended purpose and the effectiveness of the risk management techniques applied during design, manufacture and use. The definition of the terms ‘intended purpose’ (EU)/‘intended use’ (FDA) and ‘indication for use’ is described below. The ‘intended purpose’ describes the use for which the device is designed according to the data supplied by the manufacturer on the labeling, in the instructions and/or in promotional materials. The regulatory requirements, that is the essential requirements applying to the intended purpose, must be met. The application of the criteria to classification rules is governed by the intended purpose. The FDA defines ‘intended use’ as ‘the objective intent of

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the persons legally responsible for the labeling of the device’. This covers all aspects of how and for what purposes and under what circumstances the device is intended for use, i.e. the intended use describes the overall picture of use. The term ‘indication for use’ is a narrowed definition within the concept of intended purpose/use having a more precise structure. It includes a description of the disease or condition that the device will treat, prevent, cure or mitigate. A description of the patient population for which the device is intended is also included in the indication for use. The type of device will affect the elements of the statement included in the indications for use. If the device is for therapeutic use the ‘indications for use’ statement should have at minimum the following descriptions: • • • • •

identification of function (e.g. cavity filling); identification of tissue type (e.g. bone); identification of a specific organ (e.g. skeletal system); identification of a particulate target population/disease (e.g. tumor); identification of an effect on clinical outcome (e.g. bone formation).

The manufacturer of the medical device is responsible for ensuring that the medical device is used according to intended purpose/use. The GHTF has recommended each medical device to be allocated to one of four risk classes (GHTF/SG1/N15 Principles of Medical Device Classification). The lowest risk devices are categorized as class A. Class B devices are categorized as moderate low-risk and class C devices as moderate high-risk products. Devices with the highest risk are class D devices. The levels of scrutiny and evidence requirements become more demanding as the risk class of the device increases. This classification follows the product classification of the Medical Device Directive in the EU (classes I, IIa, IIb and III). In the USA, medical devices are separated into three classes (I, II, III). The factors affecting device classification are, for example, the duration of the device in contact with the body (transient, short-term or long-term use), the degree of invasiveness, and possible medicinal products delivered to the patient. Medical device classification in several countries/ areas is presented in Table 4.1. Classified products are usually divided into groups that describe the characteristics of similar products. Defined terms are collected to nomenclatures. The most common nomenclatures are GMDN (Global Medical Device Nomenclature code) and UMDNS (Universal Medical Device Nomenclature System). The main differences in the regulatory requirements of the indication areas concern safety and effectiveness. The key to the requirements is the risk management process. The differences in requirements for three different indication areas (dental applications, implants and wound care) in which bioactive glass products can be used are discussed within the following chapters in accordance with the GHTF guidelines. The classification presented below is of a general nature, and may be altered in individual cases.

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Table 4.1 Medical device classification in different countries/areas Country/area

GHTF EU USA Australia Brazil Canada China India Japan2 Mexico Russia3 Taiwan

Low risk

Moderate low risk

A I1 I I1 I I I A I I 1 I

B IIa

Moderate risk

Moderate high risk

High risk

C IIb

D III III III IV IV III D IV III 3, 4 III

II IIa II II

IIb III III II

B II

C III II

2a

2b II

Notes: 1 Separate requirements for sterile devices and devices with measuring function under class I. 2 Two separate groups in moderate low-risk class (specified controlled MDs and controlled MDs). 3 Separate class 4 for product with high risk to environment, individuals and public health.

4.3.1 Dental applications Dental medical devices are invasive products placed either surgically or nonsurgically into the oral cavity. Medical devices for dental applications are usually seen as lower-risk products than similar other medical devices. The dental medical devices containing bioactive glass are usually classified to either moderate lowrisk (class B) or moderate high-risk (class C) devices depending on the intended use. The level of required preclinical data for dentin devices is the same as for the implantable tissue/bone devices (ISO 10993–1). The need for clinical studies depends on the device and existing clinical data. Implantable dental bioactive glass products used to secure teeth to the bone are class C products (e.g. sinus elevation). Other implantable dental bioactive glass products, such as dental filling materials, are class B products. For example, in Europe an explicit prior authorization with regard to conformity is required for these products before being placed on the market. Elements of the regulatory requirements of dental bioactive glass devices with manufacturer’s and authority’s responsibilities are summarized in Table 4.2.

4.3.2 Implants According to the Medical Device Directive (EU), an implantable device is any device intended to be totally introduced into the human body or to replace an

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Table 4.2 Bioactive glass devices for dental applications and wound care: elements of regulatory requirements with manufacturer’s and authority’s responsibilities Element

Responsibility of manufacturer

Responsibility of authority

Quality management system (QMS)

Class B: Establish and maintain a QMS (full or without design). Class C: Establish and maintain a full QMS.

Ensure appropriate QMS is in place prior to marketing authorization.

Post-market surveillance

Establish and maintain an adverse event reporting procedure.

Ensure appropriate adverse event reporting procedure is in place.

Technical documentation

Class B: Establish Summary Technical Document (STED) and have available for review upon request. Class C: Prepare and submit a STED for review.

Class B: When submission is requested by the authority, a premarket review of the STED for conformity to Essential Principles. Class C: Premarket review of the STED for conformity to Essential Principles.

Declaration of conformity

Establish, sign and ensure availability for review.

Verify compliance with requirements.

epithelial surface or the surface of the eye, by surgical intervention, and which is intended to remain in place after the procedure. Any device intended to be partially introduced into the human body through surgical intervention and to remain in place after the procedure for at least 30 days is also considered as an implantable device. The level of preclinical studies required is high to ensure the safety of the implantable products. A clinical investigation is a requirement for implantable devices unless its use is justified on the basis of existing clinical data. The implantable medical devices that contain bioactive glass are of the highest risk (class D) devices (e.g. bone void fillers). Implantable bioactive glass products are generally implanted for more than 30 days and the products undergo a chemical change in the body or have a biological effect or are mainly absorbed. These products are thus classified as belonging to the most critical devices. For example, in Europe an explicit prior authorization with regard to conformity is required for these products before placed on the market. Elements of the regulatory requirements of implantable bioactive glass devices with manufacturer’s and authority’s responsibilities are summarized in Table 4.3.

4.3.3 Wound care Medical devices that come into contact with injured skin are classified according to intended use as follows: •

a mechanical barrier, for compression or for absorption of exudates;

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Table 4.3 Implantable bioactive glass devices: elements of regulatory requirements with manufacturer’s and authority’s responsibilities Element

Responsibility of manufacturer

Responsibility of authority

Quality management system (QMS)

Establish and maintain a full QMS.

Post-market surveillance Technical documentation

Establish and maintain an adverse event reporting procedure. Prepare and submit a STED for review.

Declaration of conformity

Establish, sign and ensure availability for review.

Ensure appropriate QMS is in place prior to marketing authorization. Ensure appropriate adverse event reporting procedure is in place. An in-depth premarket review of the STED to determine conformity to Essential Requirements. Verify compliance with requirements.

• •

used principally with wounds that have breached the dermis and can only heal by secondary intent; managing the micro-environment of a wound.

Devices acting as a mechanical barrier, for compression or for absorption of exudates, are low risk products (class A). If bioactive glass is added to these lowrisk products (e.g. a bandage) the intended purpose of the product is changed, as bioactive glass does not act as a mechanical barrier or for compression/absorption. The device must therefore be reclassified to either the moderate low (B) or the moderate high-risk class (C). According to the guidance, a device that can be classified according to two or more classes is allocated to the highest class. The products intended for use with wounds that breach the dermis are class C devices. Other wound products are categorized as class B. The required preclinical studies for products in contact with a breached surface are lower than for the dental and implantable devices and the need for clinical studies depends on the device and existing clinical data. For elements of regulatory requirements of bioactive glass products used in wound care see Table 4.2.

4.4

Market approval process in some geographical areas

Medical device manufactures need to comply with the requirements of the country where the product is sold, i.e. the EU, the USA, Canada, Australia, Brazil, Japan, China, Russia and Taiwan all have their own regulatory processes, and these processes are, to make things even more demanding, continuously changing. New regulations are implemented to ensure the safety and effectiveness of the products

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as new information appears. A manufacturer selling products in several countries may have various processes to implement (i.e. in registration and adverse events). One part of the changes is the harmonization of these processes. An example of the harmonization ongoing is the Summary Technical Document (STED) format for regulatory submissions in the USA, Australia, Canada, the EU and Japan developed by the GHTF. New regulations also usually concern products that are already registered, so the manufacturer needs to be aware of the changing demands. The market approval processes of the EU and the USA are presented in more detail, as they can be used in several countries as part of the registration process. In addition, a few examples of the processes in other geographical areas are presented.

4.4.1 European Union The European Medicines Agency (EMEA) has the responsibility of protecting human health by evaluating and supervising medical products. The Council Resolution of 1985 set forth a regulatory scheme leading to technical harmonization and standardization in Europe. From the medical device perspective this ‘New Approach’ harmonization of the regulatory requirements in EU resulted in the Medical Device Directive 93/42/EEC. This directive is designed to ensure that medical devices placed on the market are safe and reliable within the European Economic Area. The directive has been amended several times, the most recent amendment being Medical Device Directive 2007/47/EC. Compliance with the new regulations has been mandatory since March 21, 2010 and the directives have to be implemented as national laws. As bioactive glass products are medical devices, this chapter will concentrate only on the Medical Device Directive (MDD). In addition to the legally binding requirements there are several guidance documents to ensure uniform application of the requirements. Examples of these guidance are MEDDEV and NB-MED documents and consensus statements. Also, the harmonized standards have a significant role in supporting the aims of the directive. The MDD covers a wide range of products. The compliance of a product with the requirements of the directive is declared by placing the CE marking on the product and supplying a Declaration of Conformity with it. The routes of compliance depend on the classification of the product. Products are classified to four categories: class I, class IIa, class IIb and class III. These classes are determined by the risk the device presents. The rules for classification and the alternative conformity assessment procedures for each class are stated in Annex IX of the MDD. The conformity assessment procedures are stated in Annexes II, III, IV, V, VI and VII of the MDD; e.g. while placing the CE mark on an implantable bioactive glass product in class III, the manufacturer can follow procedure stated in Annex II, or in Annex III and Annex V, or in Annex III and IV of the MDD. Class I products (except those that are custom-made or intended for clinical

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investigations) follow the procedure referred to in Annex VII of the MDD. Annex X sets requirements for clinical evaluation. Regardless of the classification of the product the manufacturer has to comply with the essential requirements presented in Annex I of the MDD. The manufacturer is obliged to maintain a technical file for the product or product family and is also responsible for issuing and keeping a file for declarations of conformity for CE-marked devices. All manufacturers must be registered with the competent authority in the member state where they are resident. In case the manufacturer is not based within the EEA, an authorized representative must be appointed. In addition, the medical device may need to be notified to authorities. Class II and class III devices require the involvement of a ‘notified body’ to evaluate the compliance. Notified bodies are organizations designated by the national governments of the EU member state as being competent to approve products, i.e. to make independent judgments about the compliance of a product with the requirements of the MDD. These reviews are generally undertaken at the manufacturer’s facility.

4.4.2 United States The Food and Drug Administration (FDA) is an agency within the US Department of Health and Human Services. It consists of six product centers, one research center and two offices. FDA’s Center for Devices and Radiological Health (CDRH) is responsible for regulating firms who manufacture, repackage, relabel, and/or import medical devices sold in the United States. FDA is responsible for protecting public health by assuring the safety, effectiveness and security, for example of medical devices. Regulatory control for medical devices was added to law (Federal Food, Drug and Cosmetic Act) in 1976. The most recent amendment concerning medical devices is from 2007. All foreignproduced products offered for importation in USA need to be in compliance with the law. In the USA, medical devices are classified into class I, II, and III, with increasing regulatory requirements from class I to class III. Most class I devices are exempt from Premarket Notification 510(k); most class II devices require Premarket Notification 510(k); and most class III devices require Premarket Approval (PMA). Exempted devices can be seen as one classification group with the lowest level of risk and regulatory control. All products requiring 510(k) or PMA are reviewed by FDA before being placed on the US market. FDA also has authority to inspect the facilities of a manufacturer who has a medical device on the US market. The basic regulatory requirements that manufacturers of medical devices distributed in the USA must comply with are:

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Establishment registration (electronically): annual registration of the establishment/facility with the FDA. Medical device listing (electronically): list of medical devices that are made in the registered establishment/facility. Premarket Notification 510(k), unless exempt, or Premarket Approval (PMA): approval of the medical device before being marketed in USA (510(k) approval or PMA). Investigational Device Exemption (IDE) for clinical studies conducted in the USA: for investigational device meant to be used in a clinical study in order to collect safety and effectiveness data required to support a Premarket Approval (PMA) application or a Premarket Notification 510(k) submission to FDA. Quality System (QS) regulation: manufacturing facilities are suspect to FDA inspections to assure compliance with the QS requirements. Labeling requirements: includes labels on the device as well as descriptive and informational literature that accompany the device. Medical Device Reporting (MDR): incidents in which a device may have caused or contributed to a death, serious injury or certain malfunctions must be reported to FDA. The MDR regulation is a mechanism for FDA and manufacturers to identify and monitor significant adverse events involving medical devices. Electronic reporting is preferred.

The Division of Small Manufacturers Assistance provides technical and other non-financial assistance to small manufacturers of medical devices.

4.4.3 Other regions The market approval process in other areas than the EU or the USA may benefit from approvals in the EU and/or the USA. ISO 13485 certification with certificates of MDD and/or 510(k)/PMA approval can be used as the base information in the market approval processes in several other countries. In some countries the requirements are almost equivalent to the EU or the USA. The support of a sponsor, distributor, sales representative etc. should be utilized as the requirements (laws/acts) are often in the national language and only a short guidance version is available in English. The authority, defined guidelines and current link to the requirements in a few geographical areas have been collected to Table 4.4.

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Therapeutic Goods Administration (TGA)

Brazilian National Health Surveillance Agency (ANVISA) Health Canada

Australia

Brazil

The Pharmaceuticals and Medical Devices Agency (PMDA) Federal Commission for Protection of Sanitary Risks (COFEPRIS) Federal Service on Surveillance in Healthcare and Social Development of Russian Federation Swiss Agency for Therapeutic Products

Japan

Pharmaceutical Affairs Law/Guidelines for Registration of Medical Devices

Russian Federal Law on Technical Regulation (2010)/Instruction No. 237 on registration procedures for foreignmade medical equipment and devices Law on Therapeutic Products/ Medical Device Ordinance (MDO)

Therapeutic Goods Act 1989/Australian regulatory guidelines for medical devices (ARGMD) Law No. 6360 (1976), Decree 74.094/97/ Resolution RDC-185 (2001) Food and Drugs Act/Medical Device Regulations 1998 Regulations for the Supervision and Administration of Medical Devices/ Initial registration of import product Medical Devices Regulation Act/ Guidance Document on Common Submission Format for Registration of Medical Devices in India (2010) Pharmaceutical Affairs Law (PAL) Mexican General Health Law

Laws/guidelines

Local representative

http://www.doh.gov.tw/ EN2006/index_EN.aspx

http://www.swissmedic.ch/ index.html?lang=en

http://www.roszdravnadzor.ru/ registration/zarub/contact/ zarub_contact.html

Local office/ distributor/consult

Local office/ distributor

http://www.pmda.go.jp/ english/index.html http://www.cofepris.gob.mx/ wb/cfp/ingles

http://cdsco.nic.in

Manufacturer/ importer/responsible agent Market Authorization Holder (MAH) Local office/distributor (‘Registration holder’)

http://www.anvisa.gov.br/eng/ index.htm http://www.hc-sc.gc.ca/ dhp-mps/md-im/index-eng.php http://eng.sfda.gov.cn/eng/

http://www.tga.gov.au/ industry/devices.htm

Local office/ distributor Manufacturer/ distributor/importer Local office/distributor (‘Legal agent’)

Local office/ distributor ‘Sponsor’

Registration

Caution: This is not an exhaustive list of the factors involved in the country-wise regulatory requirements; the requirements are complex and should be considered carefully in the light of a specific application.

Taiwan

Switzerland

Russia

Department of Health (DOH).

Central Drugs Standard Control Organization (CDSCO)

India

Mexico

State Food and Drug Administration (SFDA)

China

Canada

Authority

Country

Table 4.4 Authority, defined guidelines and current link to the requirements in a few geographical areas

Regulatory aspects of bioactive glass

4.5

103

References

GHTF SG5/N2R8 Clinical Evaluation GHTF/SG5/N4 Post-Market Clinical Follow-Up Studies GHTF/SG1/N15 Principles of Medical Device Classification MEDDEV 2.7.1 Evaluation of clinical data: A guide for manufacturers and notified bodies MEDDEV 2.12–1 The Medical Devices Vigilance System, European Commission guidelines MEDDEV 2.12–2 Guidelines on post-market clinical follow-up NB-MED/2.5.1 Technical documentation NB-MED/2.12 Post-Marketing Surveillance (PMS) 93/42/EEC Medical Device Directive (last amended by 2007/47/EC) 21 CFR Part 801 Labeling 21 CFR Part 803 Medical Device Reporting 21 CFR Part 807 Subpart E Establishment registration and device listing for manufacturers and initial importers of devices 21 CFR Part 814 Premarket Approval of Medical Devices 21 CFR Part 820 Quality System Regulation (QSR) 21 CFR Part 860 Medical Device Classification Procedures FDA Medical Devices, Device Advice: How to Prepare a Traditional 510(k) FDA Medical Devices, Device Advice: Premarket Approval (PMA)

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5 Bioactive glass and glass-ceramic scaffolds for bone tissue engineering X. CHATZISTAVROU, University of Erlangen-Nuremberg, Germany, P. NEWBY, Imperial College London, UK and A. R. BOCCACCINI , University of Erlangen-Nuremberg, Germany and Imperial College London, UK

Abstract: Melt-derived bioactive glasses and glass-ceramics are class A bioactive materials, exhibiting surface bioreactivity in contact with physiological fluids, making them significant in developing bone tissue engineering (BTE) scaffolds. Their biological performance relates to the effect of released ionic dissolution products on osteogenesis and angiogenesis. This chapter discusses the fabrication of three-dimensional, highly porous bioactive glass scaffolds for BTE using the foam replica technique, considering two types of scaffolds in detail: BTE scaffolds fabricated from boron-containing bioactive glass and polymer coated scaffolds of enhanced mechanical behavior, adequate bioactivity and potential drug delivery capability. The presence of a polymer coating leads to significant toughening, and the polymer phase can be used to act as a carrier for biomolecules, growth factors and antibiotics. Optimization of BTE scaffolds is key to the advancement of BTE, and the utilization of bioactive glasses as suitable vehicles for the controlled release of certain metallic ions, which stimulate specific cellular responses, is highlighted as important for future research. New compositions of silicate bioactive glasses incorporating determined therapeutic ions, combined with biopolymers, will become significant in BTE as materials of choice for scaffold developments. Key words: bioactive glass scaffolds, bone tissue engineering, polymer-coated bioactive glass scaffolds, mechanical properties, boron-containing bioactive glasses, drug delivery scaffolds.

5.1

Introduction

Bone tissue engineering (BTE) scaffolds to repair critical-size bone defects are fabricated from bioactive materials that are able to react with physiological fluids to form tenacious bonds to bone. The use of a suitable bioactive scaffold, in combination with relevant cells and signaling molecules, should promote the regeneration of new vascularized bone tissue. Bone regeneration is one of the key areas in the tissue-engineering field attracting considerable research efforts [1–3]. The most common bioactive materials used for BTE scaffolds are bioceramics, including special compositions of silicate glasses (bioactive glasses) and glassceramics, as well as hydroxyapatite (HA) and related amorphous or crystalline calcium phosphates [4]. Developing composite materials for tissue engineering is also an attractive approach [5, 6], since composite properties can be engineered to suit the mechanical and physiological demands of the host tissue. Not only the 107 © Woodhead Publishing Limited, 2011

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combination of the ‘right’ biomaterials but also the structure and morphology of the scaffold, characterized by a highly interconnected, three-dimensional (3D) pore network as well as tailored surface characteristics, determine the suitability of a scaffold for a given application. For BTE, high porosity and pore interconnectivity are required, in order to facilitate the attachment and proliferation of bone cells, the in-growth of new bone tissue and vascularization [7]. Bioactive silicate glasses, originally developed by Hench and co-workers 40 years ago [8], represent very attractive materials for the development of 3D BTE scaffolds due to their proven in vitro and in vivo biological response. The application of bioactive glasses in BTE is an expanding research field, with more than 100 papers being published each year (based on a search in Web of Science® using the combined keywords ‘bioactive glass’ and ‘scaffold’), in comparison to less than 30 published 5 years ago. The specific application of melt-derived bioactive silicate glasses and glassceramics in BTE is the subject of the present chapter. The chapter is organized in the following manner: Section 5.2 includes a concise overview on the requirements for tissue engineering (TE) scaffolds, discussing aspects of materials selection and properties required, Section 5.3 treats specifically the use of silicate bioactive glasses and glass-ceramics in BTE, while Section 5.4 discusses the most common technologies for producing BTE scaffolds from bioactive glasses. Sections 5.5 and 5.6 discuss two cases in point of current research interest, namely boron-containing bioactive glass scaffolds and polymer-bioactive glass composite scaffolds, respectively. The chapter finishes with the conclusion and scope for future research (Section 5.7).

5.2

Requirements for bone tissue scaffolds

The basic scaffold design requirements for BTE have been identified [1, 4, 5, 9], and they are summarized in Fig. 5.1. The different criteria required for TE scaffolds can be split into five areas as summarized below [9]. •

•

•

Biocompatibility: The materials needs to enable cell attachment, proliferation and differentiation once implanted, but they should not produce any toxic by-products and must not cause an inflammatory response once the material is implanted. Osteoconduction and osteoproduction: These properties are particularly important in the case of BTE because they are related to the material ability to bond to bone and the ability to recruit cells surrounding the implantation site to differentiate into new bone growth. This area also looks at the materials ability to provide a structural framework for the formation of new tissue growth. Biodegradability: The material must degrade at a rate that matches the rate of new tissue formation at the implant site, meaning that the rate of material degradation should be controllable if the material is to be used in a range of tissue engineering applications.

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5.1 Requirements for bone tissue engineering scaffolds.

•

•

Mechanical competence: The mechanical properties should closely mimic that of the tissue that the scaffold is trying to replace in order to provide support during the regeneration process. Interconnected porous structure: BTE scaffolds need to have high porosity (>80%) and a pore size >100 µm. These are the preferred conditions for vascularization, cell penetration and tissue in growth [5, 7].

5.3

Bioactive glasses and glass-ceramics in bone tissue engineering

The most widely applied bioactive glasses consist of a silicate network incorporating sodium, calcium and phosphorus in different relative proportions in the SiO2-Na2O-CaO-P2O5 system [8]. The original bioactive glass composition universally known as 45S5 Bioglass® (in wt%: 45% SiO2, 24.5% Na2O, 24.5% CaO and 6% P2O5) has received approval from the US Food and Drug Administration (FDA), and it has found applications in clinical treatments of periodontal diseases as bone filler as well as in middle ear surgery [10]. This glass has also been the subject of intensive research for the development of BTE scaffolds [11]. Other silicate compositions contain no sodium or have additional elements in the silicate network such as fluorine, magnesium, strontium, iron, silver or zinc [12–16]. In addition, a range of silicate glass-ceramics is being investigated to fabricate BTE scaffolds [17–19]. Borosilicate and borate glasses are also receiving increasing attention for the fabrication of BTE [20, 21], as discussed further below. The typical feature common to all bioactive glasses is their positive interaction with living tissues, in particular bone tissue, which is termed ‘bioactivity’ [10, 22,

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23]. Bonding to bone is established by the precipitation of a calcium-deficient, carbonated apatite surface layer on the bioactive glass surface when in contact with relevant physiological fluids [10]. When bioactive glasses are applied in vivo, apatite crystals precipitate in the vicinity of collagen fibrils inducing enhanced bone cell attachment and strong bonding to bone. Both micron-sized and nanoscale bioactive glass particles are considered in BTE [24, 25], and also the fabrication of composite materials, e.g. a combination of biodegradable polymers and bioactive glass [26] is being exploited, as discussed further below. There are several other reasons to select bioactive glasses for BTE scaffolds, which are related to the specific effect of dissolution products of bioactive glasses on cellular behavior [27, 28]. For example, it has been shown that dissolution products from 45S5 Bioglass® upregulate the expression of genes that control osteogenesis [27, 29, 30], leading thus to higher rate of bone formation in comparison to hydroxyapatite [31]. There is also increasing evidence that bioactive glass particles added to a biopolymer construct can enhance the angiogenic potential of the scaffold, i.e. increase the secretion of vascular endothelial growth factor (VEGF) in vitro and the enhancement of vascularization in vivo [32–34]. These studies suggest that Bioglass® BTE scaffolds might stimulate neo-vascularization [34, 35], which is required for the development of large tissue-engineered constructs. Bioactive glasses can also serve as platform for the local delivery of selected ions, which can act to control specific cell functions, for example Co addition to suppress cell hypoxia [36]. This is in addition to the well-known development of bioactive glasses with antibacterial properties, for example incorporating Ag [37, 38], and as a delivery platform for other therapeutic ions for the development of multifunctional scaffolds [39]. The range of bioactive glass compositions and morphologies exhibiting these attractive properties for BTE has increased over the years, with the development of new synthesis and fabrication methods. This includes the development of sol-gel based techniques [40] as well as 3D manufacturing methods for bioactive glass and glass-ceramic scaffolds [11, 17–19], the production of bioactive glass nanofibres [41] and nanoparticles [24, 25, 42] as well as a great variety of composites combining bioactive glass and biopolymers [5, 26].

5.4

Bioactive glass-based scaffolds: fabrication technologies

The use of bioactive glasses for BTE scaffold fabrication requires manufacturing complicated 3D porous shapes and at the same time retention of the specific degree of bioactivity. Bioactive glass and glass-ceramic scaffolds exhibiting highly porous structure are being fabricated by a variety of techniques described in detail elsewhere [4]. The fabrication methods can be broadly divided into two groups; (1) methods relying on glass powders and (2) sol-gel approaches. Only powder-based methods will be considered in this chapter.

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The foam replica technique is one of the most popular methods being applied to fabricate BTE scaffolds from glass powders. It was used for Bioglass® for the first time in 2006 [11] and it is currently one of the methods of choice for fabricating BTE scaffolds leading to foam-like structures of high porosity (>90%) and high pore interconnectivity [43]. The technique involves the use of polymeric sponges as sacrificial templates to prepare cellular bioactive glass (and glass-ceramic) structures of various pore sizes, porosities and chemical compositions. The sacrificial template, e.g. a polyurethane foam, is initially soaked in a glass powder suspension until the foam struts are homogeneously coated with a high concentration of particles. Binders are usually added to the initial suspension in order to prevent cracking of the struts during drying and upon the subsequent heat treatment. In the following step, the polymer template is burnt out by a controlled heat treatment and the glass or glass-ceramic structure is finally densified by sintering at high temperatures. If the fabrication of the scaffold proceeds from a bioactive glass powder, such as in the foam replica method [11], then the viscosity–temperature and crystallization characteristics of the glass must be known because they define the sintering conditions. Depending on the polymer sponge used as template, pore sizes between 200 µm and 3 mm can be obtained. Numerous types of bioactive glass-ceramic [11, 17–19] foams have been produced by the replica method using polyurethane sponges as templates. Figure 5.2a shows the macroscopic pore structure of a Bioglass®-based glass-ceramic scaffold (SEM image) fabricated by this technique [11] whilst Fig. 5.2b is a higher magnification SEM image showing the crosssection of the scaffold struts. Table 5.1 includes a summary of bioactive glass and glass-ceramic scaffolds that have been fabricated by the foam replica technique indicating also the mechanical properties achieved and other typical scaffold characteristics investigated. Another method developed to produce porous scaffolds from bioactive glass powder is the sacrificial porogen technique. This method involves the preparation of a composite comprising a sacrificial phase mixed with glass particles. The sacrificial phase is extracted (usually thermally) from the partially consolidated matrix to generate pores within the microstructure. The mechanical strength of structures made by the sacrificial template method is usually higher than that of scaffolds fabricated by the replica method. However porosity and pore interconnectivity are substantially lower than in scaffolds made by the replica technique. Another advantage of the foam replica method is the possibility of developing scaffolds with graded or layered porosity [47, 48]. These structures are created by modifying the shape of the starting polyurethane sponges as discussed elsewhere [47]. To improve the mechanical properties of highly porous scaffolds made from bioactive glasses, crystallization of the scaffold struts, effectively developing glass-ceramics, is a suitable approach [17–19]. In the case of porous scaffolds extensive sintering of the bioactive glass is required in order to densify the struts

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5.2 SEM images showing the pore structure of a Bioglass® scaffold fabricated by the foam replica method (a) and the microstructure of the strut cross-section (b).

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81.5

85

48

68

54

Calcium silicate (CaSiO3, WT)

Glass 13–93, Composition: 53SiO2-6Na2O-12K2O-5MgCaO-4P2O5 wt%

Sol-gel bioglass powder. Stoichiometric ratio SiO2:CaO:P2O5 = 70:26:4

Glass: 50SiO2–22.6CaO–5.9 Na2–4P2O5–12K2–5.3MgO– 0.2B2O3 (wt%)

Molar composition: 45% SiO2, 3% P2O5, 26% CaO, 7% MgO, 15% Na2O, 4% K2O

89–90

45S5 Bioglass®

100–500

240

50

100–500

300–500

510–720

Porosity (%) Pore size (µm)

Composition

0.4*

0.40

–

11

0.3

0.27–0.42

Compressive strength (MPa)

Reference

Apatite formation after 28 d in SBF * Scaffolds pre-treated in SBF – Study proliferation of human osteoblasts

Apatite formation after 4 d in SBF

Apatite formation after 1 week in SBF

Apatite formation after 7d in SBF MC3T3-E1 attachment and proliferation

Apatite formation after 14 d in SBF HBDC attachment-viability

[45] *[46]

[44]

[19]

[18]

[17]

Apatite formation after [11] 28 d in SBF osteoblastlike cells (MG 63) attachment-proliferation*

Study of bioactivity SBF or cell culture

Table 5.1 Different silicate scaffolds fabricated by the foam replica technique, and their principal characteristics

Ca3Mg(SiO4)2 and Ca2MgSi2O9

Amorphous

Quartz (SiO2), wollastonite (CaSiO3), apatite

Amorphous

α-CaSiO3

Na2Ca2Si3O9

Crystalline phases

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and create a structure able to withstand mechanical loading during manipulation by surgeons and during application in vivo. In most cases reported, silicate bioactive glasses crystallize to wollastonite, quartz, apatite, calcium silicate, sodium calcium silicate or combeite crystalline phases (or combinations of these phases). It has been also confirmed recently that glass crystallization, for example in Bioglass®-derived glass-ceramic scaffolds, does not suppress bioactivity, it only retards the formation of the surface hydroxyapatite layer when the scaffold is immersed in body fluid [11, 49]. In this regard, a very important characteristic of Bioglass®-derived glassceramics is the transformation of the Na2Ca2Si2O9 crystalline phase to an amorphous matrix containing hydroxyapatite crystallites after immersion in simulated body fluid for 28 days [11]. The kinetic of HA formation can be tailored through the fabrication process by changing the sintering conditions. Fine crystals of Na2Ca2Si2O9 grow and almost complete densification of the scaffold struts occur when scaffolds are sintered at temperatures >1000°C. These conditions confer the scaffold the best possible compressive strength for the given porosity. It should therefore be emphasized that the final goal in the design of BTE scaffolds is attaining adequate structural integrity to warrant mechanical support of the surrounding tissue whilst maintaining bioactive behavior. Then, in the later stages of the tissue regeneration process the scaffold should biodegrade at a set rate matched to the rate of formation of new bone tissue, as mentioned above. Bioactive glass and glass-ceramic scaffolds fabricated by the replica technique (Table 5.1) are able to provide this function and they remain the subject of extensive research efforts considering their promising properties in the context of BTE. In the following sections two exemplary cases will be considered to illustrate the development of BTE scaffolds based on bioactive glasses and their fabrication by the foam replica manufacturing technique: (1) bioactive glass scaffolds made from a novel boron-containing glass, and (2) polymer-coated 45S5 Bioglass® based scaffolds.

5.5

Scaffolds from boron-containing bioactive glass

5.5.1 Fabricating bioactive glass scaffolds with different compositions Many parameters are involved in the selection of the bioactive glass for BTE scaffolds, which are dictated by the target application and the processing characteristics chosen for the scaffold fabrication. The main factors to be considered in relation to the scaffold requirements discussed in Section 5.2 are: the extent of bioactive response (bioreactivity), the kinetics of biodegradability and osteoconductivity as well as appropriate mechanical strength and structural integrity over a period of time [9]. In addition, in the case of silicate systems, the chemical composition will have an effect on the viscosity–temperature relationship

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and the crystallization potential of bioactive glasses, which are two important considerations in relation to the ability to process scaffolds of complex porous morphology from glass powders. Finding optimal chemical compositions of bioactive glasses for BTE scaffolds requires knowledge of the effect of addition or suppression of different oxides, in relation to the basic 45S5 Bioglass® composition, on relevant properties [50, 51]. It is well known that the properties of silicate glasses can be adjusted in a more or less controlled fashion by changes in chemical composition [52]. For conventional glasses, the property and composition optimization is supported by well-established relationships, and these are also relevant for bioactive glasses [50, 51, 53]. The composition also affects the processing characteristics of scaffolds considering that scaffold fabrication demands a wide working range and, preferably, a low tendency to devitrify. It has been discussed that bioactive glass compositions intended for sintered porous bodies should have preferably devitrification temperatures above 800°C and they should be chosen within the calcium silicate (e.g. wollastonite) primary crystalline phase, as these compositions can better sustain higher temperatures without crystallizing [50]. Generally, the working range of glasses increases with rising silica content and with a simultaneous decrease in in vitro bioactivity. The bioactivity has been found to depend more on the total amount of alkalis and alkaline earths than on the types of oxides, i.e. sodium and potassium, or calcium and magnesium, in each group. However it has been observed that with incorporation of potassium oxide, magnesium oxide or boron oxide, the forming properties of the glasses can be adjusted [53, 54]. In particular, the incorporation of boron oxide represents an interesting approach owing to the documented osteogenic and angiogenic effect of B2O3 containing bioactive glasses, which is attributed to the effect of the boron ion [21]. Boron-containing silicate glasses are therefore being increasingly considered to fabricate BTE scaffolds, as discussed next.

5.5.2 Boron-containing bioactive glass scaffolds The incorporation of boron in silicate bioactive glasses has been reported to lead to enhanced bone formation [21]. In addition, silica-free borate glasses have been shown to possess low chemical durability, and to convert rapidly to calcium phosphate (or hydroxyapatite) in physiological media [55] and to bond directly to bone in a manner comparable to the silicate-based 45S5 Bioglass®. Boron containing silicate glasses are being currently investigated considering the reported positive effect of the B ion but also based on the fact that these glasses can be heat treated (sintered) at lower temperatures than 45S5 Bioglass® and do not crystallize [20, 44, 56]. The borosilicate bioactive glass ‘code 0106’ with nominal composition (in wt%): 50 SiO2, 22.6 CaO, 5.9 Na2O, 4 P2O5, 12 K2O, 5.3 MgO and 0.2 B2O3 developed at the Process Chemistry Centre, Åbo Akademi University, Turku

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Finland [57] has been recently used to fabricate BTE scaffolds by the foam replica method. Powder of this glass with average particle size of 36 µm was used. The heat treatment involved sintering for 3 to 10 hours at temperatures in the range 630 to 680°C (considerably lower than temperatures used for 45S5 Bioglass® scaffolds mentioned above) [44]. The heat treatment leading to scaffolds with the most promising characteristics involved sintering at 670°C for 5 hours. The morphology of the scaffolds is shown in Fig. 5.3. The scaffold struts were seen to be well sintered, which should lead to favorable mechanical properties. The average pore size of the fabricated scaffolds was 240 µm, while the average porosity was 68%. This porosity value is lower than that of scaffolds fabricated previously using Bioglass® [49], which exhibited porosity of up to 90%. However the pore size is within the desired range (>200 µm) and the pore structure is highly interconnected (Fig. 5.3). In vitro bioactivity studies in simulated body fluid (SBF) showed that a carbonate hydroxyapatite (HCAp) layer was deposited on scaffolds after only 4 days of immersion in SBF, demonstrating the high in vitro bioactivity, which is comparable to that of 45S5 Bioglass®. The FTIR characterization results shown in Fig. 5.4 confirm that scaffolds exhibit marked bioactive behavior given by the significant and rapid formation of the HCAp surface layer [44]. Figure 5.5 shows SEM images illustrating the transformation of the microstructure of samples during the formation of HCAp. After one day’s immersion in SBF the onset of a degradation process is clearly visible and the development of a uniform, thick HCAp layer on the surface of the scaffolds is confirmed after only four days of immersion in SBF. X-Ray diffraction analysis showed that as-sintered scaffolds consisted of an amorphous structure. The fabrication of bioactive scaffolds from bioactive glasses in powder form, to retain an amorphous structure, is quite unusual as most known bioactive glasses crystallize during the high-temperature fabrication process, as mentioned in the

5.3 SEM images showing the morphology of borosilicate bioactive glass scaffolds at low (left) and high (right) magnifications (sintering temperature and time: 670°C, 5 hours, respectively) [44] (reproduced with permission of IOP).

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5.4 FTIR spectra of borosilicate bioactive BTE scaffolds before and after immersion in SBF for different soaking times [44]. The transmittance spectrum of HAp is also presented for comparison (reproduced with permission of IOP).

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5.5 SEM images showing the surface of borosilicate bioactive scaffolds after one and four days of immersion in SBF (modified from [44], reproduced with permission of IOP).

previous section. An amorphous phase is less stable than its crystalline counterpart in terms of Gibbs free energy, which is expected to result in a better bioactive behavior in comparison to partially crystallized scaffolds. The degradation rate of the boron-containing scaffolds in relevant in vitro or in vivo conditions has not been investigated to date, and it remains a key study to confirm the positive effect

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of boron (released during scaffold degradation) on both osteogenic and angiogenic potential predicted by other authors [20, 21].

5.6

Polymer-coated composite scaffolds

5.6.1 Advantages of composite systems The major disadvantage of bioactive glasses is their low fracture toughness and brittleness. For applications such as BTE, bioactive glasses are often used in combination with biodegradable polymers to achieve the best possible mechanical and biological performance [5, 6]. In the development of composites for tissue engineering scaffolds, two main approaches are being followed; the first approach considers the incorporation of bioactive glass particles as inclusions into polymer structures, e.g. foams [5]; and the second approach considers the incorporation of polymer coatings onto a 3D porous bioactive glass based scaffold [58]. In this section specific examples of the second approach will be discussed while composite scaffolds based on biopolymers containing bioactive glass particulate inclusions are presented in Chapter 7. The introduction of a polymer into the structure of bioactive glass scaffolds, whether as a coating or by infiltrating into the strut structure, has been investigated by several authors [58, 59]. It is hypothesized that by introducing a polymer into the structure of struts, the polymer can fill micro-cracks, bridge large cracks and form a polymer-bioactive glass composite network, behaving in a similar way to collagen fibers in bone. It is well established that collagen fibers bridge cracks in bone, to increase bone fracture toughness [62, 63]. If the polymer just forms a layer that rests on top of the scaffold strut and does not infiltrate the material structure then the polymer will not improve the mechanical properties as well as if the polymer infiltrated the micro-cracks and remaining pores of the scaffold struts. This behavior is illustrated schematically in Fig. 5.6a and 5.6b. Processes developed to fabricate both polymer coated bioactive glass scaffolds and polymer-bioactive glass scaffolds with interpenetrating network microstructure are based on infiltrating a sintered (or partially sintered) bioactive glass or glassceramic scaffold with the biodegradable polymer in solution [58]. A novel method recently developed to coat 3D scaffolds with polymers is Matrix Assisted Pulsed Laser Evaporation (MAPLE) [64]. This technique is often preferred over other film deposition methods since it provides high control over film characteristics. It was shown that this technique can be used to produce PDLLA (poly(D,L lactide))coated Bioglass® scaffolds [65]. The alternative approach of fabricating hybrid polymer–ceramic composite scaffolds, e.g. exploiting the molecular mixing of inorganic and organic phases for example in sol-gel based techniques, has also been explored [66], however for the sake of brevity, these hybrid materials will not be considered in the present chapter.

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5.6 Schematic diagram illustrating how a polymer lies on the surface of a cracked (or partially sintered) scaffold strut (coating) (a); or how it can infiltrate the micro-cracks on the surface of the struts (b).

Table 5.2 Compressive strength values of a variety of polymer-coated bioactive glass composites and relevant bone tissues Inorganic phase

Polymer coating

Type

Compressive strength (MPa)

Reference

45S5 Bioglass® 45S5 Bioglass® 45S5 Bioglass®

No coating P(3HB) P(3HB) microspheres PDLLA – –

Scaffold Scaffold Scaffold

0.1–0.3 ~1.5 ~0.4

[11] [67] [68]

Scaffold Bone Bone

0.2–0.65 4–12 130–180

[59] [69] [69]

45S5 Bioglass® Cancellous bone Cortical bone

Chen et al. [59] developed Bioglass®-based scaffolds coated with PDLLA for the first time. In those polymer-coated scaffolds bioactivity was not impaired as demonstrated by tests in simulated body fluid. Polyhydroxyalkanoate (P(3HB) has been investigated as an alternative coating material for tissue-engineering scaffolds [67], which leads to a significant improvement of the work of fracture in compression, as discussed below. Although polymer/bioactive glass composite scaffolds developed so far have not quite reached the mechanical requirements of the surrounding host tissue, as shown in Table 5.2, they are much closer to the desired mechanical properties than uncoated scaffolds. Polymers have the added advantage that they bind the structure of the scaffold together making the scaffold tougher and providing it with extra mechanical stability during the in vivo tissue regeneration process. In addition, the polymer can have other functions in the scaffold, such as being a carrier for therapeutic drugs, growth factors and even antibacterial metal ions; this should increase the functionality of the scaffolds [70]. There are several polymer–bioceramic composites that demonstrate the evolution of this hypothesis and are reviewed in the literature [58]. In the following section, BTE scaffolds in

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the system poly(3-hydroxybutyrate) coated Bioglass® will be discussed as a typical example of this type of composite.

5.6.2 Bioglass® scaffolds coated with poly(3-hydroxybutyrate) Poly(3-hydroxybutyrate) (P(3HB)), a member of the polyhydroxyalkanoate family, is being considered for the development of composites for tissue engineering [71]. This polymer is naturally thermoplastic and is produced by many types of micro-organisms [72]. Bretcanu et al. [73] first used this bacteria derived polymer to infiltrate 45S5 Bioglass®-derived scaffolds. P(3HB), when used as a thin coating material, does not affect the interconnectivity of the scaffold porous structure and coated scaffolds were seen to maintain a high porosity of 85% [67, 73]. It was also found that when applied as a coating on a 3D porous Bioglass® scaffold, P(3HB) does not form a fully homogeneous layer on the struts, as shown in Fig. 5.7. However, this lack of coating connectivity is useful for the scaffold to retain a high bioactivity because the underlying Bioglass® surface will be in direct contact with the biological environment promoting the occurrence of the typical surface bioreactions of bioactive glasses [10]. When the compressive strength of P(3HB)-coated Bioglass® scaffolds is compared to that of uncoated Bioglass®-derived scaffolds, there is a considerable improvement; from 0.74MPa to ∼1.5MPa [67], which makes up for the reduction in scaffold porosity. This composite is comparable in terms of mechanical properties (and again a vast improvement on the plain biomaterial), to other polymer-coated bioactive glass composites described above, as illustrated in Table 5.2. There is increasing interest in using these scaffolds to produce drug delivery devices, e.g. incorporating growth factors, antibiotics and other additives into the polymer coating to improve the scaffold multi-functionality.

5.7 SEM images showing the microstructure of a 45S5 Bioglass®/ P(3HB) composite scaffold (a); and the discontinuous nature of the P(3HB) coating on the scaffold (b) [73].

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Another method of incorporating the polymer onto a scaffold is in the form of microspheres, as discussed by Francis and Meng et al. [68]. This approach has two effects on the overall performance of the composite scaffold, the first being increased functionality of the scaffold as the microspheres can be filled with antibiotics or growth factors, and secondly it should improve the properties of the scaffold in the same manner as the coating with P(3HB) films discussed above. The microspheres are formed using the solid-in-oil-water emulsion technique, which produces microspheres with a mean diameter ranging from 1.5 to 2.0 µm. Pre-formed 45S5 Bioglass® scaffolds produced using the foam replication technique as described above are coated with P(3HB) microspheres by pipetting drops of a slurry containing the microspheres onto the scaffolds. It should be noted that the incorporation of the additive (the antibiotic Gentamicin in the case reported by Francis and Meng et al. [68]) into the microspheres occurs before the microspheres are used to coat the scaffolds. Figure 5.8 indicates that the microsphere coating of the scaffolds is extensive but not completely homogeneous: there are some gaps in the layer of microspheres, and this will help the composite scaffold maintain its high level of bioactivity because strut surfaces will be directly exposed to the biological environment. Similarly, the addition of the polymer coating (in the form of microspheres) was seen to reduce the overall porosity of the scaffold from ∼90% down to ∼70%; this reduction in porosity has a positive effect on the compressive strength

5.8 SEM image showing the microstructure of a 45S5 Bioglass® scaffold coated with P(3HB) microspheres [68].

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of the scaffolds, which doubles compared to uncoated scaffolds [68]. The addition of the microspheres also increased the work of fracture of the composite scaffolds, i.e. they are more resistant to crack propagation compared to uncoated scaffolds [68]. It was also demonstrated that addition of P(3HB) microspheres did not inhibit the formation of hydroxyapatite on the surface of the scaffold in contact with SBF, which is important as this surface bioreaction is the main function provided by the 45S5 Bioglass® scaffold to ensure bone cell attachment [11]. Consequently, it can be concluded that there is not much difference between the two different applied techniques of incorporating P(3HB) into Bioglass®-based scaffolds, as the two methods, namely film and microsphere coating, lead to scaffolds with increased compressive strength, and in both cases multifunctional BTE scaffolds incorporating P(3HB) that retain the basic 45S5 Bioglass® bioactivity can be fabricated.

5.7

Conclusions

An overview on the development of BTE scaffolds based on (melt-derived) bioactive glasses and glass-ceramics has been provided in this chapter highlighting the manufacturing of 3D scaffolds by the foam replica technique. The increasing interest in the application of bioactive glasses and glass-ceramics for fabricating BTE scaffolds is based on the attractive biological properties of bioactive glasses, which include not only the classical concept of bioactivity based on the bioreactivity of the material in contact with relevant fluids, but also the recognized effect of ionic glass dissolution products on osteogeneis and angiogenesis. The utilization of bioactive glasses as platforms for the controlled release of certain metallic ions, beyond Si, Ca, and P, to stimulate specific cellular responses can be highlighted as an important research area likely to concentrate research efforts in the near future. In this chapter, two specific types of scaffolds were discussed in some detail: (1) BTE scaffolds fabricated from boron-containing bioactive glass, which leads to non-crystalline scaffolds of high bioactivity with the potential advantage of the release of boron to promote bone formation; and (2) polymer-coated scaffolds of enhanced mechanical behavior and adequate bioactivity as basic structures for the future development of multifunctional scaffolds with drug delivery capability. In this context, a significant toughening effect by polymer incorporation, especially in scaffolds exhibiting interpenetrating network microstructure, represents an important improvement in scaffold design. The addition of a polymer phase might have extra functions, e.g. the biodegradable polymer can act as carrier for biomolecules, growth factors and antibiotics, hence increasing the capability of the BTE constructs. The optimization of scaffolds for BTE strategies is still one of the key tasks to be accomplished in order to realize the promise that tissue engineering holds to

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improve the life of millions of humans suffering bone and muscoloskeletal diseases. Bioactive glasses, and the possibility of new compositions incorporating specific therapeutic ions and their combination with biopolymers to form composites, will continue to play a significant role in the field as one of the materials of choice for scaffold developments.

5.8

References

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35. Deb S., Mandegaran R., Di Silvio L., A porous scaffold for bone tissue engineering/45S5 Bioglass® derived porous scaffolds for co-culturing osteoblasts and endothelial cells. J Mater Sci Mater Med 2010; 21:893–905. 36. Azevedo M., Jell G., Hill R., Stevens M. M., Novel hypoxia mimicking bioactive materials for tissue engineering. Tissue Eng A 2008; 14:889–889 (Meeting Abstract: P288). 37. Bellantone M., Williams H. D., Hench L. L., Broad-spectrum bactericidal activity of Ag2O-doped bioactive glass. Antimicrob Agents Chemother 2002; 46:1940–1945. 38. Vitale-Brovarone C., Miola M., Balagna C., Verne E., 3D-glass-ceramic scaffolds with antibacterial properties for bone grafting. Chemical Eng J 2008; 137:129–136. 39. Mouriño V., Boccaccini A. R., Bone tissue engineering therapeutics: controlled drug delivery in three-dimensional scaffolds. J R Soc Interface 2010; 7:209–227. 40. Li R., Clark A. E., Hench L. L., An investigation of bioactive glass powders by sol-gel processing. J App Biomater 1991; 2 (4):231–239. 41. Quintero F., Pou J., Comesana R., Lusquinos F., Riveiro A., Mann A. B., Hill R. G., Wu Z. Y., Jones J. R., Laser spinning of bioactive glass nanofibers. Adv Funct Mater 2009; 19:1–7. 42. Brunner T. J., Grass R. N., Stark W. J., Glass and bioglass nanopowders by flame synthesis. Chem Commun 2006; 13:1384–1386.74. 43. Zenati R., Fantozzi G., Chevallier J., Mourad A., Porous Bioglass and Preparation Thereof. French Patent no FR2005/001921, 2005. 44. Mantsos, T., Chatzistavrou X., Roether J. A., Hupa L., Arstila H., Boccaccini A. R., Non-crystalline composite tissue engineering scaffolds using boron-containing bioactive glass and poly(D,L-lactic acid) coatings. Biomed Mater 2009; 4:055002 (12 pp). 45. Renghini C., Komlev V., Fiori F., Verne E., Baino F., Micro-CT studies on 3-D bioactive glass-ceramic scaffolds for bone regeneration. Acta Biomater 2009; 5: 1328–1337. 46. Vitale-Brovarone C., Verne E., Robiglio L., Appendino P., Bassi F., Martinasso G., Muzio G., Canuto, R. Development of glass-ceramic scaffolds for bone tissue engineering: characterization, proliferation of human osteoblasts and nodule formation. Acta Biomater 2007; 3:199–208. 47. Bretcanu O., Samaille C., Boccaccini A. R., Simple methods to fabricate Bioglass®derived glass-ceramic scaffolds exhibiting porosity gradient, J Mater Sci 2008; 43:4127–4134. 48. Vitale-Brovarone C., Baino F., Verne E., Feasibility and tailoring of bioactive glassceramic scaffolds with gradient of porosity for bone grafting. J Biomater Appl 2010; 24:693–712. 49. Boccaccini A. R., Chen Q. Z., Lefebvre L., Gremillard L., Chevalier J., Sintering, crystallisation and biodegradation behaviour of Bioglass®-derived glass-ceramics. Faraday Discuss 2007; 136:27–44. 50. Arstila H., Vedel E., Hupa L., Hupa M., Factors affecting crystallization of bioactive glasses. J Eur Ceram Soc 2007; 27:1543–1546. 51. Andersson Ö. H., Karlsson K. H., On the bioactivity of silicate glass. Journal of Non-Crystalline Solids 1991; 129: 145–151. 52. Westerund T., Hatakka L., Karlsson K. H., A model for optimizing. Glass batch compositions. J Amer Ceram Soc 1983; 66 (8):574–579. 53. Brink, M., The influence of alkali and alkaline earths on the working range for bioactive glasses. J Biomed Mater Res 1997; 36:109–117.

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54. Itälä A., Koort J., Ylänen H. O., Hupa M., Aro H. T., Biologic significance of surface microroughening in bone incorporation of porous bioactive glass implants. J Bio Mater Res A 2003; 67 (2):496–503. 55. Day D. E., White J. E., Brown R. F., McMenamin K. D., Transformation of borate glasses into biologically useful materials. Glass Technol 2003; 44 (2):75–81. 56. Rahaman M. N., Day D. E., Brown R. F., Fu Q., Jung S. B., Nanostructured bioactive glass scaffolds for bone repair, Advances in Bioceramics and Porous Ceramics, ed. R. Narayan and P. Colombo (Westerville, OH: American Ceramic Society) 2008; pp. 211–222. 57. Vedel E., Zhang D., Arstila H., Hupa L., Hupa M., Predicting physical and chemical properties of bioactive glasses from chemical composition. Part IV: tailoring composition with desired properties. Glass Technol: Eur J Glass Sci Technol A 2009; 50:9–16. 58. Yunos M. D., Bretcanu O., Boccaccini A. R., Polymer–bioceramic composites for tissue engineering scaffolds. J Mater Sci 2008; 43:4433–4442. 59. Chen Q. Z., Boccaccini A. R., Poly(D,L-lactide) coated 45S5 Bioglass®based scaffolds: processing and characterisation. J Biomed Mater Res 2006; A 77: 445–452. 60. Miao X., Tan L. P., Tan L. S., Huang X., Porous calcium phosphate ceramics modified with PLGA-bioactive glass. Mater Sci Eng C 2007; 27:274–280. 61. Kim H. W., Knowles J. C., Kim H. E., Hydroxyapatite porous scaffold engineered with biological polymer hybrid coating for antibiotic Vancomycin release. J Mater Sci Mater Med 2005; 16:189–195. 62. Wang X., Bank R. A., Tekoppele J. M., Agrawal C. M., The role of collagen in determining bone mechanical properties. J Orthopaed Res 2001; 19:1021–1026. 63. Nalla R. K., Kinney J. H., Ritchie R. O., Effect of orientation on the in vitro fracture toughness of dentin: the role of toughening mechanisms. Biomaterials 2003; 24:3955–3968. 64. Califano V., Bloisi F., Vicari L. R. M., Bretcanu O. and Boccaccini A. R., Matrixassisted pulsed laser evaporation of poly(D,L-lactide) for biomedical applications: effect of near infrared radiation. J Biomed Opt 2008; 13:014028. 65. Califano V., Bloisi F., Vicari L. R. M., Yunos D. M., Chatzistavrou X., Boccaccini A. R., Matrix assisted pulsed laser evaporation (MAPLE) of Poly(D,L lactide) (PDLLA) on three-dimensional Bioglass® structures. Adv Eng Mater 2009; 11: 685–689. 66. Catauro M., Raucci M. G., De Gaetano F., Buri A., Marotta A., Ambrosio L., Sol-gel synthesis, structure and bioactivity of polycaprolanctone/CaO.SiO2 hybrid material. J Mater Sci Mater Med 2004; 15:991–995. 67. Bretcanu O., Misra S., Roy I., Renghini C., Fiori F., Boccaccini A. R., Salih V., In vitro biocompatibility of 45S5 Bioglass®-derived glass-ceramic scaffolds coated with poly(3-hydroxybutyrate). J Tissue Eng Regen Med 2009; 3:139–148. 68. Francis L., Meng D., Knowles J. C., Roy I., Boccaccini A. R., Multifunctional P(3HB) microsphere/45S5 Bioglass®-based composite scaffolds for bone tissue engineering. Acta Biomater 2010; 6:2773–2786. 69. Gibson L. J., Ashby M. F., Cellular solids: structure and properties. 2nd edn. Oxford: Pergamon 1999, pp. 429–452. 70. Baroli B., From natural bone graft to tissue engineering therapeutics: brainstorming on pharmaceutical formulative requirements and challenges. J Pharm Sciences 2009; 98:1317–1375.

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71. Misra S. K., Valappil S. P., Roy I., Boccaccini A. R., Polyhydroxyalkanoate (PHA)/ inorganic phase composites for tissue engineering applications. Biomacromolecules 2006; 7 (8):2249–2258. 72. Chen G. Q., Wu Q., The application of polyhydroxyalkanoates as tissue engineering materials. Biomaterials 2005; 26:6565–6571. 73. Bretcanu O., Chen Q. K., Misra S., Boccaccini A. R., Roy I., Verne E., Brovarone C. V., Biodegradable polymer coated 45S5 Bioglass-derived glass-ceramic scaffolds for bone tissue engineering. Glass technology-Europ. J Glass Sci Technol Part A 2007; 48 (5):227–234.

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6 Nanoscaled bioactive glass particles and nanofibres M. EROL , Istanbul Technical University, Turkey and A. R. BOCCACCINI, University of Erlangen-Nuremberg and Imperial College London, UK

Abstract: Nanoscale bioactive glasses of silicate composition are gaining increasing attention due to their superior bioactivity, enhanced osteoconductivity and antibacterial properties when compared to conventional (micron-sized) bioactive glasses. In this chapter, we present an overview of the technology, characterization and applications of nanoparticle and nanofibrous bioactive silicate glasses. Novel fabrication technologies are presented, covering sol-gel routes, microemulsion techniques, the gas phase synthesis method (flame spray synthesis), laser spinning and electro-spinning and the synergistic effects that determine the final properties of these materials are discussed. The advantages of nanoscaled bioactive glasses compared to conventional bioactive glasses are also discussed based on available literature evidence. The focus of the chapter is the diverse application areas of nanoscale bioactive glasses ranging from tissue engineering scaffolds to drug delivery and dentistry, considering also polymer/ bioactive glass nanocomposites as promising materials for implants, bone fillers and bioactive coatings for orthopaedic applications. The nanofeatures characteristic of this type of bioactive glass are discussed and the possibilities of expanding the use of these nanomaterials in other nanotechnology approaches aiming at advanced biomedical applications (nanomedicine) are also highlighted. Key words: bioactive glass nanoparticles, bioactive glass nanofibres, nanocomposites, tissue engineering, bioactive glass coatings, nanomedicine.

6.1

Introduction

Silicate bioactive glasses were developed for the first time by Hench and co-workers in 1969 [1]. These highly surface reactive inorganic materials are able to bond to bone tissue in a physiological environment [2]. The most widely used group of bioactive glasses for applications in the biomedical field, related to the first chemical composition developed by Hench et al. [1], consist of a silicate network incorporating sodium, calcium and phosphorus in different proportions. The traditional 45S5 bioactive glass (45S5 Bioglass®) of composition in wt%: 45 SiO2, 24.5 Na2O, 24.5 CaO and 6 P2O5, has received approval from the US Food and Drug Administration (FDA) for clinical use in the treatment of periodontal diseases as bone filler as well as in middle ear surgery [2]. A wide range of bioactive glass compositions is now available that have additional elements incorporated such as fluorine, magnesium, strontium, iron, silver, boron, potassium or zinc [3–9]. More recently, a range of other clinical applications of bioactive glasses was proposed, for example in periodontology [10, 11], endodontology [12, 13] or as coating on 129 © Woodhead Publishing Limited, 2011

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metallic orthopaedic implants [14, 15]. Moreover bioactive glasses are finding application in the development of scaffolds for tissue engineering and regenerative medicine [16, 17–20]. In these approaches, both micron-sized and nanoscale particles are considered for fabrication of composite scaffolds, e.g. combining biodegradable polymers and bioactive glass [16, 18, 21]. In this context, bioactive silicate glasses have advantages in comparison to other bioactive ceramics, e.g. sintered hydroxyapatite, considering that their dissolution products have been shown to upregulate the expression of genes controlling osteogenesis [17]. There is also emerging evidence in the literature showing the potential angiogenic effects of bioactive glasses, i.e. increased secretion of vascular endothelial growth factor (VEGF) in vitro and enhancement of vascularisation in vivo [22–24]. The specific properties of bioactive glasses mentioned above can be influenced and possibly controlled to a greater extent in the nanoscale, e.g. developing nanoparticles or nanofibres, for a range of applications including coatings of biomedical devices, fillers in composite materials for biodegradable implants, dental fillers, tissue engineering scaffolds or drug delivery vehicles. Nanotechnology approaches are being proposed for a wide variety of medical engineering and biomedical applications in what is termed nanomedicine. Nanoscience and nanotechnology are particularly attractive in tissue engineering since the interactions between cells and biomaterial surfaces occur firstly in the nanoscale and the components of biological tissues have nanoscale dimensions [25, 26]. In the nanometer scale, key properties determining the cell–biomaterial interaction such as surface area, surface roughness, hydrophilicity and wettability, which influence cell adhesion and provide bonding properties to host (bone) tissue for long-term functionality, are completely different from the conventional (micrometer) scale [27, 28]. Nanoscaled biomaterials have in general a low defect concentration and a high ratio of surface area to volume [29, 30], which results in greater surface energy and surface bioreactivity compared with micro-scale biomaterials. In the case of bioactive glass nanoparticles, a superior in vitro bioactive behaviour in comparison with µm-scale particulate glasses is expected, mainly due to enhanced textural properties (higher surface area) than the micrometer-sized counterparts. It has been reported that the larger specific surface area of the nanoparticles allows not only for a faster release of ions but also a higher protein adsorption, which improves bioactivity [27, 31]. In addition, the use of nanosized bioactive glass particles enhances osteoblast adhesion, proliferation and differentiation, and induces an increase in the biomineralization process [27, 32]. Furthermore, the greater specific surface area of the nanosized bioactive glass particles when used as filler in biopolymers will lead to higher interface effects resulting in improved mechanical properties of the materials, compared with micrometer size particles, provided a homogeneous dispersion of the nanoparticles in the polymer matrices is achieved [33–37]. A range of techniques has emerged recently to fabricate nanoscale bioactive glasses including sol-gel [31, 38, 39], laser spinning [40], microemulsion [41] and

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gas-phase synthesis [42–45]. The produced nanoscale bioactive glasses are being proposed for the design of numerous nanomaterials for biomedical applications, including: combination of nanofibres or nanoparticles with polymeric matrices to produce nanocomposites [33, 36, 37, 46], incorporation of nanoparticles or nanofibres into porous 3D scaffolds [34, 47, 48], nanoparticle coatings on implant surfaces [49–51] and production of non-porous materials containing nanoparticles in the form of gels, injectable materials or hard devices [52]. In this chapter, we present the current state of-the-art fabrication technologies for nanoscaled bioactive glasses (nanoparticles and nanofibres) and the application of the produced nanomaterials in the biomedical field. Section 6.2 discusses the specific nanoscale-related features of bioactive glasses. Different synthesis methods for bioactive glass nanoparticles and nanofibres are reviewed in Section 6.3. In Section 6.4, the application of bioactive glass nanoparticles and nanofibres is comprehensively reviewed. Finally in Section 6.5, a summary of the topic is presented and areas for future research are highlighted.

6.2

Characteristics of nanoscale bioactive glasses

The higher specific surface area of nanoscale bioactive glasses (nanoparticles and nanofibres) is expected to enable a faster solubility of the material (higher ion release rate) and also higher protein adsorption and thus enhanced bioactivity. Faster deposition or mineralization of tissues such as bone or teeth is possible when the tissues are in contact with nanoscale particles [31, 53]. This effect is related to the bone structure, which can be considered as a nanostructured composite mixture of collagen fibrils and carbonate hydroxyapatite nanocrystals [31, 53]. Mimicking the nanofeatures of bone on the surface of synthetic bone implants, for example, has been shown to increase bone-forming cell adhesion and proliferation [31]. Cells in their natural environment are surrounded by nanostructures in contact with other cells and with the extra-cellular matrix (ECM), formed by biomolecules configured in different geometrical arrangements and (nano)structures. In the case of bone tissue applications, nanoscale-related features of the materials, such as textural properties, surface energy and chemistry, nanotopography and wettability, control protein interactions modulating subsequent osteoblast adhesion and long-term functionality [54, 55]. Moreover nanotopography (roughness, shape/size of surface features, geometric vs. random distribution, etc.) affects cell interactions and it is expected to alter cellular behaviour when compared to conventional (µm-sized) topography [54, 56, 57]. At the nanoscale, experimental evidence indicates that cell types react specifically and in a differential manner to topographical surfaces [55]. Moreover, nanotopography enhances osteoblastic differentiation which could also promote stability and change the biomechanical environment for healing [58]. Surface energy is another important factor that regulates cell response to biomaterials and it can be altered by incorporating

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nanoparticles or nanofibres. Surface energy modulates protein adsorption, which further regulates cell adhesion, cell spreading and proliferation [59], with higher surface energy enhancing the interaction between an implant surface and its biological environment [60, 61]. A direct relationship exists between roughness and surface energy of materials [62]. Moreover it has been demonstrated that the polar component of surface energy increases significantly with roughness and that cell adhesion enhancement is related to the degree of roughness and hydrophobicity [63]. It is also evident that the hydrophobicity or hydrophilicity of a surface can significantly alter cell behaviour. Hydrophilic materials that exhibit higher surface energy are preferred for cell attachment and proliferation [64]. Moreover, the wettability of a surface can allow for ranking materials with regards to their hydrophobic or hydrophilic category. For example, the surface wettability of alumina and titanium were shown to be enhanced by reducing particle size and this resulted in improved attachment and proliferation of preosteoblasts [65, 66]. As reported in the literature, features affected in the nanoscale, such as surface energy, nanotopography and wettability, have direct effects on cell orientation, morphology and cytoskeleton arrangements [54, 55, 67]. Although most previous results have been obtained on TiO2, alumina and hydroxyapatite, the findings should be directly applicable to nanoscaled bioactive glasses, as discussed in this chapter. For bone tissue engineering purposes, where scaffolds made of polymer/bioactive glass composites are applicable [16, 18, 68], the use of nanoscale bioactive glasses is expected to improve both the scaffolds’ mechanical and biological properties. As mentioned above, the surface bioreactivity of nanoparticles is higher than that of µm-sized particles. In addition, bioactive glass nanoparticles will induce nanostructured features on scaffold surfaces, which are likely to improve osteoblast cell attachment and subsequent cell behaviour, following the discussion above. Other advantages of the reduced size of inorganic particles include the possibility to use them to reinforce polymers in the form of nanofibres, and to process thin bioactive coatings and nanoscaled injectable systems [69].

6.3

Fabrication of bioactive glass nanoparticles and nanofibres

6.3.1 Sol-gel methods The sol-gel process is commonly applied in the context of silicate systems and other oxides and it is a technology widely used for inorganic material synthesis, for example for the fabrication of thin films, powders, nanoparticles and fibres [34, 70–72]. Li et al. [70] showed the sol-gel synthesis of silicate bioactive glasses using metal alkoxides as precursors. Typical precursors for sol-gel bioactive glasses are tetraethyl orthosilicate, calcium nitrate and triethylphosphate. After hydrolysis and polycondensation reactions a gel is formed, which after calcination at 600 to 700°C

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forms the glass. Recent work on fabricating bioactive silicate glass nanoparticles by a sol-gel process has been carried out by Hong et al. [38]. A schematic diagram showing the sol-gel synthesis process developed by Hong et al. [34, 38] is presented in Fig. 6.1. The morphology and size of bioactive glass nanoparticles could be tailored by varying the production conditions and the feeding ratio of reagents [39, 52]. Not only the size but also the morphology of bioactive glass nanoparticles (NBG) is relevant in order to achieve the desired behaviour when applied in the biomedical field. For example, the effects of different morphologies on the in vitro bioactivity of nanosized bioactive glass particles in the system CaO-P2O5-SiO2 has been investigated [73]. In addition, surface-modified bioactive glass nanoparticles have been developed by sol-gel to improve their dispersibility by using a wet mechanical grinding technique [74]. It was reported that a layer of silane could prevent agglomeration of sol-gel derived bioactive glass nanoparticles [74]. It is usually difficult to synthesize silicate glasses in a nanosize scale with multicomponent chemical composition, e.g. by addition of specific metallic ions.

6.1 Schematic diagram for the sol-gel synthesis process of bioactive glass nanoparticle. Drafted according to the methods described in refs. [34, 38].

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Delben et al. [75] have developed sol-gel derived bioactive glass nanoparticles doped with silver with a mean particle size of 100 nm. In an expansion of the sol-gel method to develop nanoscale bioactive glass, the method has been combined with electrospinning to fabricate silicate fibres [76]. Electrospinning of sol-gel precursors can result in bioactive glass fibres with diameters <100 nm [76–78]. The diameter and morphology of nanofibres can be controlled to some extent by varying the amount and type of additives and the applied electric field. The resulting bioactive glass nanofibres are flexible due to their small diameter and they are useful for developing porous tissue engineering scaffolds. Figure 6.2 shows electron microscopy images of bioactive glass nanofibres prepared by electrospinning of a silicate sol, according to Kim et al. [76].

6.2 Analysis of bioactive glass nanofibres after electrospinning and heat treatment at 700°C. SEM images of the nanofibres of different average diameters (630 nm, 220 nm, and 84 nm in sequence) (a–c) with varying sol concentration (1, 0.5, and 0.25 M in sequence); TEM image of the nanofibres with average diameter = 84 nm (d) (reprinted from ref. [76] with permission).

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The average diameters of the fibres varied between 84 nm and 630 nm depending on the initial sol concentration. Moreover Lu et al. [78] have developed bioactive glass nanofibres in the CaO-SiO2 system (70 mol% SiO2, 30 mol% CaO) by the combined sol-gel and electrospinning method. Clearly the sol-gel technique is powerful and versatile enough to synthesize a great variety of inorganic materials and it has been successfully used for the production of nanoscale bioactive glasses. However, sol-gel techniques have limitations regarding (1) the chemical compositions that can be produced and (2) the need to remove water or residual solvents.

6.3.2 Microemulsion methods A microemulsion is a thermodynamically stable transparent, isotropic dispersion of two immiscible liquids such as water and oil stabilized by surfactant molecules at the water/oil interface. In water-in-oil microemulsions, nanosized water droplets are dispersed in the continuous hydrocarbon phase and surrounded by the monolayer of surfactant molecules [79]. Microemulsion techniques are capable of producing nanosized particles with minimum agglomeration, since the reaction takes place in nanosized domains [80, 81]. Only a few reports are available on the synthesis of nanosized bioactive glass particles by microemulsion methods. Zhao et al. [41], for example, developed a microemulsion technique to fabricate nanoparticles in the system CaO-P2O5-SiO2. Spherical amorphous particles were obtained in the 25 to 50 nm range (Fig. 6.3). It was shown that the nanoparticle diameter could be related to the molar ratio of water to surfactant (γ ) in water/oil emulsions.

6.3.3 Flame spray synthesis Flame spray synthesis belongs to the gas phase-based methods using metalorganic precursors to produce nanoparticles at temperatures above 1000°C. The basic principle of all gas phase synthesis methods is the formation of molecular nuclei, which is followed by condensation and coalescence inducing the subsequent growth of nanoparticles in high temperature regions during the process. High cooling rates (>1000 K s−1) and short residence times (1 ms) enable nanoparticle formation. In contrast to wet phase processes, gas phase synthesis allows generally higher production rates. Flame spray synthesis was originally developed for manufacturing carbon black [82], and it is nowadays routinely used to produce large amounts of silica and titania nanoparticles. Adapting the process to allow the use of organic liquid precursors loaded with metals instead of gaseous precursors led to increased versatility [83–85]. In this context, the metal carboxylate system is a very convenient precursor, because it allows the synthesis of oxide nanoparticles of almost any composition [85]. The process enables production of numerous nanoparticulate-mixed oxides with high chemical homogeneity. Moreover, and depending on the composition, fast quenching can

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6.3 TEM image showing the diameter distribution of bioactive glass particles produced by microemulsion method (reprinted from ref. [41] with permission).

preserve the amorphous state of the material [42, 43]; consequently, by using flame spray synthesis, the preparation of nanoparticles of different bioactive glass compositions has become possible. Mixtures of 2-ethylhexanoic acid salts of calcium and sodium, hexamethyldisiloxane, tributyl phosphate and fluorobenzene to introduce fluorine have been employed [42].

6.3.4 Laser spinning methods Laser spinning enables the production of glass fibres with a wide range of diameters (from the nanometre to micrometre scale) [86]. Large quantities of nanofibres can be produced with specific, controllable chemical compositions without the necessity of any chemical additives or post-heat treatments. The process is very efficient; nanofibres are produced in several microseconds. The laser spinning technique essentially involves the quick heating and melting of a small volume of the precursor material up to high temperatures using a high power laser. At the same time, a supersonic gas jet is injected into the melt volume

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to blow the molten material. The molten material is then quickly stretched and cooled by the supersonic gas jet [86]. Long fibres with high length-to-diameter ratios can be produced, and the obtained material is in amorphous form because of the high cooling speed. Glass fibres in the form of a disordered net of intertwined amorphous micro- and nanofibres have been produced [87, 88]. More recently, Quintero et al. [40] developed the first bioactive glass nanofibres by the laser spinning technique. They produced bioactive glass nanofibres in the 45S5 Bioglass® and 52S4.6 silicate compositions with average diameter in the range 200 to 300 nm. Bioactivity testing demonstrated that the nanofibres were covered by a foamy and porous layer of amorphous calcium phosphate after immersion in SBF for five days. The laser spinning technique is thus proposed as an effective method of producing bioactive glass nanofibres in desired compositions. As a summary of the fabrication methods discussed in this section, their advantages and disadvantages are listed in Table 6.1. Table 6.1 Methods of fabrication for nano-scaled bioactive glasses showing their advantages and disadvantages Shape of the nano-structure

Method of fabrication

Advantages

Disadvantages

Particle

Sol-gel

Controlled porosity, controlled pore size and surface area, homogeneous products with high purity, simple operation, low cost of precursors

Problems with residual solvent and water, high calcination temperature. Relatively time consuming and limited to some compositions

Microemulsion

Ability to synthesize nanosized particles of organic and inorganic composition with minimum agglomeration, low processing temperature

Low production yield and the use of a large quantity of oil and surfactant phases

Gas phase synthesis method

High chemical homogeneity, High processing higher production rates, no temperatures need to additional source of energy

Sol-gel and electrospinning

Homogeneous products with high purity, simple operation

Difficulties in controlling many parameters that affect the fibre diameter

Laser spinning

Higher production rates, controllable chemical compositions, very fast process, high chemical homogeneity, long fibres with extraordinarily high length to diameter ratios

High processing temperatures

Fibre

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6.4

Applications of nanoscale bioactive glasses

6.4.1 Composites for tissue engineering scaffolds Inorganic phases can be added to different polymer matrices in the form of micron-sized or nanoscale particles or fibres to form bioactive composite materials [16]. The size of the filler particles affects the effective mechanical and biological properties of the composites. For example, the introduction of nanoscale fillers with desired morphology should lead to a higher increase of mechanical strength and stiffness in comparison to neat polymers and to composites with micronsize reinforcement, provided the nanoparticles or nanofibres are dispersed homogeneously in the matrix [89]. Additionally, in the case of bioactive silicate glass nanoparticles and nanofibres, they will induce higher alkalinity when compared to commercially available (µm-sized) bioactive glass particles [90]. The larger specific surface area of the nanoparticles will increase interface-related effects, contributing to improved bioactivity. Further advantages of the use of nanoscale fillers in composites for medical applications, in particular in relation to surface effects for tissue engineering scaffolds, were discussed in Section 6.2. Synthetic polymer/nanoparticulate bioactive glass composites Poly(3-hydroxybutyrate) (P(3HB))/nanoparticulate bioactive glass composites with different filler concentrations have been fabricated by solvent casting [30]. The addition of nanoparticles was shown to have a significant stiffening effect in comparison with the µm-sized counterparts. Moreover surface effects induced by the nanoparticles (nanotopography) considerably improved total protein adsorption compared to the unfilled polymer and to composites containing micron-sized bioactive glass particles [30]. In addition, cell proliferation investigations using osteoblast-like cells confirmed enhanced cytocompatibility of the P(3HB)/bioactive glass composites [30]. Misra et al. [91] have also reported that the ALP activity of MG-63 cells on nanoparticulate bioactive glass/P(3HB) composites was considerably higher than that on the control surface confirming the suitability of the composites for bone tissue engineering. Zheng et al. [92] developed composites using poly(hydroxybutyrate-2-co2-hydroxyvalerate) (PHBV) containing biomimetically synthesized nanosized bioactive glass (BMBG) (CaO-P2O5-SiO2). The porous composites were shown to be bioactive, and cell attachment studies indicated that the material has attractive biomineralization and cell biocompatibility [92]. Composites combining poly(L-lactic acid) and sol-gel-derived bioactive glass-ceramic (BGC) nanoparticles were fabricated by Hong et al. [34] using thermally induced phase separation. It was shown that composites containing BGC nanoparticles with lower phosphorous and higher silicon content had higher bioactivity than that of the BGC with lower silicon and higher phosphorous content [34]. The effect of nanoparticle content on the properties of nanocomposite

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scaffolds has been also investigated [35] and an improvement of the mechanical properties was measured. A similar system was developed by El-Kady et al. [93] using sol-gel derived bioactive glass nanoparticles and poly(L-lactide) (PLA). In this system, it was shown that the scaffold’s pore size decreased with the increase of glass nanoparticles content. A recent approach to improve the mechanical properties of nanoparticulate bioactive glass/PLLA composites by using solvent evaporation has been reported by Liu et al. [36, 37]. It was shown that surface modification of nanosized bioactive glass particles by grafting organic molecules or polymers is a convenient solution to improve the mechanical properties of the composites. Moreover surface modified bioactive glass/PLLA composites exhibited much better cell proliferation ability than non-modified bioactive glass/ PLLA composites or pure PLLA [36, 37]. Natural polymer/bioactive glass nanocomposites Polysaccharides (starch, chitin, chitosan) and proteins (silk, collagen) are candidate natural polymers for preparing nanocomposites for biomedical applications. Peter et al. [94, 95] have synthesized α-chitin/sol-gel derived bioactive glass-ceramic nanoparticle and chitosan/sol-gel derived bioactive glassceramic nanoparticle composite scaffolds by using lyophilization technique. Macroporous composite scaffolds with pore size in the range 150 to 300 µm were fabricated (Fig. 6.4) [95]. In vitro studies showed the deposition of apatite and the attachment of osteoblast-like cells (MG-63) on the surface of the composite scaffolds [94, 95]. More recently, the same authors [96] have fabricated chitosan– gelatine/nano-sized bioactive glass-ceramic nanocomposite scaffolds by through freezing and lyophilization technique. It was reported that high surface area of

6.4 SEM images showing the macro porous microstructure of composite scaffold synthesized from chitosan/sol-gel derived bioactive glass-ceramic nanoparticle by using lyophilization technique (a). Pore size ranged from 150 to 300 µm, and nBGC particles were on the chitosan matrix (b) (reproduced from ref. [95] with permission).

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bioactive glass-ceramic nanoparticles resulted in increasing protein adsorption especially adhesive proteins [96]. New porous bioactive nanocomposites combining sol-gel derived bioactive glass nanoparticles (BG), collagen (COL), hyaluronic acid (HYA) and phosphatidylserine (PS) have been fabricated by a combination of sol-gel and freeze-drying methods [97]. A bioactive nanocomposite was also synthesized by cross-linking collagen and HYA by using 1-ethyl-3-(3-dimethylaminopropyl) carbodiimide (EDC) and N-hydroxysuccinimide (NHS). Biomineralization, degradation in SBF and mechanical strength of the EDC/NHS-cross-linked BG-COL-HA-PS composite scaffolds were better than those of the scaffolds without HYA, PS and the cross-linking process. PS and HYA can contribute to regulating the biomineralization process, inducing HA to precipitate on the surface of the composites. The in vivo bone regeneration ability of the EDC/NHS-cross-linked BG-COL-HA-PS composite scaffolds was investigated by Xie et al. [98] using a rabbit radius defect model. X-ray and histological studies showed the bone regeneration ability of both nanocomposites and for nanocomposites combined with growth factors (BMP). Moreover, nanocomposites containing BMP showed better ectopic bone formation ability [98]. In related studies, Luz and Mano [99] have proposed a new composite membrane combining chitosan with sol-gel derived bioactive glass nanoparticles based on both ternary (SiO2-CaO-P2O5) and binary (SiO2-CaO) systems. It was reported that bioactive glass nanoparticles could be distributed homogeneously in the biodegradable polymeric membrane because of their reduced size, thus promoting bioactivity [99]. Nanocomposites containing bioactive glass nanofibres Kim et al. [47] have developed PLA composites filled with sol-gel-derived bioactive glass as a nanoscale composite fibre using electrospinning (ES). The in vitro bioactivity and osteoblast responses of the developed nanocomposites were studied [48]. These nanocomposites showed excellent bioactivity, inducing CaP precipitation within 24 hours of immersion in SBF. The results have also been confirmed by Noh et al. [100] in a similar study. Kim et al. [101] also developed BGNF-collagen nanocomposite both in the form of a thin membrane and as macroporous scaffold. BGNF-collagen nanocomposites were seen to exhibit high bioactivity, which was assessed by the rapid formation of bonelike apatite minerals on their surfaces when immersed in SBF. Moreover, the nanocomposites enhanced the adhesion and growth of human osteoblast-like cells [101]. In related research, Lee et al. [102] have produced poly(e-caprolactone) (PCL)/ sol-gel-derived BGNF nanocomposites. The glass nanofibres were distributed homogeneously in the PLC matrix, showing a much rougher surface than the pure PCL. The precipitated apatite covered the surface of the nanocomposite membrane almost completely after immersion in SBF for 14 days. Osteoblastic cells

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6.5 Surface and cross-sectional images of PCL/bioactive glass nanocomposites containing 20 wt% BGP (a, b) and 20 wt% BGNF (c, d) (reproduced from ref. [46] with permission).

(MC3T3-E1) were seen to spread better and to grow with many cytoplasmic extensions, showing improved proliferation behaviour in comparison to those on the pure PCL membrane [102]. Jo et al. [46] have fabricated (PCL)/sol-gel derived BGNF composites exhibiting a highly homogeneous BGNF distribution (Fig. 6.5). This microstructure resulted in a significant improvement of the biological and mechanical properties of the PCL/BGNF composites, compared to that of the micron-sized ones. Multifunctional composite scaffolds The desired combination of biocompatibility of biodegradable polymers and bioactivity of bioceramics can be achieved by preparation of porous polymer/ ceramic composites by different methods, as mentioned above [16]. However, the need for advanced scaffold systems has prompted the addition of different functionalities into the materials to be able to closely mimic the natural bone’s structure and properties. Properties such as bioactivity, mechanical competence, electrical or magnetic conduction, growth factor and drug delivery, antioxidative effects and antibactericidal behaviour are being considered for designing new multifuncational scaffolds. Misra et al. [103] have prepared P(3HB)/micro scaled Bioglass® 3D composite scaffolds using the conventional solvent casting/

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particulate leaching technique and by employing commercially available sugar cubes as the porogen preform. Functionalized P(3HB)/micro scaled Bioglass® 3D composite scaffolds were fabricated by the systematic addition of appropriate amounts of vitamin E and multiwalled carbon nanotubes (MWCNTs) in the polymer solution and by sonicating the mixture before impregnating it onto the sugar sacrificial preforms. Studies indicated that each additive was able to contribute to the overall performance of the composite, i.e. Bioglass® particles imparting bioactivity, vitamin E improving protein adsorption and finally MWCNTs inducing electrical conductivity. P(3HB) composite scaffolds were also produced, using nanoscale bioactive glass particles to enhance the antimicrobial properties. Bactericidal studies revealed that nanoscale bioactive glass particles ensure both biocompatibility and enhanced antimicrobial properties. It was also concluded that the higher specific surface area of NBG particles played a vital role in imparting the desired antimicrobial activity [103]. The novel properties of nanoscaled bioactive glasses make them good candidates for the further development of advanced multifunctional 3D scaffolds. However, it is clear that the present status of research and development in the field of multifunctional 3D scaffolds is still at the starting point for overcoming some of the limitations of the biomaterials currently used in bone tissue engineering. In future, it should be possible to design multifunctional 3D composites with incorporation of NBG (nanoparticles, nanofibres or both) for a variety of healthcare applications beyond the general field of tissue engineering therapeutics [104], i.e. ranging from drug delivery to biosensing devices.

6.4.2 Applications in dentistry Bioactive glasses have been used in dentistry for de- and remineralization of dentin, root canal disinfection, restorative dental applications, augmenting alveolar ridges and for treating periodontal pockets and dentin hypersensitivity [10–13, 105–107]. Such favourable properties of bioactive glasses as bioactivity, ability to mineralize dentine and their antimicrobial effects make them interesting materials for dental applications. Bioactive glasses are for example promising candidates for the remineralization of human dentin and have potential as a filler component in mineralizing restorative materials [108]. However, the application of bioactive glass as a remineralization agent in dental practice has been limited owing to its relatively long reaction times. An alternative approach is to decrease glass particle size and thus to increase the surface area, which should enhance dissolution of ions from the glass resulting in an acceleration of the remineralization of dentin. Vollenweider et al. [90] investigated the in vitro remineralization capabilities of bioactive glass nanoparticles (45S5 composition) fabricated by flame spray synthesis (see Section 6.3.3). It was reported that NBG treatments of 10 and 30 days resulted in a markedly higher

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mineral content compared to equivalent treatments with suspensions of conventional (µm-size) bioactive glass particles. The mineral content of dentin remineralized for 30 days with NBG was equivalent to natural dentin. Kernen et al. [109] also reported the rapid remineralization rate of flame spray-derived nanoparticulate 45S5 bioactive glass. It was found that NBG-treated discs incubated for 7 days showed a degree of mineralization comparable to the native control disc [109]. Bioactive glasses are also interesting materials in dentistry because of their antimicrobial effect in closed systems [106, 107]. The antimicrobial effect of bioactive glasses has been attributed to their ability to raise pH in an aqueous environment [110]. Studies with conventional bioactive glasses showed that their antibacterial efficacy in human teeth is inferior to that of calcium hydroxide, the gold standard material [105]. Some attempts have been made to increase the antimicrobial efficacy of bioactive glasses for example incorporating silver into the glass network [111–113]. As an alternative way, Waltimo et al. [108] used nanometric 45S5 bioactive glass to improve the antibacterial properties against different Enterococcus faecalis strains in a direct exposure model. The antibacterial effect of nanoparticulate bioactive glass appears to be directly linked to its high surface area, and thus the resulting release of ionic components in solution [108]. Gubler et al. [114] also investigated the antibacterial efficiency of nanometric bioactive glasses with the compositions 28S5, 45S5 and 77S. Studies showed that in addition to the high surface area of bioactive glass nanoparticles, the antimicrobial effect is also related to the amount of sodium and thus the resulting alkaline environment. Microorganisms remaining or re-entering the root canal system are the main cause of post-treatment disease following root canal treatment [115]. Alkaline capacity and the antibacterial properties of bioactive glasses make them potential candidates for root canal disinfection treatments. The possibility of using 45S5 bioactive glass nanoparticles as root canal disinfectants has been investigated [13]. It was reported that not only the specific surface area, but also the total mass of material per volume in bioactive glass slurries is important for root canal disinfectant treatments [13]. More recently, Mortazavi et al. [116] have synthesized 58S, 63S, and 72S bioactive glass nanopowders using the sol-gel technique, and investigated the antibacterial effects of bioactive gel glass nanoparticles on aerobic bacteria. It was found that 58S and 63S bioactive glass nanopowders had antibacterial activity even at concentrations lower than those currently used in clinical applications. However, 72S bioactive glass nanopowder showed no antibacterial effect because of the high SiO2 content. It was concluded that bioactive glass nanopowders with antibacterial properties could be considered for the treatment of oral bone defects and disinfection of the root canal [116]. Besides the alkaline capacity and antibacterial properties, a certain level of radiopacity is necessary for a clear distinction between the material and the surrounding tissues in clinical practice. To the authors’ knowledge, Mohn et al.

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6.6 Transmission electron microscopy images of classical nanometric bioactive glass (a) and bioactive glass (flame derived) containing 50 wt% bismuth oxide (b) (reproduced from ref. [45] with permission).

[45] were the first to develop flame spray-derived bioactive glass nanoparticles including bismuth oxide as a radiopacifier. Flame spray synthesis resulted in novel bioactive glass nanoparticles with sticking bismuth oxide (2 to 7 nm) onto the surface (Fig. 6.6) [45]. The studies revealed that bioactive glasses with bismuth oxide have high in vitro bioactivity, alkaline capacity and radiopacity, which make them potential bioactive root canal dressing or filling materials [45]. The various primary positive results regarding the bioactivity and antibacterial properties of bioactive glass nanoparticles make them attractive materials in dentistry, notably in dentin regeneration and root canal disinfection, as briefly discussed in this section. Furthermore, bioactive glass particles can be potential candidates as nanofillers in synthesizing new dentin composites to improve the bioactivity, radiopacity and mechanical properties of the composite systems.

6.4.3 Bioactive coatings and other orthopaedic applications The clinical applications of bioactive glasses have been limited largely to nonload bearing parts due to their inferior mechanical properties. In order to solve the lack of strength and fracture toughness of bulk bioactive glasses for loadbearing applications, coatings of these materials on metallic prostheses have been developed [14, 15]. Several techniques have been investigated for the preparation of glass coatings, such as enamelling [117], plasma spraying [118], ion beam sputtering [119], sol-gel [120], pulsed-laser deposition (PLD) [121], electrophoretic deposition (EPD) [122–124] and laser cladding [125]. Several

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factors influence the properties of bioactive glass coatings such as coating thickness, crystallinity, chemical purity, porosity and adhesion [126]. Novel bioactive glass nanoparticles represent a significant opportunity to develop improved bioactive and nanostructured orthopaedic coatings. As discussed above, bioactive glass nanoparticles have better bioactivity, antibacterial properties and enhanced bone-forming cell adhesion and proliferation compared with micronsized particles. Recently, Fathi et al. [49] prepared bioactive glass nanopowders for coating 316L stainless steel by the sol-gel technique. The particle size of the bioactive glass was less than 100 nm. Nanopowders were seen to attach very well to the stainless steel substrate and crack-free homogeneous coatings were obtained. The most important finding was the improved corrosion resistance, biocompatibility and bone bonding ability of the metallic substrate [49]. Sol-gel derived bioactive glass nanoparticles have also been used to coat different materials, with the intention of combining improved mechanical properties and high bioactivity in one material [33, 49]. Bioactive glass nanoparticle coating using the sol-gel technique has been applied for example on the struts of porous HA scaffolds by Esfahani et al. [33], in order to improve the mechanical properties of the scaffold. It was shown that the compressive strength of scaffolds increased and a new crystalline phase was detected with the increase in sintering temperature. According to Esfahani et al. [33], crystallization occurred in bioactive glass nanoparticles resulting in an improvement of the mechanical properties. Fathi et al. [50] prepared a sol-gel derived bioactive glass nanopowder coating for treating oral bone defects. It was reported that crack-free and homogeneous bioactive glass coatings were achieved with no observable defects. In vitro studies showed that the bioactive glass coating induced the formation of a weakly crystalline hydroxyapatite-rich layer on the bioactive glass surface as an indication of bioactivity [50]. Furthermore, novel nanoscaled coatings can be used for dental hard tissues to improve the aesthetic appearance and to protect them against ageing processes such as wear and cracking [127]. In related investigations, Couto et al. [51] have fabricated polycation (chitosan) and bioactive glass nanoparticle multilayer coatings by the layer-by-layer (LbL) technique. In this method, a glass panel was first dipped in a 1% (v/v) acetic acid solution containing 0.4% (w/v) of chitosan, then into water, then into a 0.4% (w/v) bioactive glass aqueous solution and finally into water again. The multilayer was achieved by repeating the dipping process sequentially. The scheme of this process is shown in Fig. 6.7 [51]. Spherical nanoparticles with sizes in the range 30 to 100 nm were dispersed homogeneously on the surface of the multilayered coatings. The organic component, chitosan, provided viscoelastic properties to the final coating, and the bioactive glass induced the required bioactivity for improving attachment of the coating to the bone. The coating was evaluated by immersing specimens in simulated body fluid, and the formation of a surface layer of hydroxyapatite, as a marker of bioactive behaviour, was confirmed.

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6.7 Scheme of the layer-by-layer coating procedure for production of chitosan/nano-bioactive glass (BG) coatings (a). After each cycle it is expected that a complete nanostructured bi-layer will be formed, resulting in a multilayered film adsorbed at the surface after repeated layering, according to Couto et al. (b) (adapted from ref. [51]).

6.4.4 Applications in drug delivery and nanomedicine Research efforts are increasingly devoted to the development of new drug-delivery systems with greater efficiency, lower toxicity, controlled and prolonged drug release capability, predictable therapeutic response and safety [128]. Drug-delivery systems based on particulate carriers will benefit from nanoscale dimensions, as particles need to reach the given location in the body based on the size of vessels of the human circulatory system in order to transport the desired drug molecules to the targeted cells or tissues and to release them in a controlled manner [128, 129]. Ordered mesoporous bioactive glasses and composites of bioactive glasses and polymers have been proposed as delivery systems for antibiotics and other antibacterial agents, anti-inflammatory drugs, fluoride ions, vascular endothelial growth factor, proteins and peptides [20, 22–24, 130–136]. Ionic products of bioactive glass dissolution also increase the proliferation of human osteoblasts and induce insulin-like growth factor II mRNA expression as well as protein synthesis [117, 137]. It is now well established that ionic dissolution products released from bioactive glass upregulate seven families of genes that control osteogenesis [17, 138, 139] and promote angiogenesis [22–24, 132]. Besides these advantages,

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however, the drug-storage capacity of conventional bioactive glasses is generally relatively low and the drug delivery is difficult to control [140]. An alternative way of overcoming these disadvantages is the synthesis of nanoscaled bioactive glasses because these offer a number of distinct advantages for drug adsorption and delivery over conventional (µm-sized) particles. Nanoparticles have in general relatively higher intracellular uptake compared to microparticles as discussed above [141]. The efficiency of uptake of nanoparticles has been reported to be 15to 250-fold greater than that of µm-sized particles [142]. Due to their sub-micron size, nanoparticles can penetrate deep into tissues through fine capillaries, and are generally taken up efficiently by cells [143]. This behaviour enables the efficient delivery of therapeutic agents to target sites in the body [144]. Nanosized particles in particular can affect the efficiency of the immune response [145]. Moreover, nanoparticles show larger storage capacity than microparticles because of their high surface area. Yun et al. [146] have synthesized mesoporous bioactive glass (MBG) nanospheres of SiO2-CaO-P2O5 composition, which exhibit large specific surface area and pore volume. The size of the nanospheres was shown to depend on the amount of CaO incorporated and it can be controlled over the range of diameters from 20 to 200 nm (Fig. 6.8) [146]. In vitro studies indicated that the MBG nanospheres with high amount of CaO show good in vitro bone-forming bioactivity as well as favourable in vitro biocompatibility. Cytotoxicity tests of MBG nanospheres with a high amount of CaO revealed that they have no negative effect on macrophage cell behaviour. It was concluded that MBG nanospheres are promising materials for drug delivery applications [146]. Further research is focusing on the development of nano-carriers to enhance the effective targeting in specific cancer regions without destroying or affecting the viability of nearby normal tissues [147]. It was shown that ferromagnetic bioactive glasses and glass-ceramics containing magnetite could be used for hyperthermia treatment of cancer [148]. Recently, Wang et al. [149] developed a novel magnetic degradable material, adding Fe ions to bioactive glass (Na2O-CaO-P2O5-SiO2) as thermoseed for hyperthermia cancer therapy using the sol-gel method. The particle size was measured in the range 50 to 100 nm and spherical and rod shape particles were developed; the morphology was dependent on the amount of iron ion added. In a biocompatibility test, these magnetic bioactive glasses had no significant influence on cell viability and mediated low cytotoxicity when cultured with fibroblasts. The material was cultured in vitro with either human Caucasian lung carcinoma (A549) or normal HFL1 fibroblast cells to examine the hyperthermia effect. In vitro cell culture studies clearly indicated that after exposure to an alternating magnetic field, the cell number of human Caucasian lung carcinoma cells (A549) significantly decreased, while normal HFL1 fibroblasts were still alive without severe damage. It was stated that magnetic degradable bioactive glass incorporated with Fe ions would be a potential candidate for tumour hyperthermia treatment in future [149].

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6.8 FEG-SEM images of MBG-Ca4 (a), MBG-Ca8 (b), MBG-Ca16 (c) and a comparison of their particle sizes as a function of the amount of calcium (d), according to Yun et al. (reproduced from ref. [146] with permission).

Smart hydrogels as injectable scaffolds are gaining relevance for several tissue engineering applications since they enable the encapsulation of cells and bioactive agents in a biodegradable matrix to be delivered through minimally invasive procedures [150, 151]. To the authors’ knowledge, Couto et al. [52] were the first to develop injectable biodegradable materials with bioactive glass nanoparticles in order to produce thermo-responsive hydrogels for orthopaedic reconstructive and regenerative medicine applications. In their study, chitosan-β-glycerophosphate salt formulation was combined with sol-gel derived bioactive glass nanoparticles to synthesize novel thermo-responsive hydrogels. The inner structure of the hydrogels was characterized by using cryogenic scanning electron microscopy (cryoSEM) and it was found that bioactive glass nanoparticles were well dispersed in the organic matrix (Fig. 6.9) [52]. In vitro bioactivity tests showed that the bioactive glass nanoparticles incorporated in the chitosanbased thermo-responsive system induced the formation of bone-like apatite clusters that are well integrated in the hydrogel organic structure. It was also observed that the density of the apatite precipitates increased with increasing © Woodhead Publishing Limited, 2011

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6.9 CryoSEM images of the surface fracture of the BG-30% hydrogel: global view, where the arrows indicate some clusters of BG particles (a); magnified image, where the arrows indicate small agglomerates of BG nanoparticles deposited on the hydrogel walls (b): in the upper right-hand side there is part of a cluster of nanoparticles involved with chitosan (reprinted from ref. [52] with permission).

NBG content and soaking time in SBF, which indicates that the system could promote a positive contact with surrounding tissue upon injection in a bone defect. The high surface area, enhanced bioactivity and antibacterial properties of nanoscaled bioactive glasses make them promising materials for slow and targeted drug delivery systems. However, although significant advantages of nanoscaled bioactive glasses are expected in combined drug delivery and regenerative medicine approaches, as discussed above, only limited work has been reported to date. Moreover a large amount of biological information is necessary to fully understand the drug-delivery functions of these novel systems. It is clear that using bioactive glass nanoparticles in drug delivery systems and nanomedicine is in its early years and comprehensive research should be carried out to measure and assess potential exposure risks of nanoparticles to patients. In the future, the challenge would be to develop the next generation of advanced smart biomaterials with the integration of extended functions such as antioxidative and anticancer functions and in several of these novel applications nanoscale bioactive glasses will be the materials of choice. Table 6.2 summarizes recent investigations on the biomedical application areas of nanoscale bioactive glasses. The novel properties of nanoscale bioactive glasses make them promising materials for a variety of healthcare applications, many of which are bound to expand in the near future as more in-depth and relevant knowledge from basic research is generated.

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Method of fabrication

Solvent casting

Thermally-induced phase-separation

Solvent evaporation

Freezing/lyophilization

Freeze-drying

Freeze-drying

Sol-gel/ES

Sol-gel/ES

Sol-gel/ES

Sol-gel/ES

Solvent casting/particulate leaching technique

Flame spray synthesis

Composition

P(3HB)/NBG

PLA/NBG

PLLA/surface modified NBG

Chitin/NBG and chitosan/NBG

NBG-COL-HYA-PS

EDC/NHS-crosslinked NBG-COL-HA-PS

PLA/BGNF

Collagen/BGNF

PCL/BGNF

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PCL/BGNF

P(3HB)/NBG

NBG

NBG particles

3D porous scaffold

Porous matrix

Membrane

Membrane and porous scaffold

3D porous scaffold

Porous scaffold

Porous scaffold

Porous scaffold

Porous scaffold

3D porous scaffold

2-D films

End product shape

Rapid remineralization rate

Enhanced biocompatibility and antimicrobial properties

Significant improvement of the biological and mechanical properties and in vivo animal test results showed bone-forming ability

Rough surface improved proliferation behaviour

High bioactivity and good cell adhesion and growth

Excellent bioactivity and good osteoblast response

Ability of bone regeneration

Better biomineralization, mechanical strength, cell attachment and proliferation ability

Improvement in cell adhesion and proliferation and increasing protein adsorption

Less nanoparticle aggregation, improved mechanical properties, bioactivity, cell adhesion and growth

Improved bioactivity and mechanical properties

Enhanced bioactivity, cell adhesion and growth

Key results achieved

[109]

[103]

[46]

[102]

[101]

[47, 48]

[98]

[97]

[94, 95, 96]

[36, 37]

[34, 35]

[30, 91]

Reference

Table 6.2 Selected biomedical studies carried out on nanoscaled bioactive glasses, covering a wide range of both in vivo and in vitro investigations

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Flame spray synthesis

Flame spray synthesis

Sol-gel

Sol-gel

Sol-gel

LbL technique

Sol-gel

Sol-gel

Combining sol-gel derived NBG with chitosan–βglycerophosphate salt formulation

NBG

NBG

NBG

NBG

NBG

Chitosan/NBG

NBG

DP-NBG with Fe

Chitosan/NBG

Hydrogel

Sphere and rod shape particles

Nanosphere

Nanoparticle multilayer coatings

Nanopowder coating

Porous scaffold

Nanopowder coating

NBG particles

NBG particles

[108, 114]

[33]

[49]

[45]

Promoting a positive contact with surrounding tissue upon injection in a bone defect

Potential candidate for tumour hyperthermia treatment

Good in vitro bone-forming bioactivity and biocompatibility

Provided viscoelastic properties and bioactivity to the final coating

[52]

[149]

[146]

[51]

Promote bone formation in osseous defects and [50] bone grafting to improve the long-term prognosis of dental implants

Improved mechanical properties without using any polymeric materials

Improvement of the corrosion resistance, biocompatibility and bone bonding ability of the metallic substrate

High in vitro bioactivity, alkaline capacity and radiopacity

Improved antibacterial properties

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Conclusions

The preparation of bioactive glasses in nanoparticle and nanofibre form has recently become feasible by advances in a variety of processing methods. Nanoscale particulate and nanofibre bioactive glasses have shown advantages over conventional (micron-sized) bioactive glasses due to their large surface area and enhanced solubility as well as bioreactivity. In addition, their application leads to nanotopographic surface features in biomaterials, e.g. scaffolds and coatings. These nanomaterials and their advantages have also inspired researchers to investigate a range of new biomedical applications. In this chapter, new developments in the fabrication techniques of nanoscale bioactive glass particles and fibres and the biomedical applications of these novel materials have been reviewed and comprehensively discussed. Substantial advantages of nanoscale bioactive glasses compared to conventional (µm-scale) bioactive glasses were demonstrated in particular for bone tissue engineering, dentistry, orthopaedic coatings, antibacterial materials and drug delivery systems. Current research shows that it is possible to design improved 3D composite scaffolds by incorporating nanoscaled bioactive glasses (and other additives), which exhibit multifunctionalities such as bioactivity, electrical conduction and antibacterial behaviour. These 3D multifunctional composite scaffolds will have applications in tissue engineering therapeutics including drug delivery and biosensing functions. Moreover, the intrinsic properties of nanoscaled bioactive glasses make them promising materials in dentistry, especially for dentin regeneration, root canal disinfection and as nanofillers in new dentin composites to improve the bioactivity, radiopacity, antibacterial effect and mechanical properties of the materials. Other applications of nanoscale bioactive glasses in nanomedicine are being explored. The new challenge in this field is to develop the next generation of smart biomaterials with the integration of extended functions such as antioxidative and anti-cancer functions, where nanoscaled bioactive glasses may represent a key component for novel biomedical devices. It is anticipated that further research efforts will lead to relevant basic knowledge and fundamental information on these novel nanomaterials, e.g. cell biologydriven research, which will prompt the further expansion of their biomedical applications.

6.6

Acknowledgment

ME gratefully acknowledges the financial support from The Scientific and Technological Research Council of Turkey (TUBITAK), Turkey.

6.7

References

1. Hench L. L., Splinter R. J., Allen W. C., Greenlee T. K. Bonding mechanisms at the interface of ceramic prosthetic materials. J Biomed Mater Res 1971; 5(6):117–141.

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128. Zhang C., Li C., Huang S., Hou Z., Cheng Z., Yang P., Peng C., Lin J. Self-activated luminescent and mesoporous strontium hydroxyapatite nanorods for drug delivery. Biomaterials 2010; 31:3374–3383. 129. Slowing I. I., Trewyn B. G., Giri S., Lin V. S. Y. Mesoporous silica nanoparticles for drug delivery and biosensing applications. Adv Funct Mater 2007; 17:1225–1236. 130. Ladron de G. F. S., Ragel C. V., Vallet-Regi M. Bioactive glass-polymer materials for controlled release of ibuprofen. Biomaterials 2003; 24:4037–4043. 131. Domingues Z. R., Cortes M. E., Gomes T. A., Diniz H. F., Freitas C. S., Gomes J. B., Faria A. M. C., Sinisterra R. D. Bioactive glass as a drug delivery system of tetracycline and tetracycline associated with betacyclodextrin. Biomaterials 2004; 25:327–333. 132. Leach J. K., Kaigler D., Wang Z., Krebsbach P. H., Mooney D. J. Coating of VEGFreleasing scaffolds with bioactive glass for angiogenesis and bone regeneration. Biomaterials 2006; 27:3249–3255. 133. Bergeron E., Marquis M. E., Chrétien I., Faucheux N. Differentiation of preosteoblasts using a delivery system with BMPs and bioactive glass microspheres. J Mater Sci: Mater Med 2007; 18:255–263. 134. Li S., Nguyen L., Xiong H., Wang M., Hu T. C., She J. X., Serkiz S. M., Wicks G. G., Dynan W. S. Porous-wall hollow glass microspheres as novel potential nanocarriers for biomedical applications. Nanomedicine: Nanotechnology, Biology, and Medicine 2010; 6:127–136. 135. Wu C., Ramaswamy Y., Zhu Y., Zheng R., Appleyard R., Howard A., Zreiqat H. The effect of mesoporous bioactive glass on the physiochemical, biological and drugrelease properties of poly(DL-lactide-co-glycolide) films. Biomaterials 2009; 30:2199–2208. 136. Chen Q. Z., Rezwan K., Armitage D., Nazhat S. N., Boccaccini A. R. The surface functionalization of 45S5 Bioglass-based, glass-ceramic scaffolds and its impact on bioactivity, J Mater Sci: Mater Med 2006; 17:979–987. 137. Xynos I. D., Edgar A. J., Buttery L. D. K., Hench L. L., Polak J. M. Ionic products of bioactive glass dissolution increase proliferation of human osteoblasts and induce insulin-like growth factor II mRNA expression and protein synthesis. Biochem Bioph Res Co 2000; 276:461–465. 138. Xynos I. D., Edgar A. J., Buttery L. D., Hench L. L. Gene-expression profiling of human osteoblasts following treatment with the ionic products of Bioglass® 45S5 dissolution. J Biomed Mater Res 2001; 55:151–157. 139. Hench L. L. Genetic design of bioactive glass. J Euro Ceram Soc 2009; 29:1257– 1265. 140. Vallet-Regi M., Balas F., Arcos D. Mesoporous materials for drug delivery, Angewandte Chemie International Edition 2007; 46:7548–7558. 141. Desai M. P., Labhasetwar V., Walter E., Levy R. J., Amidon G. L. The mechanism of uptake of biodegradable microparticles in Caco-2 cells is size dependent. Pharm Res Adv 1997; 1568–1573. 142. Desai M. P., Labhasetwar V., Amidon G. L., Levy R. J. Gastrointestinal uptake of biodegradable microparticles: effect of particle size. Pharm Res 1996; 13:1838–1845. 143. Vinagradov S. V., Bronich T. K., Kabanov A. V. Nanosized cationic hydrogels for drug delivery: preparation, properties and interactions with cells. Adv Drug Del Rev 2002; 54:223–233. 144. Panyama J., Labhasetwar V. Biodegradable nanoparticles for drug and gene delivery to cells and tissue. Adv Drug Del Rev 2003; 55:329–347.

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145. Witasp E., Kupferschmidt N., Bengtsson L., Hultenby K., Smedman C., Paulie S., Garcia-Bennett A. E., Fadee B. Efficient internalization of mesoporous silica particles of different sizes by primary human macrophages without impairment of macrophage clearance of apoptotic or antibody-opsonized target cells. Toxicol Appl Pharmacol 2009; 239:306–319. 146. Yun H., Kim S., Lee S., Song I. Synthesis of high surface area mesoporous bioactive glass nanospheres. Mater Lett in press. 147. Kukowska-Latallo J. F., Candido K. A., Cao Z., Nigavekar S. S., Majoros I. J., Thomas T. P., Balogh L. P., Khan M. K., Baker J. R. Jr. Nanoparticle targeting of anticancer drug improves therapeutic response in animal model of human epithelial cancer. Cancer Res 2005; 65:5317–5324. 148. Bretcanu O., Spriano S., Vitale C. B., Verne E. Synthesis and characterization of coprecipitation-derived ferrimagnetic glass-ceramic. J Mater Sci 2006; 41: 1029–1037. 149. Wang T. W., Wu H., Wang W. R., Lin F. H., Lou P. J., Shieh M. J., Young T. H. The development of magnetic degradable DP-Bioglass for hyperthermia cancer therapy. J Biomed Mater Res A 2007; 83A(3):828–837. 150. Roy I., Gupta M. N. Smart polymeric materials: emerging biochemical applications. Chem Biol 2003; 10:1161–1171. 151. Phong A. T., Sarin L., Hurtb R. H., Webster T. J. Opportunities for nanotechnologyenabled bioactive bone implants. J Mater Chem 2009; 19:2653–2659.

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7 Bioactive glass containing composites for bone and musculoskeletal tissue engineering scaffolds S. VERRIER, AO Research Institute Davos, Switzerland, J. E . GOUGH, University of Manchester, UK and A. R. BOCCACCINI , University of Erlangen-Nuremberg, Germany and Imperial College London, UK

Abstract: Composites developed by combining bioactive glasses and biodegradable polymers are attractive materials for use in scaffolds for musculoskeletal tissue engineering due to their ability to be tailored for different applications. Scaffolds must be osteoconductive and osteoinductive in order to guide and encourage new bone formation. Additionally, neo-vascularization of the construct is required. This chapter focuses on bioactive glass-containing composites for tissue engineering, with emphasis on the in vitro and in vivo performance of these scaffolds. A wide range of cell types (primary or cell lines) from different origins have been considered for studies in vitro. Several studies have also been carried out in specific animal models (in vivo). Recent studies report the angiogenic potential of such composites in both cases. The evidence indicates that these composite materials are promising and of high interest for bone tissue engineering and musculoskeletal tissue regeneration. However, in vitro and in vivo understanding of these scaffolds is still limited, especially regarding their long-term degradation and ion release effects on the biological environment. Key words: bioactive glasses, composites, scaffolds, tissue engineering, in vitro, in vivo, angiogenesis, musculoskeletal tissue.

7.1

Introduction

One of the key areas within tissue engineering (TE) and regenerative medicine gaining increasing attention is related to bone and musculoskeletal tissue engineering and regeneration [1]. It is well known that critical size bone defects that occur due to trauma or disease are very difficult to repair via the natural growth of host tissue. A common TE approach to restore function to diseased or damaged bone tissue is to design combinations of functional cells and biodegradable scaffolds made from engineered biomaterials [1–3]. In musculoeskeletal tissue regeneration, for example intervertebral disc regeneration, scaffolds made of suitable engineered biomaterials are also being investigated in combination with relevant cells to tackle the increasing medical need in this area [4]. In common tissue engineering strategies, scaffolds are highly porous biomaterial engineered structures that serve as temporary 3D templates for cell adhesion, 162 © Woodhead Publishing Limited, 2011

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proliferation, migration, and ultimately for supporting the formation of new tissue [2, 3]. The viability of this approach for growing new functional bone depends not only on the osteogenesis potential of the construct but also on the ability to induce rapid vascularization upon implantation [5, 6]. The geometry of the biomaterial scaffold is very important for the success of this strategy: when cells have attached to the surface there must be enough space and open channels to allow for nutrient ingress, waste delivery, protein transport and vascular growth to occur, functions that are obtainable with an interconnected network of pores. Suitable biomaterials for the development of bone TE scaffolds are those exhibiting bioactive properties [1]. Bioactive materials show high surface reactivity, leading to strong chemical interactions with relevant physiological fluids that induce the development of tenacious bonds to bone through the biological interaction of collagen with the material surface. In this way, bioactive materials can transfer loads to and from living bone. The most investigated bioactive materials for bone TE are bioceramics, including specific compositions of silicate glasses and glass-ceramics, as well as hydroxyapatite (HA) and related calcium phosphate ceramics [7]. Silicate bioactive glasses were first developed by Hench and co-workers 40 years ago [8]. The first bioactive glass composition, universally known as 45S5 Bioglass® (BG) (composition in wt%: 45% SiO2, 24.5% Na2O, 24.5% CaO and 6% P2O5), has the approval of the US Food and Drug Administration (FDA). This glass has been successfully used, for example, in clinical treatments of periodontal diseases as bone filler as well as in middle ear surgery [7]. Other clinical applications of bioactive glasses have been proposed, for example in dentistry [9] or as coating on metallic orthopaedic implants [10, 11]. In the last 40 years tens of different bioactive glass compositions for different biomedical applications have been developed, and several traditional and modern applications of bioactive glasses are described in other chapters of this book. The common characteristic of these glasses is their high surface reactivity and their ability to bond to bone in a physiological environment [7]. More recently, the application of bioactive glasses in bone TE and regenerative medicine has received marked impulse from the TE research community [12–15] and it can be stated that bone TE represents one of the most exciting future clinical applications of bioactive glasses, e.g. to fabricate optimal scaffolds with osteogenic and angiogenic potential [16]. In the context of bone TE, bioactive silicate glasses have several attractive advantages in comparison to other bioactive ceramics, e.g. sintered hydroxyapatite. One important finding in 2000 [17] is related to the effect of dissolution products from bioactive glasses on the upregulation of the expression of genes that control osteogenesis, which is being actively investigated [16, 18–20]. Further in vivo and in vitro studies using 45S5 BG particles have shown encouraging results regarding the potential angiogenic effects of Bioglass®, i.e. increased secretion of vascular endothelial growth factor (VEGF) in vitro and enhancement of vascularization in vivo [21–23]. Bioactive glasses can also serve as vehicle for the controlled

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delivery of selected ions, which can act to control specific cell functions. In this context, bioactive glass compositions incorporating additional elements in the silicate network such as magnesium, strontium, boron, iron, silver, potassium or zinc are being developed [24–28]. Like most inorganic materials, the major disadvantage of bioactive glasses is their limited fracture strength and low fracture toughness (i.e. high brittleness). This problem can be tackled by combining bioactive glasses with biopolymers forming bioactive composite materials [29]. In the most common approach, bioactive glass particles or fibres are incorporated in biodegradable polymer matrices [29–32], this being the type of composite to be discussed in the present chapter, as illustrated in Fig. 7.1. Other possibilities include the coating of (porous) polymer scaffolds with bioactive glass particles [33] or the coating and impregnation of 3D bioactive glass scaffolds with biodegradable polymers (discussed in Chapter 5 of this book). Since the requirements for optimal scaffolds are manifold [34], it can be stated that a combination of degradable polymers and bioactive glasses represents an optimal approach in terms of achievable mechanical and biological performance towards improved TE scaffolds [29, 30]. A wide range of biodegradable synthetic polymers is available for developing composite scaffolds in combination with bioactive glasses [35, 36]. For example, poly(lactic-acid) (PLA), poly(glycolic-acid) (PGA) and their co-polymers poly(lactid-co-glycolic-acid) (PLGA) [37, 38] are being highly investigated. These polymers have extensive FDA approval history. Also

7.1 Schematic diagram showing combinations of biodegradable polymers and bioactive glasses to form composites [46].

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polyhydroxyalkanoates (PHAs), a family of microbial polyesters, are increasingly finding application for composite scaffolds development, in particular poly(3hydroxybutyrate) and poly(3-hydroxybutyrate-co-hydroxyvalerate) [39, 40]. PCL is another popular polymer being combined with bioactive glasses for composite scaffolds [41]. In addition, composite scaffolds are also being developed combining natural polymers, including collagen, chitosan, gelatine or alginate, with bioactive glasses (e.g. [42, 43]). For bone tissue engineering, porosity of ∼90% and pore size >100 µm are desirable, as well as high pore interconnectivity, in order to facilitate the attachment and proliferation of cells and the ingrowth of new tissue into the scaffold, as well as to enable mass transport of oxygen, nutrition and waste products [34, 44]. This chapter will discuss the materials science and technology of composites based on the combination of biodegradable polymers (specifically synthetic polymers) and bioactive glass particles for bone tissue engineering scaffolds. To demonstrate the versatility of the materials in musculoskeletal tissue engineering, the development and application of polymer/Bioglass® scaffolds for intervertebral disc regeneration will be also presented. In Section 7.2 the basics of the composite materials approach to tissue engineering scaffolds are described, including a discussion of typical composite scaffold fabrication technologies, scaffold microstructure and relevant properties achieved. Section 7.3 discusses the specific results of the in vitro and in vivo applications of these scaffolds in relation to bone regeneration and intervetebral disc regeneration, while Section 7.4 presents a discussion of the latest developments on cell/tissue response of bioactive composite scaffolds, highlighting the effect of the presence of bioactive glass particles within polymer matrices on cell growth and differentiation. The chapter finishes with a summary and the scope for future developments in the field (Section 7.5).

7.2

Composite materials approach to tissue engineering scaffolds

7.2.1 Advantages of composite materials Composite materials are made by combining two or more chemically distinct materials (metallic, ceramic, or polymeric) in the micron or nanoscale, which are separated by an interface. Fillers of different morphologies can be used to fabricate composite materials with biopolymer matrices, such as particulates, short fibres, continuous fibres and nanofillers (e.g. nanoparticles, nanofibres) [45] (Fig. 7.1). For applications in tissue engineering scaffolds, composites must exhibit a set of desired properties, such as adequate mechanical strength, tailored initial elastic modulus close to the elastic modulus of the tissue of concern and controlled degradation behaviour in vivo. Composite scaffolds must degrade at a predefined rate and must retain their structural integrity in vivo providing the necessary

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mechanical support for cell attachment and proliferation. Polymers are generally flexible exhibiting relatively low compression strength and stiffness, being in cases too flexible to meet the mechanical demands in surgery and in in vivo situations. On the other hand, bioactive glasses are too stiff and brittle compared to bone. Several reasons have been put forward for suggesting the combination of biodegradable polymers and bioactive glasses for tissue engineering applications [29–32]. In addition to the possibility of increasing the compression strength and stiffness of polymer scaffolds by incorporating bioactive glass inclusions, the availability of effective processing methods for porous polymer structures can also be exploited. For example, polymers can be easily fabricated into complex shapes and porous structures by a variety of well-established techniques [46]. Moreover, the addition of bioactive phases to bioresorbable polymers can be used to control polymer degradation behaviour, e.g. by buffering the pH of the nearby solution and hence controlling the acidic degradation of the polymer, in particular in case of polylactic acid. Here, dissolution products from bioactive glasses can alter the autocatalytic effect of the acidic end groups resulting from hydrolysis of the polymer chains. In addition, bioactive glass inclusions contribute to water absorption in the scaffold due to the increased number of internal interfaces formed between the polymer and the hydrophilic bioactive glass particles [47]. The incorporation of bioactive glass into a biodegradable polymer matrix will allow the composite to interact with the surrounding bone tissue inducing the formation of a strong bond with bone via the growth of a carbonate hydroxyapatite layer, as mentioned above. Composite materials represent thus a convenient alternative for fabricating tissue engineering scaffolds considering that their properties can be engineered to suit the mechanical and physiological demands of the host tissue, which can be controlled by varying the volume fraction, morphology and arrangement of the bioactive glass inclusions [29, 31, 32]. Two most commonly used bioactive glass inclusions in biopolymers for biomedical composites are fibres and particulates [29]. With the availability of bioactive glass nanoparticles [48] (see also Chapter 6) nanocomposites are also being considered for bone tissue engineering [49]. It has been shown that increased volume fraction and higher surface area-to-volume ratio of inclusions favour bioactivity [29]. For certain applications the incorporation of (nano)fibres is preferred to particles [50]. In addition, composite mechanical properties are determined by the inclusion shape and size as well as by the quality of the distribution of the particles or fibres in the matrix. Moreover, as in all composite systems, the reinforcement-matrix interfacial bonding is of major importance in influencing the final composite properties. In the case of highly porous scaffolds, particularly of interest for bone tissue engineering, is the scaffold porosity, e.g. pore volume, size, shape, orientation and connectivity, that strongly affects the mechanical properties and structural integrity of scaffolds. The porosity structure is related mainly to the fabrication method used, as discussed next.

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7.2.2 Processing methods Tissue engineering scaffolds must mimic the numerous functions of the natural extracellular matrix, including providing support for cell adhesion and migration, and organizing cells into 3D structure [34]. Scaffolds for bone and muscoloeskeletal tissue engineering are highly porous, thus porosity and pore structure significantly affect the characteristics of a scaffold [29, 51]. A minimum pore size is required for tissue ingrowth, and high 3D interconnectivity is necessary for access of nutrients, transport of waste products, better cell spreading and vascularization [29–31]. The pore structure and properties of scaffolds are dictated by the manufacturing process employed. There has been a broad range of scaffold manufacturing techniques developed recently, which can be applied for biopolymer/bioactive glass composites, including well known methods such as the use of porogens, chemical segregation, microsphere sintering, solvent casting, particulate (salt or sugar) leaching and thermally induced phase separation (TIPS), as well as a series of computer-assisted rapid prototyping techniques, such as three-dimensional printing and fused deposition modelling. In addition, electrospinning is being increasingly considered for use in forming nanofibrous structures. Each of the developed techniques has the ability to produce scaffolds with a different pore architecture, but they also have limitations with respect to specific properties than can be achieved, as discussed elsewhere [29, 34, 52]. Solvent casting with and without particle leaching [53, 54], thermally induced phase separation (TIPS) [55, 56] and solid free form fabrication methods [2, 57, 58] have been applied successfully to manufacture synthetic biopolymer-ceramic composite scaffolds. Solvent casting for production of composite scaffolds involves dissolving the polymer in an organic solvent, mixing with bioactive glass particles and casting the solution into a predefined 3D mould [54]. The solvent is subsequently allowed to evaporate. This is a simple processing technique not requiring specialized or expensive equipment. The disadvantages of the method include poor pore interconnectivity especially at low porosities and a difficulty in generating large (3D) structures (over 3 mm thick). Composite constructs are also fabricated by combining solvent casting, particle leaching and microsphere packing methods [54]. Polymer microspheres are first formed from traditional water oil/water emulsions, and scaffolds are then developed by mixing solvent, salt particles (porogens), bioactive glass particles and pre-hardened microspheres [59]. The sintered scaffolds present a wellintegrated microstructure with a porosity of 40%. The mechanical properties of these composites were found to be similar to those of cancellous bone. Biodegradable polymer scaffolds with very high porosities (∼97%) can be produced using the thermally induced phase separation (TIPS) technique [31]. The obtained scaffolds exhibit pores with anisotropic tubular morphology and high pore interconnectivity. The TIPS process has been used to produce composite

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scaffolds based on PLGA and PDLLA foams containing Bioglass® particles [31, 47, 60]. Micrographs showing the pore structure of PDLLA and PDLLA/ Bioglass® scaffolds developed by the TIPS method are shown in Fig. 7.2 [60]. Scaffolds obtained by TIPS usually exhibit oriented tubular pores of diameters of several hundred microns (>100 µm) and an isotropic pore network of smaller pore size (∼10 µm) connecting the large tubular pores. Solid freeform fabrication (SFF) techniques including 3D printing, selective laser sintering, multi-phase jet solidification, and fused deposition modelling (FDM) have been developed to manufacture tissue scaffolds for bone tissue engineering with specific structure and properties [61, 62]. SFF allows a high degree of pore interconnectivity and controlled morphology. The methods enable the fabrication of tailored scaffolds by incorporating patient-specific information as well as a designed microenvironment. Tissue geometry can be extracted from patient’s computed tomography (CT) or magnetic resonance imaging (MRI) data and reconstructed as a 3D model. However, a shortcoming of these methods is increased fabrication time and the requirement of relatively complex equipment compared with the methods described above.

7.2 Scanning electron microscopy images of the transversal section of TIPS foams, showing the typical homogeneous regions of (a) pure PDLLA foam, (b) PDLLA/2 vol% Bioglass® foam, (c) PDLLA/15 vol% Bioglass® foam (reproduced from [60] with permission from Elsevier).

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7.2.3 Mechanical properties Bioactive glass inclusions incorporated in biodegradable polymer matrices positively affect the mechanical properties leading to reinforcement of the scaffold structure. The enhancement of mechanical properties (compression strength, elastic constants) depends strongly on particle shape and size distribution; as well as on the quality of the inclusion distribution in the matrix and, crucially, on the strength of the inclusion-matrix interface. Current porous scaffolds incorporating bioactive glass particles (size <10 µm) are however at least one order of magnitude weaker than cancellous bone and orders of magnitude weaker than cortical bone [29]. The increase in stiffness and strength on addition of bioactive glass fillers is below expectations, which is probably due to the lack of a strong interfacial bonding between the bioactive glass particles and the polymer matrix. The bonding at the interface can be improved by using surfactants chemisorbed on the bioactive glass particle surfaces. It has been suggested that using surface functionalized nanoparticles, featuring a higher specific surface and thus a higher interface area, might increase the interfacial bonding strength [29]. However, the increase of interfacial bonding and introduction of surfactants can have a negative effect on degradation kinetics and cytotoxicity of the composites. The inclusion of bioactive glass nanoparticles into biopolymer matrices is being explored with the dual objective of improving their mechanical properties as well as of incorporating nanotopographic features that mimic the nanostructure of natural bone [49, 63–65]. In addition, composite scaffolds can be developed with extra functionalities beyond being simply a mechanical support for cells and new tissue, including intelligent surfaces capable of providing both chemical and physical signals to guide cell attachment and spreading.

7.3

In vitro and in vivo evaluation

7.3.1 Bioactive glass containing composites: current approaches As mentioned above, composite materials described in this chapter, based on the association of bioactive glasses with biopolymers, are of great interest owing to the considerable versatility and the wide range of properties they offer [29]. Polymer composition and bioactive glass content, as well as microstructure and macroscopic morphology (e.g. porosity, pore shape, pore size and orientation), can be adjusted and tailored to obtain the required mechanical performance and biological response. In addition, the requirements for degradation in vitro and in vivo and bio-integration with host tissue vary and can be adapted to specific requirements. Numerous composite materials are being developed and their performance as scaffolds suitable for bone tissue engineering applications are being investigated both in vitro and in vivo [66]. An overview of the responses

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observed when cells were cultured on several types of polymer/bioactive glass composites is given in Table 7.1, while Table 7.2 summarizes the in vivo responses.

Table 7.1 Summary of in vitro evaluation of biopolymer/BG scaffolds Polymer matrix

Cell type

Outcomes/remarks

References

PLGA

Rat-Mesemchymal Stem Cells (MSC)

Cell growth and differentiation

[67,68]

Mice fibroblast cell line (L929)

Cell adhesion and growth, ↑ VEGF synthesis, BG contents dependent

[69, 70]

Human osteosarcoma cell line (Saos-2)

Cell growth, differentiation, matrix synthesis + mineralization

[71, 72]

PGA

Rat fibroblast cell line (208F)

↑ VEGF synthesis

[21]

PDLLA

Human primary osteoblast/MSC

Cell growth and differentiation, matrix synthesis + mineralization, BG contents dependent

[73, 74, 75]

Human foetal osteoblast

Differentiation, matrix synthesis [76] + mineralization

Human osteosarcoma cell line (MG63)

Cell adhesion and growth, material colonization, BG contents dependent

[77]

PLLA

Human osteosarcoma cell line (MG63)

Cell adhesion and growth, material colonization, BG contents dependent

[77]

PLA

Mouse pre-OB cell line (MC3T3-E1)

Cell adhesion, low growth, no differentiation, P bioglass

[78]

PCL

Human osteosarcoma cell line (Saos-2)

Cell growth

[79]

Mouse pre-OB cell line (MC3T3-E1)

PCL + BG powder vs fibres. PCK Fibre > powder > PCL for cell viability

[41]

Human osteosarcoma cell line (MG63)

10% BG => less growth than for 0%BG. Material colonization

[80]

Human osteosarcoma cell line (MG63) Human osteosarcoma cell line (MG63)

Cytocompatibility, BG contents dependent (10%>20%>30%) Negative effect of 20% BG on cell growth, positive effect of VitE + BG

[63]

Human osteosarcoma cell line (HOS-TE85)

BG scaffold – 2 porosities – cell growth on composite > neat BG

[82]

P(3HB)

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Table 7.1 Continued Polymer matrix

ECM comp

Cell type

Outcomes/remarks

References

Human osteosarcoma cell line (MG63)

Composite films – cell attachment, proliferation and differentiation

[83]

Human MSC

BG < BG + Coll +/− HYA<
[84]

Notes: BG: Bioactive glass PLGA: poly-(D/L-lactic-co-glycolic) acid PGA: polyglycolic acid PLA: polylactic acid PLLA: poly-(L-lactic) acid PDLLA: poly-(D/L-lactic) acid PHB: polyhydroxybutyrate PCL: poly-caprolactone BMSC: Bone marrow stromal cells OB: osteoblast ECM: Extra cellular matrix VEGF: Vascular endothelial growth factor ECM comp: Extra cellular matrix component (e.g. type I collagen, hyaluronic acid, demineralized bone matrix) PS: phosphatidylserine P(3HB): poly(3-hydroxybutyrate).

Table 7.2 Summary of in vivo evaluations of biopolymer/Bioglass® composites Polymer matrix

Model

Outcomes/remarks

References

PLGA

Mousesubcutaneous

Inflammation/encapsulation, peripheral granulation tissue, blood vessel infiltration

[69]

Rat-subcutaneous

Fibrovascular tissue, blood vessel invasion

[96]

Rat calvaria

Bone formation. Blood vessel invasion [97] but < as VEGF coated composite

PGA

Rat-subcutaneous

Cells + blood vessel infiltration

PDLLA

RabbitNo inflammation, bone formation, subcutaneous-femur implant resorption

PLLA

Nude micesubcutaneous

[21] [98]

Precellurarized implant (MSC). Bone [75] formation, coll I synthesis, no significant differences between BG contents (continued )

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Table 7.2 Continued Polymer matrix

Model

Outcomes/remarks

References

PLA

Sheep tibia

+/− Inflammation, bone formation also [99] within implant pores, implant resorption + osteolysis

PCL

Rabbit humerous

ECM comp added. OB activity + bone formation in implant and at bone interface. Positive effect of ECM comp

Rat calvaria

BG nano-fibres induced bone formation [41] >> PCL alone and empty defect

P(3HB)

Rat-subcutaneous

Thin encapsulation, fibroblasts growth, collagen, newly formed capillaries

[80]

ECMcomp

Rat calvaria

Collagen sponges – higher vascularization, higher faster bone formation compared to neat collagen

[100]

[79]

Notes: BG: Bioactive glass PLGA: poly-(D/L-lactic-co-glycolic) acid PGA: polyglycolic acid PLA: polylactic acid PLLA: poly-(L-lactic) acid PDLLA: poly-(D/L-lactic) acid PHB: polyhydroxybutyrate PCL: poly-caprolactone BMSC: Bone marrow stromal cells OB: osteoblast ECM: Extra cellular matrix VEGF: Vascular endothelial growth factor ECM comp: Extra cellular matrix component (e.g. type I collagen, hyaluronic acid, demineralized bone matrix) PS: phosphatidylserine P(3HB): poly(3-hydroxybutyrate).

Bioglass® (type 45S5) is the most investigated silicate bioactive glass system for bone tissue engineering [21, 33, 67–69, 71–77, 85–87] and, as shown in Tables 7.1 and 7.2, the majority of the composites comprising Bioglass® are prepared with either PLGA or PDLLA biodegradable polymer matrices. Several studies have investigated the biological response of different cell types to composite materials made of polylactide-co-glycolide and bioactive glass. For example, Lu et al. [71] found extensive growth of SaOS-2 (human osteosarcoma cell line) at the surface and within the pores of PLGA/BG composites in comparison to neat PLGA and TCP scaffolds. Higher ALP activity and collagen I synthesis were also observed. Lu et al. [72] also described a dose–response effect according to BG content, for both physico-chemical characteristics and cellular

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response points of view. After 7 days of culture on PLGA, PLGA/BG 10 wt% or PLGA/BG 50 wt%, SaOS-2 cell proliferation was found to be higher on the wt10% material compared to the 0 or 50 wt% BG composites. Although similar ALP activity was found in the 0 and 10 wt% BG materials, mineralization was higher for the PLGA/BG 10 wt% scaffold. Similar observations were made by Yao et al. [68], using rat MSC. Further studies [74–77] using PDLLA scaffolds also showed a dose-related influence of the bioactive glass filler on cellular behaviour. Tsigkou et al. [76], for example, investigated differentiation of human foetal osteoblast cultured on PDLLA, PDLLA containing 5 wt% BG and 40 wt% BG. Despite the fact that the cultures were performed in absence of osteogenic factor supplements, significantly higher cell maturation and differentiation on the composite materials were measured when compared to the cell culture plastic (control) and PDLLA alone. Moreover, even if significant differences were observed between cells grown on materials containing 5 wt% or 40 wt% BG (ALP activity and osteoblastic gene expression), the osteocalcin secretion and matrix mineralization were found to be higher on PDLAA/BG 5 wt% compared to PDLLA/BG 40 wt%. Using comparable scaffolds (PDLLA alone, PDLLA/BG5 wt% and PDLLA/ BG40 wt%), Yang et al. [75] found very similar results using human MSC or Stro-1 positive fractions of MSC. Namely, significant increase of ALP activity was determined for the material containing 5 wt% BG compared to the neat scaffold or the 40 wt% BG scaffold. As for the previous mentioned study, it was shown that a BG high content has a negative effect on the biological response compared to a lower amount (40 wt% v. 5 wt%). But looking at the influence of pre-treatment in medium containing 20% Fetal Calf Serum (FCS), Yang et al. [75] came to the conclusion that this might be due to the higher and extended ion release in the culture medium in the case of higher amounts of Bioglass® present. In the in vivo part of their study however, the materials were not pre-treated before implantation, and no significant differences were observed whether PDLLA contained BG or not. More recently, a few studies reported the use of polycaprolactone (PCL) associated to bioactive glass particulates [41, 79]. In their work using MC3T3-E1 cell line (mouse pre-osteoblast), Jo et al. [41] showed a significant higher number of cells on the composite containing bioactive glass nanofibres compared to the PCL only or containing bioactive glass powder. The same trend was observed when the composite materials were implanted in a rat calvarial defect. The same composite (PCL/BG) was also used by Erdemli et al. [79], to which a third component, namely bone demineralized matrix (BDM) or calcium sulphate (CS), was added. From the in vitro and in vivo studies carried out, no major beneficial effect of these new scaffolds on Saos-2 cell proliferation was found, however in their in vivo experiment, the authors observed a better tendency of bone ingrowth when composites supplemented with BMD were implanted [79]. Another group of polymers, poly(3-hydroxybutyrate) (P(3HB)) has recently gained the attention of materials scientists interested in bone tissue engineering

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[39]. P(3HB) belongs to the family of polyhydroxyalkanoate, which exhibits a much longer degradation time compared to other polymers groups such as for example PLA or PLGA. Chen and co-workers [88] already studied the potential of P(3HB) polymers for tissue engineering application and found adequate biodegradability properties and good biocompatibility with several cell lines. From 2006, Misra and colleagues [89] started to investigate the use of P(3HB) extensively as the polymeric phase of composites for tissue engineering scaffolds. Bioactive glasses were added to P(3HB) in different amounts, forms and sizes (e.g. nanofibres, micro- and nanoparticulates). Looking at the protein adsorption on the different materials, the same team [81] described a significant supremacy of the composites containing 20 and 30 wt% nanoscale Bioglass® compared to the other samples (neat or micro-scale). However, following MG63 cell proliferation, they showed a rather negative effect of increasing amounts of Bioglass® incorporated (10, 20 or 30 wt%) compared to the neat material, and also for both nano- or micro-scale particles, confirming the observations of several groups mentioned in the previous paragraph. Cell growth was observed in all cases, following the proliferation profile of cells on positive control surfaces (cellculture plastic). Later, further in vitro studies [80, 81], showed that, besides increasing cell proliferation, the addition of anti-oxidant Vitamin E to the P(3HB)/ Bioglass® composite also increased the hydrophilicity (and therefore protein adsorption) on these scaffolds. The antibacterial properties of these composites have been also investigated [80]. When implanted subcutaneously in the abdominal region of Sprague Dawley rats for one week, the two types of materials (neat or containing Bioglass®) induced the formation of a thin capsule containing proliferating fibroblasts, collagen fibres, some macrophages and more interestingly, capillaries sprouts, indicating the ability of those scaffolds to enable cell colonization, and neo-vascularization, which is a key aspect in tissue engineering.

7.3.2 Bioactive glass containing composites: new developments in angiogenesis As described previously in this chapter, a variety of composite materials have been developed and show promising potential for bone tissue engineering applications. However, during the past 10 years, even if tremendous progress has been made in the development of adequate matrices (scaffolds) for bone tissue engineering, materials scientists together with cell biologists have been essentially concentrating their efforts on the osteoconductivity and osteoinductivity properties of the scaffolds. But, unlike organ transplants where there is a pre-existing vasculature, synthetic tissue engineered bone constructs are devoid of this. Back in 1963, Trueta [90], published the importance of the vasculature in osteogenesis; in a review published in 2003, Carano and Filvaroff [5] underlined the critical importance of vascularization following the implantation of a tissue-engineered scaffold for its survival, integration to the surrounding tissues and its functionality

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(i.e. supporting and promoting neo-tissue formation). In recent years, researchers have started to address this problem by improving the rather poor angiogenic potential of tissue-engineered scaffolds by either changing their physico-chemical properties or by supplementation with angiogenic factors [91–93]. Indeed, recent in vitro studies [21, 69, 94–96] (see also ref. [23] for a recent review) using different cell types (e.g. fibroblastic cell line, intestinal epithelial cells) cultured in the presence of bioactive glasses (e.g. 45S5 Bioglass®) in different forms (e.g. coated cell culture dishes, PGLA based composites) have shown the potential angiogenic effect of Bioglass®. Specifically, a significant up-regulation of angiogenic growth factors such as VEGF (Vascular Endothelial Growth Factor) or bFGF (basic Fibroblast Growth Factor) has been found in Bioglass®-containing samples compared to Bioglass®-devoid samples at both protein secretion and gene expression levels (Table 7.3). The effect of 45S5 Bioglass® on VEGF secretion was first shown by Day et al. [21] using a rat fibroblast cell line (208F). While investigating the effect of the incorporating bioactive glass particles (0.01 to 10% w/v) into polyglycolic acid (PGA) scaffolds on cell adhesion and proliferation, the authors observed a strong inhibition of cell proliferation for percentages higher than 0.2% w/v of BG. However, VEGF secretion increased after 24, 48 and 72 hours of incubation for percentages below 0.2% w/v. In related investigations using a human myofibroblast cell line [96], or mice fibroblast [69], an up-regulation of VEGF secretion was also observed when cells were cultured in PLGA microporous spheres [96] or foams [69] containing Bioglass®, when compared to neat materials. Some other studies [94] used conditioned medium from fibroblastic cells cultured on such BG containing composites to treat endothelial cells. After 11 days of culture, a higher cell proliferation accompanied with formation of tubular structures was observed using fibroblast conditioned medium when compared to classical angiogenic medium [94]. Following the same aim, Leu and Leach [101] investigated the direct effect of bioactive glasses on neo-angiogenesis by using composite materials made of bovine collagen and 45S5 Bioglass®. Human microvascular endothelial cells (HMVEC) showed a higher cell proliferation rate compared to basic medium, which was accompanied with a higher tubular structure formation and VEGF mRNA up-regulation. However, again, a BG dose-dependent effect could be observed with 1.2 mg BG/scaffolds showing optimal responses while 0.6 mg or lower and 6 mg or higher BG concentration showed a rather negative effect. More recently, in their in vivo study using a rat calvaria implantation model, Leu et al. [100] demonstrated the angiogenic potential of BG-containing collagen scaffolds, leading to greater neo-vascularization and further bone healing when compared to the neat collagen sponges. These findings confirmed those made earlier by Day et al. [21, 69], who, using PLGA based composites, showed the presence of a significantly higher level of blood vessel invasion of materials containing bioactive glass compared to PLGA only, in both rat and mouse subcutaneous implantation models.

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Table 7.3 Angiogenic indicators stimulated in response to bioactive glass Composite system Cell type

Outcomes/remarks

References

BG-coated culture Rat fibroblasts (208F) VEGF secretion wells up-regulated

[21]

Human fibroblasts (CCD-18Co)

VEGF up-regulation

[94]

Human fibroblasts (CCD-18Co)

VEGF, bFGF up-regulation, [94, 102] endothelial cells proliferation

Mouse fibroblasts (L929)

Disks – VEGF up-regulation, BG dose effect, cell proliferation

Human fibroblasts (CCD-18Co)

Porous microspheres – VEGF [96] up-regulation

Human microvascular endothelial cells (HMVEC)

Scaffolds – cell proliferation

Alginate/BG

Human fibroblasts (CCD-18Co)

Alginate encapsulation + BG, [95] or BG coating – VEGF ↑, less cells, dose BG effect => conditioned medium => ↑ of human endothelial-, and microvascular endothelial cells proliferation

ECM-comp

Human endothelial cell and HMVEC

BG scaffolds + collagen – VEGF up-regulation, cell proliferation

[101]

Endothelial cells/ fibroblast co-culture

BG treated aortic rings – endothelial tubule formation

[101]

PLGA/BG

[69]

[97]

Notes: BG: Bioactive Glass PLGA: poly-(D/L-lactic-co-glycolic) acid PGA: polyglycolic acid PLA: polylactic acid PLLA: poly-(L-lactic) acid PDLLA: poly-(D/L-lactic) acid PHB: polyhydroxybutyrate PCL: poly-caprolactone BMSC: Bone Marrow Stromal Cells OB: osteoblast ECM: Extra Cellular Matrix VEGF: Vascular Endothelial Growth Factor ECM comp: Extra Cellular Matrix component (e.g. type I collagen, hyaluronic acid, demineralized bone matrix) PS: phosphatidylserine P(3HB): poly(3-hydroxybutyrate).

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7.3.3 Polymer/Bioglass® composites in musculoskeletal tissue engineering: intervertebral disc regeneration The versatility of biopolymer/Bioglass® composites in terms of structure and properties achieved enables their application in the wide musculoskeletal tissue engineering field. The case of intervertebral tissue engineering will be described here as a case in point to demonstrate the applicability of Bioglass®-containing composites in the field. In recent developments, Helen and Gough [103–105] investigated the response of annulus fibrosus cells from the intervertebral disc according to the Bioglass® particulate content in PDLLA matrix composites. Intervertebral discs (IVDs) are cartilage-like in structure and function and are situated between vertebrae, which provide a cushioning effect and allow mobility in the vertebral column. Due to the previously-reported collagenstimulating effects of bioactive glasses [106] it was hypothesized that including bioactive glasses in a composite may stimulate collagen production and benefit the repair of the collagen-rich annulus (outer structure of the IVD). Initially films were prepared containing 0, 5 and 30 wt% Bioglass® and the responses of bovine annulus fibrosus (bAFs) cells were characterized [103]. Sulphated glycosaminoglycan production was found to be higher on the scaffolds containing Bioglass® compared to PDLLA alone. 30 wt% Bioglass® inclusion gave the highest level of sGAG production after three weeks of in vitro culture, but by 4 weeks there was no significant difference between 5 and 30 wt%. The inclusion of 30 wt% Bioglass® particles also resulted in higher collagen production over the 4-week culture period. This increase in extracellular matrix production was then investigated in 3D porous scaffolds produced using TIPS [104]. These foams are highly porous with pore sizes in the range of 100 µm diameter with interconnected pores of approximately 10 to 50 µm diameter (similar to those shown in Fig. 7.2). SEM images in Fig. 7.3 show the extensive cellular colonization of bAFs cells of the scaffolds, increasing with Bioglass® content [104]. Cell numbers were increased over a 4-week culture period on foams containing 5 or 30 wt% Bioglass® as was sGAG and collagen production, with 30 wt% resulting in the highest levels of collagen production. Western blotting analysis showed that both collagen types I and II were produced, both of which are essential components of the IVD. The response of human AF (hAF) cells was also determined [105]. Similar results were obtained with the bAFs where the inclusion of Bioglass® particles increased the sGAG and collagen production. Immunostaining revealed presence of both collagen types I and II. More recently in unpublished data, hAF cells increased expression of SOX-9, collagen I and II and human mesenchymal stem cells were shown to increase expression of SOX-9, collagen I, II and aggrecan in response to inclusion of Bioglass® particles in the porous foams. These results suggest that bioactive glasses may be beneficial in regenerating the annulus of degenerate IVDs.

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7.3 SEM micrographs of bovine annulus fibrosus cells (bAFs) cultured on PDLLA/Bioglass® composite scaffolds with different Bioglass® content at different time points: neat PDLLA foams ((a): 2 weeks), PDLLA/5 wt% Bioglass® composite ((b): 2 weeks, (c): 4 weeks) showing extensive cellular colonization (reproduced from [104] with permission from Elsevier).

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Discussion

The promising and well-explored field of tissue engineering aims to repair a diseased or a damaged tissue, by restoring both its structural and functional entities. Suitable biomaterials are required to confer to the engineered tissue adapted structural features, while cells are necessary to restore the functional activity of a given tissue. In the field of biomaterials for tissue engineering purposes, synthetic and naturally occurring polymers combined with bioactive glasses, forming composites, offer a wide range of freedom in terms of physical and chemical properties, thus the composites are highly versatile for bone and musculoskeletal tissue engineering [29, 30]. Bone and musculoskeletal tissue engineering scaffolds are generally three-dimensional, showing highly interconnected porosity in which the pore size can be adapted. In the studies cited in this chapter, which focus only on biopolymer/bioactive glass composites, two aspects have been looked at. On the one hand, the importance of the material itself, in terms of design, structure or mechanical properties and their effects on the cell/tissue response, has been highlighted. On the other hand, the positive influence of bioactive glass particulates present in the different polymer matrices on cell growth and differentiation has been considered in detail. Researchers have shown that the presence of these particulates can stimulate cell proliferation and differentiation. However, differences in the biological responses have been observed according to cell types studied, as well as to the method used for composite preparation, which leads not only to different pore morphologies but also to different bioactive glass exposition to the surrounding fluid environment. Variations in the level of exposure of the bioactive glass particulates to the biological environment, which influences the composite degradation and the ion exchange mechanism at the interface between the composite and the surrounding environment, are also of importance. Mineral deposition has been shown to slow down the polymer-based scaffold degradation and concurrently increase its mechanical properties. The presence of a mineral phase within a polymer scaffold always induces an increase of the ALP activity and improves the overall osteo-inductive and osteo-conductive properties of a composite. However a BG dose-dependent effect response has been underlined in several investigations using different cell types. Osteoblastic differentiation and fibroblastic angiogenic factor secretion appeared to occur at an optimal content between ‘too low’ and ‘too high’. The optimal content however was found lower for fibroblastic cell types compared to cells from the osteoblastic lineage. Several hypotheses to explain this effect have been suggested, including the effect of BG particle inclusion on the material wettability, or micro-/macro-surface topography, both influencing protein adsorption and subsequent cell adhesion mechanisms. Another significant parameter that has also been suggested to affect the biological behaviour of scaffolds is the pH of the culture media and cell environment, which tends to acidify owing to polymer degradation. The inclusion of bioactive glass

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particulates within biodegradable polymer scaffolds will not only slow down the scaffold resorption, but also the alkalinity of the BG dissolution should neutralize the acidic degradation products of many polymers. In the same way, the formation of a calcium phosphate layer on the surface of BG goes through a series of ion exchange steps and precipitation/dissolution reactions, which will induce an increase of the local pH. Xynos et al. [17, 20] have investigated the influence of ion release during BG degradation in vitro, and found an important direct influence of Bioglass® dissolution products on human osteoblast cell proliferation. A summary of previous work on the gene expression effects of dissolution products of bioactive glasses has been published recently [16]. Further studies with Bioglass®-containing composites have shown encouraging results regarding the potential angiogenic effects of Bioglass® dissolution products, i.e. increased secretion of vascular endothelial growth factor (VEGF) in vitro and enhancement of vascularization in vivo, suggesting that scaffolds containing controlled concentrations of Bioglass® might stimulate neo-vascularization, which is beneficial to large tissue engineered constructs (Table 7.3) [21]. A recent review [23] has comprehensively discussed the current experimental evidence (in vitro and in vivo) of the angiogenic effect of bioactive glasses in the context of tissue engineering.

7.5

Conclusions and future trends

The repair of large bone defects remains a major clinical problem, and after blood, bone is the most implanted tissue. Because of the multi-aspect of its function (biological, structural and mechanical), bone also constitutes one of the most challenging domains of tissue engineering and regeneration. Musculoskeletal tissue engineering is, in general, an area of research requiring the availability of suitable scaffolds. A number of biodegradable polymers and bioactive ceramic combinations have been studied in this field of tissue engineering. In this group of materials, composites formed by combining bioactive glasses and biodegradable polymers seem to represent the materials of choice due to their adequate properties and versatility to be tailored to different applications. However, even if most of these composites can be tailored to meet the required structural/mechanical properties, and to show adequate biodegradability and cytocompatibility, the question of bioactivity of these materials is still under examination. It is important for scaffolds for bone tissue engineering to be osteoconductive and osteoinductive, in order to guide and encourage neo-bone formation. But also neo-vascularization of the tissue-engineering construct is required, representing a critical contribution to the success of regeneration and growth of new tissues as it provides cells with oxygen and nutrients. In this chapter we reviewed the relevant field, focusing on bioactive glass-containing composites for bone and musculoskeletal tissue engineering, reporting on the in vitro and in vivo performance of this family of scaffold materials. Since bioactive glasses have been developed in 1969 and have exhibited successful clinical applications as bone filler material, an increasing amount of studies have

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been underlining the interest of these BG particulates in tissue engineering. The bioactivity of BG-based composites has been shown using a wide range of cell types (primary or cell lines) from different animal origins in vitro, but also in vivo. More recent studies have reported the angiogenic potential of such composites in both in vitro and in vivo approaches. However, the extended (clinically relevant) in vitro and in vivo understanding of these composite scaffolds is still limited, especially regarding their long-term behaviour related to the kinetics of degradation and ion release effects from bioactive glasses on the biological environment. In future studies, the specific influence of metallic ions released during degradation of the bioactive glass component of the scaffolds on bone formation and angiogenesis must be clarified and quantitatively investigated at a fundamental level. The influence of using micron-sized or nanoscale bioactive glass fillers in composites on bone and blood vessel formation will need to be investigated too. Further investigations of the scaffold bioactivity in relation to surface modification strategies, for example through the use of protein adsorption or plasma treatment, to provide more cues to cell attachment and response, will be relevant. The results of such research will enable a better understanding of the synergetic effect of bioactive glass on osteogeneis and angiogenesis, leading to control of the mineralization and neo-vascularization of the construct. This knowledge will also support the design of specific compositions of bioactive glasses to be combined with biopolymers to form improved composite constructs, which will exhibit the required architecture and microstructure (porosity content, size and orientation of the pores, amount of BG, size distribution of BG particles, type of polymer, surface topography) as well as suitable (time-dependent) mechanical properties. The scaling up of such constructs would also be a point of interest as most of the current studies are made in small-size materials. Larger scaffolds will have an influence on the nutrient diffusions of potential pre-seeded cells, but also on the specific surface area available, influencing the amount of released components. Overall, BG-based composite materials offer very promising properties that are of interest in bone tissue engineering and musculoskeletal tissue regeneration in general and future significant advances in the field are expected based on ongoing research efforts worldwide. This fact anticipates that development of composite scaffolds will remain a major area of application for bioactive glasses.

7.6

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7.7

Appendix: list of abbreviations

BG: Bioactive glass BMSC: Bone Marrow Stromal Cells ECM comp: Extra Cellular Matrix component (e.g. Type I collagen, hyaluronic acid, demineralized bone matrix) ECM: Extra Cellular Matrix OB: osteoblast P(3HB): poly(3-hydroxybutyrate) PCL: poly-caprolactone PDLLA: poly-(D/L-lactic) acid PGA: polyglycolic acid PHB: polyhydroxybutyrate PLA: polylactic acid PLGA: poly-(D/L-lactic-co-glycolic) acid PLLA: poly-(L-lactic) acid PS: phosphatidylserine VEGF: Vascular Endothelial Growth Factor

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8 Use of bioactive glasses as bone substitutes in orthopaedics and traumatology J. HEIKKILÄ , Sports Clinic and Hospital Mehiläinen Turku, Finland

Abstract: Bioactive glass is one of many biomaterials suggested for bone replacement. Its advantages include non-toxicity, biocompatibility and bioactivity. Additionally, through a surface reaction the glass forms a Ca’P-layer after implantation within the body, through which the host bone grows together with the glass. Through continuous reaction and layer formation the glass is finally absorbed and replaced with bone. Fractures and cavitary bone tumors at metaphyseal bone regions are among the best clinical applications for bioactive glass. It has also been used in diaphyseal bone defects and fractures, and in spinal surgery with good clinical results. The possible clinical applications are reported in this chapter. There have been some good prospective randomized studies into using bioactive glass in metaphyseal bone areas; however, there is a need for more research into using bioactive glass within orthopaedics and traumatology. Key words: bioactive glass, bone substitutes, bone defect, fractures, metaphyseal area, clinical, orthopaedics and traumatology.

8.1

Introduction

The search for suitable bone substitute materials has been going on for more than 120 years [1]. The aim has been to find a material that can be used safely and effectively in place of autogenous bone, since the harvesting of autogenous bone from the iliac crest is always an invasive accessory operation causing pain, bleeding and morbidity for the patient. Furthermore, postoperative rehabilitation is usually more difficult. The operation time is lengthened and a second surgeon is usually needed for the harvesting procedure. In modern orthopaedics and traumatology there is a trend towards smaller incisions, arthroscopic procedures, and shorter operation times, which are all facets of minimal invasive surgery, i.e. surgery that disturbs the life of the patient as little as possible. In this respect, the use of bone substitutes is a necessity. Hospital directors want to ensure that operating theaters are used as effectively as possible: the use of bone substitutes offers interesting possibilities in this regard.

8.1.1 History of bone substitutes The first attempts to identify possible substitutes for autogenous bone were made by a Dutch scientist, Job van Meekeren, who tried to use fresh canine skull bone as a bone grafting material as early as 1668 [2]. It was not until 1880 that Maceven 189 © Woodhead Publishing Limited, 2011

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first used fresh bone from a human cadaver to replace the damaged part of a bone [3]. Following on from this, there developed a widespread clinical use of allograft bone, and fresh frozen allograft bone is still in clinical use today [4]. In a way it can be regarded as the gold standard bone substitute material. The first artificial material used for this purpose was calcium sulphate, CaSO4: Dreesman was already using this to fill cavities in 1883 [5]. This material is surprisingly still on the market. In 1889 Senn used antiseptic decalcified bone [1], an idea that was further developed 75 years later by Urist [6], and in 1920 Albee introduced artificial tricalcium phosphate [7]. Deproteinized bovine bone was first used by Orell in 1934 [8]; he named it os purum, clean bone. In 1952 Bauermeister introduced a similar concept, Kieler bone, which was also made from deproteinized bovine bone. Demineralized bone was further studied by Urist in 1965 [9], and through these studies he discovered bone morphogeneic proteins in 1971 [10]: these proteins represent one future direction in the development of bone substitute materials. The idea that certain types of glass might be bioactive and could chemically bond with bone was first introduced by Hench in 1967; he then proved this to be true in 1971 [11]. Hench glass is still in clinical use today. Coralline hydroxylapatite was introduced in 1974 by Roy and Linnehan [12], and in 1977 Jarcho proved that dense hydroxylapatite could bond with bone [13]. At the same time the first attempt to use composite materials for the purpose was made by Mittelmeier and his colleagues, who combined demineralized bone matrix with collagen [1]. Gross and Strunz studied bioactive glass-ceramic in Germany [15], while Kokubowas was conducting similar investigations in Japan at the same time [16]. Osborn and Furlong further developed hydroxylapatite, and in 1991 they manufactured synthetic hydroxylapatite [17], which is used today for coating metal prosthesis and external fixation pins. Shortly afterwards, in 1992, Van Blitterswijk developed bioactive polymer in Leiden [18]. Meanwhile, in California, Constanz worked on the formation of coralline calcium carbonate in corals, and realized that it was possible to mimic coral formation in the laboratory. As a result, an injectable form of calcium phosphate was developed, which sets into fully hard dahllite 12 hours after injection into the body [19]. At Åbo Akademi University in Turku, Finland, Professor Kaj Karlsson developed a series of different glasses in the late 1980s, of which S53P4 bioactive glass seemed to have the most potential. Andersson published his thesis concerning these materials in 1991 [20]; it was the first thesis on bioactive glass to emerge from Turku, and since then several others have been published. The university project started by the Turku Biomaterials Group has now finally been commercialized, and the material is marketed by BonAlive Biomaterials Ltd.

8.1.2 Ideal bone substitute materials The ideal bone substitute material would be biomechanically stable, non-toxic and biocompatible. It would be replaced by healing bone, and would contain

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inductive factors to promote bone healing. Additionally, it must not induce any risks for the patient, should be easy to use, inexpensive, and readily available. With these prerequisites it is obvious that there is not a single material that would fulfill all these requirements. Implanting materials in the tissues can be regarded as forming a chronic wound in the tissue, and the material characteristics of the implant and implantation site regulate the histological and tissue response towards implantation. The response is also affected by biomechanical factors and the possible movement of the implants, which may affect the long-term results of the implanted material [21]. Bone defects can be divided into non-segmental, segmental, and articular joint defects [22]: each of these is best reconstructed with a specific type of bone substitute. There is no one material that is suitable for use in all of these defects.

8.1.3 History of bioactive glass in orthopaedics The concept of bioactive glass and its use in bone surgery was introduced in 1967 by Hench, who hypothesized that glass containing calcium and phosphorus would be biocompatible: he used the glass phase diagram as the basis for the choice of composition. In 1969 the US Army Medical Research and Development Command approved an experimental study plan with two aims. The first aim was to achieve a direct chemical bond between biomaterial and bone. The second was to achieve an understanding of the reactions occurring between bone and implanted materials. The first animal experiments were performed in 1969, with the first results published in 1971. The original hypothesis was proven to be correct [11], and in addition the first bone bonding material was designed. The development of these new materials started in various centers immediately after the first reports by the Florida group. A German group led by Gross and Strunz developed a bone-bonding glass-ceramic (GC), Ceravital®, and reported the first results in 1980 [15]. Ceravital® consists of a glass phase into which apatite crystals are embedded. It has a significantly higher bending strength compared with glasses [23]. An even stronger GC material, Cerabone®, was manufactured at Kyoto University in Japan by Professor Kokubo and his colleagues in 1982 [16, 24]. It has also been called AW G-C (apatite wollastonite glass-ceramic). It is crystallographically composed of apatite and wollastonite crystals embedded in a glassy matrix. Since its development, this material has been used clinically even in load-bearing situations [25]. Ilmaplant®-L1 [26] and Bioverit® are glass-ceramics developed in the former East Germany. The first has a composition resembling that of Ceravital® [15], but crystallographically it consists of apatite and wollastonite embedded in glass. The structure is similar to that of Cerabone® [23]. The second, Bioverit®, was developed at Friedrich Schiller University in Jena. It differs from the materials mentioned above, and includes remarkable amounts of fluoride and aluminum oxide. Crystallographically three phases can be detected: apatite, phlogopite and glass. Bioverit® has been claimed to be bioactive.

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At Åbo Akademi, Professor Kaj Karlsson developed his series of glasses, and selected those that were most interesting in terms of clinical applications. In his group, Dr. Andersson developed a mathematical analysis to evaluate glasses and he selected S53P4 as the glass that was both most suitable and most interesting for clinical use, and published this in his thesis in 1991 [27]. Since then, at least 40 theses on bioactive glass have been published by the Turku Biomaterials Group [20, 28–30].

8.2

Glass surface reactions

According to Hench [11], bone bonding can be presented as a complex series of reactions in the glass and on the glass surface (Fig. 8.1). The chemical reactions on the glass surface are based on leaching, dissolution, and precipitation as reviewed by Hench and Andersson [31] and Ducheyne et al. [32]. After implantation in simulated or in vivo body fluids the first rapid reaction on the glass surface is the exchange of Na+ or K+ with H+ or H3O+ from solution. The leaching is facilitated by the fact that alkaline and earth alkaline cations are not part of the network, but are only modifiers. This reaction produces an alkaline microenvironment in which the alkalinity of the solution results from the breaking of -Si-O-Si- bonds, mainly by hydroxyl ions. This dissolution occurs only locally at the glass surface and results in the formation of silanol (SiOH) groups at the glass–solution interface. The hydrated silica groups condense and repolymerize

8.1 Surface reactions and Si-rich and Ca,P-layer formation at the surface of bioactive glasses. First a layer of glass with the thickness of a couple of microns is dissolved (I). Secondly, Na, Ca, P and also Si are leached from the surface of the glass (II) and a Si-rich layer is formed through repolymerization (III). Ca and P from the solution and partly leached from the glass precipitate in the surface of the Si-rich layer (IV).

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with the silanol groups forming a SiO2- rich layer on the surface. This precipitation is facilitated by the migration of Ca2+ and PO43− groups to the surface through SiO2-rich layer and there is further incorporation of Ca2+ and PO43− from the solution into this layer. This step occurs within minutes of the implantation of the reactive bioactive glasses (Fig. 8.1) [20]. The SiO2-rich layer increases in thickness due to diffusion-controlled alkali ion exchange and the final thickness of this layer is dependent on its composition. The amorphous CaO-P2O5-rich film grows by incorporation of soluble calcium phosphate from the supersaturated solution. The nucleation of a CaO-P2O5-rich film on top of the SiO2-rich layer and formation of bone-like carbonated calciumdeficient apatite on the glass surface has been observed by Fourier Transform Infrared spectroscopy (FTIR) within 10 hours of the implantation. Crystallization of the amorphous CaO-P2O5 film also occurs by incorporation of OH−, CO32− and F− anions from solution to form a mixed hydroxyl, carbonate, fluorapatite layer, probably through similar reactions as those that cause the calcium-deficient apatite dahllite to crystallize on the bone [31, 32]. The reactions described above result in a 100 to 120 µm thick layer, which is rich in SiO2 and a layer about 30 µm thick of hydroxycarbonate apatite (HCA) [33].

8.2.1 Tissue reactions on the surface of bioactive glass The biological processes involved in bone formation and bonding on the glass surface can be divided into six stages (Table 8.1). The SiO2-HCA layer is formed in a similar manner both in vitro and in vivo. However, in vitro a Ca,P layer is formed on the glass surface, whereas in vivo it is formed within the surface of the Si-rich layer. In vivo, biological compounds such as proteins and collagen adhere to the glass surface. This phase can occur even in the absence of cells or growth factors. Apatite is formed by crystallization around collagen which has been observed to be trapped in the growing HCA layer [34]. Bioactive glasses and in particular the formed HCA layer favor the chemotaxis of osteoblasts on the glass surface [21]. These cells attach and differentiate and commence their secretion, and more matrix is formed. Simultaneously and also causatively, collagen fibers and Table 8.1 The different stages of bond formation between glass surface and bone Stage 1.

Biological structures adsorbed in the SiO2-HCA layer

Stage 2.

Precipitation of Ca and P from the glass and solution between the biological structures, especially collagen

Stage 3.

Attachment of stem cells (and preosteoblasts) to the surface layer

Stage 4.

Differentiation of cells on the glass surface

Stage 5.

Matrix formation

Stage 6.

Crystallization of matrix

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mucopolysaccharides agglomerate around and within the apatite crystals and bond with them (Table 8.1). The reactions in vivo produce an interface with a chemical connection between glass surface and bone. The phenomenon is called bonding osteogenesis: bone bonding. The chemical bonding is further strengthened by the adhesion of proteins and collagen to the Ca,P-layer. Hench and Paschall [35] published a paper on the adhesion of poly-L-alanine to a bioactive glass surface. Organic material and collagen have been observed in vitro by Wilson and Nolletti to be trapped within the growing HCA layer [34]. The same result has been shown to occur in vivo by various authors [36–38].

8.3

The bonding of bioactive glass and bone formation

In summary, bone repair after implantation of bioactive glass and the reactions that occur simultaneously on the surface of the glass can be explained in the following way. Bone minerals have a rapid turnover, and crystal formation occurs epitaxially on bone. The biological part of the matrix, consisting of collagens, non-collagenous proteins, and glycosaminoglycans, modulates crystal formation. Similarly, crystallization of Ca,P on the bioactive glass surface also occurs epitaxially; it is also likely that the same biological structures modulate the precipitation and are trapped between the crystals. The carbonated HA on the glass surface closely resembles the calcium deficient carbonated HA, dahlite found in bone. Taking these chemical and biological reactions into account, it is easy to understand the formation of the bond between bioactive glass and bone. A similar type of crystal formation at the surface of bioactive glass enables bone to bond biochemically with bioactive glass. Nevertheless, it should be noted that the bone formation cannot simply be a case of coalescence of two epitaxially-growing crystal fronts: the biological structures, collagen, proteins, and glycosaminoglycans affect bond formation, probably through similar processes to those that occur during bone crystal formation.

8.3.1 The bonding strength of bioactive glass The bonding between bone and bioactive glass has been verified with push-out tests, which have shown that the strength of the bond between bioactive glasses and bone ranges from 16 to 23 N/mm2. Using the same test method, the push-out strength when there is no contact between bone and implant (e.g. with inactive glass or metals) was 0.5 N/mm2, and for materials with physical contact from 2 to 4 N/mm2 [39] (Fig. 8.2). It is postulated that when the bond between the bone and bioactive glass forms, the silica-rich layer is the weakest part. However this is not supported by push-out tests, in which the fracture line occurs at random in the

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8.2 Schematic presentation of bone response towards three types of implant materials: (a) a fibrous capsula encapsulates the biotolerant implant material; (b) intimate contact without bone bonding occurs at the interface between bioinert materials and bone; (c) intimate contact with chemical bone bonding between bioactive surface and bone. Note the gradual transformation between bone and implant material.

bone close to the glass, running a short distance into the reaction layer and within the glass [40]. It has been reported that the bond strength between bone and bioactive glass is higher than that of hydroxyapatite and bone [41, 42].

8.4

The biocompatibility of bioactive glasses

Bioactive glasses behave as biocompatible materials that are independent of the implantation site. In bone, they act mainly osteoconductively and there is no real evidence showing any osteoinductive capacity in bioactive glasses [43, 44]. The interaction at the interface cannot be predicted simply by theoretical analysis or calculation. Even the biocompatibility of a given material depends on the implantation site and loading [45]. Only bioactive materials are able to build up a chemical bond with the host bone: they do not create a chronic wound in the bone, while other implant materials cause an inflammatory reaction in accordance with their biocompatibility. Bone substitute material can be osteoinductive or osteoconductive. The two phenomena occur simultaneously in vivo and are not individual phenomena, but are affected by a multitude of factors besides the material itself. The stability of the implanted site is important, as are the biological implantation bed and the immunological reactions caused by the implant material [46]. Banked bone, cortical segments, hydroxylapatites and bioactive glasses and glass-ceramics are osteoconductive: they favor the proliferation of osteoblasts, but do not cause phenotype conversion of cells. With bioactive glasses, fibroblasts are transformed into a nonmitotic state while exposed to the surface of the glass [47].

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It has also been reported that calcium phosphates influence alkaline phosphatase activity as well as the incorporation of 3H-thymidine in the repairing bone. Furthermore, it has been suggested that there is likely to be a solution-mediated effect on cell proliferation and differentiation [48]. This activation probably occurs as a result of calcium-mediated events associated with intracellular dissolution of phagocytized material. Thus these types of materials do not cause bone formation at soft-tissue sites, but do promote bone formation when implanted in bone tissue. Hench has a similar view on the osteoconduction of bioactive glass, suggesting that bioactive glasses have an effect that differs from simple osteoconduction and is related to the soluble silicon released by the glass. In addition, the osteostimulatory effect of bioactive glass granules has been observed when filling rabbit bone defects with bioactive glass granules of the appropriate size [44].

8.5

The strength of bioactive glass

The clinical use of bioactive glasses and glass-ceramics has been limited by the structural weakness of the glass. In the case of metal prosthesis, there have been a number of technical problems in the coating process, such as crystallization of the coating material during the process and the limited thermal working range of glass, and these have been barrier to the formation of coatings that could function as the bulk glass. Bioactive glasses and glass-ceramics can be used clinically as space fillers or for regenerative purposes. They can protect other materials from corrosion within the body, replace or augment tissues, and replace functioning parts. Bioactive glasses resemble other glasses in terms of their mechanical strength. The modulus of elasticity is high, but due to the amorphous structure of glass, they are brittle and rigid, have low fracture toughness and are mechanically weak. Critical crack size is small, and the tensile bending strength is 40 to 60 MPa depending on the composition of the glass. When used as fillers, as buried implants or in unloaded applications, these mechanical properties are not harmful [24, 31]. The mechanical properties of bioactive glass together with those of cancellous and cortical bone are presented in Table 8.2.

Table 8.2 Mechanical strength of cancellous and cortical bone, and glass Cancellous bone

Cortical bone

Glass

2–12 MPa

100–230 MPa

785 MPa

Tensile strength

10–20 MPa

50–150 Mpa

35–175 MPa

Young’s modulus

0.05–0.5 MPa

7–30 MPa

70 000 MPa

Compressive strength

Pulling strength Poisson’s ratio

94 MPa 0.2–0.3

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Bone formation

8.6.1 Histology Four main types of tissue can be found at the interface of implanted materials: soft tissue, osteoid, chondroid and bone [15]. Three types of tissue will develop at the interface between bone and bioactive glass: bone, fibrous, and bone marrow tissue. Only mild inflammatory responses have been observed around implanted bioactive glass with a reaction similar to that which occurs in empty defects after surgery. Some round cells and occasionally polynuclear cells can be seen early after implantation, but this reaction will subside rapidly. The reaction layer increases in thickness with time, and will be gradually replaced by bone.

8.6.2 Cartilage repair There is no evidence that cartilage tissue can grow on the glass surface. It seems, however, that a growing bone layer can support cartilage-like tissue, at least at the margins of osteochondral defects, as long as initial subchondral bone repair has occurred. Hyaline-like cells can be observed in the cartilage in the process of repair when bone is growing on the surface of bioactive implants [29]. For traumatic cartilage lesions and for osteochondritis dissecans, one potentially advantageous approach would be to reconstruct subchondral bone first in order to lay the foundations for cartilage healing. It is obvious that large osteochondral defects require a solid surface or a support on which cartilage cells are able to spread, allowing cartilage repair to occur [49–53].

8.6.3 Resistance to toxic effects Polymethylmethacrylate (PMMA) interrupts bone formation at the interface of bone and glass; however glass can withstand this effect better than hydroxylapatite. The surface reaction layer of a bioactive glass shows more resistance to the toxic effect of PMMA. One reason for this might be that the structure of HCA formed on the bioactive glass surface bears a closer resemblance to bone mineral than to synthetic HA [54]. In addition, the alkaline surface of bioactive glass [55] might be better able to buffer the toxic effects. The hydration layer of bioactive glass is one further possible explanation for the resistance of bioactive glass to the toxic effect of PMMA.

8.6.4 Bone formation pattern in defects Experimental defects filled with bioactive glass granules show rapid bone growth on the granules. The pattern of bone growth is centripetal starting from the margins of the defect and growing inwards. The granules did not disturb bone formation in experimental defects, and indeed the filler effect of bone together with bioactive

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glass was better than that of bone alone. The centripetal growth pattern has also been observed in clinical use, when benign tumors and metaphyseal fracture defects have been filled with bioactive glass granules. Bone formation from the periphery of the defect [1, 56] is a feature common to all osteoconductive materials, but the combination of bone substitutes with bone inductors might enable more rapid healing of the defects as suggested by Miller et al. [56]. For bioactive glass, an osteoconductive growth pattern [58] and the time-related increase in the amount of bone at the interface has been reported [23, 42, 49]. The bone formation pattern at the interface of bioactive glass resembles intramembranous ossification, while enchondral ossification is absent. This finding leads to the conclusion that bioactive glass does not induce massive bone formation, but instead causes osteoblasts and osteocytes to spread along the glass surface, and that the material is mainly osteoconductive. This view has also is also shared by Hench and Paschall [35] and Ono et al. [57]. The mRNA measurements of type I and III collagen using Northern hybridization provide further evidence for this hypothesis [43]. Nevertheless, some indirect signs that bioactive glass also has bone inductive properties were observed during the experiments discussed. Similar observations have also been published by Ducheyne and Cuckler [60] and Schepers et al. [44].

8.6.5 Bone formation imaging methods In order to image the morphology of the interface various methods must be used. A combination of bulk glass, reaction layer, Ca,P-layer, crystallized Ca,P and mineralizing bone form a complex structure. Histological stains visualize the biological structures. Using toluidine blue and the von Kossa method, the reaction layer can also be stained in a particular way. Although the true nature of these stainings in the reaction layer remain unexplained, these staining characteristics are directly related to the bone bonding of the glass. The von Kossa method stains calcium, and it is likely that there is a layer of free calcium, deep in the reaction layer, that can be visualized using this method (Fig. 8.3). Toluidine blue stains the reaction layer differently and stains the acid part of the layer blue (Fig. 8.4). The staining intensity reflects the surface area of the structures, which has been shown to be important for bone bonding [61]. The formation of the reaction and bone bonding can be analysed by light microscopy and simple histochemical stainings. These histological methods are simpler and less expensive than electron microscopy (EM), scanning electron microscopy (SEM) or energy dispersive X-ray analysis (EDXA), and certainly simpler than push-out tests. Toluidine blue reveals that the interface has a double-layered staining character, and a similar staining pattern occurs when the von Kossa method is used: the former indicates the surface area of the reaction layer and the latter the calciumion distribution. Both are necessary for bonding. Van Gieson, a collagen stain,

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8.3 Von Kossa staining of bioactive glass (bg) to bone (b) interface (x 75). Note the double-layered structure. Calcium phosphate at the outer part of the silica gel layer stains dark brown (long arrows); the base of the silica gel layer stains dark brown (short arrows).

8.4 Histology of bone (b) to bioactive glass (bg) interface. Toluidine blue (x 150) staining two parallel stripes (arrows) indicating high surface area and acid proteoglycans.

failed to stain the reaction layer, indicating that no collagen is present on the reaction layer side of the interface. The bone-bonding phenomenon of bioactive glass occurs owing to a set of complex reactions, which are interdependent. Consequently, the methods used to study the rate of bone bonding must also be manifold. SEM is not able to image the biological structures, but produces a highly magnified image of the glass, the reaction layer and mineralized bone. EDXA detects the elements at selected spots or linearly, and reflects the chemical

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continuum between bone and the implant. Nevertheless, these methods cannot prove that bone bonding occurs: they simply suggest that it does. To definitively show and measure chemical bonding, push-out tests [20] must be combined with these methods.

8.7

Clinical use for benign bone tumors

For the treatment of benign bone tumors many different methods have been used: curettage with or without bone grafting, percutaneous sclerotherapy, corticosteroid injections, total or subtotal excision with or without reconstruction and irradiation. Some tumors disappear spontaneously, for example after fracture healing. Autograft has remained the gold standard filler material. It is osteoinductive without causing negative effects at the grafting site. However, sometimes the quantity of autogenous bone is insufficient, and especially in old or osteoporotic patients the density of bone is too low for packing in the large bone defects. It necessitates a second incision and further operations and can thus cause morbidity and complications. Among many other materials, bioactive glass granules have been used as bone substitute material. Benign bone tumors are contained lesions with intact bone boundaries, surrounded by sclerotic bone. The bone is normally sufficiently strong even after evacuating the tumor. Thus osteosynthesis or external support from outside the bone is not needed. For these reasons granules can be used in any benign bone tumor as filler material. The filler effect during the operation is good, and when granules are moistened with blood or saline implantation into the defect becomes easier [62]. It has been shown that the thickness of cortical bone at the implantation site shows a greater increase when bioactive glass granules are used than when autogenous bone is applied, highlighting the osteostimulative effect of bioactive glass [63]. The use of autogenous bone causes the volume of the cavity to diminish more rapidly than the use of bioactive glass. The growth pattern seems to start from the margins of the defect when osteoconductive bioactive glass is implanted and from the whole defect when autogenous bone is used. After 36 months there seems to be no difference between the two materials in terms of cavity volume, even in the case of large tumors, but there is a significant difference at 12 and 24 months [64]. Bioactive glass is well tolerated and does not put patients at higher risk of infections: no material-related adverse effects have been observed during or after clinical use [62–64]. The inflammatory reaction seems to be similar to that observed when autogenous bone is used. Furthermore, when examining blood samples no differences are observed between these materials [65]. Bioactive glass can be used with good results even in children, without disturbing the remodeling capacity of the bone even in cases where preoperative deformity is present [66]. This is true as long as the growth line is not disturbed either during the operations or by the tumor itself.

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In certain cases repeat operations have been performed 1 or 2 years after the original operation. The bone between the granules seemed to be physically harder than the surrounding bone. Bioactive glass has also been used in repeat operations, in occurrences of recurrent tumors. The granules disappear gradually by surface reaction and dissolution and by osteoclastic activity in 1 to 4 years depending on the cavity size [64].

8.8

Bioactive glass and infection

Bioactive glass has also been shown to be bacteriostatic in experimental and clinical otorhinological use [67, 68]. Also in the treatment of fractures in the metaphyseal bone region, the granules have shown bacteriostatic properties. In certain cases with a superficial wound infection a revision of the wound has been necessary, leaving the granules open. The wound has healed well without deep infection, despite the direct contact between infected tissues and glass granules [62]. There is some evidence that bioactive glass is advantageous even in chronic osteomyelitic cases [69]. In a retrospective multicenter study, data from 11 patients has been collected and analysed. Postoperative, traumatic or bloodborne chronic osteomyelitis was treated in the spine or lower extremity using bioactive glass granules. The most common cultured pathogens were Staphylococcus aureus and gram-negative bacilli. The cavitary defect or the vicinity of the spinal implant was filled after revision with bioactive glass granules. The results of the surgical intervention and implantation were surprisingly good. It has been postulated that the result is related to the leaching of alkaline earth ions leading to a rapid increase in pH around the glass particles. Bearing the above-mentioned in mind it seems that the aims of the study started by Professor Hench and approved by the US Army Medical Research and Development Command during the Vietnam War seems to have been fulfilled. However, it must be borne in mind that these results are preliminary, and no prospective randomized studies exist concerning the use of bioactive glass in treating osteomyelitis; nor are there any studies dealing with the use of any other materials with the same aim.

8.9

Bioactive glass in cancellous bone and metaphyseal fractures

Fractures in the metaphyseal area often cause compression in the subchondral cancellous bone area, resulting in deformity in the joint line. This type of fracture includes for example distal radius and proximal tibia fractures. After reduction of the joint space a defect in the cancellous bone usually remains, which requires a bone graft. Osteosynthesis is also needed after reduction. In experimental studies [29] bioactive glass provided good support for the growing new bone. In clinical

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study, with now more than 10 years of follow-up, results in the bioactive glass group were identical to those in the autogenous bone group [62]. No significant differences were observed between the two demographically identical groups with regard to clinical results, redepression of articular surface measured by plain films and computerized tomography (CT), valgus alignment, tibio-femoral (TF) angle, mechanical axis of lower limb (DMA) or subjective evaluation by the patient. Bioactive glass granules were incorporated into surrounding bone at three months and remodeled thereafter [62, 64]. According to this prospective randomized study it seems that bioactive glass can replace autograft and can be used as bone graft substitute material in metaphyseal fractures. During implantation of the granules meticulous care must be taken to make sure the granules do not enter intra-articular space.

8.10

Bioactive glass in diaphyseal bone fractures

A few studies exist that have examined the use of bioactive glass granules in the treatment of long bone fractures [70]; such as tibia or femoral shaft fractures. These fractures often need repositioning and rigid fixation. Sometimes a large amount of autogenous bone is needed to fill voids caused by these fractures and the use of bone substitute material is advantageous. There is some evidence that bioactive glass causes a thickening of cortical bone in the vicinity of the implantation site, which might help in diaphyseal bone repair. There is normally not as good a vascular supply in diaphyseal bone as there is in metaphyseal bone, especially after the soft tissue damage caused by the fracture. It is therefore to be expected that the reactions of implanted glass are slower, i.e. ossification and the resorption of the granules will take place more slowly.

8.11

Bioactive glass in spinal surgery

To treat instability or in fracture cases spinal fusion is needed to ensure stability between the vertebrae. This can be achieved using anterior or posterior or posterolateral interspinal fusion [71]. One feature common to all methods is the need to build bone in an extra-osseal location. Vascularity can be compromised, and this can negatively affect bone formation. There are experimental and clinical studies concerning bone formation using bioactive glass in posterolateral fusion, which show that a good fusion rate has been achieved. In a clinical study bioactive glass was examined in posterolatreal fusion of spinal fracture patients. The study was designed in such a manner that each patient served as its own control, while autogenous bone was used in the left posterolateral fusion and bioactive glass granules in the right. Similar fusion rate and bone formation was observed in both sides using computerized tomography as the analysis method. It is more difficult to create bone at an extra-osseal than at an intra-osseal implantation site: this is also true when autogenous bone is used. The remodeling

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of bone tends to restore the original anatomical bone structures and to lead to resorption of the graft, unless the load is carried through the grafted area.

8.12

Arthroplasty

The use of bioactive glass as a coating for metal prosthesis does not appear to be technically possible. However, bioactive glass might be suitable for use in revision arthroplasties, together with autogenous or allogenous bone when filling defects adjacent to metal implants: bactriostatic property of glass is advantageous in this respect, although care must be taken to prevent granules from entering into the joint. No studies have been carried out in this regard.

8.13

Summary of applications in orthopaedics and traumatology

Repairing defects or fractures in metaphyseal areas are an ideal use for bioactive glass. There seems to be no contraindication for use in pediatric orthopaedic surgery or traumatology. In spinal surgery it can be used with good results in posterolateral spondylodesis. The results of the use of granules in the cages for interbody fusion have not been reported. In arthroplasty it is possible to use granules together with autogenous or allograft bone as an extender, though again, no clinical results have been reported. The applicability of bioactive glasses are presented in summary in Table 8.3.

Table 8.3 Clinical applications of bioactive glass granules in orthopaedics Metaphyseal fractures

+++

Diaphyseal fractures

+

Benign bone tumors

+++

Malignant bone tumors

–

Spinal surgery ALIF, PLIF

?

Posterolateral

++

Pediatric use Benign tumors

+++

Trauma

+++

Arthroplasty

+/−

Notes: +++, ++ – proven suitable; + – suitable, not proven; ? – not known; +/– – might be helpful as additive.

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8.14

Future trends

The advantage of bioactive glass is the formation of hydroxylapatite within the body. The disadvantages are related to the long resorption time. For this reason bioactive glass is not suitable for clinical application where the filled area needs to be operated on, drilled or resected shortly after the primary operation. Resorption in large defects can take several years, and in smaller defects with diameter approximately 1 cm, up to a year. Any attempt to drill or resect through the implanted area too early will result in splittering and crushing of the formed bone by the harder fragments of unresorbed glass causing damage to the surrounding bone. Economic considerations are leading to an increase in the use of bone substitute materials instead of autogenous bone. The costs of the time in the operating theater are being calculated more accurately. In a situation whereby the cost of extra time in the operating theater is equivalent to the cost of bone substitute material, bone substitute materials will be favored: in a clinical situation these are preferable not only for the patient but also for the surgeon. There are currently a large number of synthetic bone substitute materials on the market. Allograft bone has so far been the most frequently used material. Local hospital bone banks are in the process of changing to more centralized bone banks. At the same time the use of bone substitute materials has increased [72]. The various materials are in competition, but at the same time the marketing of one material supports that of the others. It can reasonably be expected that certain materials will be used in a certain applications, i.e. one single material is not suitable for all applications. It is also to be expected that composite materials [73, 74] and materials embedded with growth factors and even prefabricated bone will appear in the near future [75].

8.15

References

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9. Urist M. R., Strates B. S. Bone formation in implants of partially and wholly demineralized matrix. Clin Orthop 1970; 71:271–78. 10. Urist M. R., Strates B. S. Bone morphogenetic protein. J Dent Res 1971; 50:1392–406. 11. Hench L. L., Splinter R. J., Allen W. C., Greenlee T. K. Bonding mechanism at the interface of ceramic prosthetic materials. J Biomed Mat Res Symp 1971; No 2 (Part 1):117–43. 12. Roy D. M., Linnehan S. K. Hydroxyapatite formed from coral skeleton carbonate by hydrothermal exchange. Nature 1974; 247:220–22. 13. Jarcho M. Calcium phosphate ceramics as hard tissue prosthetics. Clin Orthop 1981; 19:152–54. 14. Mittelmeier H., Mittelmeier W. Moderne Entwicklung von Knochenersatzmaterialen. Heft zur Unfallheilkunde, pp. 69–84, Heft 216. Hrsg. A. H. Huggler, E. H. Kuner. Berlin 1991. 15. Gross U. M., Strunz V. Surface staining of sawed sections of undecalcified bone containing alloplastic implants. Stain Technol 1977; 52(4):217–19. 16. Kokubo T., Shigematsu M., Nagashime Y., Tashiro T., Nakamura T., Yamamuro T., Higashi S. Apatite- and wollastonite-containing glass-ceramic for prosthetic application. Kyoto University: Bull Inst Chem Res 1982; 60:260–68. 17. Furlong R. J., Osborn J.-F. Fixation of hip prosthesis by hydroxyapatite ceramic coatings. J Bone Joint Surg 1991; 73B:741–45 18. van Blitterswijk C. A., Bakker D., Leenders H., Brink J. V. D., Hesseling S. C., Bovell Y. P., Radder A. M., Sakkers R. J., Gaillard M. L., Heinze P. H., Beumer G. J. Interfacial reactions leading to bone-bonding with PEO/PBT copolymers (Polyactive®). In P. Ducheyne, T. Kokubo, C. A. van Blitterswijk (eds) Bone-Bonding Biomaterials. Holland: Reed Healthcare Communications 1992; pp. 13–30. 19. Constanz B. R., Ison I. C., Fulmer M. T., Poser R. D., Smith S. T., VanWagoner M., Ross J., Goldstein S. A., Jupiter J. B., Rosenthal D. I. Skeletal repair by in situ formation of the mineral phase of bone. Science 1995; 267:1796–99. 20. Andersson Ö. H. The bioactivity of silicate glass. Thesis. Åbo Akademi University, Turku, Finland 1990. 21. Hench L. L., Ethridge A. C. Biomaterials. An Interfacial Approach. New York: Academic Press, 1982; pp. 279–88. 22. Aebi M. Biologisher oder artifizieller Knochenersatz? Hefte zur Unfallheilkunde, Springler Verlag, Berlin Heft 1991; 216:1–9. 23. Gross U., Strunz V. Interface of various glasses and glass-ceramics of different solubility in the femur of the rat. J Biomed Mater Res 1980; 14:607–18. 24. Kokubo T., Ito S., Shigematsu M., Sakka S., Yamamuro T. Mechanical properties of a new type of apatite-containing glass-ceramic for prosthetic application. J Mat Sci 1985; 20:2001–4. 25. Kudo K., Miyashava M., Fujioka Y., Kamegai T., Nakano H., Seino Y., Ishikawa F., Shioyama T., Ishibashi K. Clinical application of dental implant with root of coated bioglass: short-time results. Oral Surg Oral Med Oral Pathol 1990; 70:18–23. 26. Wolf H., Lorenz T., Benser A., Schubert R., Bertoldt R., Reefschläger J., Berger G., Sauer R. In vitro detection and evaluation of pathophysiological effects of ions released from bioactive glass-ceramics. Biomaterials 3rd World congress. Summaries, abstract and programme, Kyoto, 198827; p. 69; Andersson Ö. H., Guizhi Liu, Karlsson K. H., Niemi L., Miettinen J., Juhanoja J. In vivo behaviour of glasses in the SiO2Na2O-CaO-P2O5-Al2O3-B2O3 system. J Mater Sci: Materials in Medicine 1 1990; 219–227.

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9 Bioactive glass S53P4 as a bone graft substitute in the treatment of osteomyelitis N. C. LINDFORS , Helsinki University Central Hospital, Finland

Abstract: Bioactive glasses (BAGs) are bone substitutes with bone-bonding, angiogenesis-promoting and antibacterial properties. The bioactive process that leads to bone bonding is described through reactions at the glass surface. This chapter describes the chemical reactions following implantation of the glass that lead to the formation of silanol (SiOH) groups at the glass surface and the subsequent formation of a CaO-P2O5 hydroxyapatite (HA) layer on top of the Si-rich layer. Finally, cell interactions with the HA layer are shown to initiate the bone-forming pathway. The chapter also explains and compares the antibacterial properties of different bioactive glasses, showing BAG-S53P4 to be the most effective against significant pathogens and bacteria. These findings are demonstrated in a multicentre trial involving eleven patients of osteomyelitis treated with BAG-S53P4 implants, which demonstrated the good grafting and antibacterial properties of S53P4 as an implant material and bone graft substitute. It has since been used successfully in further operations. Key words: bioactive glass, osteomyelitis, S53P4, bone subsitute, infection.

9.1

Introduction

Osteomyelitis is caused by infected micro-organisms and defines a destructive inflammatory process in bone that is often accompanied by bone destruction (Lazzarini et al., 2004). Osteomyelitis is heterogenous in its pathophysiology, clinical presentation, and management. It is often considered to be the most difficult-to-treat infectious disease. The infection can arise from a variety of aetiologies (Lew and Waldfogel, 2004). Most often it is caused by trauma, but any kind of bone or soft tissue surgery where pathogens can enter the bone may cause the infection. Haematogenous osteomyelitis has been found in children, as well as in elderly patients (Riise et al., 2008); and in diabetic patients, osteomyelitis may appear as a secondary manifestation due to vascular insufficiency and soft tissue infection (Haartemann-Heurtier and Senneville, 2008). The most common pathogens causing osteomyelitis are Staphylococcus aureus and Gram-negative bacilli (Parsson and Strauss, 2004). Acute osteomyelitis is characterized by oedema, locally decreased blood supply and pus formation. Untreated or due to treatment failure, the infection can progress to a more chronic phase, with formation of a sequestrum, a large area of 209 © Woodhead Publishing Limited, 2011

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devascularized dead bone. In treatment of chronic osteomyelitis, debridement of the dead bone is mandatory. This procedure unfortunately, often results in a poorly vascularized large bone defect, a dead space. Bacterial infection can also cause local acidosis, leading to dissolution of bone matrix mineral (Konttinen et al., 2001). Many different methods have been used to treat osteomyelitis, including antibiotic-impregnated polymethyl methacrylate (PMMA) beads, free vascularized bone grafts, local muscle flaps, granulation formation according to the technique of Papineau and the Masquelet technique (Powerski et al., 2009) or bone reconstruction based on Ilizarov technology (Parsson and Strauss, 2004).

9.2

Bone grafts in the treatment of osteomyelitis

Bacterial colonization of implanted materials is promoted by the ability of the bacteria to produce protein-specific adhesins on their surfaces. This is followed by interactions with host protein components, such as fibrinogen, fibronectin and collagen. Bacteria communicate through hormone-like compounds in biofilms, making treatment with antimicrobial agents difficult (Lew and Waldfogel, 2004). Therefore, synthetic bone grafts are generally not recommended in treatment of osteomyelitis. Debridement in combination with the local administration of antibiotics, e.g. gentamicin-loaded PMMA beads, has become the method of choice in treating osteomyelitis. However, PMMA is known to provide a favourable environment for proliferation of bacteria (Boyd and Towler, 2005). In a long-term follow-up of 100 patients treated with genta